U.S. patent application number 13/042607 was filed with the patent office on 2011-09-29 for guided mode resonant filter biosensor using a linear grating surface structure.
This patent application is currently assigned to SRU Biosystems, Inc.. Invention is credited to Brian T. Cunningham, Peter Li, Homer Pien, Jean Qiu.
Application Number | 20110237464 13/042607 |
Document ID | / |
Family ID | 46205670 |
Filed Date | 2011-09-29 |
United States Patent
Application |
20110237464 |
Kind Code |
A1 |
Cunningham; Brian T. ; et
al. |
September 29, 2011 |
Guided Mode Resonant Filter Biosensor Using a Linear Grating
Surface Structure
Abstract
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.
Inventors: |
Cunningham; Brian T.;
(Champaign, IL) ; Li; Peter; (Andover, MA)
; Qiu; Jean; (Andover, MA) ; Pien; Homer;
(Andover, MA) |
Assignee: |
SRU Biosystems, Inc.
|
Family ID: |
46205670 |
Appl. No.: |
13/042607 |
Filed: |
March 8, 2011 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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12048560 |
Mar 14, 2008 |
7923239 |
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13042607 |
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11201237 |
Aug 10, 2005 |
7371562 |
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12048560 |
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10059060 |
Jan 28, 2002 |
7070987 |
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11201237 |
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09930352 |
Aug 15, 2001 |
7094595 |
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10059060 |
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60244312 |
Oct 30, 2000 |
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60283314 |
Apr 12, 2001 |
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60303028 |
Jul 3, 2001 |
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Current U.S.
Class: |
506/39 ; 422/69;
435/288.7 |
Current CPC
Class: |
Y10S 436/805 20130101;
B01L 3/5085 20130101; G02B 5/1809 20130101; G01N 33/54373 20130101;
G01N 21/7743 20130101; G01N 21/253 20130101; G01N 21/4788 20130101;
Y10S 435/808 20130101; Y10S 436/813 20130101 |
Class at
Publication: |
506/39 ; 422/69;
435/288.7 |
International
Class: |
C40B 60/12 20060101
C40B060/12; G01N 30/00 20060101 G01N030/00; C12M 1/34 20060101
C12M001/34 |
Claims
1. A biosensor comprising: (a) a grating layer comprising a
one-dimensional or two-dimensional grating having a first surface,
wherein the grating layer is comprised of a low refractive index
material; (b) an interfacial layer on the first surface of the
grating layer; (c) a high refractive index material layer on a
surface of the interfacial layer opposite of the grating layer; and
(d) one or more specific binding substances immobilized on a
surface of the high refractive index material layer opposite of the
interfacial layer; wherein, when the biosensor is illuminated a
resonant grating effect is produced on the reflected radiation
spectrum.
2. The biosensor of claim 1, wherein the interfacial layer is
comprised of a material selected from the group consisting of
silicon oxide, silicon oxynitride, borosilicate glass,
phosphosilicate glass, pyrex, glass, and a metal oxide.
3. The biosensor of claim 1, wherein the interfacial layer is about
1 nm to about 200 nm thick.
4. The biosensor of claim 1, wherein the one or more specific
binding substances are detection label-free.
5. The biosensor of claim 1, wherein, when the biosensor is
illuminated a resonant grating effect is produced on a reflected
radiation spectrum, and wherein the cross-sectional period of the
grating layer is less than the wavelength of the resonant grating
effect.
6. The biosensor of claim 1, wherein the cross-sectional profile of
the grating layer is triangular, sinusoidal, trapezoidal,
rectangular, v-shaped, u-shaped, upside-down u-shaped, upside-down
v-shaped, stepped or square.
7. The biosensor of claim 1, wherein a narrow band of optical
wavelengths is reflected from the biosensor when the biosensor is
illuminated with a broad band of optical wavelengths.
8. The biosensor of claim 1, wherein the low refractive index
material comprises glass, plastic, polymer, or epoxy.
