WO2013171197A1 - Compact plasmon-enhanced fluorescence biosensor - Google Patents

Compact plasmon-enhanced fluorescence biosensor Download PDF

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Publication number
WO2013171197A1
WO2013171197A1 PCT/EP2013/059908 EP2013059908W WO2013171197A1 WO 2013171197 A1 WO2013171197 A1 WO 2013171197A1 EP 2013059908 W EP2013059908 W EP 2013059908W WO 2013171197 A1 WO2013171197 A1 WO 2013171197A1
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Prior art keywords
fluorescence
sensing area
chip
grating
surface plasmon
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PCT/EP2013/059908
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French (fr)
Inventor
Jakub Dostalek
Koji Toma
Wolfgang Knoll
Milan Vala
Adam PAVEL
Jiri Homola
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Ait Austrian Institute Of Technology Gmbh
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Publication of WO2013171197A1 publication Critical patent/WO2013171197A1/en

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
    • G01N21/62—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light
    • G01N21/63—Systems in which the material investigated is excited whereby it emits light or causes a change in wavelength of the incident light optically excited
    • G01N21/64—Fluorescence; Phosphorescence
    • G01N21/645—Specially adapted constructive features of fluorimeters
    • G01N21/648—Specially adapted constructive features of fluorimeters using evanescent coupling or surface plasmon coupling for the excitation of fluorescence
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
    • G01N21/75—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated
    • G01N21/77—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated by observing the effect on a chemical indicator
    • G01N21/7703—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated by observing the effect on a chemical indicator using reagent-clad optical fibres or optical waveguides
    • G01N21/774—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated by observing the effect on a chemical indicator using reagent-clad optical fibres or optical waveguides the reagent being on a grating or periodic structure
    • G01N21/7743—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated by observing the effect on a chemical indicator using reagent-clad optical fibres or optical waveguides the reagent being on a grating or periodic structure the reagent-coated grating coupling light in or out of the waveguide
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
    • G01N21/75—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated
    • G01N21/77—Systems in which material is subjected to a chemical reaction, the progress or the result of the reaction being investigated by observing the effect on a chemical indicator
    • G01N2021/7769—Measurement method of reaction-produced change in sensor
    • G01N2021/7786—Fluorescence

