WO2004092730A2

WO2004092730A2 - Method and device for detecting the presence of an analyte

Info

Publication number: WO2004092730A2
Application number: PCT/IB2004/001468
Authority: WO
Inventors: Tommi E. Vaskivuo; Ilkka Alasaarela; Ari KÄRKKÄINEN; Juha S. Tapanainen
Original assignee: Biogenon Ltd.
Priority date: 2003-04-15
Filing date: 2004-04-15
Publication date: 2004-10-28
Also published as: WO2004092730A3

Abstract

The present invention relates generally to device for detecting the presence and/or quantity of an analyte of interest in a sample, which device comprises an optical grating and a sensitive layer opposite to the grating and in close contact with it. More specifically, the present invention relates to biosensors.

Description

Method and device for detecting the presence of an analyte

FIELD OF INVENTION

The present invention relates generally to device and method for detecting the presence of an analyte. More specifically, the present invention relates to biosensors.

DESCRIPTION OF RELATED ART

Recently, the interest for sensors for biological applications such as diagnostics, studying protein-protein interactions or interactions of proteins or other biological molecules, such as DNA, RNA, lipids and pharmacological molecules with each other, or control of industrial processes such as fermentation or chromatography, has been increased markedly. Ideally this kind of sensor would be very sensitive, free of labeling of the analytes and it should yield results quickly. Furthermore, ideal apparatus for aforementioned applications should also have the capability to be multiplexed to achieve high throughput rate. At the same time applications for home and bedside diagnostic areas require analyzing methods that are easy to use and inexpensive to manufacture.

Analytical microarray systems are being looked to more and more as a viable means for meeting the increased requirements to increase the throughput rate and decrease the time of analysis. However, the complex chemical nature of proteins has made the development of protein microarray systems more challenging than that of gene chips because DNA is more stable and resistant to high and low temperatures.

The array must be robust but at the same time the contact with the receptor and the chip surface must not interfere with the binding of the receptor to its ligand.

A common way to recognize the binding of the analyte to its receptor is to detect a label attached to the analyte. The label may be for example fluorescent, luminescent or radioactive label or other kind of tag. Alternatively, a known amount of a molecule competing with the analyte for binding to same receptor is labeled and added to the sample. The intensity of fluorescent or other type of labels can be measured after non-specific labeled molecules have been removed by washing. (Schweitzer B, Kingsmore SF, Measuring proteins on microarrays. Curr Opin Biotechnol 2002 Feb;13(l):14-9.) The disadvantages of labeling techniques are the increased time, effort and costs that are associated with the labeling. Furthermore, the methodology is not readily multiplexed to achieve high-throughput analyzing devices.

Direct optical methods that enable label free detection of biological molecules exist, but they have limitations. Such detection methods include surface plasmon resonance (SPR) technology (US 5641640, Cullen, Brown & Lowe, Detection of immuno-complex formation via surface plasmon resonance on gold-coated diffraction gratings. Biosensors 1987-88;3(4):211-25), ellipsometry (Jin et al., Biosensor concept based imaging ellipsometry for visualization of biomolecular interactions. Analytical Biochemistry, 232:69-72.), reflectometry (Brecht & Gauglitz. Optical probes and transducers. Biosensors and Bioelectronics 10:923-936), resonant mirror (The Resonant Mirror: a novel optical biosensor for direct sensing of biomolecular interactions. Part 1 : Principle of operation and associated instrumentation. Cush R., Cronin J.M, Stewart W.J., Maule C.H., Molloy J., and Goddard N.J. Biosensors and Bioelectronics 8 347-353 (1993)), grating coupling sensors (Lukosz, W. Integrated Optical Chemical and Direct Biochemical Sensors. Sensors and Actuators B-Chemical 1995. 29:37-50) and resonance grating sensors (Cunningham et al., 'Label-free high-throughput optical technique for detecting biomolecular interactions', US Patent Application Publication US 2002/0127565 Al).