9. The biosensor of claim 1, wherein the high refractive index
material layer is zinc sulfide, titanium dioxide, indium tin oxide,
tantalum oxide, or silicon nitride.
10. The biosensor of claim 1, wherein the grating layer has a
cross-sectional period of about 0.01 microns to about 1 micron and
a depth of about 0.01 microns to about 1 micron.
11. The biosensor of claim 1, wherein the one or more specific
binding substances are arranged in an array of distinct
locations.
12. The biosensor of claim 11, wherein the distinct locations
define a microarray spot of about 10 to 500 microns in
diameter.
13. The biosensor of claim 1, wherein the one or more specific
binding substances are immobilized on the grating layer by physical
adsorption or by chemical binding.
14. The biosensor of claim 1, wherein the one or more specific
binding substances are selected from the group consisting of
nucleic acids, polypeptides, antigens, polyclonal antibodies,
monoclonal antibodies, single chain antibodies (scFv), F(ab)
fragments, F(ab).sub.2 fragments, Fv fragments, small organic
molecules, cells, viruses, bacteria, polymers, peptide solutions,
protein solutions, chemical compound library solutions,
single-stranded DNA solutions, double-stranded DNA solutions, RNA
solutions, and biological samples.
15. The biosensor of claim 14, wherein the biological sample is
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 prostatic fluid.
16. The biosensor of claim 1, wherein the biosensor is an internal
surface of a liquid-containing vessel.
17. The biosensor of claim 16, wherein the liquid-containing vessel
is selected from the group consisting of a microtiter plate, a test
tube, a petri dish and a microfluidic channel.
Description
PRIORITY
[0001] This application is a divisional of U.S. Ser. No.
12/048,560, filed Mar. 14, 2008 (now allowed), which is a
divisional application of U.S. Ser. No. 11/201,237, filed Aug. 10,
2005, now U.S. Pat. No. 7,371,562, 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
[0002] 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
[0003] 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.
[0004] 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)).
[0005] 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), evanescent 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
[0006] 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.
[0007] 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.
[0008] 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.
[0009] 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.
[0010] 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').sub.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
prostatic fluid.
[0011] 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.
[0012] 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.
[0013] 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.
[0014] 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.
[0015] 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.
[0016] 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.
[0017] 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.
[0018] 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.
[0019] 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.
[0020] 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
[0021] FIG. 1A shows a schematic diagram of one embodiment of an
optical grating structure used for a colorimetric resonant
reflectance biosensor. n.sub.substrate represents substrate
material. n.sub.1 represents the refractive index of a cover layer.
n.sub.2 represents the refractive index of a one- or
two-dimensional grating. n.sub.bio represents the refractive index
of one or more specific binding substances. t.sub.1 represents the
thickness of the cover layer. t.sub.2 represents the thickness of
the grating. t.sub.bio represents the thickness of the layer of one
or more specific binding substances. FIG. 1 shows a cross-sectional
view of a biosensor.
[0022] FIG. 2 shows a schematic drawing of a one-dimensional linear
grating surface structure.
[0023] FIG. 3A-B shows a two-dimensional grating comprising a
rectangular grid of squares (FIG. 3A) or holes (FIG. 3B).
[0024] FIG. 4 shows a biosensor cross-section profile utilizing a
sinusoidally varying grating profile.
[0025] 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.
[0026] 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.
[0027] 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.
[0028] 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.
[0029] FIG. 9 shows an example of a biosensor used as a
microarray.
[0030] FIG. 10A-B shows two biosensor formats that can incorporate
a colorimetric resonant reflectance biosensor. FIG. 10A shows a
biosensor that is incorporated into a microtitre plate.
[0031] FIG. 10B shows a biosensor in a microarray slide format.
[0032] FIG. 11 shows an array of arrays concept for using a
biosensor platform to perform assays with higher density and
throughput.
[0033] 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.
[0034] FIG. 13 shows a SEM photograph showing the separate grating
regions of an array of biosensor electrodes.
[0035] 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.
[0036] 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.
[0037] 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.
[0038] FIG. 17 demonstrates an example of a biosensor that occurs
on the tip of a fiber probe for in vivo detection of biochemical
substances.