Definitions

  • the present invention is related to the signal amplification and enhancing the sensitivity in fluorescence-based assays by using surface plasmon optics combined with diffractive optical elements. It enables design of a compact biosensor for the detection of trace amounts of biomolecules in liquid samples.
  • Sensitive, portable and cost-effective biosensors are intensively pursued for the detection of chemical and biological analytes in important areas of point-of-care medical diagnostics, food control, security and environmental monitoring.
  • Fluorescence-based detection is one of the mostly used methods due to its intrinsic capability for sensitive detection.
  • various approaches were researched including surface plasmon-enhanced fluorescence (PEF) (J. R. Lakowicz, Plasmonics, 1 , (2006) 5-33; Stewart, et al., Chemical Reviews, 108, (2008) 494-521 ).
  • PEF surface plasmon-enhanced fluorescence
  • SPs Surface plasmons
  • SPFS surface plasmon- enhanced fluorescence spectroscopy
  • SPCE surface plasmon-coupled fluorescence emission
  • US2003043475A1 describes a compact diffraction type lens which can converge two wavelengths of light onto their corresponding optical recording media having disc thickness values different from each other.
  • US2010091370A1 describes an optical resonator structure including a substrate; a center disc formed on the substrate; a plurality of concentric grating rings surrounding the center disc, the concentric rings spaced apart from the center disc and from one another by regions of lower index of refraction material with respect thereto
  • US2012038918A1 describes an optical element which has the dual
  • the grating side may be coated with a dielectric film.
  • EP1644871 B1 discloses an optical security element having a substrate layer in which a relief structure defined by relief parameters is shaped out in a surface region.
  • US20060018021A1 discloses an optical security element including an optical grating structure which exhibits pleochroic properties when rotated or viewed from changing observation locations.
  • the optical grating structure is formed from a plurality of selectively arranged grating elements.
  • the invention describes a method and apparatus for an amplification of fluorescence signal in assays for the detection of biomolecules through simultaneous excitation and collection of emitted fluorescence light by using surface plasmons. It is based on a planar sensor chip (biochip) with diffractive optical elements (DOE) combined with Kretschmann configuration for the excitation and back-coupling of surface plasmons into the substrate. Diffractive optical elements are prepared on the biochip surface and include a concentric relief grating (CG) and linear relief grating (LG). CG is designed to function as a holographic lens for the imaging of highly directional surface plasmon-coupled fluorescence emission from fluorophore labels to a spot below the biochip where a detector is placed.
  • CG concentric relief grating
  • LG linear relief grating
  • Additional LG element is employed for the generating of surface plasmons at the excitation wavelength in order to increase the excitation rate of fluorophores attached to biomolecules that are specifically captured on the biochip.
  • the reported approach offers the advantage of increased intensity of fluorescence signal and reduced background.
  • the present invention concerns a method and implementation of combined surface plasmon-enhanced fluorescence spectroscopy (SPFS) and surface plasmon- coupled fluorescence emission (SPCE) for sensitive detection of molecular and biological analytes. It allows simplification of the biosensor design that is suitable for portable compact devices. It relies on diffractive optical elements integrated to the biochip for the coupling of an excitation beam to surface plasmons (linear grating - LG) and for the imaging of surface plasmon-coupled fluorescence emission light to a spot below the biochip where a detector is placed (concentric grating - CG), see Fig.2.
  • SPFS surface plasmon-enhanced fluorescence spectroscopy
  • SPCE surface plasmon- coupled fluorescence emission
  • the whole biochip can be fabricated by mass production-compatible technologies such as nanoimprint lithography, hot embossing or injection molding to inexpensive polymer materials.
  • CG acts as a diffractive lens.
  • the fluorescence light intensity emitted from different areas on the biochip can be projected to spatially separated spots in the image plane and independently detected with a spatially sensitive detector (e.g., charge-coupled device - CCD).
  • the dielectric substrate may be constructed of a material that is transparent to light wavelengths that are suitable for the excitation and emission of fluorescence light in visible and near infrared part of spectrum, such as polycarbonate, quartz or glass.
  • a flat substrate comprising at least one linear grating element and at least one concentric grating element.
  • a structured polymer layer on the substrate wherein said relief concentric grating has a depth of at least 10 nm and period between 0.2 and 1 micrometer.
  • a polymer layer wherein said polymer layer is structured by using interference lithography, electron beam lithography or nanoimprint lithography.
  • a sensor chip comprising a transparent glass or plastic substrate and on the outer surface a polymer layer with at least one concentric grating element and at least one sensing area is prepared.
  • a chip further comprising a metal layer on the sensing area in the middle of the concentric grating.
  • a chip wherein said metal is gold, silver, or aluminum.
  • the metal is to be selected to support surface plasmons in the visible and near infrared part of spectrum.
  • a chip wherein said metal layer on a dielectric substrate exhibits a thickness between 30 and 70 nm and support surface plasmons.
  • the thickness of the metal layer is in such a manner to support surface plasmons at the outer metal surface that are leaky into the dielectric substrate. According to yet another aspect of the invention there is provided a chip, wherein biomolecular recognition elements for selected chemical or biological compounds are attached to the metallic surface on the sensing area.
  • the biomolecular recognition element is able to capture target analyte.
  • Various types of biomolecular recognition elements are known in the art, e.g. antibodies, aptamers, peptides, molecularly imprinted polymers, etc.
  • a biomolecular element specific to the target analyte is immobilized in the sensing area on the biochip. It needs to be assured that the biological activity of the immobilized biomolecular recognition elements is conserved and the biochip surface exhibits non- fouling properties.
  • the possibility to regenerate the biomolecular recognition elements i.e., break their complex with the analyte molecules and make them available for another use) can be considered.
  • methods for the immobilization of biomolecular recognition elements on metal films exploit physicochemical interactions such as chemisorptions, covalent binding, electrostatic coupling, and high-affinity molecular linkers in multilayer systems (e.g., streptavidin-biotin, proteins A or G, and complementary oligonucleotides) and photo-immobilization (e.g., albumin conjugated with aryldiaziridines as a photo-linker).
  • n-alkylthiols or disulfides may spontaneously self-assemble on gold into well-ordered arrays.
  • biomolecular recognition elements are covalently attached by using self- assembled monolayer (SAM), in polymer brush or hydrogel film.
  • SAM self- assembled monolayer
  • SAMs have been employed in many immobilization methods for spatially controlled attachment of biomolecular recognition elements to surfaces of sensors.
  • mixed SAMs of long-chained (nD12 and higher) n-alkylthiols terminated with functional group for further attachment of biomolecular recognition elements and short-chained alkylthiols for a non-fouling background have been developed.
  • the immobilization chemistry needs to be spatially controlled.
  • Most of the current technologies of proteins arrays are based on the surfaces and formats that were earlier developed for DNA arrays. Most DNA array production techniques were developed for glass supports, but they can be tailored to noble metal surfaces with appropriate immobilization chemistries.
  • a chip wherein fluorophores are employed as labels and after their capture at the sensing area they emit light via surface plasmons that are out-coupled to the substrate by reverse Kretschmann configuration.
  • the characteristic surface plasmon-coupled fluorescence emission cone is totally internally reflected at the bottom substrate interface and hits the concentric grating element at the top biochip surface.
  • a chip wherein the concentric grating focuses the light beam emitted from the sensing area at a narrow spot below the biochip.
  • a chip wherein the concentric grating focuses the light beam emitted from different locations on the sensing area to different spots below the biochip.
  • a chip wherein the concentric grating exhibits a relief modulation with the depth between 20 and 200 nm and the relief modulation with period between 200 nm and 1 ⁇ is chirped and serves as a diffractive lens.
  • a chip wherein the concentric grating that is coated by a metal layer to increase its diffraction efficiency.
  • a chip wherein the concentric grating exhibits sinusoidal or blazed relief modulation providing high diffraction efficiency.
  • the fluorophore labels at the sensing area are organic chromophores or quantum dots emitting fluorescence light in visible or near infrared part of spectrum.
  • a chip wherein said fluorophores are exited with a light beam at lower wavelength than their emission wavelength.
  • a chip wherein excitation light is monochromatic and it is coupled to the biochip by a linear grating, propagates in the biochip substrate and hits the sensing area.
  • a chip wherein the said excitation beam hits the sensing area at the angle at which surface plasmons are resonantly excited providing enhanced field intensity at the top of the sensing area.
  • a chip wherein the linear grating exhibits a periodic relief corrugation with the modulation depth between 20 and 200 nm and period between 200 nm and 1 ⁇ .
  • a chip wherein the periodic modulation of linear grating is sinusoidal or blazed and provides high diffraction efficiency.
  • a chip wherein the linear grating is coated by additional metal layer to enhance its diffraction efficiency.
  • an apparatus for detecting fluorescence in biochemical assays by combined surface plasmon- enhanced fluorescence and surface plasmon-coupled emission methods comprising: a) a chip as described above with at least one linear, at least one concentric gratings and layer architecture supporting surface plasmons; and