Resonant grating sensors like subwavelength structured surface (SWS) biosensors or surface-relief volume diffractive (SRVD) biosensors comprise diffractive structures produced on the substrate layer. Because depth and period of the diffractive structures are less than the resonance wavelength of light a resonant grating effect is produced on the reflected radiation spectrum when the biosensor is illuminated. Specific binding substances are immobilized on the surface of the diffractive structures. When the analyte binds the binding substance, the wavelength of the resonant grating effect changes which can be detected for example by illuminating the sensor with a white light and detecting the wavelength of the reflected beam. An example of such biosensor is described in US 2003/0027328.

To achieve good sensitivity with a resonance grating biosensor, the illumination beam has to be well collimated. When using a white light source good collimation decreases the light intensity which causes long integration times. This can be avoided by using a wavelength tunable laser, but it is expensive. Another possibility is to use a laser with a fixed wavelength and a scanning mirror. Scanning mirror is a mechanical moving part which is limited in speed, wears out and may be expensive. So SWS and SRVD biosensor is not the ideal solution for easily multiplexed sensitive biosensor.

SPR technology relates to an assay of the type wherein the presence of the analyte is detected by determining the change in the refractive index at a solid optical surface. This change is caused by the analyte involving or influencing the binding of a refractive index enhancing species to the optical surface, or release there from, respectively. SPR is observed as a dip in intensity of light reflected at a specific angle from the interface between an optically transparent material (e.g. glass) and a thin metal film, usually silver or gold. SPR depends, among the other factors, on the refractive index of the medium (e.g. a sample solution) close to the metal surface. A change of refractive index at the metal surface, such as by the adsorption or binding of material thereto, will cause a corresponding shift in the angle at which SPR occurs. To couple the light to the interface such that SPR arises, three alternative arrangements are used: a metallized diffraction grating (Wood's effect), or a metallized glass prism or a prism in optical contact with a metallized glass substrate (Kretschmann effect), or metallized waveguide (for example fiber or planar waveguide) based structures (see Jiri Homola, Sinclair S. Yee ,Gϋnter Gauglitz, "Surface plasmon resonance sensors: review" Sensors and Acuators B 54 (1999) 3- 15). US 5313264 describes an optical biosensor system employing the principle of SPR.

SPR assays have certain fundamental limitations that restrict the technical performance thereof. One major limiting factor is the sensitivity, or the signal strength. The SPR response depends on the volume and refractive index of the bound analyte, which volume is limited by mass transfer, reaction kinetic and equilibrium parameters. Since the SPR-measurement response is proportional to the change in refractive index caused, when e.g. protein molecules are adsorbed to the surface and displace water there from, the refractive index difference between the protein and the buffer solution puts a theoretical limit to the strength of the response that may be obtained. SPR-based immunoassays for substances of low molecular weight or substances occurring at low concentrations are problematic due to the very small changes in refractive index caused when the analyte binds to or dissociates from the antibody-coated sensing surface.

One disadvantage of SPR technology is that multiplexing is complicated and expensive and no commercially available high-throughput device based on SPR exists. However, arrays for rapid and simultaneous detection of several analytes i.e. multiplexed sensor setups are needed for example in diagnostic applications. The capability to measure multiple different diagnostically relevant values from a clinical sample would have obvious benefits. Such sensor would certainly prove valuable also in research use. Miniaturized DNA arrays have been developed (for example, see US 5,412,087, 5,445,932 and 5,744,305) and are available commercially. Also multiplexed protein microarrays are being developed (for example, see US 6,475,809, and US 6,365,418). However, the detection of the biological molecules (such as DNA or proteins) is based on labeling the analytes with fluorescent or other labels. Although the miniaturization of said assays increases their throughput rate, the requirement for labeling still set limitations for their use.

Up to date, the available methods for biomolecular recognition have not produced commercially available platforms for high- throughput analysis. Therefore, there is need for new sensitive and low-cost detection technologies.

An object of the present invention is to provide a novel label free analyzing device and method for detecting the presence and/or quantity of an analyte of interest in a sample. A further object of the present invention is to provide a sensor device, which can be easily multiplexed to create sensor arrays for simultaneous detection of several analytes. This sensor device preferably utilizes the above-mentioned label- free analyzing method.

A further object of the present invention is to provide new materials for use in such label-free analyzing devices.

