(en)Methods and compositions are provided for detecting biomolecular interactions. The use of labels is not required and the methods can be performed in a high-throughput manner. The invention also provides optical devices useful as narrow band filters.

1.ApplicationNumber: US-4856008-A
1.PublishNumber: US-2008219892-A1
2.Date Publish: 20080911
3.Inventor: CUNNINGHAM BRIAN T.
LI PETER
QIU JEAN
PIEN HOMER

4.Inventor Harmonized: CUNNINGHAM BRIAN T(US)
LI PETER(US)
QIU JEAN(US)
PIEN HOMER(US)

5.Country: US
6.Claims:
(en)Methods and compositions are provided for detecting biomolecular interactions. The use of labels is not required and the methods can be performed in a high-throughput manner. The invention also provides optical devices useful as narrow band filters.

7.Description:
(en)PRIORITY
This application is a divisional application of U.S. Ser. No. 11/201,237, filed Aug. 10, 2005 (allowed), which is a continuation-in-part of U.S. application Ser. No. 10/059,060, filed Jan. 28, 2002, now U.S. Pat. No. 7,070,987. This application is also a continuation-in-part of U.S. application Ser. No. 09/930,352, filed Aug. 15, 2001, now U.S. Pat. No. 7,094,595, which claims the benefit of U.S. provisional application 60/244,312 filed Oct. 30, 2000; U.S. provisional application 60/283,314 filed Apr. 12, 2001; and U.S. provisional application 60/303,028 filed Jul. 3, 2001.


TECHNICAL AREA OF THE INVENTION
The invention relates to compositions and methods for detecting biomolecular interactions. The detection can occur without the use of labels and can be done in a high-throughput manner. The invention also relates to optical devices.
BACKGROUND OF THE INVENTION
With the completion of the sequencing of the human genome, one of the next grand challenges of molecular biology will be to understand how the many protein targets encoded by DNA interact with other proteins, small molecule pharmaceutical candidates, and a large host of enzymes and inhibitors. See e.g., Pandey & Mann, “Proteomics to study genes and genomes,” Nature, 405, p. 837-846, 2000; Leigh Anderson et al., “Proteomics: applications in basic and applied biology,” Current Opinion in Biotechnology, 11, p. 408-412, 2000; Patterson, “Proteomics: the industrialization of protein chemistry,” Current Opinion in Biotechnology, 11, p. 413-418, 2000; MacBeath & Schreiber, “Printing Proteins as Microarrays for High-Throughput Function Determination,” Science, 289, p. 1760-1763, 2000; De Wildt et al., “Antibody arrays for high-throughput screening of antibody-antigen interactions,” Nature Biotechnology, 18, p. 989-994, 2000. To this end, tools that have the ability to simultaneously quantify many different biomolecular interactions with high sensitivity will find application in pharmaceutical discovery, proteomics, and diagnostics. Further, for these tools to find widespread use, they must be simple to use, inexpensive to own and operate, and applicable to a wide range of analytes that can include, for example, polynucleotides, peptides, small proteins, antibodies, and even entire cells.
Biosensors have been developed to detect a variety of biomolecular complexes including oligonucleotides, antibody-antigen interactions, hormone-receptor interactions, and enzyme-substrate interactions. In general, biosensors consist of two components: a highly specific recognition element and a transducer that converts the molecular recognition event into a quantifiable signal. Signal transduction has been accomplished by many methods, including fluorescence, interferometry (Jenison et al, “Interference-based detection of nucleic acid targets on optically coated silicon,” Nature Biotechnology, 19, p. 62-65; Lin et al, “A porous silicon-based optical interferometric biosensor,” Science, 278, p. 840-843, 1997), and gravimetry (A. Cunningham, Bioanalytical Sensors, John Wiley & Sons (1998)).
Of the optically-based transduction methods, direct methods that do not require labeling of analytes with fluorescent compounds are of interest due to the relative assay simplicity and ability to study the interaction of small molecules and proteins that are not readily labeled. Direct optical methods include surface plasmon resonance (SPR) (Jordan & Corn, “Surface Plasmon Resonance Imaging Measurements of Electrostatic Biopolymer Adsorption onto Chemically Modified Gold Surfaces,” Anal. Chem., 69:1449-1456 (1997), (grating couplers (Morhard et al., “Immobilization of antibodies in micropatterns for cell detection by optical diffraction,” Sensors and Actuators B, 70, p. 232-242, 2000), ellipsometry (Jin et al., “A biosensor concept based on imaging ellipsometry for visualization of biomolecular interactions,” Analytical Biochemistry, 232, p. 69-72, 1995), evanascent wave devices (Huber et al., “Direct optical immunosensing (sensitivity and selectivity),” Sensors and Actuators B, 6, p. 122-126, 1992), and reflectometry (Brecht & Gauglitz, “Optical probes and transducers,” Biosensors and Bioelectronics, 10, p. 923-936, 1995). Theoretically predicted detection limits of these detection methods have been determined and experimentally confirmed to be feasible down to diagnostically relevant concentration ranges. However, to date, these methods have yet to yield commercially available high-throughput instruments that can perform high sensitivity assays without any type of label in a format that is readily compatible with the microtiter plate-based or microarray-based infrastructure that is most often used for high-throughput biomolecular interaction analysis. Therefore, there is a need in the art for compositions and methods that can achieve these goals.
SUMMARY OF THE INVENTION
It is an object of the invention to provide compositions and methods for detecting binding of one or more specific binding substances to their respective binding partners. This and other objects of the invention are provided by one or more of the embodiments described below.
One embodiment of the invention provides a biosensor. The biosensor comprises a one-dimensional grating layer comprised of a material having a high refractive index, a low refractive index material layer that supports the one-dimensional grating layer, and one or more specific binding substances immobilized on the surface of the one-dimensional grating layer opposite of the low refractive index material layer. When the biosensor is illuminated a resonant grating effect is produced on a reflected radiation spectrum. The cross-sectional period of the one-dimensional grating is less than the wavelength of the resonant grating effect. In another embodiment, the biosensor comprises a one-dimensional grating surface structure comprised of a material having a low refractive index, a high refractive index material layer that is applied on top of the low refractive index one-dimensional grating layer, and one or more specific binding substances immobilized on a surface of the high refractive index layer opposite of the one-dimensional grating surface structure comprised of a material having a low refractive index.
The cross-sectional profile of the one-dimensional grating can be triangular, sinusoidal, trapezoidal, rectangular, stepped, v-shaped, u-shaped, upside-down v-shaped, upside-down u-shaped, or square. A narrow band of optical wavelengths is reflected from the biosensor when the biosensor is illuminated with a broad band of optical wavelengths.
The low refractive index material of the biosensor can comprise glass, plastic, polymer, or epoxy. The high refractive index material can be selected from the group consisting of zinc sulfide, titanium dioxide, indium tin oxide, tantalum oxide, and silicon nitride. The one-dimensional grating can have a period of about 0.01 microns to about 1 micron and a depth of about 0.01 microns to about 1 micron.
The one or more specific binding substances can be arranged in an array of distinct locations. The distinct locations can define a microarray spot of about 50-500 microns in diameter. The one or more specific binding substances can be immobilized on the high refractive index material by physical adsorption or by chemical binding. The one or more specific binding substances can be bound to their binding partners. The one or more specific binding substances or binding partners can be selected from the group consisting of 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, protein solutions, peptide solutions, single- or double-stranded DNA solutions, RNA solutions, solutions containing compounds from a combinatorial chemical library and biological samples. The biological sample can be selected from the group consisting of, blood, plasma, serum, gastrointestinal secretions, homogenates of tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, and prostatitc fluid.
A biosensor can comprise the internal surface of a liquid-containing vessel, such as a microtiter plate, a test tube, a petri dish and a microfluidic channel.
Another embodiment of the invention provides a detection system comprising a biosensor, a light source that directs light to the biosensor, and a detector that detects light reflected from the biosensor, wherein a polarizing filter occurs between the light source and the biosensor.
Yet another embodiment of the invention provides a method of detecting the binding of one or more specific binding substances to their respective binding partners. The method comprises applying one or more binding partners to a biosensor, illuminating the biosensor with light; and detecting a peak wavelength value (PWV). If the one or more specific binding substances have bound to their respective binding partners, then the PWV is shifted.
Still another embodiment of the invention provides a method of detecting the binding of one or more specific binding substances to their respective binding partners. The method comprises applying one or more binding partners to a biosensor, wherein the high refractive index material is coated with an array of distinct locations containing the one or more specific binding substances, illuminating each distinct location of the biosensor with light; and detecting peak wavelength value (PWV) for each distinct location of the biosensor. If the one or more specific binding substances have bound to their respective binding partners at a distinct location, then the PWV is shifted.
Even another embodiment of the invention provides a method of detecting activity of an enzyme. The method comprises applying one or more enzymes to a biosensor, washing the biosensor, illuminating the biosensor with light, and detecting a PWV. If the one or more enzymes have altered the one or more specific binding substances of the biosensor by enzymatic activity, then the PWV is shifted.
Another embodiment provides a method of measuring the amount of one or more binding partners in a test sample. The method comprises illuminating a biosensor with light, detecting a PWV from the biosensor, applying a test sample comprising one or more binding partners to the biosensor, illuminating the biosensor with light; and, detecting a PWV from the biosensor. The difference in PWVs is a measurement of the amount of one or more binding partners in the test sample.
Still another embodiment of the invention provides a method of detecting the binding of one or more specific binding substances to their respective binding partners. The method comprises applying one or more binding partners comprising one or more tags to a biosensor, illuminating the biosensor with light; and detecting a PWV from the biosensor. If the one or more specific binding substances have bound to their respective binding partners, then the reflected wavelength of light is shifted. The one or more tags can be selected from the group consisting of biotin, succinimidyl-6-[a-methyl-a-(2-pyridyl-dithio)toluamido]hexanoate (SMPT), dimethylpimelimidate (DMP), and histidine. The one or more tags can be reacted with a composition selected from the group consisting of streptavidin, horseradish peroxidase, and streptavidin coated nanoparticles, before the step of illuminating the biosensor with light.
Another embodiment of the invention provides a biosensor comprising a one-dimensional or two-dimensional grating layer comprised of a material having a high refractive index, a low refractive index material layer that supports the one-dimensional or two-dimensional grating layer; a surface modification layer on a surface of the one-dimensional or two-dimensional grating layer opposite of the low refractive index material layer; and one or more specific binding substances immobilized on a surface of the surface modification layer opposite of the one-dimensional or two-dimensional grating layer. When the biosensor is illuminated a resonant grating effect is produced on a reflected radiation spectrum. The surface modification layer can be comprised of silicon oxide. The thickness of the surface modification layer can be about 5 nm to about 15 nm.
Another embodiment of the invention provides a biosensor comprising a grating layer comprising a one-dimensional or two-dimensional grating on a first surface; a interfacial layer on the first surface of the grating layer, a high refractive index material layer on the surface of the interfacial layer opposite of the grating layer, and one or more specific binding substances immobilized on a surface of the high refractive index material layer opposite of the interfacial layer. When the biosensor is illuminated a resonant grating effect is produced on a reflected radiation spectrum. The interfacial layer can be comprised of a material selected from the group consisting of silicon oxide, silicon oxynitride, borosilicate glass, phosophosilicate glass, pyrex, glass, and a metal oxide. The interfacial layer can be about 1 nm to about 200 nm thick.
Therefore, unlike surface plasmon resonance, resonant mirrors, and waveguide biosensors, the described compositions and methods enable many thousands of individual binding reactions to take place simultaneously upon the biosensor surface. This technology is useful in applications where large numbers of biomolecular interactions are measured in parallel, particularly when molecular labels will alter or inhibit the functionality of the molecules under study. High-throughput screening of pharmaceutical compound libraries with protein targets, and microarray screening of protein-protein interactions for proteomics are examples of applications that require the sensitivity and throughput afforded by this approach. A biosensor of the invention can be manufactured, for example, in large areas using a plastic embossing process, or an epoxy replication process, and thus can be inexpensively incorporated into common disposable laboratory assay platforms such as microtiter plates and microarray slides.



BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1A shows a schematic diagram of one embodiment of an optical grating structure used for a calorimetric resonant reflectance biosensor. n substrate represents substrate material. n 1 represents the refractive index of a cover layer. n 2 represents the refractive index of a one- or two-dimensional grating. n bio represents the refractive index of one or more specific binding substances. t 1 represents the thickness of the cover layer. t 2 represents the thickness of the grating. t bio represents the thickness of the layer of one or more specific binding substances. FIG. 1 shows a cross-sectional view of a biosensor.
FIG. 2 shows a schematic drawing of a one-dimensional linear grating surface structure.
FIG. 3A-B shows a two-dimensional grating comprising a rectangular grid of squares ( FIG. 3A ) or holes ( FIG. 3B ).
FIG. 4 shows a biosensor cross-section profile utilizing a sinusoidally varying grating profile.
FIG. 5 shows a biosensor cross-section profile in which an embossed substrate is coated with a higher refractive index material such as ZnS or SiN. An optional cover layer of low refractive index material, for example, epoxy or SOG is layered on top of the higher refractive index material and one or more specific binding substances are immobilized on the cover layer.
FIG. 6 shows three types of surface activation chemistry (Amine, Aldehyde, and Nickel) with corresponding chemical linker molecules that can be used to covalently attach various types of biomolecule receptors to a biosensor.
FIG. 7A-C shows methods that can be used to amplify the mass of a binding partner such as detected DNA or detected protein on the surface of a biosensor.
FIG. 8 shows a graphic representation of how adsorbed material, such as a protein monolayer, will increase the reflected wavelength of on a SRVD biosensor.
FIG. 9 shows an example of a biosensor used as a microarray.
FIG. 10A-B shows two biosensor formats that can incorporate a calorimetric resonant reflectance biosensor. FIG. 10A shows a biosensor that is incorporated into a microtitre plate.
FIG. 10B shows a biosensor in a microarray slide format.
FIG. 11 shows an array of arrays concept for using a biosensor platform to perform assays with higher density and throughput.
FIG. 12 shows a diagram of an array of biosensor electrodes. A single electrode can comprise a region that contains many grating periods and several separate grating regions can occur on the same substrate surface.
FIG. 13 shows a SEM photograph showing the separate grating regions of an array of biosensor electrodes.
FIG. 14 shows a biosensor upper surface immersed in a liquid sample. An electrical potential can be applied to the biosensor that is capable of attracting or repelling a biomolecule near the electrode surface.
FIG. 15 shows a biosensor upper surface immersed in a liquid sample. A positive voltage is applied to an electrode and the electronegative biomolecules are attracted to the biosensor surface.
FIG. 16 shows a biosensor upper surface immersed in a liquid sample. A negative voltage is applied to an electrode and the electronegative biomolecules are repelled from the biosensor surface using a negative electrode voltage.
FIG. 17 demonstrates an example of a biosensor that occurs on the tip of a fiber probe for in vivo detection of biochemical substances.
FIG. 18 shows an example of the use of two coupled fibers to illuminate and collect reflected light from a biosensor.
FIG. 19 shows resonance wavelength of a biosensor as a function of incident angle of detection beam.
FIG. 20 shows an example of the use of a beam splitter to enable illuminating and reflected light to share a common collimated optical path to a biosensor.
FIG. 21 shows an example of a system for angular scanning of a biosensor.
FIG. 22 shows SEM photographs of a photoresist grating structure in plan view (center and upper right) and cross-section (lower right).
FIG. 23 shows a SEM cross-section photograph of a grating structure after spin-on glass is applied over a silicon nitride grating.
FIG. 24 shows examples of biosensor chips (1.5×10.5-inch). Circular areas are regions where the resonant structure is defined.
FIG. 25 shows response as a function of wavelength of a biosensor that BSA had been deposited at high concentration, measured in air. Before protein deposition, the resonant wavelength of the biosensor is 380 nm and is not observable with the instrument used for this experiment.
FIG. 26 shows response as a function of wavelength comparing an untreated biosensor with one upon which BSA had been deposited. Both measurements were taken with water on the biosensor's surface.
FIG. 27 shows response as a function of wavelength of a biosensor that Borrelia bacteria has been deposited at high concentration and measured in water.
FIG. 28 shows a computer simulation of a biosensor demonstrating the shift of resonance to longer wavelengths as biomolecules are deposited on the surface.
FIG. 29 shows a computer simulation demonstrating the dependence of peak reflected wavelength on protein coating thickness. This particular biosensor has a dynamic range of 250 nm deposited biomaterial before the response begins to saturate.
FIG. 30 shows an embodiment of a biosensor. n substrate represents the refractive index of a substrate. n 1 represents the refractive index of an optional optical cover layer. n 2 represents the refractive index of a one- or two-dimensional grating. n 3 represents the refractive index of a high refractive index material such as silicon nitride. n bio represents the refractive index of one or more specific binding substances. t 1 represents the thickness of a cover layer. t 2 represents the thickness of a one- or two-dimensional grating. t 3 represents the thickness of a high refractive index material. t bio represents the thickness of a specific binding substance layer.
FIG. 31 shows reflected intensity as a function of wavelength for a resonant grating structure when various thicknesses of protein are incorporated onto the upper surface.
FIG. 32 shows a linear relationship between reflected wavelength and protein coating thickness for a biosensor shown in FIG. 30 .
FIG. 33 shows instrumentation that can be used to read output of a biosensor. A collimated light source is directed at a biosensor surface at normal incidence through an optical fiber, while a second parallel fiber collects the light reflected at normal incidence. A spectrometer records the reflectance as a function of wavelength.
FIG. 34 shows the measured reflectance spectra of a biosensor.
FIG. 35 shows dependence of peak resonant wavelength measured in liquid upon the concentration of protein BSA dissolved in water.
FIG. 36 shows dependence of peak resonance wavelength on the concentration of BSA dissolved in PBS, which was then allowed to dry on a biosensor surface.
FIG. 37A-B . FIG. 37A shows a measurement of peak resonant wavelength shift caused by attachment of a streptavidin receptor layer and subsequent detection of a biotinylated IgG. FIG. 37B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 38A-B . FIG. 38A shows results of streptavidin detection at various concentrations for a biosensor that has been activated with NH 2 surface chemistry linked to a biotin receptor molecule. FIG. 38B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 39A-B . FIG. 39A shows an assay for detection of anti-goat IgG using a goat antibody receptor molecule. BSA blocking of a detection surface yields a clearly measurable background signal due to the mass of BSA incorporated on the biosensor. A 66 nM concentration of anti-goat IgG is easily measured above the background signal. FIG. 39B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 40A-B . FIG. 40A shows a nonlabeled ELISA assay for interferon-gamma (INF-gamma) using an anti-human IgG INF-gamma receptor molecule, and a neural growth factor (NGF) negative control. FIG. 40B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 41A-B . FIG. 41A shows detection of a 5-amino acid peptide (MW=860) and subsequent cleavage of a pNA label (MW=130) using enzyme caspase-3. FIG. 41B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 42A-B . FIG. 42A shows resonant peak in liquid during continuous monitoring of the binding of three separate protein layers. FIG. 42B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 43A-B . FIG. 43A shows endpoint resonant frequencies mathematically determined from the data shown in FIG. 42 . FIG. 43B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 44A-B . FIG. 44A shows kinetic binding measurement of IgG binding. FIG. 44B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 45A-B . FIG. 45A shows kinetic measurement of a protease that cleaves bound protein from a biosensor surface. FIG. 45B shows a schematic demonstration of molecules bound to a biosensor.
FIG. 46 shows comparison of mathematical fit of parabolic and exponential functions to spectrometer data from a resonant peak. The exponential curve fit is used to mathematically determine a peak resonant wavelength.
FIG. 47 shows sensitivity of the mathematically determined peak resonant wavelength to artificially added noise in the measured spectrum.
FIG. 48 shows a resonant optical biosensor incorporating an electrically conducting material.
FIG. 49 shows a resonant reflection or transmission filter structure consisting of a set of concentric rings.
FIG. 50 shows a resonant reflective or transmission filter structure comprising a hexagonal grid of holes (or a hexagonal grid of posts) that closely approximates the concentric circle structure of FIG. 49 without requiring the illumination beam to be centered upon any particular location of the grid.
FIG. 51 shows a plot of the peak resonant wavelength values for test solutions. The avidin solution was taken as the baseline reference for comparison to the Avidin+BSA and Avidin+b-BSA solutions. Addition of BSA to avidin results in only a small resonant wavelength increase, as the two proteins are not expected to interact. However, because biotin and avidin bind strongly (Kd=10 −15 M), the avidin+b-BSA solution will contain larger bound protein complexes. The peak resonant wavelength value of the avidin+b-BSA solution thus provides a large shift compared to avidin+BSA.
FIG. 52 shows a schematic diagram of a detection system.
FIG. 53A-B shows a fabrication process used to produce the biosensor and cross-section of a one-dimensional linear grating sensor. FIG. 53A shows a silicon master wafer used to replicate the biosensor structure into a thin film of epoxy between the silicon and a sheet of plastic film. After the epoxy is cured, the plastic sheet is peeled away. To complete sensor fabrication ( FIG. 53B ), a thin film of high refractive index dielectric material such as silicon nitride, titanium oxide, tantalum oxide, or zinc sulfide is deposited over the structure.
FIG. 54A-C shows a linear grating structure ( FIG. 54A ; top view) used to produce the one-dimensional linear grating guided mode resonant filter “master” structure. First, an 8-inch diameter silicon “master” wafer is produced. The 550 nm period linear grating structure is defined in photoresist using deep-UV photolithography by stepping and repeating the exposure of a 9 mm diameter circular grating reticle over the surface of a photoresist-coated silicon wafer, as shown in FIG. 54B . FIG. 54C shows that the exposure step/repeat procedure produced patterns for two standard format 96-well microtiter plates with 8 rows and 12 columns each. The exposed photoresist was developed, and the grating structure was permanently transferred to the silicon wafer using a reactive ion etch with a depth of ˜200 nm. After etching, the photoresist was removed.
FIG. 55 shows instrumentation used to illuminate and read output of a biosensor structure. The probe head contains two optical fibers. The first fiber is connected to a white light source to cast a small spot of polarized collimated light on the biosensor surface. The second fiber collects reflected light for analysis by a spectrometer.
FIG. 56 shows reflected intensity as a function of wavelength for a one-dimensional linear grating surface biosensor structure within a microtiter plate well filled with water.
FIG. 57 demonstrates peak wavelength shift relative to a clean one-dimensional linear grating surface biosensor structure for three biosensor surface activation states. The error bars indicate the standard deviation of the shift over seven separate sensor wells.
FIG. 58A-C shows the exposure of NH 2 , PEG, and PEG-Biotin activated one-dimensional linear grating surface biosensor structures to seven concentrations of anti-biotin IgG. The NH 2 surface ( FIG. 58A ) displays low levels of nonspecific protein binding at high protein exposure concentrations, while the PEG surface ( FIG. 58B ) displays low levels of nonspecific binding. The PEG-Biotin ( FIG. 58C ) surface has a strong binding interaction with the anti-biotin IgG.
FIG. 59 shows peak wavelength value shift as a function of anti-biotin IgG concentration for PEG-Biotin activated wells after a 20-minute incubation. The plotted line indicates a least-squared fit linear function.
FIG. 60 demonstrates the effect of a surface modification layer on specific binding substance immobilization onto the surface of a biosensor.
FIG. 61 shows water stability test results for biosensors with and without an interfacial layer. The addition of an interfacial layer significantly improved stability of a biosensor in aqueous solutions.