  • an excitation source that emits a light beam incident at the linear grating and that is capble exciting fluorophores in the sensing area by surface plasmons; and c) a light detector arranged to selectively detect fluorescence light that is generated by excited fluorophores, emitted in form of surface plasmon-coupled emission, and imaged to a spot below the chip.
  • an apparatus comprising the chip as descripted above as an insertable element.
  • an apparatus wherein the chip is a disposable element
  • a method for measuring the concentration of an analyte in a sample by combined surface plasmon- enhanced fluorescence and surface plasmon-coupled fluorescence emission comprising the steps of: a) contacting a sensor chip as described above with a sample enabling the capture of target analyte on a surface and decorating the captured target analyte with fluorophore labels by using sandwich, displacement or competitive assays;
  • Figure 1 depicts the geometry supporting surface plasmon modes on a continuous metallic film that interact with a fluorophore (left) and a Jablonski diagram with plasmon-mediated transitions in the fluorophore (right).
  • FIG 2a shows a schematic of the biochip carrying diffractive optical elements (DOE) for the in-coupling of the excitation laser beam to the biochip and subsequent excitation of surface plasmons (LG) and for the imaging of surface plasmon-coupled fluorescence light to a detector (CG).
  • DOE diffractive optical elements
  • LG surface plasmons
  • CG detector
  • Figure 3 depicts schematics of the preparation of the concentric grating element (CG) by using sequential recording into a photoresist based on interference lithography
  • Figure 4 shows the dependence of the concentric grating element (CG) period on the distance from its center obtained from simulations (line) and measured on prepared biochips (squares).
  • Figure 5 depicts a schematic of UV imprint lithography employed for the transfer of LG and CG elements to the biochip.
  • Figure 6 depicts an optical setup for the observation of imaging properties of concentric (CG) element.
  • Figure 7 depicts an optical setup employed for a model immunoassay
  • Figure 8 shows the dependence of emission probability from a randomly oriented dipole via surface lossy waves (quenching), surface plasmons that are out- coupled into the substrate (SPCE) and optical waves propagating in free space.
  • CG concentric grating
  • Figure 1 1 shows simulations of +1 st order diffraction efficiency of a sinusoidal linear grating (LG) element for the normal incident TM polarized beam at the
  • Figure 13 shows measured kinetics of a-mlgG binding to the surface with immobilized mlgG (specific affinity partner) and rlgG (reference molecule) on the prepared biochip with CG and LG elements.
  • the inserted graph shows the magnified fluorescence intensity F(t) for the concentration from 30 pM to 1 nM
  • Figure 14 shows calibration curve of the developed biochip for the detection of a-mlgG analyte fitted with a linear function. The baseline noise and LOD are indicated.
  • Polydimethylsiloxane (PDMS) prepolymer and its curing agent were purchased from Dow Corning (SYLGARD® 184).
  • Poly(methyl methacrylate) (PMMA) was from Sigma-Aldrich bottles (Austria).
  • 1 ,1 '-dioctadecyl-3,3,3 ⁇ 3'-tetramethylindodicarbo cyanine, 4-chlorobenzenesulfonate salt (DiD) was from Invitrogen.
  • PEG3-OH PEG-thiol
  • PBS Phosphate buffered saline
  • PBS-Tween (PBS-T) buffer was prepared by adding 0.05% of Tween20 (Sigma-Aldrich, USA) to PBS buffer solution.
  • Anti-mouse mlgG (a-mlgG) and mouse mlgG (mlgG) were from Molecular Probes (USA).
  • a-mlgG molecules were labeled with Alexa Fluor 647 with the dye-to-protein molar ratio of 4.5. This dye exhibits the absorption and emission wavelengths of nm, respectively.
  • Rabbit IgG (rlgG) was from Abeam (USA).
  • 10 mM acetate buffer (ACT) with pH 5.5 was prepared in house.
  • EDC Ethyl-3-(3-dimethylaminopropyl) carbodiimide
  • NHS N- hydroxysuccinimide
  • Interference lithography was used for the preparation of master gratings.
  • Two different masters with linear sinusoidal relief grating (LG) and concentric relief chirped grating (CG) were prepared and characterized by atomic force
  • the gratings were subsequently etched with a developer AZ-303, rinsed with water and dried.
  • the master grating was casted to PDMS and transferred into a UV curable polymer followed by the gold deposition as shown in Fig. 5.
  • Gold layers with the thickness of 50 and 200 nm were deposited on the flat sensing area and on the surface of LG and CG elements, respectively, by sputtering (UNIVEX 450C, Leybold Systems, Germany).
  • Finite element method (FEM) grating solver DiPoG (Weierstrass Institute, Germany) was used for the calculation of the diffraction efficiency.
  • the chirped grating and the overall biochip geometry was designed by using a ray-tracing tool Zemax (Radiant Zemax, USA).
  • Zemax ray-tracing tool
  • home-developed scripts based on the CPS-model were used as described in our previous works.
  • CG element 20 nm thick PMMA layer doped with DiD dye was spin-coated on a gold-coated surface of a sensing area and dried overnight at the room temperature.
  • the sensing area with 50 nm thick gold layer was modified with a mixed thiol self- assembled monolayer (SAM), and IgG molecules were covalently attached to SAM carboxylic functional moieties.
  • SAM mixed thiol self- assembled monolayer
  • the biochip was firstly immersed in a mixture of PEG- thiol and COOH-thiol dissolved in ethanol (molar ratio of 9:1 and total concentration of 1 mM) overnight at room temperature. Afterwards the biochip was rinsed with ethanol and dried in a stream of nitrogen.
  • mlgG antibodies were immobilized by amine coupling.
  • Carboxylic terminal groups of COOH-thiol were activated by EDC and NHS solution (concentrations in deionized water of 75 and 21 mg/mL, respectively) for 15 min, followed by the incubation with mlgG dissolved in ACT buffer at the concentration of 50 g/mL for 90 min. Un reacted active ester groups of the COOH-thiol were passivated by 20 min incubation in ethanol amine solution. In a control experiment, rlgG with the same concentration was immobilized instead of mlgG.
  • the excitation light beam was made incident from the bottom of the biochip at LG.
  • the diffracted light beam propagated in the biochip and excited SPs at the sensing area with 50 mn thick gold film.
  • the fluorescence was excited with an excitation light beam directly incident at the biochip surface from the top through water.
  • the identical set of filters as in the measurement of spatial distribution of fluorescence light was used to surpress the background signal.
  • the coupling of a fluorophore with surface plasmons strongly depends on the distance from a metal surface d.
  • the fluorescence light emitted via SPs is transmitted through the metal film and forms a characteristic SPCE cone which is directional in polar angle 0 em and isotropic in azimuth angle ⁇ , see Fig. 8.
  • the biochip composed of a BK7 glass slide with a sensing area coated by a 50 nm thick gold film and attached biomolecular recognition elements for the specific capture of fluorophore-labeled target molecules on its top.
  • a relief concentric diffraction grating (CG) around the sensing area was used.
  • the grating was coated with 200 nm thick gold film acts as a diffraction lens which images the SPCE cone emitted from the sensing area to a spot below the biochip.
  • Sinusoidal relief corrugation was used with the average modulation depth of 120 nm which, according to finite element method (FEM) simulations, provides the maximum diffraction efficiency of 81 % (see Fig. 10).
  • FEM finite element method
  • additional linear grating (LG) was employed.
  • the sensing area with 50 nm thick gold film and aqueous sample on the top
  • the light beam When hitting the sensing area with 50 nm thick gold film and aqueous sample on the top, the light beam generates SPs at its surface as this angle coincides with SP resonance angle.
  • LG exhibited sinusoidal profile with the modulation depth around 1 10 nm. For 1 diffraction, simulations predict moderate maximum diffraction efficiency of 34 % (Fig. 1 1 ).
  • the imaging properties of CG were observed by the measurement of spatial distribution of fluorescence light intensity out-coupled from the biochip at different distances D below the biochip as seen in Fig. 6.
  • the sensing area of the chip carrying CG (see Fig. 12a) was coated with a PMMA layer loaded with a DiD dye (exhibiting similar characteristics as Alexa Fluor 647 used in further biosensor experiments) with the thickness 20 nm.
  • the sensing area was brought in contact with water and exposed to a normal incident laser beam at the wavelength ⁇ ⁇ illuminating the area of around ⁇ 1 mm 2 .
  • the intensity of fluorescence light exhibits characteristic ring distribution with a decreasing diameter when increasing the distance D.
  • the fluorescence beam is focused and reaches a minimum diameter at the distance of D-15 mm which is close to that predicted by simulations.
  • the diameter of the spot at which the fluorescence light is confined depends on the finess with which CG is prepared.
  • FWHM full width at half maximum
  • the laser beam was coupled to the biochip by the LG, propagated along in the chip substrate and excited SPs on the sensing area at the resonance angle Q ex . Only moderate enhancement of the fluorescence light collected by SPCE -2.3 was observed with respect to that measured for normal incident excitation beam from the top (data not shown).
  • mouse IgG was immobilized on the sensing area by using thiol SAM linker molecules and active ester chemistry.
  • a flow-cell was attached to the sensing area in order to flow liquid samples with the affinity binding partner anti-mouse IgG (a-mlgG) that was labeled with a dye Alexa Fluor 670.
  • the fluorescence signal F upon sequential flow of samples (PBST) spiked with increasing concentration of a- mlgG was measured in time by using the setup depicted in Fig. 7.
  • Fig. 13 shows the measured fluorescence kinetics F(t) for the concentrations of analyte ranging from 30 pM to 30 nM. It reveals that the affinity binding of the labeled analyte to the surface is associated with a gradual increase of the fluorescence signal and that the slope of the fluorescence signal dFldt linearly increases with a-mlgG concentration.
  • the control experiment in which identical samples were injected to the sensor with another antibody immobilized to the gold surface of sensing area (rabbit immunoglobulin G - rlgG) shows a negligible fluorescence signal indicating a highly specific response.
  • the calibration curve was obtained as shown in Fig. 14.
  • the fluorescence signal slope dFldt in the initial association phase was determined by linear fitting and plotted as a function of the concentration of a-mlgG.
  • the error bars represent the obtained standard deviation (SD) that is attributed to the chip to chip variability.
  • the limit of detection (LOD) of 1 1 pM was determined at the intersection where the sensor signal dF/dt matches 3-fold SD of the baseline fluorescence signal.