A further object of the present invention is to provide a sensor device, which can be used in several different types of assays, such as detecting the presence and/or quantity of different types of biological or pharmaceutical molecules. The detected molecules can be e.g. proteins, DNA or RNA molecules, pharmacological molecules, lipids, carbohydrates, organic molecules or inorganic molecules.

A further object of the present invention is to provide a gentle method for coupling the receptor molecules to the materials used in the sensor device. The method used can be done at relatively low temperatures and in neutral pH conditions. The mild processing conditions ensure the functional preservations of biological material attached to the sensor, allowing flexibility in the device construction. The biological receptors can be attached to the sensor material either before of after the final patterning of the sensor surface. SUMMARY OF THE INVENTION

It is well known from the prior art, that the interactions of an analyte and its receptor at a surface layer well represent the actual concentration of the analyte in the sample measured. The present invention provides means for detecting these kinds of interactions taking place close to an optical grating.

One embodiment of the present invention provides an optical grating; a sensitive layer opposite to the grating and in close contact with it, which layer is substantially porous, has substantially different refractive index than that of the grating, and is capable of binding specifically the analyte of interest throughout the layer structure and the binding of the analyte of interest to the sensitive layer causes a measurable change in the optical properties of the sensitive layer. This way the whole section of the sensitive layer can be exploited and the thickness of the layer can be substantially greater than what is generally known in the art. These changes in the optical properties are proportional to the amount of analyte bound to the sensitive layer. The binding of an analyte to the sensitive layer alters the refractive index of this layer. The alteration of the refractive index of the sensitive layer causes a change in the properties of the grating. The change in the grating properties can be seen for example in the behavior of a light beam that passes through such a grating. This change can be easily detected e.g. by measuring the change of the ratio of the light intensity of first and zeroeth order light beams that are transmitted through the grating.

In another embodiment of the invention the detection means for measuring the change in the optical properties are provided. Such means generally comprise at least a light source and a detector for measuring the transmitted or reflected light.

In another embodiment of the invention the sensitive material layer comprises a porous material capable of binding, transferring or interacting by other means with the receptor and analyte molecules. Porous material can be formed, preferably by using a hybrid aluminum oxide siloxane or hybrid tetraethoxysilane materials. In another embodiment of the invention the sensitive layer contains receptor molecules capable of binding the analyte of interest. These receptor molecules are either incorporated in the sensitive layer or they are located on top of the sensitive layer. In case the sensitive layer is porous material, the receptors may be incorporated throughout to the whole layer and are accessible for the analyte molecules at the whole layer area because of the porous structure of the sensitive layer. The area where the receptors are located may depend on the technique used to incorporate the receptor molecules. In another embodiment of the invention one or more defined areas comprising the grating and the sensitive layer are arranged on a solid surface in a form of a microarray plate or biochip. In such biochip there can be thousands of such small areas each capable of binding the analyte, preferably each one capable of binding a different kind of analyte or analytes. The optical properties of each one of these areas can be detected separately by using a suitable light and detection means. Depending on the technique used, the detection of the optical properties of each area can be done simultaneously or separately. The simultaneous detection can be done e.g. using several light sources or by splitting the light using an optical arrangement. This way a large amount of information from a sample or from multiple samples can be obtained in relatively short period of time. A detection system for the biosensor may comprise, for example, a light source that illuminates a small spot of the sensor surface, and a light detector that collects the transmitted or reflected light. These kinds of detection systems are generally well known in the art.

BRIEF DESCRIPTION OF THE DRAWINGS

Fig. 1 shows a schematic diagram of one embodiment of an optical grating

Fig. 2 shows another embodiment of a grating structure, where the period (d) of the grating can be 6 μm and height (h^ 3 μm. The refractive indexes («y and n₂) are different and can be for example 1.514 («_;) and 1.470 (n₂).

Figs. 3 A, 3B and 3C show schematic diagrams of biosensors utilizing different grating profiles: binary grating (3A), sinusoidal grating (3B), and blazed grating (3C).

Fig. 4 shows a top view of a biosensor that has an array of patches reactive towards specific analytes arranged in square formation.