DETAILED DESCRIPTION OF THE INVENTION
Subwavelength Structured Surface (SWS) Biosensor
In one embodiment of the invention, a subwavelength structured surface (SWS) is used to create a sharp optical resonant reflection at a particular wavelength that can be used to track with high sensitivity the interaction of biological materials, such as specific binding substances or binding partners or both. A colormetric resonant diffractive grating surface acts as a surface binding platform for specific binding substances.
SWSs are an unconventional type of diffractive optic that can mimic the effect of thin-film coatings. (Peng & Morris, J. Opt. Soc. Am. A , Vol. 13, No. 5, p. 993, May 1996; Magnusson, & Wang, Appl. Phys. Lett., 61, No. 9, p. 1022, August, 1992; Peng & Morris, Optics Letters , Vol. 21, No. 8, p. 549, April, 1996). A SWS structure comprises a surface-relief grating, such as a one-dimensional, two-dimensional, or three dimensional grating in which the grating period is small compared to the wavelength of incident light.
The reflected or transmitted color of this structure can be modulated by the addition of molecules such as specific binding substances, binding partners, or both, or inorganic molecules to the upper surface of the cover layer or the grating surface. The dielectric susceptibility of the added molecules results in a modification of the wavelength at which maximum reflectance or transmittance will occur.
In one embodiment, a biosensor, when illuminated with white light, is designed to reflect only a single wavelength. When specific binding substances are attached to the surface of the biosensor, the reflected wavelength (color) is shifted due to the change of the optical path of light that is coupled into the grating. By linking specific binding substances to a biosensor surface, complementary binding partner molecules can be detected without the use of any kind of fluorescent probe or particle label. The detection technique is capable of resolving changes of, for example, ˜0.1 nm thickness of protein binding, and can be performed with the biosensor surface either immersed in fluid or dried.
A detection system consists of, for example, a light source that illuminates a small spot of a biosensor at normal incidence through, for example, a fiber optic probe, and a spectrometer that collects the reflected light through, for example, a second fiber optic probe also at normal incidence. Because no physical contact occurs between the excitation/detection system and the biosensor surface, no special coupling prisms are required and the biosensor can be easily adapted to any commonly used assay platform including, for example, microtiter plates and microarray slides. A single spectrometer reading can be performed in several milliseconds, thus it is possible to quickly measure a large number of molecular interactions taking place in parallel upon a biosensor surface, and to monitor reaction kinetics in real time.
This technology is useful in applications where large numbers of biomolecular interactions are measured in parallel, particularly when molecular labels would alter or inhibit the functionality of the molecules under study. High-throughput screening of pharmaceutical compound libraries with protein targets, and microarray screening of protein-protein interactions for proteomics are examples of applications that require the sensitivity and throughput afforded by the compositions and methods of the invention.
A schematic diagram of an example of a SWS structure is shown in FIG. 1 . In FIG. 1 , n substrate represents a substrate material. n 1 represents the refractive index of an optional cover layer. n 2 represents the refractive index of a two-dimensional grating. N bio represents the refractive index of one or more specific binding substances. t 1 represents the thickness of the cover layer above the two-dimensional grating structure. t 2 represents the thickness of the grating. t bio represents the thickness of the layer of one or more specific binding substances. In one embodiment, are n2>n1. (see FIG. 1 ). Layer thicknesses (i.e. cover layer, one or more specific binding substances, or a grating) are selected to achieve resonant wavelength sensitivity to additional molecules on the top surface The grating period is selected to achieve resonance at a desired wavelength.
One embodiment of the invention provides a SWS biosensor. A SWS biosensor comprises a one-dimensional or two-dimensional grating, a substrate layer that supports the grating, and one or more specific binding substances immobilized on the surface of the grating opposite of the substrate layer.
A one-dimensional or two-dimensional grating can be comprised of a material, including, for example, zinc sulfide, titanium dioxide, tantalum oxide, and silicon nitride. A cross-sectional profile of the grating can comprise any periodically repeating function, for example, a “square-wave.” A grating can be comprised of a repeating pattern of shapes selected from the group consisting of continuous parallel lines squares, circles, ellipses, triangles, trapezoids, sinusoidal waves, ovals, rectangles, and hexagons. A sinusoidal cross-sectional profile is preferable for manufacturing applications that require embossing of a grating shape into a soft material such as plastic, or replicating a grating surface into a material such as epoxy. In one embodiment of the invention, the depth of the grating is about 0.01 micron to about 1 micron and the period of the grating is about 0.01 micron to about 1 micron.
A SWS biosensor can also comprise a one-dimensional linear grating surface structure, i.e., a series of parallel lines or grooves. See e.g., FIG. 54 . A one-dimensional linear grating is sufficient for producing the guided mode resonant filter effect. While a two-dimensional grating has features in two lateral directions across the plane of the sensor surface that are both subwavelength, the cross-section of a one-dimensional grating is only subwavelength in one lateral direction, while the long dimension can be greater than wavelength of the resonant grating effect. A one-dimensional grating biosensor can comprise a high refractive index material which is coated as a thin film over a layer of lower refractive index material with the surface structure of a one-dimensional grating. See FIG. 53 . Alternatively, a one dimensional grating biosensor can comprise a low refractive index material substrate, upon which a high refractive index thin film material has been patterned into the surface structure of a one-dimensional grating. The low refractive index material can be glass, plastic, polymer, or cured epoxy. The high refractive index material must have a refractive index that is greater than the low refractive index material. The high refractive index material can be zinc sulfide silicon nitride, tantalum oxide, titanium dioxide, or indium tin oxide, for example.
FIG. 53 shows a biosensor cross-sectional profile, in which the one-dimensional grating cross-section is rectangular. Other cross section profiles of the one dimensional linear grating structure will also produce the guided mode resonance effect. These include, for example, triangular or v-shaped, u-shaped, upside-down v- or u-shapes, sinusoidal, trapezoidal, stepped and square. Any regularly repeating periodic function will provide a guided mode resonant effect.
Additionally, a one-dimensional linear grating master structure is easy to produce using commercially available gratings, and large-scale grating master structures with uniform performance can be produced by deep-ultraviolet (DUV) photolithography. Using sub-micron microreplication of a master sensor surface structure on continuous sheets of plastic film, a biosensor can be produced inexpensively over large surface areas. A one-dimensional grating biosensor of the invention can be fabricated by creating a “master” wafer in silicon that is used as a template for producing the sensor structure on plastic by a high-definition microreplication process. The ability to produce a high-sensitivity biosensor in plastic over large surface areas enables incorporation of the biosensor into large area disposable assay formats such as microtiter plates and microarray slides. The incorporation of a plastic biosensor into the bottoms, for example, of bottomless 96-well microtiter plates, allows for the use of a biosensor plate to perform, for example, multiple protein-protein binding assays in parallel. The detection sensitivity of a plastic-substrate biosensor is equivalent to glass-substrate biosensors. A biosensor structure can incorporated into standard microtiter plates and used to perform affinity assays based on measuring the biochemical interaction between a specific binding substance immobilized on the biosensor surface and binding partners within a test sample. A biosensor can also be incorporated into other disposable laboratory assay formats, such as microarray slides, flow cells, and cell culture plates. Incorporation of a biosensor into common laboratory formats is desirable for compatibility with existing microarray handling equipment such as spotters and incubation chambers.
A one-dimensional linear grating biosensor surface contains an optical structure that, when illuminated with collimated white light, is designed to reflect only a narrow band of wavelengths. The narrow wavelength band is described as a wavelength “peak.” The “peak wavelength value” (PWV) changes when biological or other material is deposited or removed from the biosensor surface. A readout instrument illuminates distinct locations on the biosensor surface with collimated white light, and collects collimated reflected light. The collected light is gathered into a wavelength spectrometer for determination of PWV.
One dimensional linear gratings have resonant characteristics where the illuminating light polarization is oriented perpendicular or parallel to the grating period. However, a hexagonal grid of holes has better polarization symmetry than a rectangular grid of holes. Therefore, a calorimetric resonant reflection biosensor of the invention can comprise, for example, a two-dimensional hexagonal array of holes (see FIG. 3B ), a two-dimensional array of squares ( FIG. 3A ) or a one-dimensional grid of parallel lines (see FIG. 2 ). A one-dimensional linear grating has the same pitch (i.e. distance between regions of high and low refractive index), period, layer thicknesses, and material properties as the hexagonal array grating. However, light must be polarized perpendicular or parallel to the grating lines in order to be resonantly coupled into the optical structure. Therefore, a polarizing filter oriented with its polarization axis perpendicular or parallel to the one-dimensional linear grating must be inserted between the illumination source and the biosensor surface. Because only a small portion of the illuminating light source is correctly polarized, a longer integration time is required to collect an equivalent amount of resonantly reflected light compared to a hexagonal grating.
While a one-dimensional linear grating can require either a higher intensity illumination source or a longer measurement integration time compared to a hexagonal grating, the fabrication requirements for the one-dimensional linear grating structure are simpler. A two-dimensional hexagonal grating pattern is produced by holographic exposure of photoresist to three mutually interfering laser beams. The three beams are precisely aligned in order to produce a grating pattern that is symmetrical in three directions. A one-dimensional linear grating pattern requires alignment of only two laser beams to produce a holographic exposure in photoresist, and thus has a reduced alignment requirement. A one-dimensional linear grating pattern can also be produced by, for example, direct writing of photoresist with an electron beam. Also, several commercially available sources exist for producing one-dimensional linear grating “master” templates for embossing or replicating a grating structure into plastic. A schematic diagram of a linear grating structure is shown in FIG. 54 .
A rectangular grid pattern can be produced in photoresist using an electron beam direct-write exposure system. A single wafer can be illuminated as a linear grating with two sequential exposures with the part rotated 90-degrees between exposures.
A one-dimensional or two-dimensional grating can also comprise, for example, a “stepped” profile, in which high refractive index regions of a single, fixed height are embedded within a lower refractive index cover layer. The alternating regions of high and low refractive index provide an optical waveguide parallel to the top surface of the biosensor. See FIG. 5 .
For manufacture, a stepped structure is etched or embossed into a substrate material such as glass or plastic. See FIG. 53B . A uniform thin film of higher refractive index material, such as silicon nitride or zinc sulfide is deposited on this structure. The deposited layer will follow the shape contour of the embossed or etched structure in the substrate, so that the deposited material has a surface relief profile that is identical to the original embossed or etched profile. The thickness of the dielectric layer may be less than, equal to, or greater than the depth of the grating structure. The structure can be completed by the application of an optional cover layer comprised of a material having a lower refractive index than the higher refractive index material and having a substantially flat upper surface. The covering material can be, for example, glass, epoxy, or plastic.
This structure allows for low cost biosensor manufacturing, because it can be mass-produced. A “master” grating can be produced in glass, plastic, or metal using, for example, a three-beam laser holographic patterning process, See e.g., Cowan, The recording and large scale production of crossed holographic grating arrays using multiple beam interferometry, Proc. Soc. Photo - optical Instum. Eng. 503:120 (1984). A master grating can be repeatedly used to emboss a plastic substrate. The embossed substrate is subsequently coated with a high refractive index material and optionally, a cover layer.
While a stepped structure is simple to manufacture, it is also possible to make a resonant biosensor in which the high refractive index material is not stepped, but which varies with lateral position. FIG. 4 shows a profile in which the high refractive index material of the one-dimensional or two-dimensional grating, n2, is sinusoidally varying in height. To produce a resonant reflection at a particular wavelength, the period of the sinusoid is identical to the period of an equivalent stepped structure. The resonant operation of the sinusoidally varying structure and its functionality as a biosensor has been verified using GSOLVER (Grating Solver Development Company, Allen, Tex., USA) computer models.
Techniques for making two-dimensional gratings are disclosed in Wang, J. Opt. Soc. Am No. 8, August 1990, pp. 1529-44. Biosensors of the invention can be made in, for example, a semiconductor microfabrication facility. Biosensors can also be made on a plastic substrate using continuous embossing and optical coating processes. For this type of manufacturing process, a “master” structure is built in a rigid material such as glass or silicon, and is used to generate “mother” structures in an epoxy or plastic using one of several types of replication procedures. The “mother” structure, in turn, is coated with a thin film of conducive material, and used as a mold to electroplate a thick film of nickel. The nickel “daughter” is released from the plastic “mother” structure. Finally, the nickel “daughter” is bonded to a cylindrical drum, which is used to continuously emboss the surface relief structure into a plastic film. A device structure that uses an embossed plastic substrate is shown in FIG. 5 . Following embossing, the plastic structure is overcoated with a thin film of high refractive index material, and optionally coated with a planarizing, cover layer polymer, and cut to appropriate size.
A substrate for a SWS biosensor can comprise, for example, glass, plastic or epoxy. Optionally, a substrate and a two-dimensional grating or one-dimensional grating can comprise a single unit. That is, a grating and substrate are formed from the same material, for example, glass, plastic, or epoxy. The surface of a single unit comprising the grating is coated with a material having a high refractive index, for example, zinc sulfide, titanium dioxide, tantalum oxide, and silicon nitride. One or more specific binding substances can be immobilized on the surface of the material having a high refractive index or on an optional cover layer.
A biosensor of the invention can further comprise a cover layer on the surface of a two-dimensional grating or one-dimensional grating opposite of a substrate layer. Where a cover layer is present, the one or more specific binding substances are immobilized on the surface of the cover layer opposite of the grating. Preferably, a cover layer comprises a material that has a lower refractive index than a material that comprises the grating. A cover layer can be comprised of, for example, glass (including spin-on glass (SOG)), epoxy, or plastic.
For example, various polymers that meet the refractive index requirement of a biosensor can be used for a cover layer. SOG can be used due to its favorable refractive index, ease of handling, and readiness of being activated with specific binding substances using the wealth of glass surface activation techniques. When the flatness of the biosensor surface is not an issue for a particular system setup, a grating structure of SiN/glass can directly be used as the sensing surface, the activation of which can be done using the same means as on a glass surface.
Resonant reflection can also be obtained without a planarizing cover layer over a two-dimensional grating or one-dimensional grating. For example, a biosensor can contain only a substrate coated with a structured thin film layer of high refractive index material. Without the use of a planarizing cover layer, the surrounding medium (such as air or water) fills the grating. Therefore, specific binding substances are immobilized to the biosensor on all surfaces of a grating exposed to the specific binding substances, rather than only on an upper surface.
In general, a biosensor of the invention will be illuminated with white light that will contain light of every polarization angle. The orientation of the polarization angle with respect to repeating features in a biosensor grating will determine the resonance wavelength. For example, a one-dimensional linear grating biosensor structure consisting of a set of repeating lines and spaces will have two optical polarizations that can generate separate resonant reflections. Light that is polarized perpendicularly to the lines is called “s-polarized,” while light that is polarized parallel to the lines is called “p-polarized.” Both the s and p components of incident light exist simultaneously in an unfiltered illumination beam, and each generates a separate resonant signal. A biosensor structure can generally be designed to optimize the properties of only one polarization (generally the s-polarization), and the non-optimized polarization is easily removed by a polarizing filter.
In order to remove the polarization dependence, so that every polarization angle generates the same resonant reflection spectra, an alternate biosensor structure can be used that consists of a set of concentric rings. In this structure, the difference between the inside diameter and the outside diameter of each concentric ring is equal to about one-half of a grating period. Each successive ring has an inside diameter that is about one grating period greater than the inside diameter of the previous ring. The concentric ring pattern extends to cover a single sensor location—such as a microarray spot or a microtiter plate well. Each separate microarray spot or microtiter plate well has a separate concentric ring pattern centered within it. e.g., FIG. 49 . All polarization directions of such a structure have the same cross-sectional profile. The concentric ring structure must be illuminated precisely on-center to preserve polarization independence. The grating period of a concentric ring structure is less than the wavelength of the resonantly reflected light. The grating period is about 0.01 micron to about 1 micron. The grating depth is about 0.01 to about 1 micron.
In another embodiment, an array of holes or posts are arranged to closely approximate the concentric circle structure described above without requiring the illumination beam to be centered upon any particular location of the grid. See e.g. FIG. 50 . Such an array pattern is automatically generated by the optical interference of three laser beams incident on a surface from three directions at equal angles. In this pattern, the holes (or posts) are centered upon the corners of an array of closely packed hexagons as shown in FIG. 50 . The holes or posts also occur in the center of each hexagon. Such a hexagonal grid of holes or posts has three polarization directions that “see” the same cross-sectional profile. The hexagonal grid structure, therefore, provides equivalent resonant reflection spectra using light of any polarization angle. Thus, no polarizing filter is required to remove unwanted reflected signal components. The period of the holes or posts can be about 0.01 microns to about 1 micron and the depth or height can be about 0.01 microns to about 1 micron.
The invention provides a resonant reflection structures and transmission filter structures comprising concentric circle gratings and hexagonal grids of holes or posts. For a resonant reflection structure, light output is measured on the same side of the structure as the illuminating light beam. For a transmission filter structure, light output is measured on the opposite side of the structure as the illuminating beam. The reflected and transmitted signals are complementary. That is, if a wavelength is strongly reflected, it is weakly transmitted. Assuming no energy is absorbed in the structure itself, the reflected+transmitted energy at any given wavelength is constant. The resonant reflection structure and transmission filters are designed to give a highly efficient reflection at a specified wavelength. Thus, a reflection filter will “pass” a narrow band of wavelengths, while a transmission filter will “cut” a narrow band of wavelengths from incident light.
A resonant reflection structure or a transmission filter structure can comprise a two-dimensional grating arranged in a pattern of concentric circles. A resonant reflection structure or transmission filter structure can also comprise a hexagonal grid of holes or posts. When these structure are illuminated with an illuminating light beam, a reflected radiation spectrum is produced that is independent of an illumination polarization angle of the illuminating light beam. When these structures are illuminated a resonant grating effect is produced on the reflected radiation spectrum, wherein the depth and period of the two-dimensional grating or hexagonal grid of holes or posts are less than the wavelength of the resonant grating effect. These structures reflect a narrow band of light when the structure is illuminated with a broadband of light.
Resonant reflection structures and transmission filter structures of the invention can be used as biosensors. For example, one or more specific binding substances can be immobilized on the hexagonal grid of holes or posts or on the two-dimensional grating arranged in concentric circles.
In one embodiment of the invention, a reference resonant signal is provided for more accurate measurement of peak resonant wavelength shifts. The reference resonant signal can cancel out environmental effects, including, for example, temperature. A reference signal can be provided using a resonant reflection superstructure that produces two separate resonant wavelengths. A transparent resonant reflection superstructure can contain two sub-structures. A first sub-structure comprises a first one- or two-dimensional grating with a top and a bottom surface. The top surface of a one- or two-dimensional grating comprises the grating surface. The first one- or two-dimensional grating can comprise one or more specific binding substances immobilized on its top surface. The top surface of the first one- or two-dimensional grating is in contact with a test sample. An optional substrate layer can be present to support the bottom surface of the first one- or two-dimensional grating. The substrate layer comprises a top and bottom surface. The top surface of the substrate is in contact with, and supports the bottom surface of the first one- or two-dimensional grating.
A second sub-structure comprises a second one- or two-dimensional grating with a top surface and a bottom surface. The second one- or two-dimensional grating is not in contact with a test sample. The second one- or two-dimensional grating can be fabricated onto the bottom surface of the substrate that supports the first one- or two-dimensional grating. Where the second one- or two-dimensional grating is fabricated on the substrate that supports the first one- or two-dimensional grating, the bottom surface of the second one- or two-dimensional grating can be fabricated onto the bottom surface of the substrate. Therefore, the top surface of the second one- or two-dimensional grating will face the opposite direction of the top surface of the first one- or two-dimensional grating.
The top surface of the second one- or two-dimensional grating can also be attached directly to the bottom surface of the first sub-structure. In this embodiment the top surface of the second one- or two-dimensional grating will face the same direction as the top surface of the first one- or two-dimensional grating. A substrate can support the bottom surface of the second one- or two-dimensional grating in this embodiment.
Because the second sub-structure is not in physical contact with the test sample, its peak resonant wavelength is not subject to changes in the optical density of the test media, or deposition of specific binding substances or binding partners on the surface of the first one- or two-dimensional grating. Therefore, such a superstructure produces two resonant signals. Because the location of the peak resonant wavelength in the second sub-structure is fixed, the difference in peak resonant wavelength between the two sub-structures provides a relative means for determining the amount of specific binding substances or binding partners or both deposited on the top surface of the first substructure that is exposed to the test sample.
A biosensor superstructure can be illuminated from its top surface or from its bottom surface, or from both surfaces. The peak resonance reflection wavelength of the first substructure is dependent on the optical density of material in contact with the superstructure surface, while the peak resonance reflection wavelength of the second substructure is independent of the optical density of material in contact with the superstructure surface.
In one embodiment of the invention, a biosensor is illuminated from the bottom surface of the biosensor. Approximately 50% of the incident light is reflected from the bottom surface of biosensor without reaching the active (top) surface of the biosensor. A thin film or physical structure can be included in a biosensor composition that is capable of maximizing the amount of light that is transmitted to the upper surface of the biosensor while minimizing the reflected energy at the resonant wavelength. The anti-reflection thin film or physical structure of the bottom surface of the biosensor can comprise, for example, a single dielectric thin film, a stack of multiple dielectric thin films, or a “motheye” structure that is embossed into the bottom biosensor surface. An example of a motheye structure is disclosed in Hobbs, et al. “Automated interference lithography system for generation of sub-micron feature size patterns,” Proc. 1999 Micromachine Technology for Diffracting and Holographic Optics, Society of Photo - Optical Instrumentation Engineers , p. 124-135, (1999).
In one embodiment of the invention, an optical device is provided. An optical device comprises a structure similar to any biosensor of the invention; however, an optical device does not comprise one of more binding substances immobilized on the two-dimensional grating. An optical device can be used as a narrow band optical filter.
In one embodiment of the invention, an interaction of a first molecule with a second test molecule can be detected. A SWS biosensor as described above is used; however, there are no specific binding substances immobilized on its surface. Therefore, the biosensor comprises 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 biosensor is illuminated a resonant grating effect is produced on the reflected radiation spectrum, and the depth and period of the grating are less than the wavelength of the resonant grating effect.
To detect an interaction of a first molecule with a second test molecule, a mixture of the first and second molecules is applied to a distinct location on a biosensor. A distinct location can be one spot or well on a biosensor or can be a large area on a biosensor. A mixture of the first molecule with a third control molecule is also applied to a distinct location on a biosensor. The biosensor can be the same biosensor as described above, or can be a second biosensor. If the biosensor is the same biosensor, a second distinct location can be used for the mixture of the first molecule and the third control molecule. Alternatively, the same distinct biosensor location can be used after the first and second molecules are washed from the biosensor. The third control molecule does not interact with the first molecule and is about the same size as the first molecule. A shift in the reflected wavelength of light from the distinct locations of the biosensor or biosensors is measured. If the shift in the reflected wavelength of light from the distinct location having the first molecule and the second test molecule is greater than the shift in the reflected wavelength from the distinct location having the first molecule and the third control molecule, then the first molecule and the second test molecule interact. Interaction can be, for example, hybridization of nucleic acid molecules, specific binding of an antibody or antibody fragment to an antigen, and binding of polypeptides. A first molecule, second test molecule, or 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 fragment, Fv fragment, small organic molecule, cell, virus, and bacteria.
Specific Binding Substances and Binding Partners