Abstract

The present invention related to a compact biosensor chip for fluorescence assays enabling detection of trace amounts of biomolecules. It employs surface plasmons for the enhanced excitation and efficient collecting of fluorescence light in order to amplify detected fluorescence light intensity. The integration of key optical elements onto the sensor chip by using concentric and linear gratings allows simplifying of design of overall biosensor setup.

Description

COMPACT PLASMON-ENHANCED FLUORESCENCE BIOSENSOR
FIELD OF THE INVENTION
The present invention is related to the signal amplification and enhancing the sensitivity in fluorescence-based assays by using surface plasmon optics combined with diffractive optical elements. It enables design of a compact biosensor for the detection of trace amounts of biomolecules in liquid samples.
BACKGROUND OF THE INVENTION
Sensitive, portable and cost-effective biosensors are intensively pursued for the detection of chemical and biological analytes in important areas of point-of-care medical diagnostics, food control, security and environmental monitoring.
Fluorescence-based detection is one of the mostly used methods due to its intrinsic capability for sensitive detection. In order to further advance its performance, various approaches were researched including surface plasmon-enhanced fluorescence (PEF) (J. R. Lakowicz, Plasmonics, 1 , (2006) 5-33; Stewart, et al., Chemical Reviews, 108, (2008) 494-521 ). Surface plasmons (SPs) are optical waves that originate from coupled collective oscillations of electron density at metallic surfaces. These waves exhibit strongly increased intensity and highly confined profile of electromagnetic field. The coupling of fluorophore labels with surface plasmon field offers an attractive means for the amplification of fluorescence signal-to-noise-ratio in bioassays through the combination of increased fluorophore excitation rate, decreased background signal, directional fluorescence emission and enhanced fluorophore quantum yield. Up to now, two main approaches that rely on SPs propagating along continuous metallic films were pursued and implementated to PEF biosensors. In surface plasmon- enhanced fluorescence spectroscopy (SPFS) (Dostalek and Knoll, Biointerphases, 3, (2008) 12-22; Liebermann and Knoll, Colloids and Surfaces a-Physicochemical and Engineering Aspects, 171 , (2000) 1 15-130; Herrmann, et al., US 6194223 B1 ), the binding of fluorophore-labeled molecules to biomolecular recognition elements attached on a metallic sensor surface is probed by the enhanced field intensity of SPs at the wavelength matching the absorption band of used fluorophore labels. This method takes advantage of the increased excitation rate of fluorophores that is directly translated to an enhanced fluorescence signal, see Fig. 1 . In surface plasmon-coupled fluorescence emission (SPCE) (J. R. Lakowicz, et al., Biochemical and Biophysical Research Communications, 307, (2003) 435-439), high!y directiona! fluorescence light beam emitted via SPs at the emission wavelength is detected. Bulk optical prism elements (Yuk, et al., Biosensors & Bioelectronics, 25, (2010) 1344-1349; J.R.
Lakowicz, et al., US 20050053974; Gryczynski, et al., US 20090218516 A1 ), and metallic diffraction gratings (Kitson, et al., Optics Communications, 122, (1996) 147- 154) were reported for the extracting the fluorescence light trapped by SPs to the radiation propagating to specific directions in the far field. A sensor chip with series of hemispherical dielectric elements was proposed for disposable optical biosensor chips with SPCE (Yuk, et al., Biosensors & Bioelectronics, 25, (2010) 1344-1349). Owing to the confined profile of SP field at the metallic surface, only the fluorescence signal originating from fluorophores in close proximity to the metallic surface (distance up to -100 nm) is detected by SPFS and SPCE methods, leading to a greatly suppressed background.
Various dielectric diffraction elements such as concentric gratings were used for manipulation of light. However, they were not combined with metallic structures for the amplificiation of fluorescence signal by using surface plasmons. These include the following works:
US2003043475A1 describes a compact diffraction type lens which can converge two wavelengths of light onto their corresponding optical recording media having disc thickness values different from each other.
US2010091370A1 describes an optical resonator structure including a substrate; a center disc formed on the substrate; a plurality of concentric grating rings surrounding the center disc, the concentric rings spaced apart from the center disc and from one another by regions of lower index of refraction material with respect thereto US2012038918A1 describes an optical element which has the dual
functionalities of a grating and a Fresnel lens. The grating side may be coated with a dielectric film.
EP1644871 B1 discloses an optical security element having a substrate layer in which a relief structure defined by relief parameters is shaped out in a surface region.
US20060018021A1 discloses an optical security element including an optical grating structure which exhibits pleochroic properties when rotated or viewed from changing observation locations. The optical grating structure is formed from a plurality of selectively arranged grating elements. SHORT DESCRIPTION OF THE INVENTION
The invention describes a method and apparatus for an amplification of fluorescence signal in assays for the detection of biomolecules through simultaneous excitation and collection of emitted fluorescence light by using surface plasmons. It is based on a planar sensor chip (biochip) with diffractive optical elements (DOE) combined with Kretschmann configuration for the excitation and back-coupling of surface plasmons into the substrate. Diffractive optical elements are prepared on the biochip surface and include a concentric relief grating (CG) and linear relief grating (LG). CG is designed to function as a holographic lens for the imaging of highly directional surface plasmon-coupled fluorescence emission from fluorophore labels to a spot below the biochip where a detector is placed. Additional LG element is employed for the generating of surface plasmons at the excitation wavelength in order to increase the excitation rate of fluorophores attached to biomolecules that are specifically captured on the biochip. The reported approach offers the advantage of increased intensity of fluorescence signal and reduced background. The
implementation of key optical elements on the biochip surface in form of DOEs simplifies design of the overall biosensor optical setup which does not require optical matching of a sensor chip to other bulk optical elements (such as prism or microscope lens), and the biochip can be fabricated by mass production-compatible technologies such as nanoimprint lithography.
DETAILED DESCRIPTION OF THE INVENTION
The present invention concerns a method and implementation of combined surface plasmon-enhanced fluorescence spectroscopy (SPFS) and surface plasmon- coupled fluorescence emission (SPCE) for sensitive detection of molecular and biological analytes. It allows simplification of the biosensor design that is suitable for portable compact devices. It relies on diffractive optical elements integrated to the biochip for the coupling of an excitation beam to surface plasmons (linear grating - LG) and for the imaging of surface plasmon-coupled fluorescence emission light to a spot below the biochip where a detector is placed (concentric grating - CG), see Fig.2.
These elements replace traditionally used bulk optical components such as prisms and microscope lenses and avoid the necessity of their optical matching to the sensor chip. The whole biochip can be fabricated by mass production-compatible technologies such as nanoimprint lithography, hot embossing or injection molding to inexpensive polymer materials. In addition, it enables multi-analyte detection of molecular analytes by using microarrays since CG acts as a diffractive lens. Owing to this feature, the fluorescence light intensity emitted from different areas on the biochip can be projected to spatially separated spots in the image plane and independently detected with a spatially sensitive detector (e.g., charge-coupled device - CCD).
There is thus provided according to one aspect of the invention a flat dielectric substrate carrying a relief concentric grating element.
The dielectric substrate may be constructed of a material that is transparent to light wavelengths that are suitable for the excitation and emission of fluorescence light in visible and near infrared part of spectrum, such as polycarbonate, quartz or glass.
According to another aspect of the invention there is provided a flat substrate, comprising at least one linear grating element and at least one concentric grating element.
According to yet another aspect of the invention there is provided a structured polymer layer on the substrate, wherein said relief concentric grating has a depth of at least 10 nm and period between 0.2 and 1 micrometer.