[0039] FIG. 18 shows an example of the use of two coupled fibers to
illuminate and collect reflected light from a biosensor.
[0040] FIG. 19 shows resonance wavelength of a biosensor as a
function of incident angle of detection beam.
[0041] 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.
[0042] FIG. 21 shows an example of a system for angular scanning of
a biosensor.
[0043] FIG. 22 shows SEM photographs of a photoresist grating
structure in plan view (center and upper right) and cross-section
(lower right).
[0044] FIG. 23 shows a SEM cross-section photograph of a grating
structure after spin-on glass is applied over a silicon nitride
grating.
[0045] FIG. 24 shows examples of biosensor chips
(1.5.times.1.5-inch). Circular areas are regions where the resonant
structure is defined.
[0046] 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.
[0047] 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.
[0048] 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.
[0049] FIG. 28 shows a computer simulation of a biosensor
demonstrating the shift of resonance to longer wavelengths as
biomolecules are deposited on the surface.
[0050] 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.
[0051] FIG. 30 shows an embodiment of a biosensor. n.sub.substrate
represents the refractive index of a substrate. n.sub.1 represents
the refractive index of an optional optical cover layer. n.sub.2
represents the refractive index of a one- or two-dimensional
grating. n.sub.3 represents the refractive index of a high
refractive index material such as silicon nitride. n.sub.bio
represents the refractive index of one or more specific binding
substances. t.sub.1 represents the thickness of a cover layer.
t.sub.2 represents the thickness of a one- or two-dimensional
grating. t.sub.3 represents the thickness of a high refractive
index material. t.sub.bio represents the thickness of a specific
binding substance layer.
[0052] 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.
[0053] FIG. 32 shows a linear relationship between reflected
wavelength and protein coating thickness for a biosensor shown in
FIG. 30.
[0054] 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.
[0055] FIG. 34 shows the measured reflectance spectra of a
biosensor.
[0056] FIG. 35 shows dependence of peak resonant wavelength
measured in liquid upon the concentration of protein BSA dissolved
in water.
[0057] 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.
[0058] 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.
[0059] FIG. 38A-B. FIG. 38A shows results of streptavidin detection
at various concentrations for a biosensor that has been activated
with NH.sub.2 surface chemistry linked to a biotin receptor
molecule. FIG. 38B shows a schematic demonstration of molecules
bound to a biosensor.
[0060] 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.
[0061] 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.
[0062] 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.
[0063] 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.
[0064] 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.
[0065] FIG. 44A-B. FIG. 44A shows kinetic binding measurement of
IgG binding. FIG. 44B shows a schematic demonstration of molecules
bound to a biosensor.
[0066] 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.
[0067] 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.
[0068] FIG. 47 shows sensitivity of the mathematically determined
peak resonant wavelength to artificially added noise in the
measured spectrum.
[0069] FIG. 48 shows a resonant optical biosensor incorporating an
electrically conducting material.
[0070] FIG. 49 shows a resonant reflection or transmission filter
structure consisting of a set of concentric rings.
[0071] 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.
[0072] 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.sup.-15M), 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.
[0073] FIG. 52 shows a schematic diagram of a detection system.
[0074] 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.
[0075] 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 .about.200 nm. After
etching, the photoresist was removed.
[0076] 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.
[0077] 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.
[0078] 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.
[0079] FIG. 58A-C shows the exposure of NH.sub.2, PEG, and
PEG-Biotin activated one-dimensional linear grating surface
biosensor structures to seven concentrations of anti-biotin IgG.
The NH.sub.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.
[0080] 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.
[0081] FIG. 60 demonstrates the effect of a surface modification
layer on specific binding substance immobilization onto the surface
of a biosensor.
[0082] 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
[0083] Subwavelength Structured Surface (SWS) Biosensor
[0084] 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.
[0085] 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.
[0086] 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.
[0087] 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, .about.0.1 nm thickness of protein binding, and can be
performed with the biosensor surface either immersed in fluid or
dried.