One or more specific binding substances are immobilized on the one- or two-dimensional grating or cover layer, if present, by for example, physical adsorption or by chemical binding. A specific binding substance can be, for example, a nucleic acid, polypeptide, antigen, polyclonal antibody, monoclonal antibody, single chain antibody (scFv), F(ab) fragment, F(ab′) 2 fragment, Fv fragment, small organic molecule, cell, virus, bacteria, polymer, peptide solutions, single- or double-stranded DNA solutions, RNA solutions, solutions containing compounds from a combinatorial chemical library, or biological sample. A biological sample can be for example, blood, plasma, serum, gastrointestinal secretions, homogenates of tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, or prostatitc fluid.
Preferably, one or more specific binding substances are arranged in a microarray of distinct locations on a biosensor. A microarray of specific binding substances comprises one or more specific binding substances on a surface of a biosensor of the invention such that a surface contains many distinct locations, each with a different specific binding substance or with a different amount of a specific binding substance. For example, an array can comprise 1, 10, 100, 1,000, 10,000, or 100,000 distinct locations. Such a biosensor surface is called a microarray because one or more specific binding substances are typically laid out in a regular grid pattern in x-y coordinates. However, a microarray of the invention can comprise one or more specific binding substance laid out in any type of regular or irregular pattern. For example, distinct locations can define a microarray of spots of one or more specific binding substances. A microarray spot can be about 50 to about 500 microns in diameter. A microarray spot can also be about 150 to about 200 microns in diameter. One or more specific binding substances can be bound to their specific binding partners.
A microarray on a biosensor of the invention can be created by placing microdroplets of one or more specific binding substances onto, for example, an x-y grid of locations on a one- or two-dimensional grating or cover layer surface. When the biosensor is exposed to a test sample comprising one or more binding partners, the binding partners will be preferentially attracted to distinct locations on the microarray that comprise specific binding substances that have high affinity for the binding partners. Some of the distinct locations will gather binding partners onto their surface, while other locations will not.
A specific binding substance specifically binds to a binding partner that is added to the surface of a biosensor of the invention. A specific binding substance specifically binds to its binding partner, but does not substantially bind other binding partners added to the surface of a biosensor. For example, where the specific binding substance is an antibody and its binding partner is a particular antigen, the antibody specifically binds to the particular antigen, but does not substantially bind other antigens. A binding partner can be, for example, a nucleic acid, polypeptide, antigen, polyclonal antibody, monoclonal antibody, single chain antibody (scFv), F(ab) fragment, F(ab′) 2 fragment, Fv fragment, small organic molecule, cell, virus, bacteria, polymer, peptide solutions, single- or double-stranded DNA solutions, RNA solutions, solutions containing compounds from a combinatorial chemical library and biological sample. A biological sample can be, for example, blood, plasma, serum, gastrointestinal secretions, homogenates of tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, and prostatitc fluid.
One example of a microarray of the invention is a nucleic acid microarray, in which each distinct location within the array contains a different nucleic acid molecule. In this embodiment, the spots within the nucleic acid microarray detect complementary chemical binding with an opposing strand of a nucleic acid in a test sample.
While microtiter plates are the most common format used for biochemical assays, microarrays are increasingly seen as a means for maximizing the number of biochemical interactions that can be measured at one time while minimizing the volume of precious reagents. By application of specific binding substances with a microarray spotter onto a biosensor of the invention, specific binding substance densities of 10,000 specific binding substances/in 2 can be obtained. By focusing an illumination beam to interrogate a single microarray location, a biosensor can be used as a label-free microarray readout system.
Immobilization or One or More Specific Binding Substances
Immobilization of one or more binding substances onto a biosensor is performed so that a specific binding substance will not be washed away by rinsing procedures, and so that its binding to binding partners in a test sample is unimpeded by the biosensor surface. Several different types of surface chemistry strategies have been implemented for covalent attachment of specific binding substances to, for example, glass for use in various types of microarrays and biosensors. These same methods can be readily adapted to a biosensor of the invention. Surface preparation of a biosensor so that it contains the correct functional groups for binding one or more specific binding substances is an integral part of the biosensor manufacturing process.
One or more specific binding substances can be attached to a biosensor surface by physical adsorption (i.e., without the use of chemical linkers) or by chemical binding (i.e., with the use of chemical linkers). Chemical binding can generate stronger attachment of specific binding substances on a biosensor surface and provide defined orientation and conformation of the surface-bound molecules.
Several examples of chemical binding of specific binding substances to a biosensor of the invention appear in Example 8, below. Other types of chemical binding include, for example, amine activation, aldehyde activation, and nickel activation. These surfaces can be used to attach several different types of chemical linkers to a biosensor surface, as shown in FIG. 6 . While an amine surface can be used to attach several types of linker molecules, an aldehyde surface can be used to bind proteins directly, without an additional linker. A nickel surface can be used to bind molecules that have an incorporated histidine (“his”) tag. Detection of “his-tagged” molecules with a nickel-activated surface is well known in the art (Whitesides, Anal Chem. 68, 490, (1996)).
Immobilization of specific binding substances to plastic, epoxy, or high refractive index material can be performed essentially as described for immobilization to glass. However, the acid wash step can be eliminated where such a treatment would damage the material to which the specific binding substances are immobilized.
For the detection of binding partners at concentrations less than about ˜0.1 ng/ml, it is preferable to amplify and transduce binding partners bound to a biosensor into an additional layer on the biosensor surface. The increased mass deposited on the biosensor can be easily detected as a consequence of increased optical path length. By incorporating greater mass onto a biosensor surface, the optical density of binding partners on the surface is also increased, thus rendering a greater resonant wavelength shift than would occur without the added mass. The addition of mass can be accomplished, for example, enzymatically, through a “sandwich” assay, or by direct application of mass to the biosensor surface in the form of appropriately conjugated beads or polymers of various size and composition. This principle has been exploited for other types of optical biosensors to demonstrate sensitivity increases over 1500× beyond sensitivity limits achieved without mass amplification. See, e.g., Jenison et al., “Interference-based detection of nucleic acid targets on optically coated silicon,” Nature Biotechnology, 19: 62-65, 2001.
As an example, FIG. 7A shows that an NH 2 -activated biosensor surface can have a specific binding substance comprising a single-strand DNA capture probe immobilized on the surface. The capture probe interacts selectively with its complementary target binding partner. The binding partner, in turn, can be designed to include a sequence or tag that will bind a “detector” molecule. As shown in FIG. 7A , a detector molecule can contain, for example, a linker to horseradish peroxidase (HRP) that, when exposed to the correct enzyme, will selectively deposit additional material on the biosensor only where the detector molecule is present. Such a procedure can add, for example, 300 angstroms of detectable biomaterial to the biosensor within a few minutes.
A “sandwich” approach can also be used to enhance detection sensitivity. In this approach, a large molecular weight molecule can be used to amplify the presence of a low molecular weight molecule. For example, a binding partner with a molecular weight of, for example, about 0.1 kDa to about 20 kDa, can be tagged with, for example, succinimidyl-6-[a-methyl-a-(2-pyridyl-dithio)toluamido]hexanoate (SMPT), or dimethylpimelimidate (DMP), histidine, or a biotin molecule, as shown in FIG. 7B . Where the tag is biotin, the biotin molecule will binds strongly with streptavidin, which has a molecular weight of 60 kDa. Because the biotin/streptavidin interaction is highly specific, the streptavidin amplifies the signal that would be produced only by the small binding partner by a factor of 60.
Detection sensitivity can be further enhanced through the use of chemically derivatized small particles. “Nanoparticles” made of colloidal gold, various plastics, or glass with diameters of about 3-300 nm can be coated with molecular species that will enable them to covalently bind selectively to a binding partner. For example, as shown in FIG. 7C , nanoparticles that are covalently coated with streptavidin can be used to enhance the visibility of biotin-tagged binding partners on the biosensor surface. While a streptavidin molecule itself has a molecular weight of 60 kDa, the derivatized bead can have a molecular weight of any size, including, for example, 60 KDa. Binding of a large bead will result in a large change in the optical density upon the biosensor surface, and an easily measurable signal. This method can result in an approximately 1000× enhancement in sensitivity resolution.
Surface-Relief Volume Diffractive Biosensors
Another embodiment of the invention is a biosensor that comprises volume surface-relief volume diffractive structures (a SRVD biosensor). SRVD biosensors have a surface that reflect predominantly at a particular narrow band of optical wavelengths when illuminated with a broad band of optical wavelengths. Where specific binding substances and/or binding partners are immobilized on a SRVD biosensor, the reflected wavelength of light is shifted. One-dimensional surfaces, such as thin film interference filters and Bragg reflectors, can select a narrow range of reflected or transmitted wavelengths from a broadband excitation source, however, the deposition of additional material, such as specific binding substances and/or binding partners onto their upper surface results only in a change in the resonance linewidth, rather than the resonance wavelength. In contrast, SRVD biosensors have the ability to alter the reflected wavelength with the addition of material, such as specific binding substances and/or binding partners to the surface.
A SRVD biosensor comprises a sheet material having a first and second surface. The first surface of the sheet material defines relief volume diffraction structures. A sheet material can be comprised of, for example, plastic, glass, semiconductor wafer, or metal film.
A relief volume diffractive structure can be, for example, a two-dimensional grating, as described above, or a three-dimensional surface-relief volume diffractive grating. The depth and period of relief volume diffraction structures are less than the resonance wavelength of light reflected from a biosensor.
A three-dimensional surface-relief volume diffractive grating can be, for example, a three-dimensional phase-quantized terraced surface relief pattern whose groove pattern resembles a stepped pyramid. When such a grating is illuminated by a beam of broadband radiation, light will be coherently reflected from the equally spaced terraces at a wavelength given by twice the step spacing times the index of refraction of the surrounding medium. Light of a given wavelength is resonantly diffracted or reflected from the steps that are a half-wavelength apart, and with a bandwidth that is inversely proportional to the number of steps. The reflected or diffracted color can be controlled by the deposition of a dielectric layer so that a new wavelength is selected, depending on the index of refraction of the coating.
A stepped-phase structure can be produced first in photoresist by coherently exposing a thin photoresist film to three laser beams, as described previously. See e.g., Cowen, “The recording and large scale replication of crossed holographic grating arrays using multiple beam interferometry,” in International Conference on the Application, Theory, and Fabrication of Periodic Structures, Diffraction Gratings, and Moire Phenomena II , Lerner, ed., Proc. Soc. Photo-Opt. Instrum. Eng., 503, 120-129, 1984; Cowen, “Holographic honeycomb microlens,” Opt. Eng. 24, 796-802 (1985); Cowen & Slafer, “The recording and replication of holographic micropatterns for the ordering of photographic emulsion grains in film systems,” J. Imaging Sci. 31, 100-107, 1987. The nonlinear etching characteristics of photoresist are used to develop the exposed film to create a three-dimensional relief pattern. The photoresist structure is then replicated using standard embossing procedures. For example, a thin silver film is deposited over the photoresist structure to form a conducting layer upon which a thick film of nickel can be electroplated. The nickel “master” plate is then used to emboss directly into a plastic film, such as vinyl, that has been softened by heating or solvent.
The theory describing the design and fabrication of three-dimensional phase-quantized terraced surface relief pattern that resemble stepped pyramids is described: Cowen, “Aztec surface-relief volume diffractive structure,” J. Opt. Soc. Am. A, 7:1529 (1990).
An example of a three-dimensional phase-quantized terraced surface relief pattern is a pattern that resembles a stepped pyramid. Each inverted pyramid is approximately 1 micron in diameter, preferably, each inverted pyramid can be about 0.5 to about 5 microns diameter, including for example, about 1 micron. The pyramid structures can be close-packed so that a typical microarray spot with a diameter of 150-200 microns can incorporate several hundred stepped pyramid structures. The relief volume diffraction structures have a period of about 0.1 to about 1 micron and a depth of about 0.1 to about 1 micron. FIG. 8 demonstrates how individual microarray locations (with an entire microarray spot incorporating hundreds of pyramids now represented by a single pyramid for one microarray spot) can be optically queried to determine if specific binding substances or binding partners are adsorbed onto the surface. When the structure is illuminated with white light, structures without significant bound material will reflect wavelengths determined by the step height of the structure. When higher refractive index material, such as binding partners or specific binding substances, are incorporated over the reflective metal surface, the reflected wavelength is modified to shift toward longer wavelengths. The color that is reflected from the terraced step structure is theoretically given as twice the step height times the index of refraction of a reflective material that is coated onto the first surface of a sheet material of a SRVD biosensor. A reflective material can be, for example silver, aluminum, or gold.
One or more specific binding substances, as described above, are immobilized on the reflective material of a SRVD biosensor. One or more specific binding substances can be arranged in microarray of distinct locations, as described above, on the reflective material. FIG. 9 provides an example of a 9-element microarray biosensor. Many individual grating structures, represented by small circles, lie within each microarray spot. The microarray spots, represented by the larger circles, will reflect white light in air at a wavelength that is determined by the refractive index of material on their surface. Microarray locations with additional adsorbed material will have reflected wavelengths that are shifted toward longer wavelengths, represented by the larger circles.
Because the reflected wavelength of light from a SRVD biosensor is confined to a narrow bandwidth, very small changes in the optical characteristics of the surface manifest themselves in easily observed changes in reflected wavelength spectra. The narrow reflection bandwidth provides a surface adsorption sensitivity advantage compared to reflectance spectrometry on a flat surface.
A SRVD biosensor reflects light predominantly at a first single optical wavelength when illuminated with a broad band of optical wavelengths, and reflects light at a second single optical wavelength when one or more specific binding substances are immobilized on the reflective surface. The reflection at the second optical wavelength results from optical interference. A SRVD biosensor also reflects light at a third single optical wavelength when the one or more specific binding substances are bound to their respective binding partners, due to optical interference.
Readout of the reflected color can be performed serially by focusing a microscope objective onto individual microarray spots and reading the reflected spectrum, or in parallel by, for example, projecting the reflected image of the microarray onto a high resolution color CCD camera.
A SRVD biosensor can be manufactured by, for example, producing a metal master plate, and stamping a relief volume diffractive structure into, for example, a plastic material like vinyl. After stamping, the surface is made reflective by blanket deposition of, for example, a thin metal film such as gold, silver, or aluminum. Compared to MEMS-based biosensors that rely upon photolithography, etching, and wafer bonding procedures, the manufacture of a SRVD biosensor is very inexpensive.
Liquid-Containing Vessels
A SWS or SRVD biosensor of the invention can comprise an inner surface, for example, a bottom surface of a liquid-containing vessel. A liquid-containing vessel can be, for example, a microtiter plate well, a test tube, a petri dish, or a microfluidic channel. One embodiment of this invention is a SWS or SRVD biosensor that is incorporated into any type of microtiter plate. For example, a SWS biosensor or SRVD biosensor can be incorporated into the bottom surface of a microtiter plate by assembling the walls of the reaction vessels over the resonant reflection surface, as shown in FIG. 10 , so that each reaction “spot” can be exposed to a distinct test sample. Therefore, each individual microtiter plate well can act as a separate reaction vessel. Separate chemical reactions can, therefore, occur within adjacent wells without intermixing reaction fluids and chemically distinct test solutions can be applied to individual wells.
Several methods for attaching a biosensor of the invention to the bottom surface of bottomless microtiter plates can be used, including, for example, adhesive attachment, ultrasonic welding, and laser welding.
The most common assay formats for pharmaceutical high-throughput screening laboratories, molecular biology research laboratories, and diagnostic assay laboratories are microtiter plates. The plates are standard-sized plastic cartridges that can contain 96, 384, or 1536 individual reaction vessels arranged in a grid. Due to the standard mechanical configuration of these plates, liquid dispensing, robotic plate handling, and detection systems are designed to work with this common format. A biosensor of the invention can be incorporated into the bottom surface of a standard microtiter plate. See, e.g., FIG. 10 . Because the biosensor surface can be fabricated in large areas, and because the readout system does not make physical contact with the biosensor surface, an arbitrary number of individual biosensor areas can be defined that are only limited by the focus resolution of the illumination optics and the x-y stage that scans the illumination/detection probe across the biosensor surface.
Holding Fixtures
Any number of biosensors that are, for example, about 1 mm 2 to about 5 mm 2 , and preferably less than about 3×3 mm 2 can be arranged onto a holding fixture that can simultaneously dip the biosensors into separate liquid-containing vessels, such as wells of a microtiter plate, for example, a 96-, 384-, or 1536-well microtiter plate. See e.g., FIG. 11 . Each of the biosensors can contain multiple distinct locations. A holding fixture has one or more biosensors attached to the holding fixture so that each individual biosensor can be lowered into a separate liquid-containing vessel. A holding fixture can comprise plastic, epoxy or metal. For example, 50, 96, 384, or 1,000, or 1,536 biosensors can be arranged on a holding fixture, where each biosensor has 25, 100, 500, or 1,000 distinct locations. As an example, where 96 biosenors are attached to a holding fixture and each biosensor comprises 100 distinct locations, 9600 biochemical assays can be performed simultaneously.
Methods of Using SWS and SRVD Biosensors
SWS and SRVD biosensors of the invention can be used to study one or a number of specific binding substance/binding partner interactions in parallel. Binding of one or more specific binding substances to their respective binding partners can be detected, without the use of labels, by applying one or more binding partners to a SWS or SRVD biosensor that have one or more specific binding substances immobilized on their surfaces. A SWS biosensor is illuminated with light and a maxima in reflected wavelength, or a minima in transmitted wavelength of light is detected from the biosensor. If one or more specific binding substances have bound to their respective binding partners, then the reflected wavelength of light is shifted as compared to a situation where one or more specific binding substances have not bound to their respective binding partners. Where a SWS biosensor is coated with an array of distinct locations containing the one or more specific binding substances, then a maxima in reflected wavelength or minima in transmitted wavelength of light is detected from each distinct location of the biosensor.
A SRVD biosensor is illuminated with light after binding partners have been added and the reflected wavelength of light is detected from the biosensor. Where one or more specific binding substances have bound to their respective binding partners, the reflected wavelength of light is shifted.
In one embodiment of the invention, a variety of specific binding substances, for example, antibodies, can be immobilized in an array format onto a biosensor of the invention. The biosensor is then contacted with a test sample of interest comprising binding partners, such as proteins. Only the proteins that specifically bind to the antibodies immobilized on the biosensor remain bound to the biosensor. Such an approach is essentially a large-scale version of an enzyme-linked immunosorbent assay; however, the use of an enzyme or fluorescent label is not required.
The activity of an enzyme can be detected by applying one or more enzymes to a SWS or SRVD biosensor to which one or more specific binding substances have been immobilized. The biosensor is washed and illuminated with light. The reflected wavelength of light is detected from the biosensor. Where the one or more enzymes have altered the one or more specific binding substances of the biosensor by enzymatic activity, the reflected wavelength of light is shifted.
Additionally, a test sample, for example, cell lysates containing binding partners can be applied to a biosensor of the invention, followed by washing to remove unbound material. The binding partners that bind to a biosensor can be eluted from the biosensor and identified by, for example, mass spectrometry. Optionally, a phage DNA display library can be applied to a biosensor of the invention followed by washing to remove unbound material. Individual phage particles bound to the biosensor can be isolated and the inserts in these phage particles can then be sequenced to determine the identity of the binding partner.
For the above applications, and in particular proteomics applications, the ability to selectively bind material, such as binding partners from a test sample onto a biosensor of the invention, followed by the ability to selectively remove bound material from a distinct location of the biosensor for further analysis is advantageous. Biosensors of the invention are also capable of detecting and quantifying the amount of a binding partner from a sample that is bound to a biosensor array distinct location by measuring the shift in reflected wavelength of light. For example, the wavelength shift at one distinct biosensor location can be compared to positive and negative controls at other distinct biosensor locations to determine the amount of a binding partner that is bound to a biosensor array distinct location.
SWS and Electrically Conducting Material
An optional biosensor structure can further enable a biosensor array to selectively attract or repel binding partners from individual distinct locations on a biosensor. As is well known in the art, an electromotive force can be applied to biological molecules such as nucleic acids and amino acids subjecting them to an electric field. Because these molecules are electronegative, they are attracted to a positively charged electrode and repelled by a negatively charged electrode.
A grating structure of a resonant optical biosensor can be built using an electrically conducting material rather than an electrically insulating material. An electric field can be applied near the biosensor surface. Where a grating operates as both a resonant reflector biosensor and as an electrode, the grating comprises a material that is both optically transparent near the resonant wavelength, and has low resistivity. In one embodiment of the invention, the material is indium tin oxide, InSn x O 1-x (ITO). ITO is commonly used to produce transparent electrodes for flat panel optical displays, and is therefore readily available at low cost on large glass sheets. The refractive index of ITO can be adjusted by controlling x, the fraction of Sn that is present in the material. Because the liquid test sample solution will have mobile ions (and will therefore be an electrical conductor) it is necessary for the ITO electrodes to be coated with an insulating material. For the resonant optical biosensor, a grating layer is coated with a layer with lower refractive index material. Materials such as cured photoresist (n=1.65), cured optical epoxy (n=1.5), and glass (n=1.4-1.5) are strong electrical insulators that also have a refractive index that is lower than ITO (n=2.0-2.65). A cross-sectional diagram of a biosensor that incorporates an ITO grating is shown in FIG. 48 . n 1 represents the refractive index of an electrical insulator. n 2 represents the refractive index of a two-dimensional grating. t 1 represents the thickness of the electrical insulator. t 2 represents the thickness of the two-dimensional grating. n bio represents the refractive index of one or more specific binding substances and t BIO represents the thickness of the one or more specific binding substances.
A grating can be a continuous sheet of ITO that contains an array of regularly spaced holes. The holes are filled in with an electrically insulating material, such as cured photoresist. The electrically insulating layer overcoats the ITO grating so that the upper surface of the structure is completely covered with electrical insulator, and so that the upper surface is substantially flat. When the biosensor is illuminated with light a resonant grating effect is produced on the reflected radiation spectrum. The depth and the period of the grating are less than the wavelength of the resonant grating effect.
As shown in FIG. 12 and FIG. 13 , a single electrode can comprise a region that contains many grating periods. Building two or more separate grating regions on the same substrate surface creates an array of biosensor electrodes. Electrical contact to each biosensor electrode is provided using an electrically conducting trace that is built from the same material as the conductor within the biosensor electrode. The conducting trace is connected to a voltage source that can apply an electrical potential to the electrode. To apply an electrical potential to the biosensor that is capable of attracting or repelling a molecule near the electrode surface, a biosensor upper surface can be immersed in a liquid sample as shown in FIG. 14 . A “common” electrode can be placed within the sample liquid, and a voltage can be applied between one selected biosensor electrode region and the common electrode. In this way, one, several, or all electrodes can be activated or inactivated at a given time. FIG. 15 illustrates the attraction of electronegative molecules to the biosensor surface when a positive voltage is applied to the electrode, while FIG. 16 illustrates the application of a repelling force such as a reversed electrical charge to electronegative molecules using a negative electrode voltage.
Detection Systems
A detection system can comprise a biosensor of the invention, a light source that directs light to the biosensor, and a detector that detects light reflected from the biosensor. In one embodiment, it is possible to simplify the readout instrumentation by the application of a filter so that only positive results over a determined threshold trigger a detection.
A light source can illuminate a biosensor from its top surface, i.e., the surface to which one or more specific binding substances are immobilized or from its bottom surface. By measuring the shift in resonant wavelength at each distinct location of a biosensor of the invention, it is possible to determine which distinct locations have binding partners bound to them. The extent of the shift can be used to determine the amount of binding partners in a test sample and the chemical affinity between one or more specific binding substances and the binding partners of the test sample.