According to yet another aspect of the invention there is provided a polymer layer, wherein said polymer layer is structured by using interference lithography, electron beam lithography or nanoimprint lithography.
According to yet another aspect of the invention there is provided a sensor chip comprising a transparent glass or plastic substrate and on the outer surface a polymer layer with at least one concentric grating element and at least one sensing area is prepared.
According to yet another aspect of the invention there is provided a chip, further comprising a metal layer on the sensing area in the middle of the concentric grating.
According to yet another aspect of the invention there is provided a chip, wherein said metal is gold, silver, or aluminum.
The metal is to be selected to support surface plasmons in the visible and near infrared part of spectrum.
According to yet another aspect of the invention there is provided a chip, wherein said metal layer on a dielectric substrate exhibits a thickness between 30 and 70 nm and support surface plasmons.
The thickness of the metal layer is in such a manner to support surface plasmons at the outer metal surface that are leaky into the dielectric substrate. According to yet another aspect of the invention there is provided a chip, wherein biomolecular recognition elements for selected chemical or biological compounds are attached to the metallic surface on the sensing area.
The biomolecular recognition element is able to capture target analyte. Various types of biomolecular recognition elements are known in the art, e.g. antibodies, aptamers, peptides, molecularly imprinted polymers, etc. To enable analyte detection, a biomolecular element specific to the target analyte is immobilized in the sensing area on the biochip. It needs to be assured that the biological activity of the immobilized biomolecular recognition elements is conserved and the biochip surface exhibits non- fouling properties. The possibility to regenerate the biomolecular recognition elements (i.e., break their complex with the analyte molecules and make them available for another use) can be considered.
In general, methods for the immobilization of biomolecular recognition elements on metal films exploit physicochemical interactions such as chemisorptions, covalent binding, electrostatic coupling, and high-affinity molecular linkers in multilayer systems (e.g., streptavidin-biotin, proteins A or G, and complementary oligonucleotides) and photo-immobilization (e.g., albumin conjugated with aryldiaziridines as a photo-linker). For example, n-alkylthiols or disulfides may spontaneously self-assemble on gold into well-ordered arrays.
According to yet another aspect of the invention there is provided a chip, wherein biomolecular recognition elements are covalently attached by using self- assembled monolayer (SAM), in polymer brush or hydrogel film.
SAMs have been employed in many immobilization methods for spatially controlled attachment of biomolecular recognition elements to surfaces of sensors. To provide a desired surface concentration of biomolecular recognition elements on metal, mixed SAMs of long-chained (nD12 and higher) n-alkylthiols terminated with functional group for further attachment of biomolecular recognition elements and short-chained alkylthiols for a non-fouling background have been developed. To deliver molecular recognition elements to different sensing areas, the immobilization chemistry needs to be spatially controlled. Most of the current technologies of proteins arrays are based on the surfaces and formats that were earlier developed for DNA arrays. Most DNA array production techniques were developed for glass supports, but they can be tailored to noble metal surfaces with appropriate immobilization chemistries.
Combination of SAMs with covalent coupling of biomolecular recognition elements or non-covalent streptavidin-biotin system as a linker for attachment of biotinylated biomolecular recognition elements are most frequently used approaches to
development of protein arrays on metal.
According to yet another aspect of the invention there is provided a chip, wherein fluorophores are employed as labels and after their capture at the sensing area they emit light via surface plasmons that are out-coupled to the substrate by reverse Kretschmann configuration. The characteristic surface plasmon-coupled fluorescence emission cone is totally internally reflected at the bottom substrate interface and hits the concentric grating element at the top biochip surface.
According to yet another aspect of the invention there is provided a chip, wherein the concentric grating focuses the light beam emitted from the sensing area at a narrow spot below the biochip.
According to yet another aspect of the invention there is provided a chip, wherein the concentric grating focuses the light beam emitted from different locations on the sensing area to different spots below the biochip.
According to yet another aspect of the invention there is provided a chip, wherein the concentric grating exhibits a relief modulation with the depth between 20 and 200 nm and the relief modulation with period between 200 nm and 1 μιη is chirped and serves as a diffractive lens.
According to yet another aspect of the invention there is provided a chip, wherein the concentric grating that is coated by a metal layer to increase its diffraction efficiency.
According to yet another aspect of the invention there is provided a chip, wherein the concentric grating exhibits sinusoidal or blazed relief modulation providing high diffraction efficiency.
According to yet another aspect of the invention there is provided a chip, wherein the fluorophore labels at the sensing area are organic chromophores or quantum dots emitting fluorescence light in visible or near infrared part of spectrum.
According to yet another aspect of the invention there is provided a chip, wherein said fluorophores are exited with a light beam at lower wavelength than their emission wavelength.
According to yet another aspect of the invention there is provided a chip, wherein excitation light is monochromatic and it is coupled to the biochip by a linear grating, propagates in the biochip substrate and hits the sensing area. According to yet another aspect of the invention there is provided a chip, wherein the said excitation beam hits the sensing area at the angle at which surface plasmons are resonantly excited providing enhanced field intensity at the top of the sensing area.
According to yet another aspect of the invention there is provided a chip, wherein the excitation beam excites the fluorophore labels adhered to close proximity to the sensing area.
According to yet another aspect of the invention there is provided a chip, wherein the linear grating exhibits a periodic relief corrugation with the modulation depth between 20 and 200 nm and period between 200 nm and 1 μηι.
According to yet another aspect of the invention there is provided a chip, wherein the periodic modulation of linear grating is sinusoidal or blazed and provides high diffraction efficiency.
According to yet another aspect of the invention there is provided a chip, wherein the linear grating is coated by additional metal layer to enhance its diffraction efficiency.
According to yet another aspect of the invention there is provided an apparatus for detecting fluorescence in biochemical assays by combined surface plasmon- enhanced fluorescence and surface plasmon-coupled emission methods, comprising: a) a chip as described above with at least one linear, at least one concentric gratings and layer architecture supporting surface plasmons; and
b) an excitation source that emits a light beam incident at the linear grating and that is capble exciting fluorophores in the sensing area by surface plasmons; and c) a light detector arranged to selectively detect fluorescence light that is generated by excited fluorophores, emitted in form of surface plasmon-coupled emission, and imaged to a spot below the chip.
According to yet another aspect of the invention there is provided an apparatus, comprising the chip as descripted above as an insertable element.
According to yet another aspect of the invention there is provided an apparatus, wherein the chip is a disposable element
According to yet another aspect of the invention there is provided a method for measuring the concentration of an analyte in a sample by combined surface plasmon- enhanced fluorescence and surface plasmon-coupled fluorescence emission comprising the steps of: a) contacting a sensor chip as described above with a sample enabling the capture of target analyte on a surface and decorating the captured target analyte with fluorophore labels by using sandwich, displacement or competitive assays;
b) exposing the fluorescence label to excitation electromagnetic energy in an amount sufficient to achieve surface plasmon-enhanced excitation; and
c) measuring the emitted fluorescence light intensity in form of surface plasmon-coupled emission that is out-coupled from the biochip and imaged to a detector by using the concentric grating BRIEF DESCRIPTION OF THE DRAWINGS
Figure 1 depicts the geometry supporting surface plasmon modes on a continuous metallic film that interact with a fluorophore (left) and a Jablonski diagram with plasmon-mediated transitions in the fluorophore (right).