[0088] 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.
[0089] 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.
[0090] A schematic diagram of an example of a SWS structure is
shown in FIG. 1. In FIG. 1, n.sub.substrate represents a substrate
material. n.sub.1 represents the refractive index of an optional
cover layer. n.sub.2 represents the refractive index of a
two-dimensional grating. N.sub.bio represents the refractive index
of one or more specific binding substances. t.sub.1 represents the
thickness of the cover layer above the two-dimensional grating
structure. t.sub.2 represents the thickness of the grating.
t.sub.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.
[0091] 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.
[0092] 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.
[0093] 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.
[0094] 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.
[0095] 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.
[0096] 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.
[0097] 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 colorimetric 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.
[0098] 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.
[0099] 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.
[0100] 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.
[0101] 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.
[0102] 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.
[0103] 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, n.sub.2, 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.
[0104] 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.
[0105] 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.
[0106] 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.
[0107] 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.
[0108] 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.
[0109] 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.
[0110] 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.
[0111] 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.
[0112] 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.
[0113] 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.
[0114] 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.
[0115] 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.
[0116] 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.
[0117] 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.
[0118] 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.
[0119] 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.
[0120] 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).
[0121] 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.
[0122] 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.
[0123] 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').sub.2 fragment, Fv fragment, small organic
molecule, cell, virus, and bacteria.
Specific Binding Substances and Binding Partners
[0124] 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').sub.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
prostatic fluid.
[0125] 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.
[0126] 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.
[0127] 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').sub.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 prostatic fluid.
[0128] 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.
[0129] 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.sup.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
[0130] 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.
[0131] 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.
[0132] 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)).
[0133] 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.
[0134] For the detection of binding partners at concentrations less
than about .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.times. 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.
[0135] As an example, FIG. 7A shows that an NH.sub.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.
[0136] 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.
[0137] 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.times. enhancement in sensitivity
resolution.
[0138] Surface-Relief Volume Diffractive Biosensors
[0139] 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.
[0140] 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.
[0141] 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.
[0142] 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.
[0143] 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.
[0144] 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).
[0145] 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.
[0146] 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.
[0147] 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.
[0148] 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.
[0149] 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.
[0150] 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
[0151] 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.
[0152] 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.
[0153] 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
[0154] Any number of biosensors that are, for example, about 1
mm.sup.2 to about 5 mm.sup.2, and preferably less than about
3.times.3 mm.sup.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 biosensors 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
[0155] 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.
[0156] 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.
[0157] 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.
[0158] 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.
[0159] 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.
[0160] 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
[0161] 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.
[0162] 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.sub.xO.sub.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.sub.1 represents the refractive index of an electrical insulator.
n.sub.2 represents the refractive index of a two-dimensional
grating. t.sub.1 represents the thickness of the electrical
insulator. t.sub.2 represents the thickness of the two-dimensional
grating. n.sub.bio represents the refractive index of one or more
specific binding substances and t.sub.BIO represents the thickness
of the one or more specific binding substances.
[0163] 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.
[0164] 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
[0165] 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.
[0166] 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.
[0167] 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
[0168] 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.sub.substrate=1.50). The grating is a two-dimensional pattern of
silicon nitride squares (t.sub.2=180 nm, n.sub.2=2.01 (n=refractive
index), k.sub.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.sub.1=1.40) that covers the silicon nitride squares by
t.sub.2=100 nm.
[0169] 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.
[0170] 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.
[0171] 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.
[0172] 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.
[0173] 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.
[0174] 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.
[0175] 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
[0176] 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.
[0177] 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.
[0178] 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.
[0179] 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.
[0180] 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.
[0181] 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
[0182] 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.
[0183] 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
[0184] 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.
[0185] 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.
[0186] The step of selecting a set of resonant reflectance data can
include selecting a set of resonant reflectance data:
[0187] x.sub.i and y.sub.i for i=1,2,3, . . . n,
wherein x.sub.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.sub.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.
[0188] 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.
[0189] 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.sub.k such that:
(y.sub.k>=y.sub.i) for all i.noteq.k.