A biosensor of the invention can be illuminated twice. The first measurement determines the reflectance spectra of one or more distinct locations of a biosensor array with one or more specific binding substances immobilized on the biosensor. The second measurement determines the reflectance spectra after one or more binding partners are applied to a biosensor. The difference in peak wavelength between these two measurements is a measurement of the amount of binding partners that have specifically bound to a biosensor or one or more distinct locations of a biosensor. This method of illumination can control for small nonuniformities in a surface of a biosensor that can result in regions with slight variations in the peak resonant wavelength. This method can also control for varying concentrations or molecular weights of specific binding substances immobilized on a biosensor
Computer simulation can be used to determine the expected dependence between a peak resonance wavelength and an angle of incident illumination. A biosensor structure as shown in FIG. 1 can be for purposes of demonstration. The substrate chosen was glass (n substrate =1.50). The grating is a two-dimensional pattern of silicon nitride squares (t 2 =180 nm, n 2 =2.01 (n=refractive index), k 2 =0.001 (k=absorption coefficient)) with a period of 510 nm, and a filling factor of 56.2% (i.e., 56.2% of the surface is covered with silicon nitride squares while the rest is the area between the squares). The areas between silicon nitride squares are filled with a lower refractive index material. The same material also covers the squares and provides a uniformly flat upper surface. For this simulation, a glass layer was selected (n 1 =1.40) that covers the silicon nitride squares by t 2 =100 nm.
The reflected intensity as a function of wavelength was modeled using GSOLVER software, which utilizes full 3-dimensional vector code using hybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVER calculates diffracted fields and diffraction efficiencies from plane wave illumination of arbitrarily complex grating structures. The illumination can be from any incidence and any polarization.
FIG. 19 plots the dependence of the peak resonant wavelength upon the incident illumination angle. The simulation shows that there is a strong correlation between the angle of incident light, and the peak wavelength that is measured. This result implies that the collimation of the illuminating beam, and the alignment between the illuminating beam and the reflected beam will directly affect the resonant peak linewidth that is measured. If the collimation of the illuminating beam is poor, a range illuminating angles will be incident on the biosensor surface, and a wider resonant peak will be measured than if purely collimated light were incident.
Because the lower sensitivity limit of a biosensor is related to the ability to determine the peak maxima, it is important to measure a narrow resonant peak. Therefore, the use of a collimating illumination system with the biosensor provides for the highest possible sensitivity.
One type of detection system for illuminating the biosensor surface and for collecting the reflected light is a probe containing, for example, six illuminating optical fibers that are connected to a light source, and a single collecting optical fiber connected to a spectrometer. The number of fibers is not critical, any number of illuminating or collecting fibers are possible. The fibers are arranged in a bundle so that the collecting fiber is in the center of the bundle, and is surrounded by the six illuminating fibers. The tip of the fiber bundle is connected to a collimating lens that focuses the illumination onto the surface of the biosensor.
In this probe arrangement, the illuminating and collecting fibers are side-by-side. Therefore, when the collimating lens is correctly adjusted to focus light onto the biosensor surface, one observes six clearly defined circular regions of illumination, and a central dark region. Because the biosensor does not scatter light, but rather reflects a collimated beam, no light is incident upon the collecting fiber, and no resonant signal is observed. Only by defocusing the collimating lens until the six illumination regions overlap into the central region is any light reflected into the collecting fiber. Because only defocused, slightly uncollimated light can produce a signal, the biosensor is not illuminated with a single angle of incidence, but with a range of incident angles. The range of incident angles results in a mixture of resonant wavelengths due to the dependence shown in FIG. 19 . Thus, wider resonant peaks are measured than might otherwise be possible.
Therefore, it is desirable for the illuminating and collecting fiber probes to spatially share the same optical path. Several methods can be used to co-locate the illuminating and collecting optical paths. For example, a single illuminating fiber, which is connected at its first end to a light source that directs light at the biosensor, and a single collecting fiber, which is connected at its first end to a detector that detects light reflected from the biosensor, can each be connected at their second ends to a third fiber probe that can act as both an illuminator and a collector. The third fiber probe is oriented at a normal angle of incidence to the biosensor and supports counter-propagating illuminating and reflecting optical signals. An example of such a detection system is shown in FIG. 18 .
Another method of detection involves the use of a beam splitter that enables a single illuminating fiber, which is connected to a light source, to be oriented at a 90 degree angle to a collecting fiber, which is connected to a detector. Light is directed through the illuminating fiber probe into the beam splitter, which directs light at the biosensor. The reflected light is directed back into the beam splitter, which directs light into the collecting fiber probe. An example of such a detection device is shown in FIG. 20 . A beam splitter allows the illuminating light and the reflected light to share a common optical path between the beam splitter and the biosensor, so perfectly collimated light can be used without defocusing.
Angular Scanning
Detection systems of the invention are based on collimated white light illumination of a biosensor surface and optical spectroscopy measurement of the resonance peak of the reflected beam. Molecular binding on the surface of a biosensor is indicated by a shift in the peak wavelength value, while an increase in the wavelength corresponds to an increase in molecular absorption.
As shown in theoretical modeling and experimental data, the resonance peak wavelength is strongly dependent on the incident angle of the detection light beam. FIG. 19 depicts this dependence as modeled for a biosensor of the invention. Because of the angular dependence of the resonance peak wavelength, the incident white light needs to be well collimated. Angular dispersion of the light beam broadens the resonance peak, and reduces biosensor detection sensitivity. In addition, the signal quality from the spectroscopic measurement depends on the power of the light source and the sensitivity of the detector. In order to obtain a high signal-to-noise ratio, an excessively long integration time for each detection location can be required, thus lengthening overall time to readout a biosensor plate. A tunable laser source can be used for detection of grating resonance, but is expensive.
In one embodiment of the invention, these disadvantages are addressed by using a laser beam for illumination of a biosensor, and a light detector for measurement of reflected beam power. A scanning mirror device can be used for varying the incident angle of the laser beam, and an optical system is used for maintaining collimation of the incident laser beam. See, e.g., “Optical Scanning” (Gerald F. Marchall ed., Marcel Dekker (1991). Any type of laser scanning can be used. For example, a scanning device that can generate scan lines at a rate of about 2 lines to about 1,000 lines per second is useful in the invention. In one embodiment of the invention, a scanning device scans from about 50 lines to about 300 lines per second.
In one embodiment, the reflected light beam passes through part of the laser scanning optical system, and is measured by a single light detector. The laser source can be a diode laser with a wavelength of, for example, 780 nm, 785 nm, 810 nm, or 830 nm. Laser diodes such as these are readily available at power levels up to 150 mW, and their wavelengths correspond to high sensitivity of Si photodiodes. The detector thus can be based on photodiode biosensors. An example of such a detection system is shown in FIG. 52 . A light source ( 100 ) provides light to a scanner device ( 200 ), which directs the light into an optical system ( 300 ) The optical system ( 300 ) directs light to a biosensor ( 400 ) Light is reflected from the biosensor ( 400 ) to the optical system ( 300 ), which then directs the light into a light signal detector ( 500 ). One embodiment of a detection system is shown in FIG. 21 , which demonstrates that while the scanning mirror changes its angular position, the incident angle of the laser beam on the surface changes by nominally twice the mirror angular displacement. The scanning mirror device can be a linear galvanometer, operating at a frequency of about 2 Hz up to about 120 Hz, and mechanical scan angle of about 10 degrees to about 20 degrees. In this example, a single scan can be completed within about 10 msec. A resonant galvanometer or a polygon scanner can also be used. The example shown in FIG. 21 includes a simple optical system for angular scanning. It consists of a pair of lenses with a common focal point between them. The optical system can be designed to achieve optimized performance for laser collimation and collection of reflected light beam.
The angular resolution depends on the galvanometer specification, and reflected light sampling frequency. Assuming galvanometer resolution of 30 arcsec mechanical, corresponding resolution for biosensor angular scan is 60 arcsec, i.e. 0.017 degree. In addition, assume a sampling rate of 100 ksamples/sec, and 20 degrees scan within 10 msec. As a result, the quantization step is 20 degrees for 1000 samples, i.e. 0.02 degree per sample. In this example, a resonance peak width of 0.2 degree, as shown by Peng and Morris (Experimental demonstration of resonant anomalies in diffraction from two-dimensional gratings, Optics Lett., 21:549 (1996)), will be covered by 10 data points, each of which corresponds to resolution of the detection system.
The advantages of such a detection system includes: excellent collimation of incident light by a laser beam, high signal-to-noise ratio due to high beam power of a laser diode, low cost due to a single element light detector instead of a spectrometer, and high resolution of resonance peak due to angular scanning.
Fiber Probe Biosensor
A biosensor of the invention can occur on the tip of a multi-mode fiber optic probe. This fiber optic probe allows for in vivo detection of biomarkers for diseases and conditions, such as, for example, cardiac artery disease, cancer, inflammation, and sepsis. A single biosensor element (comprising, for example, several hundred grating periods) can be fabricated into the tip of a fiber optic probe, or fabricated from a glass substrate and attached to the tip of a fiber optic probe. See FIG. 17 . A single fiber is used to provide illumination and measure resonant reflected signal.
For example, a fiber probe structure similar to that shown in FIG. 18 can be used to couple an illuminating fiber and detecting fiber into a single counterpropagating fiber with a biosensor embedded or attached to its tip. The fiber optic probe is inserted into a mammalian body, for example, a human body. Illumination and detection of a reflected signal can occur while the probe is inserted in the body.
Mathematical Resonant Peak Determination
The sensitivity of a biosensor is determined by the shift in the location of the resonant peak when material is bound to the biosensor surface. Because of noise inherent in the spectrum, it is preferable to use a procedure for determining an analytical curve—the turning point (i.e., peak) of which is well defined. Furthermore, the peak corresponding to an analytic expression can be preferably determined to greater than sub-sampling-interval accuracy, providing even greater sensitivity.
One embodiment of the invention provides a method for determining a location of a resonant peak for a binding partner in a resonant reflectance spectrum with a colormetric resonant biosensor. The method comprises selecting a set of resonant reflectance data for a plurality of colormetric resonant biosensors or a plurality of biosensor distinct locations. The set of resonant reflectance data is collected by illuminating a colormetric resonant diffractive grating surface with a light source and measuring reflected light at a pre-determined incidence. The colormetric resonant diffractive grating surface is used as a surface binding platform for one or more specific binding substances such that binding partners can be detected without use of a molecular label.
The step of selecting a set of resonant reflectance data can include selecting a set of resonant reflectance data:
x i and y i for i=1, 2, 3, . . . n,
wherein x i is where a first measurement includes a first reflectance spectra of one or more specific binding substances attached to the colormetric resonant diffractive grating surface, y i and a second measurement and includes a second reflectance spectra of the one or more specific binding substances after a plurality of binding partners are applied to colormetric resonant diffractive grating surface including the one or more specific binding substances, and n is a total number of measurements collected.
The set of resonant reflectance data includes a plurality of sets of two measurements, where a first measurement includes a first reflectance spectra of one or more specific binding substances that are attached to the colormetric resonant diffractive grating surface and a second measurement includes a second reflectance spectra of the one or more specific binding substances after one or more binding partners are applied to the colormetric resonant diffractive grating surface including the one or more specific binding substances. A difference in a peak wavelength between the first and second measurement is a measurement of an amount of binding partners that bound to the one or more specific binding substances. A sensitivity of a colormetric resonant biosensor can be determined by a shift in a location of a resonant peak in the plurality of sets of two measurements in the set of resonant reflectance data.
A maximum value for a second measurement from the plurality of sets of two measurements is determined from the set of resonant reflectance data for the plurality of binding partners, wherein the maximum value includes inherent noise included in the resonant reflectance data. A maximum value for a second measurement can include determining a maximum value y k such that:
(y k >=y i ) for all i≠k.
It is determined whether the maximum value is greater than a pre-determined threshold. This can be calculated by, for example, computing a mean of the set of resonant reflectance data; computing a standard deviation of the set of resonant reflectance data; and determining whether ((y k −mean)/standard deviation) is greater than a pre-determined threshold. The pre-determined threshold is determined by the user. The user will determine what amount of sensitivity is desired and will set the pre-determined threshold accordingly.
If the maximum value is greater than a pre-determined threshold a curve-fit region around the determined maximum value is defined. The step of defining a curve-fit region around the determined maximum value can include, for example:
defining a curve-fit region of (2w+1) bins, wherein w is a pre-determined accuracy value;
extracting (x i, , k−w<=i<=k+w); and
extracting (y i, , k−w<=i<=k+w).
A curve-fitting procedure is performed to fit a curve around the curve-fit region, wherein the curve-fitting procedure removes a pre-determined amount of inherent noise included in the resonant reflectance data. A curve-fitting procedure can include, for example:
computing g i =ln y i ;
performing a 2 nd order polynomial fit on g i to obtain g′ i defined on
(x i, , k−w<=i<=k+w);
determining from the 2 nd order polynomial fit coefficients a, b and c of for (ax 2 +bx+c)−; and
computing y′ i =e g′i .
The location of a maximum resonant peak is determined on the fitted curve, which can include, for example, determining a location of maximum reasonant peak (x p =(−b)/2a). A value of the maximum resonant peak is determined, wherein the value of the maximum resonant peak is used to identify an amount of biomolecular binding of the one or more specific binding substances to the one or more binding partners. A value of the maximum resonant peak can include, for example, determining the value with of x p at y′ p .
One embodiment of the invention comprises a computer readable medium having stored therein instructions for causing a processor to execute a method for determining a location of a resonant peak for a binding partner in a resonant reflectance spectrum with a colormetric resonant biosensor. A computer readable medium can include, for example, magnetic disks, optical disks, organic memory, and any other volatile (e.g., Random Access Memory (“RAM”)) or non-volatile (e.g., Read-Only Memory (“ROM”)) mass storage system readable by the processor. The computer readable medium includes cooperating or interconnected computer readable medium, which exist exclusively on a processing system or to be distributed among multiple interconnected processing systems that can be local or remote to the processing system.
The following are provided for exemplification purpose only and are not intended to limit the scope of the invention described in broad terms above. All references cited in this disclosure are incorporated herein by reference.
EXAMPLE 1
Fabrication of a SWS Biosensor
An example of biosensor fabrication begins with a flat glass substrate that is coated with a thin layer (180 nm) of silicon nitride by plasma-enhanced chemical vapor deposition (PECVD).
The desired structure is first produced in photoresist by coherently exposing a thin photoresist film to three laser beams, as described in previously (Cowen, “The recording and large scale replication of crossed holographic grating arrays using multiple beam interferometry,” in International Conference on the Application, Theory, and Fabrication of Periodic Structures, Diffraction Gratings, and Moire Phenomena II , J. M. Lerner, ed., Proc. Soc. Photo - Opt. Instrum. Eng., 503, 120-129, 1984; Cowen, “Holographic honeycomb microlens,” Opt. Eng. 24, 796-802 (1985); Cowen & Slafer, “The recording and replication of holographic micropatterns for the ordering of photographic emulsion grains in film systems,” J. Imaging Sci. 31, 100-107, 1987. The nonlinear etching characteristics of photoresist are used to develop the exposed film to create a pattern of holes within a hexagonal grid, as shown in FIG. 22 . The photoresist pattern is transferred into the silicon nitride layer using reactive ion etching (RIE). The photoresist is removed, and a cover layer of spin-on-glass (SOG) is applied (Honeywell Electronic Materials, Sunnyvale, Calif.) to fill in the open regions of the silicon nitride grating. The structure of the top surface of the finished biosensor is shown in FIG. 23 . A photograph of finished parts are shown in FIG. 24 .
EXAMPLE 2
A SRVD biosensor was prepared by making five circular diffuse grating holograms by stamping a metal master plate into vinyl. The circular holograms were cut out and glued to glass slides. The slides were coated with 1000 angstroms of aluminum. In air, the resonant wavelength of the grating is ˜380 nm, and therefore, no reflected color is visible. When the grating is covered with water, a light blue reflection is observed. Reflected wavelength shifts are observable and measurable while the grating is covered with a liquid, or if a specific binding substances and/or binding partners cover the structure.
Both proteins and bacteria were immobilized onto the surface of a SRVD biosensor at high concentration and the wavelength shift was measured. For each material, a 20 μl droplet is placed onto a biosensor distinct location and allowed to dry in air. At 1 μg/ml protein concentration, a 20 μl droplet spreads out to cover a 1 cm diameter circle and deposits about 2×10 −8 grams of material. The surface density is 25.6 ng/mm 2 .
For high concentration protein immobilization (biosensor 4) a 10 μl droplet of 0.8 g bovine serum albumin (BSA) in 40 ml DI H 2 O is spread out to cover a 1 cm diameter circle on the surface of a biosensor. The droplet deposits 0.0002 g of BSA, for a density of 2.5e-6 g/mm 2 . After protein deposition, biosensor 4 has a green resonance in air.
For bacteria immobilization (biosensor 2) a 20 μl droplet of NECK borrelia Lyme Disease bacteria (1.8e8 cfu/ml) was deposited on the surface of a biosensor. After bacteria deposition, the biosensor looks grey in air.
For low concentration protein immobilization (biosensor 6) a 10 μl droplet of 0.02% of BSA in DI H 2 O (0.8 g BSA in 40 ml DI H 2 O) is spread out to cover a 1 cm diameter circle. The droplet deposits 0.000002 g of BSA for a density of 2.5e-8 g/mm 2 . After protein deposition, biosensor 6 looks grey in air.
In order to obtain quantitative data on the extent of surface modification resulting from the above treatments, the biosensors were measured using a spectrometer.
Because a green resonance signal was immediately visually observed on the biosensor upon which high concentration BSA was deposited (biosensor 4), it was measured in air. FIG. 25 shows two peaks at 540 nm and 550 nm in green wavelengths where none were present before protein deposition, indicating that the presence of a protein thin film is sufficient to result in a strong shift in resonant wavelength of a surface relief structure.
Because no visible resonant wavelength was observed in air for the slide upon which a low concentration of protein was applied (biosensor 6), it was measured with distilled water on surface and compared against a biosensor which had no protein treatment. FIG. 26 shows that the resonant wavelength for the slide with protein applied shifted to green compared to a water-coated slide that had not been treated.
Finally, a water droplet containing Lyme Disease bacteria Borrelia burgdorferi was applied to a grating structure and allowed to dry in air (biosensor 2). Because no visually observed resonance occurred in air after bacteria deposition, the biosensor was measured with distilled water on the surface and compared to a water-coated biosensor that had undergone no other treatment. As shown in FIG. 27 , the application of bacteria results in a resonant frequency shift to longer wavelengths.
EXAMPLE 3
Computer Model of Biosensor
To demonstrate the concept that a resonant grating structure can be used as a biosensor by measuring the reflected wavelength shift that is induced when biological material is adsorbed onto its surface, the structure shown in FIG. 1 was modeled by computer. For purposes of demonstration, the substrate chosen was glass (n substrate =1.50). The grating is a two-dimensional pattern of silicon nitride squares (t 2 =180 nm, n 2 =2.01, k 2 =0.001) with a period of 510 nm, and a filling factor of 56.2% (i.e. 56.2% of the surface is covered with silicon nitride squares while the rest is the area between the squares). The areas between silicon nitride squares are filled with a lower refractive index material. The same material also covers the squares and provides a uniformly flat upper surface. For this simulation, a glass layer was selected (n 1 =1.40) that covers the silicon nitride squares by t 2 =100 nm. To observe the effect on the reflected wavelength of this structure with the deposition of biological material, variable thicknesses of protein (n bio =1.5) were added above the glass coating layer.
The reflected intensity as a function of wavelength was modeled using GSOLVER software, which utilizes full 3-dimensional vector code using hybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVER calculates diffracted fields and diffraction efficiencies from plane wave illumination of arbitrarily complex grating structures. The illumination may be from any incidence and any polarization.
The results of the computer simulation are shown in FIG. 28 and FIG. 29 . As shown in FIG. 28 , the resonant structure allows only a single wavelength, near 780 nm, to be reflected from the surface when no protein is present on the surface. Because the peak width at half-maximum is ˜1.5 nm, resonant wavelength shifts of 0.2 nm will be easily resolved. FIG. 28 also shows that the resonant wavelength shifts to longer wavelengths as more protein is deposited on the surface of the structure. Protein thickness changes of 2 nm are easily observed. FIG. 29 plots the dependence of resonant wavelength on the protein coating thickness. A near linear relationship between protein thickness and resonant wavelength is observed, indicating that this method of measuring protein adsorption can provide quantitative data. For the simulated structure, FIG. 29 shows that the wavelength shift response becomes saturated when the total deposited protein layer exceeds 250 nm. This upper limit for detection of deposited material provides adequate dynamic range for any type of biomolecular assay.
EXAMPLE 4
Computer Model of Biosensor
In another embodiment of the invention a biosensor structure shown in FIG. 30 was modeled by computer. For purposes of demonstration, the substrate chosen was glass n substrate =1.454 coated with a layer of high refractive index material such as silicon nitride, zinc sulfide, tantalum oxide, or titanium dioxide. In this case, silicon nitride (t 3 =90 nm, n 3 =2.02) was used. The grating is two-dimensional pattern of photoresist squares (t 2 =90 nm, n 2 =1.625) with a period of 510 nm, and a filling factor of 56.2% (i.e. 56.2% of the surface is covered with photoresist squares while the rest is the area between the squares). The areas between photoresist squares are filled with a lower refractive index material such as glass, plastic, or epoxy. The same material also covers the squares and provides a uniformly flat upper surface. For this simulation, a glass layer was selected (n 1 =1.45) that covers the photoresist squares by t 2 =100 nm. To observe the effect on the reflected wavelength of this structure with the deposition of a specific binding substance, variable thicknesses of protein (n bio =1.5) were added above the glass coating layer.
The reflected intensity as a function of wavelength was modeled using GSOLVER software, which utilizes full 3-dimensional vector code using hybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVER calculates diffracted fields and diffraction efficiencies from plane wave illumination of arbitrarily complex grating structures. The illumination may be from any incidence and any polarization.
The results of the computer simulation are shown in FIG. 31 and FIG. 32 . The resonant structure allows only a single wavelength, near 805 nm, to be reflected from the surface when no protein is present on the surface. Because the peak width at half-maximum is <0.25 nm, resonant wavelength shifts of 1.0 nm will be easily resolved. FIG. 31 also shows that the resonant wavelength shifts to longer wavelengths as more protein is deposited on the surface of the structure. Protein thickness changes of 1 nm are easily observed. FIG. 32 plots the dependence of resonant wavelength on the protein coating thickness. A near linear relationship between protein thickness and resonant wavelength is observed, indicating that this method of measuring protein adsorption can provide quantitative data.
EXAMPLE 5
Sensor Readout Instrumentation
In order to detect reflected resonance, a white light source can illuminate a 1 mm diameter region of a biosensor surface through a 400 micrometer diameter fiber optic and a collimating lens, as shown in FIG. 33 . Smaller or larger areas may be sampled through the use of illumination apertures and different lenses. A group of six detection fibers are bundled around the illumination fiber for gathering reflected light for analysis with a spectrometer (Ocean Optics, Dunedin, Fla.). For example, a spectrometer can be centered at a wavelength of 800 nm, with a resolution of ˜0.14 nm between sampling bins. The spectrometer integrates reflected signal for 25-75 msec for each measurement. The biosensor sits upon an x-y motion stage so that different regions of the biosensor surface can be addressed in sequence.
Equivalent measurements can be made by either illuminating the top surface of device, or by illuminating through the bottom surface of the transparent substrate. Illumination through the back is preferred when the biosensor surface is immersed in liquid, and is most compatible with measurement of the biosensor when it is incorporated into the bottom surface of, for example, a microwell plate.
EXAMPLE 6
Demonstration of Resonant Reflection
FIG. 34 shows the resonant reflectance spectra taken from a biosensor as shown in FIG. 1 using the instrumentation described in Example 5. The wavelength of the resonance (λ peak =772.5 nm) compares with the resonant wavelength predicted by the computer model (λ peak =781 nm), and the measured reflectance efficiency (51%) is comparable to the predicted efficiency (70%). The greatest discrepancy between the measured and predicted characteristics is the linewidth of the resonant peak. The measured full-width at half maximum (FWHM) of the resonance is 6 nm, while the predicted FWHM is 1.5 nm. As will be shown, the dominant source of the larger measured FWHM is collimation of the illumination optics, which can easily be corrected.
As a basic demonstration of the resonant structure's ability to detect differences in the refractive index of materials in contact with its surface, a biosensor was exposed to a series of liquids with well-characterized optical properties. The liquids used were water, methanol, isopropyl alcohol, acetone, and DMF. A biosensor was placed face-down in a small droplet of each liquid, and the resonant wavelength was measured with a fiber illumination/detection probe facing the biosensor's back side. Table 1 shows the calculated and measured peak resonant wavelength as a biosensor surface is exposed to liquids with variable refractive index demonstrating the correlation between measured and theoretical detection sensitivity. As shown in Table 1, the measured resonant peak positions and measured resonant wavelength shifts are nearly identical to the predicted values. This example demonstrates the underlying sensitivity of the biosensor, and validates the computer model that predicts the wavelength shift due to changes in the material in contact with the surface.