Figure 2a shows a schematic of the biochip carrying diffractive optical elements (DOE) for the in-coupling of the excitation laser beam to the biochip and subsequent excitation of surface plasmons (LG) and for the imaging of surface plasmon-coupled fluorescence light to a detector (CG). Figure 2b depicts the side view of the biochip and shows the propagation of the excitation and fluorescence beams.
Figure 3 depicts schematics of the preparation of the concentric grating element (CG) by using sequential recording into a photoresist based on interference lithography
Figure 4 shows the dependence of the concentric grating element (CG) period on the distance from its center obtained from simulations (line) and measured on prepared biochips (squares).
Figure 5 depicts a schematic of UV imprint lithography employed for the transfer of LG and CG elements to the biochip.
Figure 6 depicts an optical setup for the observation of imaging properties of concentric (CG) element.
Figure 7 depicts an optical setup employed for a model immunoassay
experiment based on the biochip with plasmon-enhanced fluorescence detection.
Figure 8 shows the dependence of emission probability from a randomly oriented dipole via surface lossy waves (quenching), surface plasmons that are out- coupled into the substrate (SPCE) and optical waves propagating in free space. Gold layer with the thickness of 50 nm sandwiched between a substrate (np) and water (ns) with a dipole emitting at the wavelength Aem=670 nm are assumed. Figure 9 shows simulated angular dependence of the fluorescence intensity F of SPCE that is back-coupled to the substrate from a randomly oriented dipole at the distance d=20 nm from the gold surface at the wavelength Aem=670 nm (solid line), and the electric field intensity enhancement | E/E0|2 at d=20 nm provided by the excitation of SPs at the wavelength λΘΧ=633 nm (dashed line).
Figure 10 shows simulations of -1 st order diffraction efficiency of a sinusoidal concentric grating (CG) element (at the fixed location with , 1=346 nm) depending on the grating depth for the TM polarized light beam incident at the emission angle θΘηι=71 deg and emission wavelength of Aem=670 nm.
Figure 1 1 shows simulations of +1st order diffraction efficiency of a sinusoidal linear grating (LG) element for the normal incident TM polarized beam at the
wavelength of λΘΧ=633 nm and the period of , 1=437 nm dependent on the grating depth.
Figure 12a) shows top view of the prepared biochip carrying LG and CG diffractive elements (left) and measured spatial distribution of SPCE cone that is out- coupled from the biochip by using a CG element at increasing distance below the biochip D=1 , 5, 10 and 15 mm (right), b) The cross-section of the fluorescence spot at the distance of D=15 mm from the bottom of the sensor for the angular width of individual CG segments are δ=3 and 10 deg.
Figure 13 shows measured kinetics of a-mlgG binding to the surface with immobilized mlgG (specific affinity partner) and rlgG (reference molecule) on the prepared biochip with CG and LG elements. The inserted graph shows the magnified fluorescence intensity F(t) for the concentration from 30 pM to 1 nM
Figure 14 shows calibration curve of the developed biochip for the detection of a-mlgG analyte fitted with a linear function. The baseline noise and LOD are indicated.
EXAMPLES
1 . Materials and methods
1 .1 . Materials
Polydimethylsiloxane (PDMS) prepolymer and its curing agent were purchased from Dow Corning (SYLGARD® 184). Poly(methyl methacrylate) (PMMA) was from Sigma-Aldrich Handels (Austria). 1 ,1 '-dioctadecyl-3,3,3\3'-tetramethylindodicarbo cyanine, 4-chlorobenzenesulfonate salt (DiD) was from Invitrogen. This dye exhibits the absorption and emission bands at wavelengths of Aab=644 nm and em=665 nm, respectively, and it was dispersed at the concentration of 700 nM in a toluene with PMMA (1 .4 wt.%). Dithiolalkanearomatic PEG 6-COOH (COOH-thiol) and
dithiolalkanearomatic PEG3-OH (PEG-thiol) are from SensoPath Technologies (USA). Phosphate buffered saline (PBS) with pH 7.4 was obtained from Calbiochem
(Germany). PBS-Tween (PBS-T) buffer was prepared by adding 0.05% of Tween20 (Sigma-Aldrich, USA) to PBS buffer solution. Anti-mouse mlgG (a-mlgG) and mouse mlgG (mlgG) were from Molecular Probes (USA). a-mlgG molecules were labeled with Alexa Fluor 647 with the dye-to-protein molar ratio of 4.5. This dye exhibits the absorption and emission wavelengths of
Figure imgf000011_0001
nm, respectively. Rabbit IgG (rlgG) was from Abeam (USA). 10 mM acetate buffer (ACT) with pH 5.5 was prepared in house. 1 -Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N- hydroxysuccinimide (NHS) were from Pierce (USA). Ethanolamine (Sigma-Aldrich, USA) was dissolved in water at 1 M concentration with the pH of the solution adjusted to 8.5 with sodium hydroxide.
1 .2. Biochip preparation
Interference lithography (holography) was used for the preparation of master gratings. Two different masters with linear sinusoidal relief grating (LG) and concentric relief chirped grating (CG) were prepared and characterized by atomic force
microscope. LG exhibited the period of /\=436 nm and depth of 100 nm. CG was composed of 120 or 36 segments carrying chirped gratings with the varied period between \=365 and 313 nm at the distance from the center of CG of 4.7 and 8.5 mm, respectively, and with the average depth of 1 10 nm. Fig. 3 shows the preparation of CG and LG elements by using interference lithography. Triangular sections on a photoresist-coated substrate was sequentially exposed to an interference field with an angular step 6 (5=10 or 3 deg) in order to approximate the concentric grating by a set of chirped linear gratings. The interference pattern formed at the intersection of a collimated and converging laser-beams was tuned by a cylindrical lens in order to match the desired dependence of the period that was determined from simulations
(see Fig. 4). After the exposure of the photoresist to the interference field, the gratings were subsequently etched with a developer AZ-303, rinsed with water and dried. In order to prepare multiple replicas, the master grating was casted to PDMS and transferred into a UV curable polymer followed by the gold deposition as shown in Fig. 5. Gold layers with the thickness of 50 and 200 nm were deposited on the flat sensing area and on the surface of LG and CG elements, respectively, by sputtering (UNIVEX 450C, Leybold Systems, Germany).
1 .3. Simulations
Finite element method (FEM) grating solver DiPoG (Weierstrass Institute, Germany) was used for the calculation of the diffraction efficiency. The chirped grating and the overall biochip geometry was designed by using a ray-tracing tool Zemax (Radiant Zemax, USA). For the simulation of surface plasmon-coupled emission, home-developed scripts based on the CPS-model were used as described in our previous works.
1 .4. Surface modification
For the observation of imaging properties of CG element, 20 nm thick PMMA layer doped with DiD dye was spin-coated on a gold-coated surface of a sensing area and dried overnight at the room temperature. In order to function as a biosensor, the sensing area with 50 nm thick gold layer was modified with a mixed thiol self- assembled monolayer (SAM), and IgG molecules were covalently attached to SAM carboxylic functional moieties. The biochip was firstly immersed in a mixture of PEG- thiol and COOH-thiol dissolved in ethanol (molar ratio of 9:1 and total concentration of 1 mM) overnight at room temperature. Afterwards the biochip was rinsed with ethanol and dried in a stream of nitrogen. mlgG antibodies were immobilized by amine coupling. Carboxylic terminal groups of COOH-thiol were activated by EDC and NHS solution (concentrations in deionized water of 75 and 21 mg/mL, respectively) for 15 min, followed by the incubation with mlgG dissolved in ACT buffer at the concentration of 50 g/mL for 90 min. Un reacted active ester groups of the COOH-thiol were passivated by 20 min incubation in ethanol amine solution. In a control experiment, rlgG with the same concentration was immobilized instead of mlgG.
1 .5. Optical setup for testing of biochip performance
Diffracted and out-coupled SPCE beam propagating below the biochip was made scattered at the rough surface of a diffuser which was placed at a distance D from the chip. The spatial distribution of scattered fluorescence light was imaged to an electron multiplying charge-coupled device (EM-CCD iXon+885, Andor Technology, Ireland) by a camera lens (UNIFOC 58, Schneider Kreuznach, Germany) as seen in Fig. 6. A set of filters including notch filter (XNF-632.8-25.0M, CVI Melles Griot, Germany) and band-pass filter band-pass filter (670FS10-25, LOT-Oriel, Germany) was used in order to suppress the background. For the investigation of surface plasmon-enhanced intensity of the out-coupled SPCE light beam, the excitation light beam was made incident from the bottom of the biochip at LG. The diffracted light beam propagated in the biochip and excited SPs at the sensing area with 50 mn thick gold film. As a reference, the fluorescence was excited with an excitation light beam directly incident at the biochip surface from the top through water.
