Canadian Patents Database / Patent 2615417 Summary

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(12) Patent Application: (11) CA 2615417
(54) English Title: PHOTONIC CRYSTAL BIOSENSOR STRUCTURE AND FABRICATION METHOD
(54) French Title: STRUCTURE DE BIOCAPTEUR A CRISTAL PHOTONIQUE ET METHODE DE FABRICATION
(51) International Patent Classification (IPC):
  • G01N 21/25 (2006.01)
  • G01N 33/53 (2006.01)
  • G01N 33/543 (2006.01)
  • G01N 21/55 (2006.01)
(72) Inventors :
  • CUNNINGHAM, BRIAN T. (United States of America)
(73) Owners :
  • SRU BIOSYSTEMS, INC. (United States of America)
(71) Applicants :
  • SRU BIOSYSTEMS, INC. (United States of America)
(74) Agent: MBM INTELLECTUAL PROPERTY LAW LLP
(45) Issued:
(86) PCT Filing Date: 2006-06-29
(87) PCT Publication Date: 2007-01-08
Examination requested: 2011-06-13
(30) Availability of licence: N/A
(30) Language of filing: English

(30) Application Priority Data:
Application No. Country/Territory Date
11/177,707 United States of America 2005-07-08

English Abstract



The invention provides sensor compositions and method of making sensors.


French Abstract

L'invention concerne des compositions de capteurs et des procédés de fabrication de capteurs.


Note: Claims are shown in the official language in which they were submitted.


CLAIMS:
We claim:
1. A sensor comprising a nanoporous material, having a low refractive index,
supported
on a bottom surface by a substrate, and coated on a top surface with a high
dielectric
constant dielectric coating;
wherein, the high dielectric constant dielectric coating or the high
dielectric
constant dielectric coating in combination with the nanoporous material form a
sub-
wavelength period grating structure;
wherein, when the sensor is illuminated a resonant grating effect is produced
on a
reflected radiation spectrum; and

wherein the depth and period of the sub-wavelength period grating structure
are
less than the wavelength of the resonant grating effect.
2. The sensor of claim 1, wherein a narrow band of optical wavelengths is
reflected from the
sensor when the sensor is illuminated with a broad band of optical
wavelengths.
3. The sensor of claim 1, wherein the refractive index of the nanoporous
material is from
about 1.1 to about 2.2.
4. The sensor of claim 1, wherein the refractive index of the nanoporous
material is from
about 1.1 to about 1.5.

5. The sensor of claim 1, wherein the period of the sub-wavelength period
grating structure is
about 50 nm to about 1,500 nm and the depth of the sub-wavelength period
grating structure is
about 50 nm to about 900 nm.
6. The sensor of claim 1, wherein the nanoporous material is porous silica
xerogel, porous
aerogels, porous hydrogen silsesquioxane, a B staged polymer, porous methyl
silsesquioxane,
porous poly(arylene ether), or combinations thereof.
7. The sensor of claim 1, wherein the substrate comprises glass, plastic or
epoxy.
8. The sensor of claim 1, wherein the refractive index of the dielectric
coating is about 1.8 to
about 3Ø
9. The sensor or claim 1, wherein the dielectric coating comprises tin oxide,
tantalum
pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or a combination
thereof.
10. The sensor of claim 1, wherein the refractive index of the substrate is
about 1.4 to about
1.6.

28


11. The sensor of claim 1, wherein the thickness of the dielectric coating is
about 30 nm to
about 700 nm and the thickness of the nanoporous material is about 10 nm to
about 5,000 nm.
12. The sensor of claim 1, wherein the dielectric coating has a cover layer on
its top surface.
13. The sensor of claim 1, wherein the sensor further comprises one or more
specific binding
substances immobilized on the high dielectric constant dielectric coating.
14. The sensor of claim 12, wherein the sensor further comprises one or more
specific binding
substances immobilized on the cover layer.
15. The sensor of claim 13, wherein the one or more specific binding
substances do not
comprise a detectable label.
16. The sensor of claim 13, wherein the one or more specific binding substance
are bound to
their binding partners.
17. The sensor of claim 16, wherein the one or more specific binding
substances and the
binding partners do not comprise a detectable label.
18. The sensor of claim 13, wherein the one or more specific binding
substances are arranged
in an array on the high dielectric constant dielectric coating.
19. The sensor of claim 14, wherein the one or more specific binding
substances are arranged
in an array on the cover layer.
20. A sensor comprising a waveguiding structure formed by a waveguiding film
covering a
substrate, wherein the waveguiding film has a refractive index higher than the
refractive index
of the substrate, and a diffraction grating is contained with in the
waveguiding structure,
wherein the diffraction grating is comprised of a nanoporous material having a
low dielectric
constant.
21. The sensor of claim 20, wherein the refractive index of the nanoporous
material is from
about 1.1 to about 1.5.
22. The sensor of claim 20, wherein the refractive index of the nanoporous
material is from
about 1.1 to about 2.2.
23. The sensor of claim 20, wherein the nanoporous material is porous silica
xerogel, porous
aerogels, porous hydrogen silsesquioxane, a B staged polymer, porous methyl
silsesquioxane,
porous poly(arylene ether), or combinations thereof
24. The sensor of claim 20, wherein the substrate comprises glass, epoxy, or
plastic.
29


25. The sensor of claim 20, wherein the waveguiding film comprises tin oxide,
tantalum
pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or a combination
thereof.
26. The sensor of claim 20, wherein the waveguiding film comprises a polymer.
27. The sensor of claim 20, wherein the sensor further comprises one or more
specific binding
substances immobilized on the waveguiding film.
28. The sensor of claim 27, wherein the one or more specific binding
substances do not
comprise a detectable label.
29. The sensor of claim 28, wherein the one or more specific binding substance
are bound to
their binding partners.
30. The sensor of claim 29, wherein the one or more specific binding
substances and the
binding partners do not comprise a detectable label.
31. The sensor of claim 27, wherein the one or more specific binding
substances are arranged
in an array on the high refractive index dielectric coating.


Note: Descriptions are shown in the official language in which they were submitted.


CA 02615417 2008-01-08

Atty. Docket No. 05--246-B
"Photonic crystal biosensor structure and fabrication method"
GOVERNMENT INTERESTS
This invention was made with Government support under Grant Number B1:SO4-
27657 awarded by the National Science Foundation. The Government has certain
rights
in the invention.