TABLE 1





Calculated
Measured










Peak

Peak




Wavelength

Wavelength

Solution
n
(nm)
Shift (nm)
(nm)
Shift (nm)










Water
1.333
791.6
0
786.08
0

Isopropyl
1.3776
795.9
4.3
789.35
3.27

Acetone
1.3588
794
2.4
788.22
2.14

Methanol
1.3288
791.2
−0.4
785.23
−0.85

DMF
1.4305
802
10.4
796.41
10.33






























Similarly, a biosensor is able to measure the refractive index difference between various buffer solutions. As an example, FIG. 35 shows the variation in peak wavelength with the concentration of bovine serum albumin (BSA) in water. Resonance was measured with the biosensor placed face-down in a droplet of buffer, and rinsed with water between each measurement.
EXAMPLE 7
Immobilized Protein Detection
While the detection experiments shown in Example 6 demonstrate a biosensor's ability to measure small differences in refractive index of liquid solutions, the biosensor is intended to measure specific binding substances and binding partners that are chemically bound to the biosensor surface. In order to demonstrate a biosensor's ability to quantify biomolecules on its surface, droplets of BSA dissolved in PBS at various concentrations were applied to a biosensor as shown in FIG. 1 . The 3 μl droplets were allowed to dry in air, leaving a small quantity of BSA distributed over a ˜2 mm diameter area. The peak resonant wavelength of each biosensor location was measured before and after droplet deposition, and the peak wavelength shift was recorded. See FIG. 37 .
EXAMPLE 8
Immobilization of One or More Specific Binding Substances
The following protocol was used on a calorimetric resonant reflective biosensor to activate the surface with amine functional groups. Amine groups can be used as a general-purpose surface for subsequent covalent binding of several types of linker molecules.
A biosensor of the invention is cleaned by immersing it into piranha etch (70/30% (v/v) concentrated sulfuric acid/30% hydrogen peroxide) for 12 hours. The biosensor was washed thoroughly with water. The biosensor was dipped in 3% 3-aminopropyltriethoxysilane solution in dry acetone for 1 minute and then rinsed with dry acetone and air-dried. The biosensor was then washed with water.
A semi-quantitative method is used to verify the presence of amino groups on the biosensor surface. One biosensor from each batch of amino-functionalized biosensors is washed briefly with 5 mL of 50 mM sodium bicarbonate, pH 8.5. The biosensor is then dipped in 5 mL of 50 mM sodium bicarbonate, pH 8.5 containing 0.1 mM sulfo-succinimidyl-4-O-(4,4′-dimethoxytrityl)-butyrate (s-SDTB, Pierce, Rockford, Ill.) and shaken vigorously for 30 minutes. The s-SDTB solution is prepared by dissolving 3.0 mg of s-SDTB in 1 mL of DMF and diluting to 50 mL with 50 mM sodium bicarbonate, pH 8.5. After a 30 minute incubation, the biosensor is washed three times with 20 mL of ddH 2 O and subsequently treated with 5 mL 30% perchloric acid. The development of an orange-colored solution indicates that the biosensor has been successfully derivatized with amines; no color change is observed for untreated glass biosensors.
The absorbance at 495 nm of the solution after perchloric acid treatment following the above procedure can be used as an indicator of the quantity of amine groups on the surface. In one set of experiment, the absorbance was 0.627, 0.647, and 0.728 for Sigma slides, Cel-Associate slides, and in-house biosensor slides, respectively. This indicates that the level of NH 2 activation of the biosensor surface is comparable in the activation commercially available microarray glass slides.
After following the above protocol for activating the biosensor with amine, a linker molecule can be attached to the biosensor. When selecting a cross-linking reagent, issues such as selectivity of the reactive groups, spacer arm length, solubility, and cleavability should be considered. The linker molecule, in turn, binds the specific binding substance that is used for specific recognition of a binding partner. As an example, the protocol below has been used to bind a biotin linker molecule to the amine-activated biosensor.
Protocol for Activating Amine-Coated Biosensor with Biotin
Wash an amine-coated biosensor with PBS (pH 8.0) three times. Prepare sulfo-succinimidyl-6-(biotinamido)hexanoate (sulfo-NHS-LC-biotin, Pierce, Rockford, Ill.) solution in PBS buffer (pH 8) at 0.5 mg/ml concentration. Add 2 ml of the sulfo-NHS-LC-biotin solution to each amine-coated biosensor and incubate at room temperature for 30 min. Wash the biosensor three times with PBS (pH 8.0). The sulfo-NHS-LC-biotin linker has a molecular weight of 556.58 and a length of 22.4 Å. The resulting biosensors can be used for capturing avidin or strepavidin molecules.
Protocol for Activating Amine-Coated Biosensor with Aldehyde
Prepare 2.5% glutaraldehyde solution in 0.1 M sodium phosphate, 0.05% sodium azide, 0.1% sodium cyanoborohydride, pH 7.0. Add 2 ml of the glutaraldehyde solution to each amine-coated biosensor and incubate at room temperature for 30 min. Wash the biosensor three times with PBS (pH 7.0). The glutaraldehyde linker has a molecular weight of 100.11. The resulting biosensors can be used for binding proteins and other amine-containing molecules. The reaction proceeds through the formation of Schiff bases, and subsequent reductive amination yields stable secondary amine linkages. In one experiment, where a coated aldehyde slide made by the inventors was compared to a commercially available aldehyde slide (Cel-Associate), ten times higher binding of streptavidin and anti-rabbit IgG on the slide made by the inventors was observed.
Protocol for Activating Amine-Coated Biosensor with NHS
25 mM N,N′-disuccinimidyl carbonate (DSC, Sigma Chemical Company, St. Louis, Mo.) in sodium carbonate buffer (pH 8.5) was prepared. 2 ml of the DSC solution was added to each amine-coated biosensor and incubated at room temperature for 2 hours. The biosensors were washed three times with PBS (pH 8.5). A DSC linker has a molecular weight of 256.17. Resulting biosensors are used for binding to hydroxyl- or amine-containing molecules. This linker is one of the smallest homobifunctional NHS ester cross-linking reagents available.
In addition to the protocols defined above, many additional surface activation and molecular linker techniques have been reported that optimize assay performance for different types of biomolecules. Most common of these are amine surfaces, aldehyde surfaces, and nickel surfaces. The activated surfaces, in turn, can be used to attach several different types of chemical linkers to the biosensor surface, as shown in Table 2. While the amine surface is used to attach several types of linker molecules, the aldehyde surface is used to bind proteins directly, without an additional linker. A nickel surface is used exclusively to bind molecules that have an incorporated histidine (“his”) tag. Detection of “his-tagged” molecules with a Nickel activated surface is well known (Sigal et al., Anal Chem. 68, 490 (1996)).
Table 2 demonstrates an example of the sequence of steps that are used to prepare and use a biosensor, and various options that are available for surface activation chemistry, chemical linker molecules, specific binding substances and binding partners molecules. Opportunities also exist for enhancing detected signal through amplification with larger molecules such as HRP or streptavidin and the use of polymer materials such as dextran or TSPS to increase surface area available for molecular binding.