In a model bioassay experiment, the SPCE signal that was focused by the CG element was collected by a microscope objective (NA=0.85, NT38-340, Edmund Optics, Germany) and its intensity F was detected by a photomultiplier tube (H6240- 01 , Hamamatsu, Japan) which was connected to a counter (53131A, Agilent, USA) (Fig.7). The identical set of filters as in the measurement of spatial distribution of fluorescence light was used to surpress the background signal.
2. RESULTS AND DISCUSSION
2.1 Surface plasmon-mediated fluorescence excitation and emission
As seen in simulations based on Chance-Prock-Silbey model as shown in Fig.
8, the coupling of a fluorophore with surface plasmons strongly depends on the distance from a metal surface d. For a gold surface in contact with an aqueous medium (ns=1 .33) and a fluorophore represented as a randomly oriented dipole, the simulations reveal that the majority of emitted light intensity at the wavelength around Aem=670 nm is collected by propagating SPs at the distance around d~20 nm. Below this distance, the emission is strongly quenched by Forster energy transfer while at higher distances d>50 nm the majority of light intensity is emitted to waves propagating in free space.
For a thin gold film attached to a dielectric substrate, the fluorescence light emitted via SPs is transmitted through the metal film and forms a characteristic SPCE cone which is directional in polar angle 0em and isotropic in azimuth angle φ, see Fig. 8. In Fig. 9, we simulated the dependence of the fluorescence intensity F on polar angle Θ assuming a randomly oriented dipole to be located on the top of 20 nm thick spacer layer {d=20 nm) on a 50 nm thick gold layer. The SPCE intensity peaks at the polar angle 0em=72 deg that is the plasmon resonance angle at the wavelength Aem=670 nm. For the identical layer structure, electric field intensity enhancement |E/E0|2 at d=20 nm was calculated upon the excitation of SPs by a light beam hitting the gold surface from the substrate under the angle of incidence θβχ at the excitation wavelength λθχ=633 nm. The results in Fig. 9 reveals that strong field intensity builds up at the angle θβχ=74 deg where the resonant condition for the excitation of SP at λΘΧ is fulfilled.
2.2 Biochip design and development
As seen in the Fig. 2, the biochip composed of a BK7 glass slide with a sensing area coated by a 50 nm thick gold film and attached biomolecular recognition elements for the specific capture of fluorophore-labeled target molecules on its top. In order to collect the fluorescence light intensity emitted in form of SPCE cone at the fluorophore emission wavelength Aem=670 nm from the sensing area, a relief concentric diffraction grating (CG) around the sensing area was used. The grating was coated with 200 nm thick gold film acts as a diffraction lens which images the SPCE cone emitted from the sensing area to a spot below the biochip. The SPCE cone propagating at polar angles between 0em=67-77 deg (see Fig. 9) in the glass substrate is totally internally reflected at the bottom glass surface, hits the surface of CG element and is diffracted to a converging wave that focuses at a desired distance D below the biochip where a detector is placed.
The dependence of the grating period A on the distance from the grating center rwas determined by simulations (shown in Fig. 4) in order to diffract the incident SPCE beam to the -1st order that propagates away from the surface through the glass substrate and converges at D=10 mm. Sinusoidal relief corrugation was used with the average modulation depth of 120 nm which, according to finite element method (FEM) simulations, provides the maximum diffraction efficiency of 81 % (see Fig. 10). In order to couple the excitation beam at the wavelength matching the absorption band of fluorophore, additional linear grating (LG) was employed. LG with the period of \=437 nm was used which allowed the coupling of normal incident monochromatic beam to a wave that propagates along in the glass substrate and totally internally reflects at bottom and top interfaces with the angle θθχ=74 deg. When hitting the sensing area with 50 nm thick gold film and aqueous sample on the top, the light beam generates SPs at its surface as this angle coincides with SP resonance angle. LG exhibited sinusoidal profile with the modulation depth around 1 10 nm. For 1 diffraction, simulations predict moderate maximum diffraction efficiency of 34 % (Fig. 1 1 ).
2.3 Imaging properties of CG and LG elements
The imaging properties of CG were observed by the measurement of spatial distribution of fluorescence light intensity out-coupled from the biochip at different distances D below the biochip as seen in Fig. 6. The sensing area of the chip carrying CG (see Fig. 12a) was coated with a PMMA layer loaded with a DiD dye (exhibiting similar characteristics as Alexa Fluor 647 used in further biosensor experiments) with the thickness 20 nm. In the used setup, the sensing area was brought in contact with water and exposed to a normal incident laser beam at the wavelength λθχ illuminating the area of around ~1 mm2. As seen in Fig. 12a, the intensity of fluorescence light exhibits characteristic ring distribution with a decreasing diameter when increasing the distance D. The fluorescence beam is focused and reaches a minimum diameter at the distance of D-15 mm which is close to that predicted by simulations. The diameter of the spot at which the fluorescence light is confined depends on the finess with which CG is prepared. As Fig. 12b shows, full width at half maximum (FWHM) of the focused beam crossection of 0.7 mm was observed for the angular width of individual CG segment of 6=3 deg which is smaller than 2.3 mm measured for the wider segments with 6=10 deg. Let us note that additional broadening of the focused fluorescence beam area is caused by finite area on the sensing area that is illuminated by the excitation beam and by chromatic abberation of the CG lens. The fluorescence intensity detected in the focal plane at the wavelength ^em=670 nm is increased when dyes dispersed on the surface are excited with the enhanced intensity of surface plasmons at the excitation wavelength λΘΧ=633 nm. In further experiments, the excitation of the fluorescence light via surface plasmons at the wavelength λθχ=633 nm was used. In this configuration, the laser beam was coupled to the biochip by the LG, propagated along in the chip substrate and excited SPs on the sensing area at the resonance angle Qex. Only moderate enhancement of the fluorescence light collected by SPCE -2.3 was observed with respect to that measured for normal incident excitation beam from the top (data not shown). This is due to the relatively low diffraction efficiency of the LG element (at least a factor of 0.34) and attenuation of the excitation beam by multiple reflections at the surface between BK7 glass substrate and 200 nm thick gold (at least a fact of 0.6) which reduced the excitation light intensity by a factor >5. This values is in agreement with the the electric field intensity enhancement |E/Eo|2=17 predicted by the simulations in Fig. 9.
2.4 Demonstration of model bioassay with SPFS/SPCE readout
In order to demonstrate the potential of developed biochip with combined SPFS and SPCE for sensitive detection of biomolecules, a model immunoassay was carried out. In this experiment, mouse IgG was immobilized on the sensing area by using thiol SAM linker molecules and active ester chemistry. A flow-cell was attached to the sensing area in order to flow liquid samples with the affinity binding partner anti-mouse IgG (a-mlgG) that was labeled with a dye Alexa Fluor 670. The fluorescence signal F upon sequential flow of samples (PBST) spiked with increasing concentration of a- mlgG was measured in time by using the setup depicted in Fig. 7. Each sample was flowed through the sensor for 10 min followed by the 10 min rinsing. Fig. 13 shows the measured fluorescence kinetics F(t) for the concentrations of analyte ranging from 30 pM to 30 nM. It reveals that the affinity binding of the labeled analyte to the surface is associated with a gradual increase of the fluorescence signal and that the slope of the fluorescence signal dFldt linearly increases with a-mlgG concentration. The control experiment in which identical samples were injected to the sensor with another antibody immobilized to the gold surface of sensing area (rabbit immunoglobulin G - rlgG) shows a negligible fluorescence signal indicating a highly specific response. From obtained kinetics measured in triplicate, the calibration curve was obtained as shown in Fig. 14. For each analyte concentration, the fluorescence signal slope dFldt in the initial association phase was determined by linear fitting and plotted as a function of the concentration of a-mlgG. The error bars represent the obtained standard deviation (SD) that is attributed to the chip to chip variability. The limit of detection (LOD) of 1 1 pM was determined at the intersection where the sensor signal dF/dt matches 3-fold SD of the baseline fluorescence signal.