BACKGROUND OF THE INVENTION:
Label-free optical sensors based upon surface structured photonic crystals
have
recently been demonstrated as a highly sensitive method for performing a wide
variety of
biochemical and cell-based assays. See, e.g., Cunningham, et al., Label-Free
Assays on
the BIND System. Journal of Biomolecular Screening, 2004. 9:481-490. These
sensors
reflect only a narrow band of wavelengths when illuminated with white light at
normal
incidence, where positive shifts of the reflected peak wavelength value (PWV)
indicate
the adsorption of detected material on the sensor surface. See, e.g.,
Cunningham, et al.,
Colorimetric resonant reflection as a direct biochemical assay technique.
Sensors and
Actuators B, 2002. 81:316-328. By spatially confining incident photons at the
resonant
wavelength, a high optical field is generated at the sensor surface that
extends a short
distance into a test sample, much like an evanescent field. The high degree of
spatial
confinement of resonant photons within the device structure leads to a strong
intei-action
between the structure and adsorbed biomaterial, and to the ability to perform
high
resolution imaging of protein and cell attachment. See, e.g., Li, et al., A
new meth,od for
label-free imaging of biomolecular interactions. Sensors and Actuators B,
2004. 99:6-13.
Previously, photonic crystal optical biosensors have been fabricated from
continuous sheets of plastic film using a process in which the periodic
surface structure is
replicated from a silicon master wafer using a UV-cured polymer material. See,
e.g.,
Cunningham, et al., A plastic colorimetric resonant optical biosensor for
multiparallel
detection of label-free biochemical interactions. Sensors and Actuators B,
2002. 85:219-
226. This patterned polymer can be subsequently coated with a high refractive
index Ti02
layer that is generally thinner than the height of the surface structure. Such
devices have
been demonstrated for a wide variety of biochemical and cell-based assays,
with a mass
density sensitivity resolutiori less than 0.1 pg/mm2 and a large dynamic range
enabling
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CA 02615417 2008-01-08

Atty. Docket No. 05-246-B
single cell detection. See, e.g., Lin et al., A label-free biosensor-based
cell attac;hment
assay for characterization of cell surface molecules. Sensors and Actuators B,
Accepted
April 2005. In general, optimization of device sensitivity requires increasing
the
interaction of the electromagnetic field intensity distribution with the
molecules deposited

atop the photonic crystal surface. Therefore, selection of optical materials
and design of
the surface structure topology is aimed at extending the electromagnetic field
profile from
the interior regions of the photonic crystal (where they cannot interact with
adsorbed
material) to the region adjacent to the photonic crystal that includes the
liquid test
sample.

Methods are needed in the art to increase the sensitivity of these and other
types
of sensors and to decrease the cost of their manufacture.

SUMMARY OF THE INVENTION
One embodiment of the invention provides a sensor comprising a nanoporous
material, having a low refractive index, supported on a bottom surface by a
substrate, and
coated on a top surface with a high dielectric constant dielectric coating.
The high
dielectric constant dielectric coating or the high dielectric constant
dielectric coating in
combination with the nanoporous material form a sub-wavelength period grating
structure. When the sensor is illuminated a resonant grating effect is
produced on a
reflected radiation spectrum and the depth and period of the sub-wavelength
period
grating structure are less than the wavelength of the resonant grating effect.
A narrow
band of optical wavelengths can be reflected from the sensor when the serisor
is
illuminated with a broad band of optical wavelengths. The refractive index of
the
nanoporous material can be from about 1.1 to about 2.2. In another embodiment,
the
refractive index of the nanoporous material can be from about 1.1 to about
1.5. The
period of the sub-wavelength period grating structure can be about 50nm to
about 1,500
nm and the depth of the sub-wavelength period grating structure can be about
50 nm to
about 900 nm. The nanoporous material can be porous silica xerogel, porous
ae:rogels,
porous hydrogen silsesquioxane, a B staged polymer, porous methyl
silsesquioxane,
porous poly(arylene ether), or combinations thereof. The substrate can
comprise glass,
plastic or epoxy. The refractive index of the dielectric coating can be about
1.8 to about
3Ø The dielectric coating can comprise tin oxide, tantalum pentoxide, zinc
sulfide,
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CA 02615417 2008-01-08

Atty. Docket No. 05 -246-B
titanium dioxide, silicon nitride, or a combination thereof. The refractive
index of the
substrate can be about 1.4 to about 1.6. The thickness of the dielectric
coating can be
about 30 nm to about 700 nm and the thickness of the nanoporous material can
be about
nm to about 5,000 nm. The dielectric coating can have a cover layer on its top
5 surface. The sensor can further comprise one or more specific binding
substances
immobilized on the high dielectric constant dielectric coating. The sensor can
further
comprise one or more specific binding substances immobilized on the cover
laye:r. The
one or more specific binding substances can be free of detection labels. The
one or more
specific binding substance can be bound to their binding partners. The one o;r
more
10 specific binding substances and the binding partners can be free of
detection labels. The
one or more specific binding substances can be arranged in an array on the
high dielectric
constant dielectric coating. The one or more specific binding substances are
arrariged in
an array on the cover layer.
Another embodiment of the invention provides a sensor comprising a
waveguiding structure formed by a waveguiding film covering a substrate,
wherein the
waveguiding film has a refractive index higher than the refractive index of
the substrate,
and a diffraction grating contained with in the waveguiding structure, wherein
the
diffraction grating is comprised of a nanoporous material having a low
dielectric
constant. The refractive index of the nanoporous material can be from about
1.1 to about
1.5. In another embodiment of the invention the refractive index of the
nanoporous
material can be from about 1.1 to about 2.2. The nanoporous material can be
porous
silica xerogel, porous aerogels, porous hydrogen silsesquioxane, a B staged
polymer,
porous methyl silsesquioxane, porous poly(arylene ether), or combinations
thereo:E The
substrate can comprise glass, epoxy, or plastic. The waveguiding film
comprises tin
oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or
a combination
thereof The waveguiding film comprises a polymer. The sensor further comprises
one
or more specific binding substances immobilized on the waveguiding film. The
one or
more specific binding substances can be free of detection labels. The one or
more
specific binding substance can be bound to their binding partners. The one or
more
specific binding substances and the binding partners can be free of detection
labels. The
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CA 02615417 2008-01-08

Atty. Docket No. 05=-246-B
one or more specific binding substances can be arranged in an array on the
high refractive
index dielectric coating.

Use of an extremely low refractive index material for the surface structure in
sensors substantially increases detection sensitivity. Therefore, substances
can be
measured in test samples with 2-4x lower concentrations, molecular weights, or
binding
affinities than have been possible previously.

BRIEF DESCRIPTION OF THE DRAWINGS
Figure 1 shows a schematic of a nano-replicated nanoporous photonic crystal
biosensor.

Figure 2A-B shows (A) bulk and (B) surface shift predicted by GSolver
simulations.
Figure 3A-B shows SEM images of imprinted and cured periodic
NANOGLASS structure.
Figure 4 shows the experimental response of nanoporous sensor under imniersion
in de-ionized water (DI) and isopropyl alcohol (IPA).
Figure 5 A-E shows process flow for nanoporous sensor fabrication.
Figure 6 shows a schematic of high dielectric constant nanoporous photonic
crystal sensor.

Figure 7 shows a cross-section schematic of porous glass sensor.
Figure 8 shows resonant peak of porous glass sensor exposed to deionized
water,
as predicted by RCWA simulation.

Figure 9 shows experimentally measured resonant peak of nanoporous glass
sensor immersed in deionized water.
Figure 10 shows a kinetic plot comparing PWV shifts for PPL deposited onto
both porous glass and polymer sensor designs.
Figure 11 shows a partial profile of PWV shift versus polymer thickness, where
alternating layers of PSS and PAH contribute to the total measured shift.

Figure 12 shows a kinetic plot comparing PWV shifts for protein A deposited
onto both porous glass and polymer sensor designs.

Figure 13 shows binding kinetics of three animal IgGs to protein A measured
with
a nanoporous glass sensor.