Fig. 5 shows a top view of a biosensor that has patches reactive towards specific analytes arranged in circular formation.

Fig. 6 shows binding of an analyte to protein receptor in a single pore in the sensitive layer Fig. 7 shows binding of an analyte to DNA receptor in a single pore in the sensitive layer

Fig. 8 shows a method for amplifying the effect the analyte binding by using a streptavidin-coated nanoparticle

Fig. 9 shows a method that can be used to immobilize a biotinylated receptor to the streptavidin coated sensitive layer

Fig. 10 shows a method that can be used to measure enzyme activity with a biosensor.

Fig. 11 shows examples of the surface chemistry that can be used to immobilize the receptors to the sensor.

Fig. 12 shows an example of a biosensor in a tip of an optical fiber.

Fig. 13 shows an example of multiple biosensors coupled to a waveguide.

Fig. 14 shows a scanning electron microscope photograph of a grating structure that can be used in a biosensor.

Fig. 15 shows a scanning electron microscope photograph of a biosensor after a sensitive layer has been spinned on top of the grating.

Fig. 16 shows an photograph of a biochip which contains biosensors that each have a sensitive area of 3 mm x 3mm.

Fig. 17 shows measurements of transmitted zeroeth-order/first-order light intensity ratios after applying 10 μl of leptin hormone (concentration 1 μg/1) over a biosensor capable of specifically detecting leptin.

Fig. 18 shows a setup that can be used to measure the effects that take place in the sensitive layer of the grating structure. A modulated light source (e.g. laser) can be used to illuminate the grating. Zeroeth and first order light beams are detected using two different light detectors. Fig. 19 shows an alternative way to measure multiple orders of light beams with one light detector.

Figs. 20 A and 20 B show how zeroeth order light beam can be measured using a lens and a second detector (20 A). The modulation of the light beam can be also made after the grating using a spatial modulator such as chopper (20 B). Herein mentioned modulator can be constructed in a way that only zeroeth order light can pass through to the grating at a time and higher orders of light (but not zeroeth order) at another time.

Fig. 21 shows an example how a CCD row camera can be used to detect multiple orders simultaneously.

Fig. 22 shows an example using a LED or a bulb as a light source providing light to the transmitting optics.

Fig. 23 shows an example of how integrated sphere can be used with LED or bulb sources.

Fig. 24 shows how an integrated sphere can be connected to an optical fiber.

Fig. 25 shows a fiber optical configuration, which uses a fiber pigtailed laser as a source.

Fig. 26 shows an example how light source and transmitting optics can be integrated to a substrate.

Fig. 27 shows an example how an additional grating can be used to divide the zeroeth order beam for further analysis when a wideband source is used.

Fig. 28 shows how prism-grating-prism component can be used to spread the zeroeth order beam into spectrum.

Fig. 29 shows an example of a light source with a scanning filter.

Fig. 30 shows how two or more different light sources may be used with an integrated sphere. Fig. 31 shows an example how electrically modulated LED's can be used as a light source.

Fig. 32 shows how wide band light source can be used with conventional silicon photodiodes.

Fig. 33 illustrates how a biochip can be simultaneously measured using different income angles.

Fig. 34 shows the simulated diffraction efficiencies for zeroeth (TO) and first (TI) transmission orders for TE and TM polarized light as a function of grating modulation depth when grating layer refractive index is 1.600 and sensitive layer refractive index 1.514.

Fig. 35 shows the simulated diffraction efficiencies as function of the change of the refractive index (nj) of the sensitive layer when grating modulation depth is 2.5 μm.

Fig. 36 shows the change of refractive index (nj) as a function of the ratio of the intensities of the first and zeroeth diffraction orders.

Fig. 37 illustrates a measurement setup that can be used with a biochip.

Fig. 38 shows an example how a biochip can be used to detect multiple analytes with multiple biosensors from one sample.

Fig. 39 shows an example how a biochip can be used to detect an analyte from multiple samples.

Fig. 40 illustrates how microfluidistic channels can be used to deliver a sample to biosensor in a biochip.