Claims

CLAIMS:
1 . A sensor chip comprising a transparent glass or plastic substrate with at least one concentric grating element and at least one sensing area on its outer surface.
2. The chip according to claim 1 , further comprising a metal layer on the sensing area, optionally in the middle of the concentric grating.
3. The chip according to claim 1 or 2, wherein said metal is gold, silver, or aluminum and wherein said metal layer exhibits a thickness between 30 and 70 nm.
4. The chip according to any one of claims 1 to 3, wherein the concentric grating exhibits a relief modulation with the depth between 20-200 nm and wherein said relief modulation is optionally chirped and exhibits the period between 200 nm and 1 pm.
5. The chip according to any one of claims 1 to 4, wherein the concentric grating diffracts surface plasmon-coupled emission beam originating from fluorophores on the sensing area to wave converging at a spot below the chip.
6. The chip according to any one of claims 1 to 5 comprising at least one linear grating that exhibits the modulation depth higher than 10 nm and the period between 200 nm and 1 m.
7. The chip according to any of claims 1 to 6 in which the linear grating diffracts a monochromatic light beam propagating in a free space to a light beam that propagates in the dielectric substrate and hits the sensing area under an angle of incidence for which surface plasmons are resonantly excited in the sensing area.
8. The chip according to claim 7 in which the linear grating diffracts a monochromatic light beam at the wavelength that is in the absorption band of the fluorophores attached to the sensing area.
9. The chip according to any one of c!aims 1 to 8, wherein biomolecular recognition elements are attached to the metallic surface, optionally the recognition elements are covalently attached.
10. An apparatus for detecting fluorescence in biochemical assays employing fluorophore labels that combines surface plasmon-enhanced fluorescence and surface plasmon-coupled emission methods, comprising:
a) a chip according to any one of claims 1 to 9; and
b) an excitation source that emits a light beam incident at the linear grating and that is capable exciting fluorophores in the sensing area by surface plasmons; and
c) a light detector arranged to selectively detect fluorescence light that is
generated by excited fluorophores, emitted in form of surface plasmon- coupled emission, and imaged to a spot below the chip.
1 1 . The apparatus according to claim 10, comprising the chip as an insertable element, optionally as disposable element, and a permanent reader with a lightsource and a detector.
12. A method for measuring the concentration of an analyte in a sample by combined surface plasmon-enhanced fluorescence and surface plasmon-coupled emission methods comprising the steps of:
a) contacting a sensor chip according to any one of claims 1 to 9 with a sample comprising at least one fluorescence labeled analyte;
b) exposing the fluorescence label at the sensing area to excitation
electromagnetic energy in an amount sufficient to achieve surface plasmon- enhanced excitation; and
c) measuring the intensity of surface plasmon-coupled fluorescence emission from fluorophore labels.
13. A flat dielectric substrate with a polymer layer comprising at least one linear grating element and at least one concentric grating element.
14. The polymer !ayer according to claim 13, wherein said gratings have a depth of at least 10 nm and period between 0.2 and 1 μιη.
15. The polymer layer according to claim 13 or 14, wherein said gratings are prepared by using interference lithography, electron beam lithography or nanoimprint lithography.
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Cited By (10)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
WO2015074001A1 (en) * 2013-11-17 2015-05-21 Quantum-Si Incorporated Optical system and assay chip for probing, detecting and analyzing molecules
US20170082548A1 (en) * 2015-09-21 2017-03-23 International Business Machines Corporation Enhancing on-chip fluorescence detection
US9606058B2 (en) 2014-08-08 2017-03-28 Quantum-Si Incorporated Integrated device for temporal binning of received photons
US9678012B2 (en) 2014-08-08 2017-06-13 Quantum-Si Incorporated Integrated device with external light source for probing detecting and analyzing molecules
US9921157B2 (en) 2014-08-08 2018-03-20 Quantum-Si Incorporated Optical system and assay chip for probing, detecting and analyzing molecules
US10174363B2 (en) 2015-05-20 2019-01-08 Quantum-Si Incorporated Methods for nucleic acid sequencing
US10845308B2 (en) 2016-12-22 2020-11-24 Quantum-Si Incorporated Integrated photodetector with direct binning pixel
CN112649401A (en) * 2019-10-10 2021-04-13 采钰科技股份有限公司 Biological sensor
US11344200B2 (en) 2016-02-17 2022-05-31 Tesseract Health, Inc. Sensor and device for lifetime imaging and detection applications
US11391626B2 (en) 2018-06-22 2022-07-19 Quantum-Si Incorporated Integrated photodetector with charge storage bin of varied detection time

Citations (10)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US6194223B1 (en) 1997-04-14 2001-02-27 Roche Diagnostics Gmbh Method for the simultaneous determination of biomolecular interactions by means of plasmon resonance and fluoresence detection
US20030043475A1 (en) 2000-03-24 2003-03-06 Fuji Photo Optical Co., Ltd. Diffraction type lens and optical pickup apparatus using the same
US20050053974A1 (en) 2003-05-20 2005-03-10 University Of Maryland Apparatus and methods for surface plasmon-coupled directional emission
US20060018021A1 (en) 2004-07-26 2006-01-26 Applied Opsec, Inc. Diffraction-based optical grating structure and method of creating the same
GB2427022A (en) * 2005-06-08 2006-12-13 Toshiba Res Europ Ltd An encoded carrier for analysing a chemical species using surface plasmon resonance
US20090218516A1 (en) 2008-01-21 2009-09-03 University Of North Texas Health Science Center At Fort Worth Ratiometric Surface Plasmon Coupled Emission Detector
US20100091370A1 (en) 2008-10-10 2010-04-15 International Business Machines Corporation Radial bragg ring resonator structure with high quality factor
WO2011144652A2 (en) * 2010-05-18 2011-11-24 Delta Dansk Elektronik, Lys & Akustik Method for depositing sensor material on a substrate
US20120038918A1 (en) 2010-08-13 2012-02-16 Zhiwen Liu Compact Spectrometer Including a Diffractive Optical Element with Dual Dispersion and Focusing Functionality
EP1644871B1 (en) 2003-06-25 2012-04-11 OVD Kinegram AG Optical safety element and system for visualising hidden information

Patent Citations (10)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US6194223B1 (en) 1997-04-14 2001-02-27 Roche Diagnostics Gmbh Method for the simultaneous determination of biomolecular interactions by means of plasmon resonance and fluoresence detection
US20030043475A1 (en) 2000-03-24 2003-03-06 Fuji Photo Optical Co., Ltd. Diffraction type lens and optical pickup apparatus using the same
US20050053974A1 (en) 2003-05-20 2005-03-10 University Of Maryland Apparatus and methods for surface plasmon-coupled directional emission
EP1644871B1 (en) 2003-06-25 2012-04-11 OVD Kinegram AG Optical safety element and system for visualising hidden information
US20060018021A1 (en) 2004-07-26 2006-01-26 Applied Opsec, Inc. Diffraction-based optical grating structure and method of creating the same
GB2427022A (en) * 2005-06-08 2006-12-13 Toshiba Res Europ Ltd An encoded carrier for analysing a chemical species using surface plasmon resonance
US20090218516A1 (en) 2008-01-21 2009-09-03 University Of North Texas Health Science Center At Fort Worth Ratiometric Surface Plasmon Coupled Emission Detector
US20100091370A1 (en) 2008-10-10 2010-04-15 International Business Machines Corporation Radial bragg ring resonator structure with high quality factor
WO2011144652A2 (en) * 2010-05-18 2011-11-24 Delta Dansk Elektronik, Lys & Akustik Method for depositing sensor material on a substrate
US20120038918A1 (en) 2010-08-13 2012-02-16 Zhiwen Liu Compact Spectrometer Including a Diffractive Optical Element with Dual Dispersion and Focusing Functionality

Non-Patent Citations (11)

* Cited by examiner, † Cited by third party
Title
DARRYN W. UNFRICHT ET AL: "Grating-coupled surface plasmon resonance: A cell and protein microarray platform", PROTEOMICS, vol. 5, no. 17, 13 October 2005 (2005-10-13), pages 4432 - 4442, XP055070763, ISSN: 1615-9853, DOI: 10.1002/pmic.200401314 *
DOSTALEK; KNOLL, BIOINTERPHASES, vol. 3, 2008, pages 12 - 22
HEYKEL AOUANI ET AL: "Bright Unidirectional Fluorescence Emission of Molecules in a Nanoaperture with Plasmonic Corrugations", NANO LETTERS, vol. 11, no. 2, 9 February 2011 (2011-02-09), pages 637 - 644, XP055070590, ISSN: 1530-6984, DOI: 10.1021/nl103738d *
J. R. LAKOWICZ ET AL., BIOCHEMICAL AND BIOPHYSICAL RESEARCH COMMUNICATIONS, vol. 307, 2003, pages 435 - 439
J. R. LAKOWICZ, PLASMONICS, vol. 1, 2006, pages 5 - 33
JONG SEOL YUK ET AL: "Analysis of immunoarrays using a gold grating-based dual mode surface plasmon-coupled emission (SPCE) sensor chip", THE ANALYST, vol. 137, no. 11, 14 March 2012 (2012-03-14), pages 2574, XP055070751, ISSN: 0003-2654, DOI: 10.1039/c2an35143a *
KITSON ET AL., OPTICS COMMUNICATIONS, vol. 122, 1996, pages 147 - 154
KOJI TOMA ET AL: "Compact surface plasmon-enhanced fluorescence biochip", OPTICS EXPRESS, vol. 21, no. 8, 16 April 2013 (2013-04-16), pages 10121, XP055070474, DOI: 10.1364/OE.21.010121 *
LIEBERMANN; KNOLL, COLLOIDS AND SURFACES A-PHYSICOCHEMICAL AND ENGINEERING ASPECTS, vol. 171, 2000, pages 115 - 130
STEWART ET AL., CHEMICAL REVIEWS, vol. 108, 2008, pages 494 - 521
YUK ET AL., BIOSENSORS & BIOELECTRONICS, vol. 25, 2010, pages 1344 - 1349

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