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CA 02615417 2008-01-08

Atty. Docket No. 05-246-B
Figure 14 shows sensor comparison of PWV shifts for each of the different IgG -

protein A interactions.
DETAILED DESCRIPTION OF THE INVENTION

One embodiment of the invention provides a sensor that can be used to, inter
alia,
detect organic or inorganic material, such as protein, DNA, small molecules,
viruses,
cells, and bacteria, without the requirement of a label, such as fluorescent
or radioactive
labels. Photonic crystal sensors of the invention reflect only a very narrow
band of
wavelengths or one wavelength when illuminated with a broad wavelength light
source
(such as a white light or LED). The reflected color shifts to longer
wavelengths in
response to attachment of material to the sensor surface. Photonic crystal
sensor
structures of the invention provide 2-4x higher sensitivity than previously
described
structures. A key difference in the sensor structure that provides higher
sensitivity is the
replacement of a polymer sub-wavelength period grating structure with a
nanoporous low
refractive index material.

Methods for fabricating sensor structures that enable low cost manufacturing
are
also disclosed. Sensor structures of the invention have higher sensitivity
than previous
structures due to the use of nanoporous low refractive index material instead
of a polymer
sub-wavelength period grating structure. When the refractive index of the
sensor
structure directly beneath (and alternatively including) the sub-wavelength
grating
structure is reduced below the refractive index of any liquid used in a
sample, the
electromagnetic field of the photonic crystal interacts more strongly with the
test sample,
yielding a structure whose reflected wavelength is more strongly tuned by a
given
amount of adsorbed biological material. The system is capable of detecting,
e.g., a single
cell attached to its surface.
The principles of the instant invention can also be applied to, e.g.,
evanescent
wave-based biosensors and any biosensors incorporating an optical waveguide.
See, e.g.,
U.S. Pat. No. 4,815,843; U.S. Pat. No. 5,071,248; U.S. Pat. No. 5,738,825.

The sensors have utility in, inter alia, the fields of pharmaceutical
researcli (e.g.,
high throughput screening, secondary screening, quality control, cytotoxicity,
clinical
trial evaluation), life science research (e.g., proteomics, protein
interaction analysis,

DNA-protein interaction analysis, enzyme-substrate interaction analysis, cell-
protein
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CA 02615417 2008-01-08

Atty. Docket No. 05==246-B
interaction analysis), diagnostic tests (e.g., protein presence, cell
identificatior.t), and
environmental detection (bacterial and spore detection and identification).
Previous
patent applications and publications describe how the photonic crystal
biosensor surface,
in combination with a high resolution imaging instrument, can be used as a
platform for
performing many biochemical assays in parallel upon on single surface, using
only
nanoliters of sample material. See, e.g., U.S. Pat. Publ. Nos.: 2002/0168295;
2002/0127565; 2004/0132172; 2004/0151626; 2003/0027328; 2003/0027327;
2003/017581; 2003/0068657; 2003/0059855; 2003/0 1 1 3 766; 2003/0092075;
2003/0026891; 2003/0026891; 2003/0032039; 2003/0017580; 2003/0077660;
2004/0132214.

Photonic Crystal Sensors

A photonic crystal sensor of the invention can be used to create a sharp
optical
resonant reflection at a particular wavelength that can be used to track with
high
sensitivity the interaction of molecules, such as biological materials.

Photonic crystal sensors comprise a subwavelength structured surface.
Subwavelength structured surfaces are a type of diffractive optic that can
mirnic the
effect of thin-film coatings. See, e.g., Peng & Morris, "Resonant scattering
fronn two-
dimensional gratings," J. Opt. Soc. Am. A, Vol. 13, No. 5, p. 993, May 1996;
Magnusson,
& Wang, "New principle for optical filters," Appl. Phys. Lett., 61, No. 9, p.
1022, August,
1992; Peng & Morris, "Experimental demonstration of resonant anomalies in
diffraction
from two-dimensional gratings," Optics Letters, Vol. 21, No. 8, p. 549, April,
1996. A
grating of a photonic crystal sensor of the invention has a grating period
that is small
compared to the wavelength of incident light such that no diffractive orders
other than the
reflected and transmitted zeroth orders are allowed. A photonic crystal sensor
can
comprise a grating, which is comprised of or coated with a high dielectric
constant
dielectric material, sandwiched between a substrate layer and a cover layer
that fills the
grating grooves. Optionally, a cover layer is not used. The grating structure
selectively
couples light at a narrow band of wavelengths. This highly sensitive coupling
cor.ldition
can produce a resonant grating effect on the reflected radiation spectrum,
resulting in a
narrow band of reflected or transmitted wavelengths. The depth and period of
the grating
are less than the wavelength of the resonant grating effect.

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Atty. Docket No. 05-=246-B
The reflected or transmitted color of a photonic crystal sensor structure can
be
modified by the addition of molecules such as specific binding substances or
binding
partners or both to the upper surface of the cover layer or the grating
surface. The added
molecules increase the optical path length of incident radiation through the
sensor
structure, and thus modify the wavelength at which maximum reflectance or
transmittance will occur.

In one embodiment, a sensor, when illuminated with white light, is designed to
reflect only a single wavelength or a narrow band of wavelengths. When
molecules are
attached to the surface of the sensor, the reflected wavelength (color) is
shifted due to the
change of the optical path of light that is coupled into the grating. By
immobilizing
molecules, such as specific binding substances to a sensor surface,
complementary
binding partner molecules can be detected without the use of any kind of
fluorescent
probe or particle label. The detection technique can be performed with the
sensor surface
either immersed in fluid or dried.

When a photonic crystal sensor is illuminated with collimated white lig',ht
and
reflects only a narrow band of wavelengths, or a single band of wavelengths is
ref.lected.
The narrow wavelength band is described as a wavelength "peak." The "peak
wavelength value" (PWV) changes when molecules are deposited or removed from
the
sensor surface. A readout instrument illuminates distinct locations on the
sensor surface
with collimated white light, and collects collimated reflected light. The
collected light is
gathered into a wavelength spectrometer for determination of PWV.

Figure 1 shows a structure of a photonic crystal sensor of the invention. The
sensor comprises a substrate, a patterned, low-k, nanoporous material, and a
substantially
uniform, high refractive index coating. The surface of the low-k nanoporous
material is
patterned into a sub-wavelength period grating structure onto which the high
refractive
index material is deposited.
In general, a low-k dielectric material of the invention has a dielectric
constant, k,
of about 1.1 to about 3.9. Examples of low-k dielectric materials include, for
exEunple:
fluorosilicate glass (about 3.2- about 3.9); polyimides (about 3.1- about 3);
hydlrogen
silsesquioxane (HSQ) (about 2.9- about 3.2); diamond-like carbon (about 2.7-
about 3.4);
black diamond (SiCOH) (about 2.7- about 3.3); parylene-N (about 2.7); B-staged
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polymers (CYCLOTENETM and SiLKTM) (about 2.6- about 2.7); fluorinated
polyimides
(about 2.5- about 2.9); methyl silsequioxane (MSQ) (about 2.6- about 2.8);
poly(arylene
ether) (PAE) (about 2.6- about 2.8); fluorinated DLC (about 2.4- about 2.8);
paiylene-
F(about 2.4- about 2.5); PTFE (about 1.9); porous silica xerogels and aerogels
(about

1.1- about 2.2); porous hydrogen silsesquioxane (HSQ) (about 1.7- about 2.2);
porous
SiLKTM (a B staged polymer) (about 1.5- about 2.0); porous methyl
silsesquioxane
(MSQ) (about 1.8- about 2.2); porous poly(arylene ether) (PAE) (about 1.8-
about 2.2).