[0190] 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.sub.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.
[0191] 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:
[0192] defining a curve-fit region of (2w+1) bins, wherein w is a
pre-determined accuracy value;
[0193] extracting (x.sub.i, k-w<=i<=k+w); and
[0194] extracting (y.sub.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:
[0195] computing g.sub.i=ln y.sub.i;
[0196] performing a 2.sup.nd order polynomial fit on g, to obtain
g'.sub.i, defined on
[0197] (x.sub.i, k-w<=i<=k+w);
[0198] determining from the 2.sup.nd order polynomial fit
coefficients a, b and c of for (ax.sup.2+bx+c); and
[0199] computing y'.sub.i=e.sup.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.sub.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.sub.p at y'.sub.p.
[0200] 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.
[0201] 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
[0202] Fabrication of a SWS biosensor
[0203] 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).
[0204] 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
[0205] 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 .about.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.
[0206] 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 .mu.l droplet is placed onto
a biosensor distinct location and allowed to dry in air. At 1
.mu.g/ml protein concentration, a 20 .mu.l droplet spreads out to
cover a 1 cm diameter circle and deposits about 2.times.10.sup.-8
grams of material. The surface density is 25.6 ng/mm.sup.2.
[0207] For high concentration protein immobilization (biosensor 4)
a 10 .mu.l droplet of 0.8 g bovine serum albumin (BSA) in 40 ml DI
H.sub.2O 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.sup.2. After protein deposition, biosensor 4
has a green resonance in air.
[0208] For bacteria immobilization (biosensor 2) a 20 .mu.l 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.
[0209] For low concentration protein immobilization (biosensor 6) a
10 .mu.l droplet of 0.02% of BSA in DI H.sub.2O (0.8 g BSA in 40 ml
DI H.sub.2O) 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.sup.2. After protein deposition, biosensor 6 looks grey in
air.
[0210] In order to obtain quantitative data on the extent of
surface modification resulting from the above treatments, the
biosensors were measured using a spectrometer.
[0211] 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.
[0212] 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.
[0213] 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
[0214] 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.sub.substrate=1.50). The grating is a two-dimensional pattern of
silicon nitride squares (t.sub.2=180 nm, n.sub.2=2.01,
k.sub.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.sub.1=1.40) that covers the silicon
nitride squares by t.sub.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.sub.bio=1.5) were added above the glass coating layer.
[0215] 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.
[0216] 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 .about.1.5 nm, resonant wavelength shifts
of .about.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 .about.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
[0217] 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.sub.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.sub.3=90 nm, n.sub.3=2.02) was
used. The grating is two-dimensional pattern of photoresist squares
(t.sub.2=90 nm, n.sub.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.sub.1=1.45) that covers the photoresist squares by t.sub.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.sub.bio=1.5) were added above
the glass coating layer.
[0218] 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.
[0219] 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
[0220] In order to detect reflected resonance, a white light source
can illuminate a .about.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 .about.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.
[0221] 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
[0222] 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 (.lamda..sub.peak=772.5
nm) compares with the resonant wavelength predicted by the computer
model (.lamda..sub.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.
[0223] 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-US-00001 TABLE 1 Calculated Measured Peak Shift Peak Shift
Solution n Wavelength (nm) (nm) Wavelength (nm) (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
[0224] 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
[0225] 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 .mu.l droplets
were allowed to dry in air, leaving a small quantity of BSA
distributed over a .about.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
[0226] The following protocol was used on a colorimetric 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.
[0227] 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.
[0228] 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 ddH2O 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.
[0229] 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.sub.2 activation
of the biosensor surface is comparable in the activation
commercially available microarray glass slides.
[0230] 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
[0231] 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 .ANG.. The
resulting biosensors can be used for capturing avidin or
strepavidin molecules.
Protocol for Activating Amine-Coated Biosensor with Aldehyde
[0232] 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
[0233] 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.
[0234] 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)).
[0235] 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-US-00002 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 .cndot. IgG HRP
sensitivity His-tag .cndot. IgG Phage Streptavidin 2-5x Others . .