TABLE 2








Label

Bare
Surface
Linker
Receptor
Detected
Molecule

Sensor
Activation
Molecule
Molecule
Material
(Optional)



Glass
Amino
SMPT
Sm m'cules
Peptide
Enhance

Polymers
Aldehyde
NHS-Biotin
Peptide
Med Protein
sensitivity

optional to
Ni
DMP
Med Protein
Lrg Protein
1000x

enhance

NNDC
Lrg Protein
• IgG
HRP

sensitivity

His-tag
• IgG
Phage
Streptavidin

2-5x

Others . . .
cDNA
Cell

Dextran TSPS



cDNA

























EXAMPLE 9
IgG Assay
As an initial demonstration for detection of biochemical binding, an assay was performed in which a biosensor was prepared by activation with the amino surface chemistry described in Example 8 followed by attachment of a biotin linker molecule. The biotin linker is used to covalently bond a streptavidin receptor molecule to the surface by exposure to a 50 μg/ml concentration solution of streptavidin in PBS at room temperature for 2-4 hours. The streptavidin receptor is capable of binding any biotinylated protein to the biosensor surface. For this example, 3 μl droplets of biotinylated anti-human IgG in phosphate buffer solution (PBS) were deposited onto 4 separate locations on the biosensor surface at a concentration of 200 μg/ml. The solution was allowed to incubate on the biosensor for 60 min before rinsing thoroughly with PBS. The peak resonant wavelength of the 4 locations were measured after biotin activation, after streptavidin receptor application, and after ah-IgG binding. FIG. 37 shows that the addition of streptavidin and ah-IgG both yield a clearly measurable increase in the resonant wavelength.
EXAMPLE 10
Biotin/Streptavidin Assay
A series of assays were performed to detect streptavidin binding by a biotin receptor layer. A biosensor was first activated with amino chemistry, followed by attachment of a NHS-Biotin linker layer, as previously described. Next, 3 μl droplets of streptavidin in PBS were applied to the biosensor at various concentrations. The droplets were allowed to incubate on the biosensor surface for 30 min before thoroughly washing with PBS rinsing with DI water. The peak resonant wavelength was measured before and after streptavidin binding, and the resonant wavelength shifts are shown in FIG. 38 . A linear relationship between peak wavelength and streptavidin concentration was observed, and in this case the lowest streptavidin concentration measured was 0.2 μg/ml. This concentration corresponds to a molarity of 3.3 nM.
EXAMPLE 11
Protein-Protein Binding Assay
An assay was performed to demonstrate detection of protein-protein interactions. As described previously, a biosensor was activated with amino chemistry and an NHS-biotin linker layer. A goat anti-biotin antibody receptor layer was attached to the biotin linker by exposing the biosensor to a 50 μg/ml concentration solution in PBS for 60 min at room temperature followed by washing in PBS and rinsing with DI water. In order to prevent interaction of nonspecific proteins with unbound biotin on the biosensor surface, the biosensor surface was exposed to a 1% solution of bovine serum albumin (BSA) in PBS for 30 min. The intent of this step is to “block” unwanted proteins from interacting with the biosensor. As shown in FIG. 39 a significant amount of BSA is incorporated into the receptor layer, as shown by the increase in peak wavelength that is induced. Following blocking, 3 μl droplets of various concentrations of anti-goat IgG were applied to separate locations on the biosensor surface. The droplets were allowed to incubate for 30 min before thorough rinsing with DI water. The biosensor peak resonant wavelength was measured before blocking, after blocking, after receptor layer binding, and after anti-goat IgG detection for each spot. FIG. 39 shows that an anti-goat IgG concentration of 10 μg/ml yields an easily measurable wavelength shift.
EXAMPLE 12
Unlabeled ELISA Assay
Another application of a biosensor array platform is its ability to perform Enzyme-Linked Immunosorbent Assays (ELISA) without the need for an enzyme label, and subsequent interaction an enzyme-specific substrate to generate a colored dye. FIG. 40 shows the results of an experiment where a biosensor was prepared to detect interferon-γ (IFN-γ) with an IFN-γ antibody receptor molecule. The receptor molecule was covalently attached to an NH 2 -activated biosensor surface with an SMPT linker molecule (Pierce Chemical Company, Rockford, Ill.). The peak resonant wavelength shift for application of the NH 2 , SMPT, and anti-human IFN-α receptor molecules were measured for two adjacent locations on the biosensor surface, as shown in FIG. 40 . The two locations were exposed to two different protein solutions in PBS at a concentration of 100 μg/ml. The first location was exposed to IFN-γ, which is expected to bind with the receptor molecule, while the second was exposed to neural growth factor (NGF), which is not expected to bind with the receptor. Following a 30 minute incubation the biosensor was measured by illuminating from the bottom, while the top surface remained immersed in liquid. The location exposed to IFN-γ registered a wavelength shift of 0.29 nm, while the location exposed to NGF registered a wavelength shift of only 0.14 nm. Therefore, without the use of any type of enzyme label or color-generating enzyme reaction, the biosensor was able to discriminate between solutions containing different types of protein.
EXAMPLE 13
Protease Inhibitor Assay (Caspase-3)
A Caspase-3 protease inhibitor assay was performed to demonstrate the biosensor's ability to measure the presence and cleavage of small molecules in an experimental context that is relevant to pharmaceutical compound screening.
Caspases (Cysteine-requiring Aspartate protease) are a family of proteases that mediate cell death and are important in the process of apoptosis. Caspase 3, an effector caspase, is the most studied of mammalian caspases because it can specifically cleave most known caspase-related substrates. The caspase 3 assay is based on the hydrolysis of the 4-amino acid peptide substrate NHS-Gly-Asp-Glu-Val-Asp p-nitroanilide (NHS-GDEVD-pNA) by caspase 3, resulting in the release of the pNA moiety.









The NHS molecule attached to the N-terminal of the GDEVD provides a reactive end group to enable the NHS-GDEVD-pNA complex to be covalently bound to the biosensor with the pNA portion of the complex oriented away from the surface. Attached in this way, the caspase-3 will have the best access to its substrate cleavage site.
A biosensor was prepared by cleaning in 3:1 H 2 SO 4 :H 2 O 2 solution (room temperature, 1 hour), followed by silanation (2% silane in dry acetone, 30 sec) and attachment of a poly-phe-lysine (PPL) layer (100 μg/ml PPL in PBS pH 6.0 with 0.5 M NaCl, 10 hours). The NHS-GDEVD-pNA complex was attached by exposing the biosensor to a 10 mM solution in PBS (pH 8.0, room temperature, 1 hour). A microwell chamber was sealed over the biosensor surface, and cleavage of pNA was performed by addition of 100 μl of caspase-3 in 1× enzyme buffer (100 ng/ml, room temperature, 90 minutes). Following exposure to the caspase 3 solution, the biosensor was washed in PBS. A separate set of experiments using a spectrophotometer were used to confirm the attachment of the complex to the surface of the biosensor, and functional activity of the caspase-3 for removal of the pNA molecule from the surface-bound complex.
The peak resonant frequency of the biosensor was measured before attachment of the NHS-GDEVD-pNA complex, after attachment of the complex (MW=860 Da), and after cleavage of the pNA (MW=136) with caspase 3. As shown in FIG. 41 , the attachment of the peptide molecule is clearly measurable, as is the subsequent removal of the pNA. The pNA removal signal of Δλ=0.016 nm is 5.3× higher than the minimum detectable peak wavelength shift of 0.003 nm. The proportion of the added molecular weight and subtracted molecular weight (860 Da/136 Da=6.32) are in close agreement with the proportion of peak wavelength shift observed for the added and subtracted material (0.082 nm/0.016 nm=5.14).
The results of this experiment confirm that a biosensor is capable of measuring small peptides (in this case, a 5-mer peptide) without labels, and even detecting the removal of 130 Da portions of a molecule through the activity of an enzyme.
EXAMPLE 14
Reaction Kinetics for Protein-Protein Binding Assays
Because a biosensor of the invention can be queried continuously as a function of time while it is immersed in liquid, a biosensor can be utilized to perform both endpoint-detection experiments and to obtain kinetic information about biochemical reactions. As an example, FIG. 42 shows the results of an experiment in which a single biosensor location is measured continuously through the course of consecutively adding various binding partners to the surface. Throughout the experiment, a detection probe illuminated the biosensor through the back of the biosensor substrate, while biochemistry is performed on the top surface of the device. A rubber gasket was sealed around the measured biosensor location so that added reagents would be confined, and all measurements were performed while the top surface of the biosensor was immersed in buffer solution. After initial cleaning, the biosensor was activated with NH 2 , and an NHS-Biotin linker molecule. As shown in FIG. 42 , goat α-biotin antibodies of several different concentrations (1, 10, 100, 1000 μg/ml) were consecutively added to the biosensor and allowed to incubate for 30 minutes while the peak resonant wavelength was monitored. Following application of the highest concentration α-Biotin IgG, a second layer of protein was bound to the biosensor surface through the addition of α-goat IgG at several concentrations (0.1, 1, 10, and 100 μg/ml). Again, the resonant peak was continuously monitored as each solution was allowed to incubate on the biosensor for 30 minutes. FIG. 42 shows how the resonant peak shifted to greater wavelength at the end of each incubation period.
FIG. 43 shows the kinetic binding curve for the final resonant peak transitions from FIG. 42 , in which 100 μg/ml of α-goat IgG is added to the biosensor. The curve displays the type of profile that is typically observed for kinetic binding experiments, in which a rapid increase from the base frequency is initially observed, followed by a gradual saturation of the response. This type of reaction profile was observed for all the transitions measured in the experiment. FIG. 44 shows the kinetic binding measurement of IgG binding.
The removal of material from the biosensor surface through the activity of an enzyme is also easily observed. When the biosensor from the above experiment (with two protein coatings of goat anti-biotin IgG and anti-goat IgG) is exposed to the protease pepsin at a concentration of 1 mg/ml, the enzyme dissociates both IgG molecules, and removes them from the biosensor surface. As shown in FIG. 45 , the removal of bound molecules from the surface can be observed as a function of time.
EXAMPLE 15
Proteomics Applications
Biosensors of the invention can be used for proteomics applications. A biosensor array can be exposed to a test sample that contains a mixture of binding partners comprising, for example, proteins or a phage display library, and then the biosensor surface is rinsed to remove all unbound material. The biosensor is optically probed to determine which distinct locations on the biosensor surface have experienced the greatest degree of binding, and to provide a quantitative measure of bound material. Next, the biosensor is placed in a “flow cell” that allows a small (e.g., <50 microliters) fixed volume of fluid to make contact to the biosensor surface. One electrode is activated so as to elute bound material from only a selected biosensor array distinct location. The bound material becomes diluted within the flow cell liquid. The flow cell liquid is pumped away from the biosensor surface and is stored within a microtiter plate or some other container. The flow cell liquid is replaced with fresh solution, and a new biosensor electrode is activated to elute its bound binding partners. The process is repeated until all biosensor distinct locations of interest have been eluted and gathered into separate containers. If the test sample liquid contained a mixture of proteins, protein contents within the separate containers can be analyzed using a technique such as electrospray tandem mass spectrometry. If the sample liquid contained a phage display library, the phage clones within the separate containers can be identified through incubation with a host strain bacteria, concentration amplification, and analysis of the relevant library DNA sequence.
EXAMPLE 16
Mathematical Resonant Peak Determination
This example discusses some of the findings that have been obtained from looking at fitting different types of curves to the observed data.
The first analytic curve examined is a second-order polynomial, given by



y=ax
2
+bx+c


The least-squares solution to this equation is given by the cost function





φ
=


∑

i
=
1

n




(


ax
i
2

+

bx
i

+
c
-

y
i


)

2



,




the minimization of which is imposed by the constraints






∂
φ


∂
a


=



∂
φ


∂
b


=



∂
φ


∂
c


=
0.






Solving these constraints for a, b, and c yields





(



a




b




c



)

=



(




∑

x
i
4





∑

x
i
3





∑

x
i
2







∑

x
i
3





∑

x
i
2





∑

x
i







∑

x





2





∑

x
i




n



)


-
1


·


(




∑


x
i
2



y
i








∑


x
i



y
i








y
i




)

.






The result of one such fit is shown in FIG. 46 ; the acquired data are shown as dots and the 2 nd -order polynomial curve fit is shown as the solid line.
Empirically, the fitted curve does not appear to have sufficient rise and fall near the peak. An analytic curve that provides better characteristics in this regard is the exponential, such as a Gaussian. A simple method for performing a Gaussian-like fit is to assume that the form of the curve is given by
y=e ax 2 +bx+c ,
in which case the quadratic equations above can be utilized by forming y′, where y′=lny. FIG. 46 shows the result of such a fit. The visual appearance of FIG. 46 indicates that the exponential is a better fit, providing a 20% improvement over that of the quadratic fit.
Assuming that the exponential curve is the preferred data fitting method, the robustness of the curve fit is examined in two ways: with respect to shifts in the wavelength and with respect to errors in the signal amplitude.
To examine the sensitivity of the analytical peak location when the window from which the curve fitting is performed is altered to fall 10 sampling intervals to the left or to the right of the true maxima. The resulting shift in mathematically-determined peak location is shown in Table 3. The conclusion to be derived is that the peak location is reasonably robust with respect to the particular window chosen: for a shift of ˜1.5 nm, the corresponding peak location changed by only <0.06 nm, or 4 parts in one hundred sensitivity.
To examine the sensitivity of the peak location with respect to noise in the data, a signal free of noise must be defined, and then incremental amounts of noise is added to the signal and the impact of this noise on the peak location is examined. The ideal signal, for purposes of this experiment, is the average of 10 resonant spectra acquisitions.
Gaussian noise of varying degrees is superimposed on the ideal signal. For each such manufactured noisy signal, the peak location is estimated using the 2 nd -order exponential curve fit. This is repeated 25 times, so that the average, maximum, and minimum peak locations are tabulated. This is repeated for a wide range of noise variances—from a variance of 0 to a variance of 750. The result is shown in FIG. 47 .