A low-k nanoporous material is an inorganic, porous, oxide-like low dielectric
material, wherein the refractive index, n, is about 1.1 to about 2.2, and
preferably about
1.1 to about 1.5. A low-k nanoporous material can be, for example, porous
silica

xerogels and aerogels (about 1.1- about 2.2); porous HSQ (about 1.7- about
2.2); porous
SiLKTM (a B staged polymer) (about 1.5- about 2.0); porous MSQ (about 1.8-
about 2.2);
porous PAE (about 1.8- about 2.2). In one embodiment of the invention the
nanoporous
material is NANOGLASSO, which is porous Si02. Porosity is created in the Si02
thereby reducing the dielectric constant from about 3.9 to as low as 1.9.
A material with a high refractive index, suitable for the invention includes,
e.g.,
tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon
nitride, or a
combination thereof A high k dielectric material has a refractive index of
about 1.8 to
about 3Ø Refractive index, n, describes the optical characteristics of a
medium and is
defined as the ratio of the speed of light in free space over the speed of
light in the
medium. A substrate can comprise, for example, glass, plastic or epoxy.
In one embodiment of the invention a sensor is defined by the following
parameters:
nh;K About 1.8 to about 3.0
nnano About 1.1 to about 1.5
nsua About 1.4 to about 1.6
A About 200 nm to about1500 nm
d About 50 nm to about 900nm
th;K About 30 nm to about 700 nm
tnano About 10 nm to about 5000 nm

In another embodinlent of the invention, the sensor structure comprises the
following materials:

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Atty. Docket No. 05-246-B
Substrate Material Glass
Nanoporous Material Nanoglass (Honeywell International, Santa Clara, CA)
High Refractive Index Ti02
Coating
and is defined by the following parameters.
nhix 2.25
nnano 1.17
nsub 1,50
A 550nm
d 170nm
thiK 120nm
tnano 600nm
Simulations of the above embodiment were performed using GSolver (Grating
Solver Development Co., Allen, TX) and FDTD Solutions (Lumerical Solutions,
Inc.,
Vancouver, BC, Canada). The results shown in Figure 2A predict the bulk
sensitivity to
be improved by more than a factor of two over that of previous designs. Bulk
sensitivity
is determined by the bulk shift coefficient, defined and calculated for this
embodiment
below.

APWV - ~.IP,, - ~.oI - 794.4 - 779.6 ' 308.3
On - nIPA -noI 1.378-1.330
(1)
Both simulation and experimental data for designs that do not incorporate a
nanoporous
material give bulk shift coefficients of approximately 150.
Since the proposed device functions by evanescent field interactions with
materials very near the sensor surface, it is instructive to consider a
refractive index shift
not only of the entire bulk media but also of a thin layer atop the sensor.
Figure 2B
shows GSOLVER simulation results with a 20nm thick "biological coating"
modeled by
a layer with a refractive index of 1.40. While individual biological molecules
or fractions
of biological molecule monolayers do not have a defined refractive index
value, the
biological layer was modeled as a uniform thin film of defined thickness for
the sake of
illustration.

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The SEM images of the patterned NANOGLASS structure shown in Figure 3
give evidence of a successful imprinting process. Upon deposition of Ti02, the
sensitivity of the completed sensor was interrogated using de-ionized water
and isopropyl
alcohol by examining the resulting peak wavelength (PWV) shift captured with a
spectrometer on the readout instrumentation. Applying Equation 1 using the
experimental data from Figure 4, the bulk shift coefficient can be calculated
as:

OPWV A,,,A -'I,,, = 855.1-841_1 = 291.7
An nIIIA -nD, 1.378-1.330

which agrees to within -5% of that demonstrated though simulation.

A cross-sectional profile of a subwavelength grating can comprise any
periodically repeating function, for example, a "square-wave." A grating can
be
comprised of a repeating pattern of shapes such as continuous parallel lines,
squares,
circles, ellipses, triangles, trapezoids, sinusoidal waves, ovals, rectangles,
and hexagons.
A sensor can comprise a one-dimensional linear grating surface structure, i.
e., a
series of parallel lines or grooves. 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.
These include, for example, triangular or v-shaped, u-shaped, upside-down v-
or u-
shapes, sinusoidal, trapezoidal, stepped and square. The grating can also be
sinusoidally
varying in height.

An alternate sensor 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 p:revious
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. All
polarization directions of


CA 02615417 2008-01-08

Atty. Docket No. 05-:246-B
such a structure have the same cross-sectional profile. The grating period of
a concentric ring
structure is less than the wavelength of the resonantly reflected light
A sensor of the invention can further comprise a cover layer on the surface of
a
grating opposite to 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 to
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.
Resonant reflection can also be obtained without a planarizing cover layer
over
the grating. Without the use of a planarizing cover layer, the surrounding
medium (such
as air or water) fills the grating. Therefore, molecules are immobilized to
the sensor on
all surfaces of a grating exposed to the molecules, rather than only on an
upper surface.

The invention provides resonant reflection structures and transmission filter
structures. For a resonant reflection structure, light output is measured on
the samLe 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 stirongly
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 structures 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.
In one embodiment of the invention, an optical device is provided. An optical
device
comprises a structure similar to any sensor of the invention; however, an
optical device does
not comprise one or more binding substances immobilized on the grating. An
optical device
can be used as a narrow band optical filter.
Evanescent wave-based sensors can comprise a waveguiding film supported by a
substrate; between the waveguiding film (and optionally as part of the
substrate) is a
diffraction grating. See, e.g., U.S. Pat. No. 4,815,843. A low-k dielectric
material, such

as low-k nanoporous material can be used for the diffraction grating or the
combined
11


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low-k nanoporous material and substrate. The waveguide comprises waveguiding
film
and the substrate. The waveguiding film can be, e.g., tin oxide, tantalum
pentoxide, zinc
sulfide, titanium dioxide, silicon nitride, or a combination thereof, or a
polymer such as
polystryrole or polycarbonate. A diffraction grating exists at the interface
of the

waveguiding film and the substrate or in the volume of the waveguiding film.
The
diffraction grating comprises a low-k material, such as low-k nanoporous
material. The
refractive index of the waveguiding film is higher than the index of the
adjacent media
(i.e., the substrate and the test sample). The substrate can be, e.g.,
plastic, glass or epoxy.
A specific binding substance can be immobilized on the surface of the
waveguiding film
and a test sample added to the surface. Laser light propagates in the
waveguiding film by
total internal reflection. Changes in refractive index of the waveguiding film
caused by
molecules binding to it can be detected by observing changes in the angle of
the ernitted,
out-coupled light.