. cDNA Cell Dextran cDNA TSPS
Example 9
IgG Assay
[0236] 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 .mu.g/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 .mu.l
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 .mu.g/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
[0237] 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 .mu.l 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
.mu.g/ml. This concentration corresponds to a molarity of 3.3
nM.
Example 11
Protein-Protein Binding Assay
[0238] 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 .mu.g/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 .mu.l 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 .mu.g/ml yields an
easily measurable wavelength shift.
Example 12
Unlabeled ELISA Assay
[0239] 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-.gamma. (IFN-.gamma.) with an IFN-.gamma.
antibody receptor molecule. The receptor molecule was covalently
attached to an NH.sub.2-activated biosensor surface with an SMPT
linker molecule (Pierce Chemical Company, Rockford, Ill.). The peak
resonant wavelength shift for application of the NH.sub.2, SMPT,
and anti-human IFN-.alpha. 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 .mu.g/ml. The first location was
exposed to IFN-.gamma., 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-.gamma. 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)
[0240] 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.
[0241] 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.
##STR00001##
[0242] 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.
[0243] A biosensor was prepared by cleaning in 3:1
H.sub.2SO.sub.4:H.sub.2O.sub.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 .mu.g/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 .mu.l of caspase-3 in 1.times. 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.
[0244] 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 .DELTA..lamda.=0.016 nm is
5.3.times. 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).
[0245] 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
[0246] 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.sub.2, and an
NHS-Biotin linker molecule. As shown in FIG. 42, goat
.alpha.-biotin antibodies of several different concentrations (1,
10, 100, 1000 .mu.g/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 .alpha.-Biotin IgG, a second layer of protein was
bound to the biosensor surface through the addition of .alpha.-goat
IgG at several concentrations (0.1, 1, 10, and 100 .mu.g/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.
[0247] FIG. 43 shows the kinetic binding curve for the final
resonant peak transitions from FIG. 42, in which 100 .mu.g/ml of
.alpha.-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.
[0248] 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
[0249] Proteomics Applications
[0250] 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
[0251] This example discusses some of the findings that have been
obtained from looking at fitting different types of curves to the
observed data.
[0252] The first analytic curve examined is a second-order
polynomial, given by
y=ax.sup.2+bx+c
The least-squares solution to this equation is given by the cost
function
.phi. = i = 1 n ( ax i 2 + bx i + c - y i ) 2 , ##EQU00001##
the minimization of which is imposed by the constraints
.differential. .phi. .differential. a = .differential. .phi.
.differential. b = .differential. .phi. .differential. c = 0.
##EQU00002##
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 ) . ##EQU00003##
The result of one such fit is shown in FIG. 46; the acquired data
are shown as dots and the 2.sup.nd-order polynomial curve fit is
shown as the solid line.
[0253] 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.sup.ax.sup.2.sup.+bx+c,
in which case the quadratic equations above can be utilized by
forming y', where y'=ln y. 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.
[0254] 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.
[0255] 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
.about.1.5 nm, the corresponding peak location changed by only
<0.06 nm, or 4 parts in one hundred sensitivity.
[0256] 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.
[0257] 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.sup.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-US-00003 TABLE 3 Comparison of peak location as a function of
window location Shift Window Peak Location .DELTA. = -10 bins
771.25-782.79 nm 778.8221 nm .DELTA. = 0 bins 772.70-784.23 nm
778.8887 nm .DELTA. = +10 bins 774.15-785.65 nm 7778.9653 nm
[0258] 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.
[0259] Based on these results, a basic algorithm for mathematically
determining the peak location of a colorimetric resonant biosensor
is as follows:
[0260] 1. Input data x.sub.i and y.sub.i, i=1, . . . ,n
[0261] 2. Find maximum [0262] a. Find k such that
y.sub.k.gtoreq.y.sub.i for all i.noteq.k
[0263] 3. Check that maximum is sufficiently high [0264] a. Compute
mean y and standard deviation .sigma. of sample [0265] b. Continue
only if (y.sub.k- y)/.sigma.>UserThreshold
[0266] 4. Define curve-fit region of 2w+1 bins (w defined by the
user) [0267] a. Extract x.sub.i, k-w.ltoreq.i.ltoreq.k+w [0268] b.