TABLE 3



Comparison of peak location as a function of window location







Shift
Window
Peak Location





Δ = −10 bins
771.25-782.79 nm
778.8221 nm


Δ = 0 bins
772.70-784.23 nm
778.8887 nm


Δ = +10 bins
774.15-785.65 nm
7778.9653 nm
























The conclusion of this experiment is that the peak location estimation routine is extremely robust to noisy signals. The entire range of peak locations in FIG. 47 is only 1.5 nm, even with as much random noise variance of 750 superimposed—an amount of noise that is substantially greater that what has been observed on the biosensor thus far. The average peak location, despite the level of noise, is within 0.1 nm of the ideal location.
Based on these results, a basic algorithm for mathematically determining the peak location of a calorimetric resonant biosensor is as follows:
1. Input data x i and y i , i=1, . . . , n
2. Find maximum
a. Find k such that y k ≧y i for all i≠k

3. Check that maximum is sufficiently high
a. Compute mean y and standard deviation σ of sample b. Continue only if (y k − y )/σ>UserThreshold

4. Define curve-fit region of 2w+1 bins (w defined by the user)
a. Extract x i ,k−w≦i≦k+w b. Extract y,k−w≦i≦k+w

5. Curve fit
a. g i =ln y i b. Perform 2 nd -order polynomial fit to obtain g i ′ defined on x i ,k−w≦i≦k+w c. Polynomial fit returns coefficients a,b,c of form ax 2 +bx+c d. Exponentiate: y i ′=e g i ′

6. Output
a. Peak location p given by x p =−b/2a b. Peak value given by y p ′(x p )

In summary, a robust peak determination routine has been demonstrated; the statistical results indicate significant insensitivity to the noise in the signal, as well as to the windowing procedure that is used. These results lead to the conclusion that, with reasonable noise statistics, that the peak location can be consistently determined in a majority of cases to within a fraction of a nm, perhaps as low as 0.1 to 0.05 nm.
EXAMPLE 17
Homogenous Assay Demonstration
An SWS biosensor detects optical density of homogenous fluids that are in contact with its surface, and is able to differentiate fluids with refractive indices that differ by as little as Δn=4×10 −5 . Because a solution containing two free non-interacting proteins has a refractive index that is different from a solution containing two bound interacting proteins, an SWS biosensor can measure when a protein-protein interaction has occurred in solution without any kind of particle tag or chemical label.
Three test solutions were prepared for comparison:
1. Avidin in Phosphate Buffer Solution (PBS), (10 μg/ml)
2. Avidin (10 μg/ml)+Bovine Serum Albumin (BSA) (10 μg/ml) in PBS
3. Avidin (10 μg/ml)+Biotinylated BSA (b-BSA) (10 μg/ml) in PBS
A single SWS sensor was used for all measurements to eliminate any possibility of cross-sensor bias. A 200 μl sample of each test solution was applied to the biosensor and allowed to equilibrate for 10 minutes before measurement of the SWS biosensor peak resonant wavelength value. Between samples, the biosensor was thoroughly washed with PBS.
The peak resonant wavelength values for the test solutions are plotted in FIG. 51 . The avidin solution was taken as the baseline reference for comparison to the Avidin+BSA and Avidin+b-BSA solutions. Addition of BSA to avidin results in only a small resonant wavelength increase, as the two proteins are not expected to interact. However, because biotin and avidin bind strongly (Kd=10 −15 M), the avidin+b-BSA solution will contain larger bound protein complexes. The peak resonant wavelength value of the avidin+b-BSA solution thus provides a large shift compared to avidin+BSA.
The difference in molecular weight between BSA (MW=66 KDa) and b-BSA (MW=68 KDa) is extremely small. Therefore, the differences measured between a solution containing non-interacting proteins (avidin+BSA) and interacting proteins (avidin+b-BSA) are attributable only to differences in binding interaction between the two molecules. The bound molecular complex results in a solution with a different optical refractive index than the solution without bound complex. The optical refractive index change is measured by the SWS biosensor.
EXAMPLE 18
Sensor Design and Fabrication
A one-dimensional linear grating surface biosensor structure requires a grating with a period lower than the wavelength of the resonantly reflected light (R. Magnusson, and S. S. Wang, “New principle for optical filters,” Appl. Phys. Lett., 61, No. 9, p. 1022, August, 1992, S. Peng and G. M. Morris, “Resonant scattering from two-dimensional gratings,” J. Opt. Soc. Am. A , Vol. 13, No. 5, p. 993, May 1996). As shown in FIG. 53 , a one-dimensional linear grating surface structure was fabricated from a low refractive index material that was overcoated with a thin film of higher refractive index material. The grating structure was microreplicated within a layer of cured epoxy.
A one-dimensional linear grating surface structure was produced on the surface of a plastic substrate material as follows. First, an 8-inch diameter silicon “master” wafer was produced. The 550 nm period linear grating structure was defined in photoresist using deep-UV photolithography by stepping and repeating the exposure of a 9 mm diameter circular grating reticle over the surface of a photoresist-coated silicon wafer, as shown in FIG. 54 . The exposure step/repeat procedure produced patterns for two standard format 96-well microtiter plates with 8 rows and 12 columns each. The exposed photoresist was developed, and the grating structure was permanently transferred to the silicon wafer using a reactive ion etch with a depth of 200 nm. After etching, the photoresist was removed.
The grating structure was replicated onto a 0.005 inch thick sheet of polycarbonate by distributing a thin layer of epoxy between the silicon master wafer and a section of the polycarbonate sheet. The liquid epoxy conforms to the shape of the master grating, and was subsequently cured by exposure to ultraviolet light. The cured epoxy preferentially adheres to the polycarbonate sheet, and is peeled away from the silicon wafer.
Sensor fabrication was completed by sputter deposition of 120 nm tantalum oxide on the cured epoxy grating surface. Following tantalum oxide deposition, 3×5-inch microtiter plate sections were cut from the sensor sheet, activated with amine functional groups and attached to the bottoms of bottomless 96-well microtiter plates (Corning Costar®, Cambridge, Mass. and Greiner, Longwood, Fla.) with epoxy.
Surface Activation and Attachment of Receptor Molecule After high refractive index material deposition, biosensors are activated with amine functional groups to enable various bifunctional linker molecules to be attached to the surface in a known orientation. Amine activation was performed by immersion of the sensor in 10% 3-aminopropyltriethoxysilane (Pierce) solution in ethanol (Aldrich) for 1 min, followed by a brief ethanol rinse. Activated sensors were then dried at 70° C. for 10 min. Other surface activation molecules could include, for example, COOH, CHO, polymer, and Poly-phe-lysine.
A simple, colorimetric method using a modified protocol from Pierce was used to determine the density of amine groups on the surface. The amine-activated biosensor was immersed in 0.1 mM of sulfo-succinimidyl-4-O-(4,4′-dimethoxytrityl)-butyrate (s-SDTB, Pierce), solution made in 50 mM sodium bicarbonate (pH 8.5), and shaken vigorously for 30 minutes. The biosensor was then washed with deionized water and subsequently treated with 30% perchloric acid (Sigma). The solution turned orange when the biosensor was amine-activated or remained colorless otherwise. This method indicated that the surface density of the amine groups is 2×10 14 groups/cm 2 .
A one-dimensional linear grating resonant biosensor was used for the detection of a well-characterized protein-protein binding interaction. The protein-protein system selected for this study was detection of anti-biotin IgG antibody using biotin immobilized on the biosensor surface as a receptor molecule. Therefore, a protocol for immobilization of biotin on the biosensor surface was developed that utilizes a bifunctional polyethyleneglycol-N-hydrosuccinimide (NHS-PEG) linker molecule (Shearwater Polymers, Inc.) to act as an intermediary between the amine surface group and the biotin. The NHS-PEG molecule is designed specifically to enable NHS to preferentially bind to the amine-activated surface, leaving the PEG portion of the molecule oriented away from the surface. The NHS-PEG linker molecule serves to separate the biotin molecule from the biosensor surface by a short distance so it can retain its conformation, and thus its affinity for other molecules. The PEG also serves to prevent nonspecific binding of proteins to the biosensor.
After attachment of amine-activated biosensor sheets into the bottom of microtiter plates, individual microtiter wells were prepared with three different surface functional groups in order to provide sufficient experimental controls for the detection of anti-biotin IgG. First, amine-activated surfaces were studied without additional modification. The amine-activated surface is expected to bind proteins nonspecifically, but not with high affinity. Second, microtiter wells with the NHS-PEG bifunctional linker molecule were prepared. The NHS-PEG molecule is expected to provide a surface that does not bind protein. Third, microtiter wells with an NHS-PEG-Biotin linker molecule were prepared. The NHS-PEG-Biotin molecule is expected to bind strongly to anti-biotin IgG.
To activate an amine-coated sensor with biotin, 2 ml of NHS-PEG-Biotin (Shearwater) solution in TPBS (a reference buffer solution of 0.01% Tween™ 20 in phosphate buffer solution, pH 8) at 1.0 mg/ml concentration was added to the biosensor surface, and incubated at 37° C. for 1 hour. An identical procedure was used for attachment of the NHS-PEG (Shearwater) molecule without biotin. All purchased reagents were used as packaged.
96-Well Plate Scanner Instrument
A schematic diagram of the system used to illuminate the biosensor and to detect the reflected signal is shown in FIG. 55 . In order to detect the reflected resonance, a white light source illuminates a ˜1 mm diameter region of the grating surface through a 100 micrometer diameter fiber optic and a collimating lens at nominally normal incidence through the bottom of the microtiter plate. After passing through the collimating lens, incident light passes through a linear polarizing filter so that the linear grating is excited only with light that is polarized either parallel or perpendicular to the grating lines. Reflected light passes through the polarizing filter again on its way back to the detection probe. A detection fiber is bundled with the illumination fiber for gathering reflected light for analysis with a spectrometer (Ocean Optics). A series of 8 illumination/detection heads are arranged in a linear fashion, so that reflection spectra are gathered from all 8 wells in a microtiter plate column at once. The microtiter plate sits upon a motion stage so that each column can be addressed in sequence.
Reflected Resonance Signal and Response Uniformity
With water in the microtiter plate well, the p-polarized reflected resonance spectrum is shown in FIG. 56 . The measured peak wavelength value (PWV) is 857 nm, and the full-width at half-maximum (FWHM) of the resonant peak is 1.8 nm. Note that compared to structures produced using, for example, two-dimensional grid holes on a hexagonal grid, only a single resonant peak is measured, with a linewidth that is >3× lower. Most importantly, the single narrow resonant peak characteristic is uniformly obtained within every well of a 96-well microtiter plate biosensor.
The ability of a biosensor to measure shifts in optical density on its surface can be calibrated by measuring the biosensor PWV when two solutions with known refractive index values are added to the microtiter plate wells, and by calculating the PWV shift (ΔPWV) between the two solutions. To measure biosensor response uniformity across a plate, the PWV of all 96 wells were measured in water (n=1.333) and subsequently in glycerol (n=1.472). The shift coefficient, σ=ΔPWV/Δn, is defined to be a figure of merit for comparing the response of biosensors with different designs or fabrication processes. The average shift in PWV from water to glycerol across all wells was 15.57 nm, providing a shift coefficient of σ=112 nm. The standard deviation of σ across 96 sensor wells was 1.07 nm, indicating a very high degree of biosensor uniformity across a large surface area.
Protein-Protein Binding Assay
A protein-antibody affinity assay was performed to demonstrate operation of the plastic biosensor. A matrix of three separate sensor surface states (NH 2 , NHS-PEG, NHS-PEG-Biotin) were prepared and exposed to 7 concentrations of goat anti-biotin IgG (Sigma). Each matrix location was measured within a separate microtiter plate well, for a total of 21 wells measured simultaneously. Because the NHS-PEG wells are not expected to bind protein, they provide a reference for canceling common mode effects such as the effect of the refractive index of the test sample and environmental temperature variation during the course of an assay. Data are reported here without the use of any mathematical correction.
FIG. 57 plots the PWV shift—referenced to a sensor with no chemical functional groups immobilized, recorded due to attachment of NH 2 , NH 2 +(NHS-PEG), and NH 2 +(NHS-PEG-Biotin) molecules to the biosensor surface. The error bars indicate the standard deviation of the recorded PWV shift over 7 microtiter plate wells. The data indicates that the biosensor can differentiate between a clean surface, and one with immobilized NH 2 , as well as clearly detecting the addition of the NHS-PEG (MW=2000 Da) molecule. The difference between surface immobilized NHS-PEG and NHS-PEG-Biotin (MW=3400 Da) is also measurable.
FIG. 58 shows the PWV shift response as a function of time for the biosensor wells when exposed to various concentrations of anti-biotin IgG (0-80 μg/ml) and allowed to incubate for 20 minutes. The NHS-PEG surface ( FIG. 58B ) provides the lowest response, while the amine-activated surface ( FIG. 58A ) demonstrates a low level of nonspecific interaction with the anti-biotin IgG at high concentrations. The NHS-PEG-Biotin surface ( FIG. 58C ) clearly demonstrates strong specific interaction with the anti-biotin IgG—providing strong PWV shifts in proportion to the concentration of exposed anti-biotin IgG.
The PWV shift magnitudes after 20 minutes from FIG. 58C are plotted as a function of anti-biotin IgG concentration in FIG. 59 . A roughly linear correlation between the IgG concentration and the measured PWV shift is observed, and the lowest concentration IgG solution (1.25 μg/ml, 8.33 nM) is clearly measurable over the negative control PSB solution.
EXAMPLE 19
Biosensor Comprising a Surface Modification Layer
A surface modification layer can be added to a one-dimensional grating or two-dimensional grating biosensor of the invention. A surface modification layer is added to the top surface of the high refractive index material or cover layer of a biosensor and is useful for immobilization of specific binding substances to the surface of a biosensor. A surface modification layer can comprise silicon oxide, silicon oxynitride, borosilicate glass, phosophosilicate glass, pyrex, any other glass (including BK7, SF11, LaSF9, Ultran, FK3, FK5) or any other metal oxide. The thickness of the surface modification layer can be about 5 nm to about 15 nm. In one embodiment of the invention the high refractive index material is tantalum oxide.
Silicon oxide was coated onto a biosensor of the invention by DC sputtering. Other possible methods of coating include evaporation, laser ablation, chemical vapor deposition, and plasma-enhanced chemical vapor deposition. NH 2 was added to a biosensor with the surface modification layer and to a biosensor without a surface modification layer. A NH 2 reacting fluorescent dye was used to visualize NH 2 attachment to the biosensors. See FIG. 60 . The higher fluorescence intensity from the surface modified layer indicates much higher NH 2 density.
EXAMPLE 20
Biosensor Design with High Stability in Aqueous Solutions
A biosensor of the invention can be exposed to aqueous solutions during use. Stability in aqueous solutions can be added to a biosensor by adding a interfacial layer under the high refractive index material layer. For example, where a plastic grating surface, e.g., a low refractive index grating material, is coated with a high refractive index material, an interfacial layer can be added between the high refractive index material and the low refractive index material.
For example, a biosensor was constructed by adding an adhesion enhancing interfacial layer between a plastic grating surface and a high refractive index optical material. Adhesion refers to the ability of a thin film material to remain firmly attached to the material it is deposited upon over a wide range of environmental conditions. For example, for a biosensor structure used in the invention, a high refractive index thin film (such as silicon nitride, zinc sulfide, tantalum oxide, or titanium oxide) is deposited upon a grating surface structure formed from a cured epoxy material. Without an adhesion enhancing layer (called a “tie layer” or “interfacial layer”), the high refractive index material can possibly delaminate from the grating surface structure under stringent experimental conditions, such as a long-period exposure to a liquid. The tie layer material is selected to have strong adhesion properties to both the underlying material and the high refractive index deposited material. Generally, the thickness of the tie layer is selected to be extremely thin, so as not to disrupt the optical properties of the structure that it is embedded within. Tie layer thickness can range from about 1 nm to about 200 nm, for example. In this example, the tie layer was a silicon oxide layer of approximately 5 nm in thickness, the plastic material was Polyethylene (PET), and the high refractive index optical material was tantalum oxide. Alternative materials for the tie layer are silicon oxynitride, borosilicate glass, phosophosilicate glass, pyrex, any other glass (including BK7, SF11, LaSF9, Ultran, FK3, FK5) or any other metal oxide. Biosensor stability performance was improved in aqueous solutions by the addition of silicon oxide interfacial layer. See FIG. 61 . The improved biosensor stability performance allowed for the enhancement of signal to noise ratio, which translates into more sensitive detection limits.