Production of Sensors
Sensors of the invention can be produced using a flexible rubber template for
embossing the grating structure into the nanoporous material while the
nanoporous
material is in an uncured, deformable state. Unlike nonflexible solid
templates, the
flexible rubber template allows solvent vapors, generated by the nanoporous
material's
curing process, to escape. Many flexible templates can be generated from a
single silicon
wafer "master" template at low cost, and a single flexible template can be
used multiple
times to inexpensively produce many structured nanoporous sub-wavelength
grating
structures.
Sensors can be produced inexpensively over large surface areas and can also
be,
for example, incorporated into single-use standard disposable assay liquid
handling
formats such as microplates, microarray slides, or microfluidic chips.
A process flow for fabricating a photonic crystal incorporating a nanoporous
layer
is outlined in Figure 5. A patterned "master" wafer, usually silicon or glass,
which
contains features that will correspond precisely with those later imprinted
into the porous
film is designed (see Fig. 5A). The master is then used as a mold into which a
liquid
elastomer is poured, as shown in (Fig. 5B). Upon curing, the newly formed
negative
rubber "daughter" mold is carefully peeled away from the master. After
application of
12


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the porous film to the desired substrate, the daughter mold is set atop the
uncure(i film,
e.g., as depicted in (Figure 5D). With the mold in place, the porous material
is partially
cured, fully cured or not cured. The gas-permeable rubber mold allows solvent
evaporation during this curing process. Once the film can sustain a rigid
shape, the

daughter mold is peeled away and the remaining structure is allowed to fully
cure. A
completed device illustrated in (Fig. 5F) is obtained by depositing a thin
high refi active
index material uniformly across the patterned surface of the porous film.
Another approach for fabrication of a photonic crystal biosensor incorporating
a
nanoporous layer is illustrated in Figure 6. With this structure, a layer of
nanoporous
material is cured onto a substrate. Next, a high dielectric constant material
is uniformly

deposited on top of the porous layer. A high dielectric constant material has
a dielectric
constant, k, greater than about 5% higher than the k of the nanoporous
material. In one
embodiment of the invention the high dielectric constant material has a k of
greater than
about 3.5. The deposited high-k material is then patterned by e-beam or DUV
lithography, and subsequently etched to obtain the desired features. While
this sensor
design is not as cost effective due to the need for high-resolution
lithographic processes
for each device, it shows promise for obtaining sensitivity enhancements
similar to those
seen with the aforementioned sensor fabricated by imprinting.
Evanescent-wave based biosensors can also be made using the same processes as
described herein.

Specific Binding Substances and Binding Partners
One or more specific binding substances can be immobilized on a grating or=
cover
layer, if present, by for example, physical adsorption or by chemical binding.
A specific
binding substance can be, for example, an organic molecule, such as a nucleic
acid,
polypeptide, antigen, polyclonal antibody, monoclonal antibody, single chain
antibody
(scFv), F(ab) fragment, F(ab')2 fragment, Fv fragment, small organic molecule,
cell,
virus, bacteria, polymer, peptide solutions, single- or double-stranded DNA
solutions,
RNA solutions, solutions containing compounds from a combinatorial chemical
liibrary,
or biological sample; or an inorganic molecule. A biological sample can be for
example,
blood, plasma, serum, gastrointestinal secretions, homogenates of tissues or
tumors,
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Atty. Docket No. 05-246-B
synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid,
cerebrospinal fluid,
peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, or
prostatitc fluid.
Preferably, one or more specific binding substances are arranged in a
microarray
of distinct locations on a sensor. One or more specific binding substances can
be bound
to their specific binding partners. A microarray of specific binding
substances coniprises
one or more specific binding substances on a surface of a sensor 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 sensor
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 substances laid out in
an.y 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.

A microarray on a sensor 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 grating or cover layer surface. When the sensor is exposed to a test sample
comprising
one or more binding partners, the binding partners will be preferentially
attracted to
distinct locations on the microarray that comprise specific binding substances
that have
high affinity for the binding partners. Some of the distinct locations will
gather binding
partners onto their surface, while other locations will not.
A specific binding substance specifically binds to a binding partner that is
contacted with the surface of a sensor of the invention. A specific binding
substance
specifically binds to its binding partner, but does not substantially bind
other binding
partners contacted with the surface of a sensor. 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 arrtigens.
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')Z
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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, an inorganic molecule, or a
biological
sample.

One example of a microarray of the invention is a nucleic acid microarray, in
which each distinct location within the array contains a different nucleic
acid molecule.
In this embodiment, the spots within the nucleic acid microarray detect
complementary
chemical binding with an opposing strand of a nucleic acid in a test sample.

While microtiter plates are the most common format used for biochemical
assays,
microarrays are increasingly seen as a means for maximizing the number of
biochemical
interactions that can be measured at one time while minimizing the volume of
precious
reagents. By application of specific binding substances with a microarray
spotter onto a
sensor of the invention, specific binding substance densities of 10,000
specific binding
substances/in2 can be obtained. By focusing an illumination beam to
interrogate a single
microarray location, a sensor can be used as a label-free microarray readout
system.
While it is not necessary for specific binding substances or binding partners
to
comprise a detectable label, detectable labels can be used to detect specific
binding
substances or binding partners on the surface of a sensor. Where specific
binding
substances and binding partners of the instant invention are free of detection
labels, they
can still comprise other types of labels and markers for enhancement.of assay
sensitivity,
immobilization of specific binding partners to a biosensor surface,
enhancement of
binding or hybridization of specific binding substances to their binding
partners, and for
other purposes.

Immobilization of One or More Specific Bindinll Substances
Molecules can be immobilized onto a sensor is so that they will not be viashed
away by rinsing procedures, and so that binding to molecules in a test sample
is
unimpeded by the sensor surface. Several different types of surface chemistry
strategies
have been implemented for covalent attachment of molecules to, for example,
glass for
use in various types of microarrays and sensors. These same methods can be
readily
adapted to a sensor of the invention.



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One or more types of molecules can be attached to a sensor 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
molecules on a sensor surface and provide defined orientation and conformation
of the
surface-bound molecules.
Other types of chemical binding include, for example, amine activation,
alclehyde
activation, and nickel activation. These surfaces can be used to attach
several different types
of chemical linkers to a sensor surface. VVhile 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 lin.ker. 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 sur.Face is
well known in the art (Whitesides, Anal. Chem. 68, 490, (1996)).
Immobilization of specific binding substances to plastic, epoxy, or high refi-
active
index material can be performed essentially as described for immobilization to
glass.
However, an acid wash step can be eliminated where such a treatment would
damage the
material to which the specific binding substances are immobilized.

Liquid-Containing Vessels
A sensor 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 sensor that is incorporated into any type of microtiter
plate. For
example, a sensor can be incorporated into the bottom surface of a microtiter
plate by
assembling the walls of the reaction vessels over the resonant reflection
surface, 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.
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
16


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Atty. Docket No. 05-246-B
mechanical configuration of these plates, liquid dispensing, robotic plate
handling, and
detection systems are designed to work with this common format. A sensor of
the
invention can be incorporated into the bottom surface of a standard microtiter
plate.
Because the sensor surface can be fabricated in large areas, and because the
readout
system does not make physical contact with the sensor surface, an arbitrary
number of
individual sensor 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 sensor surface.

A sensor can also be incorporated into other disposable laboratory assay
formats,
such as microarray slides, flow cells, and cell culture plates. Incorporation
of a sensor
into common laboratory formats is desirable for compatibility with existing
microarray
handling equipment such as spotters and incubation chambers.
Methods of Using Sensors
Sensors of the invention can be used, e.g, study one or a number of
molecule/molecule interactions in parallel; for example, 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 sensor that has one or
more specific
binding substances immobilized on its surface. A sensor is illuminated with
light and a
maximum in reflected wavelength, or a minimum in transmitted wavelength of
1:ight is
detected from the sensor. If one or more specific binding substances have
bound to their
respective binding partners, then the reflected wavelength of light is shifted
as conipared
to a situation where one or more specific binding substances have not bound to
their
respective binding partners. Where a sensor is coated with an array of
distinct locations
containing the one or more specific binding substances, then a maximum in
reflected
wavelength or minimum in transmitted wavelength of light is detected from each
distinct
location of the sensor.
In one embodiment of the invention, a variety of specific binding substances,
for
example, antibodies, can be immobilized in an array format onto a sensor of
the
invention. The sensor is then contacted with a test sample of interest
comprising b:inding
partners, such as proteins. Only the proteins that specifically bind to the
antibodies
immobilized on the sensor remain bound to the sensor. Such an approach is
essentially a
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Atty. Docket No. 05==246-B
large-scale version of an enzyme-linked immunosorbent assay; however, the use
of an
enzyme or fluorescent label is not required.