Extract y.sub.i,k-w.ltoreq.i.ltoreq.k+w
[0269] 5. Curve fit [0270] a. g.sub.i=ln y.sub.i [0271] b. Perform
2.sup.nd-order polynomial fit to obtain g'.sub.i defined on
x.sub.i,k-w.ltoreq.i.ltoreq.k+w [0272] c. Polynomial fit returns
coefficients a, b, c of form ax.sup.2+bx+c [0273] d. Exponentiate:
y'.sub.i=e.sup.g'.sub.i
[0274] 6. Output [0275] a. Peak location p given by x.sub.p=-b/2a
[0276] b. Peak value given by y'.sub.p (x.sub.p)
[0277] 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
[0278] 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 .DELTA.n=4.times.10.sup.-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.
[0279] Three test solutions were prepared for comparison:
[0280] 1. Avidin in Phosphate Buffer Solution (PBS), (10
.mu.g/ml)
[0281] 2. Avidin (10 .mu.g/ml)+Bovine Serum Albumin (BSA) (10
.mu.g/ml) in PBS
[0282] 3. Avidin (10 .mu.g/ml)+Biotinylated BSA (b-BSA) (10
.mu.g/ml) in PBS A single SWS sensor was used for all measurements
to eliminate any possibility of cross-sensor bias. A 200 .mu.l
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.
[0283] 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.sup.-15M), 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.
[0284] 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
[0285] 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.
[0286] 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 .about.200 nm. After
etching, the photoresist was removed.
[0287] 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.
[0288] Sensor fabrication was completed by sputter deposition of
120 nm tantalum oxide on the cured epoxy grating surface. Following
tantalum oxide deposition, 3.times.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.RTM., Cambridge, Mass. and Greiner,
Longwood, Fla.) with epoxy.
Surface Activation and Attachment of Receptor Molecule
[0289] 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.degree. C. for 10 min.
Other surface activation molecules could include, for example,
COOH, CHO, polymer, and Poly-phe-lysine.
[0290] 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
.about.2.times.10.sup.14 groups/cm.sup.2.
[0291] 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.
[0292] 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.
[0293] 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.TM. 20 in phosphate buffer solution, pH 8)
at 1.0 mg/ml concentration was added to the biosensor surface, and
incubated at 37.degree. 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
[0294] 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 .about.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
[0295] 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.times. lower. Most
importantly, the single narrow resonant peak characteristic is
uniformly obtained within every well of a 96-well microtiter plate
biosensor.
[0296] 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
(.DELTA.PWV) 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, .sigma.=.DELTA.PWV/An, 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 .sigma.=112 nm. The standard deviation of
.sigma. 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
[0297] A protein-antibody affinity assay was performed to
demonstrate operation of the plastic biosensor. A matrix of three
separate sensor surface states (NH.sub.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.
[0298] FIG. 57 plots the PWV shift--referenced to a sensor with no
chemical functional groups immobilized, recorded due to attachment
of NH.sub.2, NH.sub.2+(NHS-PEG), and NH.sub.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.sub.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.
[0299] 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 .mu.g/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.
[0300] 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 .mu.g/ml, 8.33 nM) is clearly measurable over the
negative control PSB solution.
Example 19
Biosensor Comprising a Surface Modification Layer
[0301] 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.
[0302] 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.sub.2 was added to a
biosensor with the surface modification layer and to a biosensor
without a surface modification layer. A NH.sub.2 reacting
fluorescent dye was used to visualize NH.sub.2 attachment to the
biosensors. See FIG. 60. The higher fluorescence intensity from the
surface modified layer indicates much higher NH.sub.2 density.
Example 20
[0303] Biosensor Design with High Stability in Aqueous
Solutions
[0304] 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.
[0305] 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
* * * * *