The activity of an enzyme can be detected by detecting the reflected
wave;length
of light from a sensor on which one or more specific binding substances have
been
immobilized and applying one or more enzymes to the sensor. The sensor is
washed and
illuminated with light. The reflected wavelength of light is detected from the
sensor.
Where the one or more enzymes have altered the one or more specific binding
substances
of the sensor by enzymatic activity, the reflected wavelength of light is
shifted.

Additionally, a test sample, for example, cell lysates containing binding
partners
can be applied to a sensor of the invention, followed by washing to remove
ur.ibound
material. The binding partners that bind to a sensor can subsequently be
eluted from the
sensor and identified by, for example, mass spectrometry. Optionally, a phage
DNA
display library can be applied to a sensor of the invention followed by
washing to remove
unbound material. Individual phage particles bound to the sensor can be
isolated and the
inserts in these phage particles can then be sequenced to determine the
identities of the
binding partners.

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
sensor of the
invention, followed by the ability to selectively remove bound material from a
clistinct
location of the sensor for further analysis is advantageous. Sensors of the
invention are
also capable of detecting and quantifying the amount of a binding partner from
a sample
that is bound to a sensor array distinct location by measuring the shift in
reflected
wavelength of light. Additionally, the wavelength shift at one distinct sensor
location can
be compared to positive and negative controls at other distinct sensor
locations to
determine the amount of a binding partner that is bound to a sensor array
distinct
location.

In one embodiment of the invention, an interaction of a first molecule with a
second
test molecule can be detected. A sensor as described above is used; however,
there are no
specific binding substances immobilized on its surface. Therefore, the sensor
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 sensor is
illumiriated a
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resonant grating effect is produced on the reflected radiation spectrum, and
the depth and
period of the grating are less than the wavelength of the resonant grating
effect.
To detect an interaction of a first molecule with a second test molecule, a
mixture of
the first and second molecules is applied to a distinct location on a sensor.
A distinct location
can be one spot or well on a sensor or can be a large area on a sensor. A
mixture of the first
molecule with a third control molecule is also applied to a distinct location
on a sensor. The
sensor can be the same sensor as described above, or can be a second sensor.
If the sensor is
the same sensor, a second distinct location can be used for the mixture of the
first molecule
and the third control molecule. Alternatively, the same distinct sensor
location can be used
after the first and second molecules are washed from the sensor. 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
sensor or sensors 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 frigment
to an antigen, and binding of polypeptides. A first molecule, second test
molecule, or third
control molecule can be, for example, a nucleic acid, polypeptide, antigen,
polyclonal
antibody, monoclonal antibody, single chain antibody (scFv), F(ab) fragment,
F(ab')2
fragment, Fv fragment, small organic molecule, cell, virus, and bacteria.
All patents, patent applications, and other scientific or technical writings
referred to
anywhere herein are incorporated by reference in their entirety. The methocls
and
compositions described herein as presently representative of preferred
embodiments are
exemplary and are not intended as limitations on the scope of the invention.
Changes
therein and other uses will be evident to those skilled in the art, and are
encompassed
within the spirit of the invention. The invention illustratively described
herein suitably
can be practiced in the absence of any element or elements, limitation or
limitatioins that
are not specifically disclosed herein. Thus, for example, in each instance
herein any of

the terms "comprising", "consisting essentially of', and "consisting of' can
be replaced
with either of the other two terms. The terms and expressions which have been
employed
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are used as terms of description and not of limitation, and there is no
intention in the use
of such terms and expressions of excluding any equivalents of the features
shovvn and
described or portions thereof, but it is recognized that various modifications
are possible
within the scope of the invention claimed. Thus, it should be understood that
al though
the present invention has been specifically disclosed by embodiments and
optional
features, modification and variation of the concepts herein disclosed are
considered to be
within the scope of this invention as defined by the description and the
appended claims.
In addition, where features or aspects of the invention are described in terms
of
Markush groups or other grouping of alternatives, those skilled in the art
will recognize
that the invention is also thereby described in terms of any individual member
or
subgroup of members of the Markush group or other group.
Examples
Example 1
Computer Simulation
Rigorous Coupled Wave Analysis (RCWA) and Finite Difference Time Domain
(FDTD) simulations were used to predict the resonant wavelength and bulk
refractive
index sensitivity of a one-dimensional surface photonic crystal biosensor. The
device
incorporates a low-index (n=1.17) nanoporous dielectric surface structure in
place of the
polymer (n=1.39) surface structure reported previously. A soft contact
embossing
method was used to create a surface-structured low-index porous film on glass
substrates
with a depth and period that are identical to the previous polymer structures
to etiable a
side-by-side sensitivity comparison. The sensitivity of porous glass
biosensors was
compared to nonporous polymer biosensors through methods that characterize
sensitivity
to bulk refractive index and surface-adsorbed material. Finally, a protein
binding assay
comparison was performed to demonstrate sensor stability and the ability to
functionalize
the device for selective detection.
The polymer and porous glass sensors were modeled and simulated using two
software packages. First, a 2-D diffraction grating analysis tool (GSOLVER)
employing
the RCWA algorithm provides a quick and simple method for initial sensor
modeling.
Second, FDTD (Lumerical) provides a much more versatile and powerful tool that
can
calculate any field component at any temporal or spectral location for an
arbitrary optical


CA 02615417 2008-01-08

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device illuminated by an arbitrary source. See, e.g., Kunz & Luebbers, The
Finite
Difference Time Domain Method for Electromagnetics. 1993, Boca Raton: CRC
Press.
FDTD was used to verify RCWA results and to gain deeper insight into the
effects of
modifying the sensor structure.
RCWA and FDTD simulations both indicated that replacement of the patterned
UV-cured polymer of previous devices with a material of lower refractive index
would
produce a two-fold increase in the bulk shift coefficient. The resonant
wavelength of the
porous glass sensor immersed in DI H20 was predicted by RCWA to be 844.3 nm
with a
full-width at half-maximum (FWHM) of approximately 2 nm, as shown in Figure 8.
Simulation predicts further improvements in the bulk shift coefficient with
slight
modifications to the sensor geometry.
The bulk sensitivity test using DI H20 and IPA was performed on 23 porous
glass
sensors and 11 polymer sensors. The average PWV shifts were 13.6 2.4 nm and
5.1
1.5 nm for the porous glass and polymer sensors, respectively. The bulk shift
coefficient

(APWV/An) of the porous glass sensor is measured to be 2.7 1.2 times greater
than that
of the polymer device. Measurements of the porous glass device in DI H20 give
an
average PWV of 829.5 16.5 nm and FWHM of 3.5 2.5 nm. One of the measured
spectra is illustrated in Figure 9, where the response has been normalized to
a perfect
reflector to account for any instrumentation losses. The lower reflection
efficiency and
broader FWHM measured from the replicated devices can be attributed to small
but
measurable material losses and to imperfections observed in the replicated
structure. The
large variability of measured spectral characteristics is due, at least in
part, to using
several slightly different (though nominally identical) master patterns and to
a lack of
automation of the replication process.

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Example 2
Sensor Fabrication

A sol-gel derived low-index nanoporous silica thin-film (See, e.g., U.S.
Patent No.
6,395,651) was incorporated into a sensor in place of the UV-cured epoxy used
in
previous designs. Since the low-index material cures by heat rather than UV
exposure, it
was necessary to develop a new fabrication process. It was desirable to retain
a loiv-cost
imprinting method, though it was obvious that a plastic substrate could not
sustain the
requisite high temperatures f'or porous glass annealing. One possible approach
to sol-gel
glass imprinting was to use a polydimethylsiloxane (PDMS) mold and a glass
substrate.
See, e.g., Parashar, et al., Nano-replication of diffractive optical elements
in sol-gel
derived glasses. Microelectronic Engineering, 2003. 67-8: p. 710-719.
The sub-wavelength grating structure of the low-k biosensor was fabricated
using
a combination of lithography, molding, and imprinting processes. Sylgard 184
PDMS
(Dow Corning) daughter molds are first replicated from a silicon master wafer
patterned
with a positive image of the surface structure desired in the finished sensor.
To facilitate
release of the PDMS mold f'rom the silicon wafer, the wafer was surface
treated with a
release layer of dimethyldichlorosilane (Repel Silane, Amersham Biosciences).
See, e.g.,
Beck et al., Improving stamps for 10 nm level wafer scale nanoimprint
lithogr=aphy.
Microelectronic Engineering, 2002. 61-2: p. 441-448. The PDMS replicas are
ther.t used
to imprint a thin film of uncured NANOGLASS (Honeywell Elec. Mat.), a low-
index
sol-gel glass, spun-on to a glass substrate. Once the low-index dielectric
becomes rigid,
the flexible PDMS mold is removed and the sol-gel glass is fully cured by
further baking.
The sensor structure is completed by evaporating 175nm of Ti02 onto the
patterned
surface. A subsequent surface treatment with dimethyldichlorosilane encourages
bio-
adsorption and promotes sensor stability. A schematic illustrating the cross-
section of the
device is shown in Figure 7.

The polymer structure is similar to that described in a previous publication.
See,
e.g., Cunningham et al., A plastic colorimetric resonant optical biosensor for
multiparallel detection of label-free biochemical interactions. Sensors and
Actuators B,
2002. 85: p. 219-226. Both structures use a 550nm period and 170mn imprint
depth,
though the polyester/polymer and low-index porous glass devices use 120nm and
165mn
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Ti02 coatings, respectively. The two devices will be referred to as the
"polymer" and
"porous glass" sensors throughout the remainder of the examples. The polymer
devices
were provided as an array of sensors aligned and attached to bottomless 96-
well standard
microtiter plates (SRU Biosystems). The porous glass devices are fabricated on
75mm x
25mm x Imm glass microscope slides. Adhesive rubber wells (Research
International
Corp.) are attached to the glass surface to provide liquid containment for 5-6
sensors on
each slide.
Deionized water (DI H20, n= 1.333) and Isopropyl Alcohol (IPA, n = 1.378)
were used to determine the bulk shift coefficient of each sensor. First, DI
H20 was
pipetted onto the surface of the sensor and the PWV was measured. The
configuration of

the readout instrument has been reported previously. See, e.g., Cunningham et
al.,
Colorimetric resonant reflection as a direct biochemical assay technique.
Sensors and
Actuators B, 2002. 81:316-328. A broad wavelength light source was coupled to
an
optical fiber that illuminates a-2 mm diameter region of the photonic crystal
surface
from below the substrate at normal incidence. Reflected light was collected by
a second
optical fiber that is bundled next to the illuminating fiber, and measured by
a
spectrometer. An automated motion stage enables parallel collection of
reflectance data
at timed intervals from many wells in order to acquire kinetic information.
Next, the surface was thoroughly dried and the previous step was repeated for
IPA. The bulk shift coefficient between DI H20 and IPA was then be calculated
as the
change in PWV divided by the change in bulk refractive index.

Example 3
PPL Bio-Adhesion Test
Sensitivity to surface-adsorbed material was characterized by the detection of
a
single layer film of Poly(Lys, Phe) (PPL; Sigma-Aldrich; MW = 35,400 Da)
prepared to
a 0.5 mg/ml solution with 0.01 M phosphate buffered saline (PBS; pH = 7.4)
applied to
the sensor surface. At a sampling interval of 1 minute, the bio-adhesion test
commenced
with the pipetting of PBS into the test wells. After 10 minutes, the buffer
was replaced
with PPL solution and was allowed to stabilize for 30 minutes. The wells were
then
washed three times and filled with PBS for the final 30 minutes of data
acquisition.

23


CA 02615417 2008-01-08

Atty. Docket No. 05-246-B
PPL was deposited on 5 porous glass and 9 polymer sensors. Figure 10 cornpares
the kinetic plots of each device, showing a-4x increase in surface sensitivity
f:or the
porous glass sensor. The first step establishes a baseline, the second
corresponds to the
rapid surface adsorption and. saturation of PPL, and the final third of the
curve illustrates
the monolayer stability after eliminating weakly or non-specifically bound
molecules by
rinsing with PBS buffer. The PWV shifts generated during PPL immobilization
onto the
porous glass sensor saturate more slowly than that measured using the polymer
devices.
Clearly, the porous glass sensor surface is significantly less conducive to
protein
monolayer adsorption. Further surface chemistry optimization should mitigate
this effect.
Nonetheless, the porous glass sensor exhibits excellent stability after
unbound molecules
are washed away.

Example 4
Multilayer Polymer Test

In order to characterize the differential sensitivity as a function of
distance from
the sensor surface, a series of polymer electrolyte monolayers were deposited
on the
sensors. By alternating between positively and negatively charged polymer
layers, a
stack of uniform, self-limiting polymers may be formed on the sensor while it
is
continuously monitored on the detection instrument. See, e.g., Cunningham et
al.,
Enhancing the surface sensitivity of colorimetric resonant optical biosensors.
Sensors
and Actuators B, 2002. 87(2):365-370. Three different polyelectrolytes were
deposited
onto the sensor surface: anionic poly(sodium 4-styrenesulfonate) (PSS; MW =
70,000
Da), cationic poly(ethyleniminie) (PEI; MW = 60,000 Da), and cationic
poly(allylamine
hydrochloride) (PAH; MW = 70,000 Da). The polyelectrolytes were purchased from
Sigma-Aldrich. A 0.9 M NaCI buffer solution (Sigma-Aldrich) was prepareci with
deionized water. The polyelectrolytes were dissolved in the buffer solutior.k
at a
concentration of 5 mg/ml. At a 1 minute sampling interval, the multilayer
surface
sensitivity characterization was performed in 5 minute steps. First, NaCI
buffe;r was
pipetted into the sensor wells. Next, the buffer was removed and replaced by
PEI
solution. The wells were then washed 3 times and filled with buffer. The
previous 2
steps were repeated for PSS and PAH until 7 PSS-PAH layers had been deposited
atop
the single PEI layer.

24


CA 02615417 2008-01-08
Atty. Docket No. 05-246-B
The 14 alternating layers of PSS and PAH described previously each cause a
measurable shift in the detected PWV as they are adsorbed onto the surface.
Figure 11
gives a spatial profile of PWV shift versus polymer thickness, where each PWV
shiift was
measured in buffer after the wash step. Each monolayer of polyelectrolyte is
approximately 4.4 nm thick and has a refractive index of 1.49. See, e.g.,
Picart et al.,
Determination of structural parameters characterizing thin films by optical
methods: A
comparison between scanning angle reflectometry and optical waveguide
lightmode
spectroscopy. Journal of Chemical Physics, 2001. 115(2): p. 1086-1094. The
porous glass
sensor exhibits an average surface sensitivity --1.5x that of the polymer
sensor. However,
note that each of the first 2 layers (-9nm) deposited onto the porous glass
device cause a
PWV shift with twice the magnitude of each of the remaining layers, while no
such effect
is observed for the polymer device.

Example 5
Bioassay: Protein A - IgG

To demonstrate selective detection by the proposed device, a bioassay was
performed that characterizes the affinity of human, sheep and chicken IgG for
protein A.
Protein A (Pierce Biotechnology) was prepared with 0.01 M PBS to a
concentration of
0.5 mg/ml. The buffer was filtered with a 0.22 m syringe filter (Nalgene)
before use.
Human, sheep, and chicken immunoglobulin-G (IgG) serums (Sigma-Aldrich) were
diluted in 0.01 M PBS to a concentration of 0.5 mg/ml. Allowing thirty minutes
between
each step and sampling at a one minute interval, PBS solution was first
pipetted irito the
sensors wells. Next, the buffer was replaced by protein A solution. The well
was then
rinsed 3 times and filled with buffer. After the signal stabilized, the buffer
in three of the
wells was replaced by human, sheep, or chicken IgG, while the fourth was left
as a
reference containing only the buffer. Finally, the IgGs were removed and the
wells were
again rinsed and filled with PBS for the final 30 minutes of data acquisition.

Protein A was introduced into 15 porous glass and 16 polymer sensor wells. The
resulting PWV shift after the wash step was -4x greater for the porous glass
de;vices.
Figure 13 illustrates the measured binding kinetics of human, sheep, and
chicken IgG
with protein A for the porous glass sensor, while Figure 14 gives an endpoint
PWV shift
comparison (relative to a reference well without IgG) between the two devices
for each


CA 02615417 2008-01-08

Atty. Docket No. 05-246-B
antibody. Protein A surface adsorption saturated much more quickly on the
polymer
sensor surface, similar to that observed in the PPL bio-adhesion test. The
porous glass
device exhibits increasingly greater sensitivity for antibodies with higher
affinity for
protein A. Human IgG binding was detected with twice the sensitivity, while
Chicken
IgG, lacking any specificity for protein A (See, e.g., Richman et al., The
bind'ing of
Staphylococci protein A by the sera of different animal species. Journal of
Immunology,
1982. 128: p. 2300-2305), results in an equivalent response and provides a
measure of
non-specific binding.

A photonic crystal biosensor is designed to couple electromagnetic energy to
biological material depositecl upon its surface from a liquid test sample.
While the device
itself consists of a low refractive index surface structure and a high
refractive index
dielectric coating, the liquid test sample that fills in the surface structure
must also be
considered an integral part of the sensor - and the only dynamic component
that can
induce a change of resonant wavelength. The motivation for incorporating an
extremely

low refractive index material into the photonic crystal biosensor structure is
to bias the
electromagnetic field of the resonant wavelength to interact more strongly
with the liquid
test sample and less strongly with the interior regions of the photonic
crystal that do not
interact with surface adsorbed material.
The use of spin-on low-k dielectric materials leverages off the large
investments
made in the integrated circuit manufacturing community, who require rapid
processing,
structural stability, and exclusion of liquid penetration. A unique aspect of
this work is
the use of an imprinting method to accurately impart a submicron surface
structui-e to a
nanoporous glass film without the use of photolithography. The presence of the
iinprint
tool on the surface of the low-k film during the initial stage of the curing
process dlid not
alter the refractive index of the final cured structured film. The imprinting
niethod
enables substantial cost to be incurred only in the production of the "master"
silicon
wafer, which is in turn used to produce a nearly unlimited number of
"daughter" PDMS
imprinting tools. Each PDMS tool can be used to produce a large number of
sensor
structures without damage to the tool because little force is needed to make
the spun-on
liquid low-k layer conform to the tool. After imprinting, the low-k dielectric
film is
cured rapidly on hotplates, using methods that are easily automated. The use
of a flexible
26


CA 02615417 2008-01-08

Atty. Docket No. 05-246-B
imprinting tool was found to be advantageous over imprinting from the silicon
inaster
wafer directly, as the PDMS mold was easier to release from the partially
cured low-k
film, and was capable of allowing permeation of volatile solvent released
during the cure
process. Although only 1x3-inch microscope slide regions were imprinted in the
work
shown here, the imprinting method can be scaled to larger surface areas to
enable
production of sensor areas large enough to cover an entire 96-well or 384-well
standard
microplate (approximately 3x5-inches).

An interesting and useful result found during comparison of porous glass
sensor
structures with polymer sensor structures is the disparity in sensitivity
gains between bulk
refractive index sensitivity and surface-adsorbed layer sensitivity. While
computer
models accurately predict the -2x sensitivity increase measured for PWV shift
induced
by a bulk refractive index change of the solution covering the porous glass
sensor
surface, a-4x increase of PWV shift was consistently measured for thin layers
of
adsorbed material. By measuring the PWV shift as a function of thickness
usiiig the

polymer multilayer experiment (Figure 11), we are able to characterize the
strength of
interaction of the coupled electromagnetic field as a function of distance
away from the
sensor surface. For the porous glass sensors, the interaction is particularly
strong for the
first few monolayers of adsorbed polymer, while the relationship between
polymer
thickness and PWV is highly linear for each adsorbed monolayer on the polymer
sensor
structure. The interaction between the test sample and the resonant
electromagnetic field
distribution is highly complex, as detected material can adsorb to the
horizontal and
vertical surfaces of the structure, where a characteristic field profile
extends into the
sample from each surface. Surface-based detection sensitivity is enhanced
beyond the
improvements in bulk sensitivity for the porous glass biosensor. Because the
majority of
biomolecular interactions are expected to occur within the first few
nanometers from the
sensor surface, the surface sensitivity is of greatest importance for
increasing sensitivity
in the context of surface-based biochemical assays.

27

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Title Date
Forecasted Issue Date Unavailable
(86) PCT Filing Date 2006-06-29
(87) PCT Publication Date 2007-01-08
(85) National Entry 2008-01-08
Examination Requested 2011-06-13
Dead Application 2013-07-02

Payment History

Fee Type Anniversary Year Due Date Amount Paid Paid Date
Filing $400.00 2008-01-08
Maintenance Fee - Application - New Act 2 2008-06-30 $100.00 2008-01-08
Registration of Documents $100.00 2008-08-20
Reinstatement - failure to complete $200.00 2008-08-20
The completion of the application $200.00 2008-08-20
Maintenance Fee - Application - New Act 3 2009-06-29 $100.00 2009-06-05
Maintenance Fee - Application - New Act 4 2010-06-29 $100.00 2010-06-04
Maintenance Fee - Application - New Act 5 2011-06-29 $200.00 2011-06-10
Request for Examination $800.00 2011-06-13
Current owners on record shown in alphabetical order.
Current Owners on Record
SRU BIOSYSTEMS, INC.
Past owners on record shown in alphabetical order.
Past Owners on Record
CUNNINGHAM, BRIAN T.
Past Owners that do not appear in the "Owners on Record" listing will appear in other documentation within the application.

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