DETAILED DESCRIPTION OF THE INVENTION

1) DEFINITIONS For a better understanding the abbreviations and concepts used hereinbefore and hereinafter are the following:

The word 'receptor' used herein denotes any molecule capable of specifically binding the analyte of interest. The receptor molecules can be e.g. biological molecules, such as proteins, peptides, polyclonal or monoclonal antibodies, single chain antibodies (scFv), antibody binding fragments (Fab), antigens, DNA or RNA molecules (nucleic acids), ligands, lipids or carbohydrate molecules or like, pharmacological molecules, small organic molecules, or other organic or inorganic molecules. The receptor molecules can also be cell organelles, viruses, bacteria or cell either in part or in whole or other biological samples. The receptor is attached onto the sensitive layer or it is incorporated into the layer by physical adsorption or by chemical binding. The word 'receptor' can also denote any molecule that is a target or a substrate for an enzyme or other molecule. Any enzymes or other molecules can modify 'receptors' that come in contact with it.

In one embodiment of the invention, the receptors can be molecules that are cleavable targets for one or more enzymes and a biosensor can be used to measure enzyme activities. If these enzymes are present in the analyte that is measured with the biosensors, the enzymes cleave their target molecules that are attached to the sensitive surface.

The word 'analyte' used herein denotes the molecule or molecule species in the sample, which is to be measured. The analyte can comprise any molecule or macromolecule, e.g. a protein, peptide, antibody, nucleic acid, cellular organelle, virus, bacteria or cell. The analyte can also be, or it is originated from, a biological sample e.g. blood, plasma, serum, gastrointestinal secretions, homogenates of tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen, prostatic fluid, tears or lymphatic fluid. The analyte or the sample possibly comprising the analyte is to be applied onto the sensor device. The sample may or may not contain the analyte of interest.

The word 'grating' used herein denotes an one-dimensional or a two- dimensional optical diffraction grating.

The expression 'diffraction efficiency' (D.E.) used herein denotes the intensity of a diffracted beam divided by the sum of all the diffracted beams. Thus diffraction efficiency is always a number between 0 and 1. The word 'substrate' used herein denotes a base layer made of glass, silicon or other material onto which other layers, such as grating and the sensitive layer, are deposited.

The word 'optical waveguide' used herein denotes an optical conductor that provides a path to guide light, such as optical fiber or a planar waveguide.

The word 'SPR' used herein denotes the surface plasmon resonance method, a commercially available label-free method to detect binding of biological molecules to the receptors attached to the surface of the sensor. Characteristic for SPR is an optical surface, which comprises a thin metal film close to the optically transparent material.

The word 'microarray' used herein denotes a detection platform where one or more sensor units comprising a grating and a sensitive layer capable of detecting different analytes have been miniaturized into a compact device. The word 'biochip' generally has the same meaning. The word 'biosensor' used herein denotes a single unit comprising a grating and the sensitive layer opposite the grating and in close contact with it. Preferably receptor molecules are incorporated into the sensitive layer or they are attached onto the layer. A biosensor may contain further components, such as substrate, a cover layer, or a buffer. Light source can be located on either side of the biosensor. The word 'biochip' used herein denotes a sensor chip or platform which comprises one or more biosensors i.e. multiplexed biosensor. The word 'microarray' generally used in the art means the same and the terms can be used interchangeably. Although the expressions 'biosensor' and 'biochip' used herein refer to biological applications, they should not be considered to limit the scope of the invention. As described in this application, the invention can also be adapted to other kinds of applications, such as detection of organic or inorganic molecules.

The word 'light source' used herein denotes an apparatus that provides light to the biosensor. The light source can produce light with only single wavelength, such as laser, or the light source can produce light with multiple wavelengths, such as light emitting diode (LED) or a light incandescent bulb.

The word 'spatial modulator' used herein denotes a light modulator for example a liquid chrystal device (LCD), a digital micro-mirror modulator (DMD), an optical chopper, an acousto-optical modulator (AOM), an electro-optical modulator (EOM) or other alike. The word 'filter' used herein denotes the optical transmission filters which are used to filter out some wavelength regions of the transmitted light beam. A special kind of filter is Fabry-Perot interference filter (FPI-filter) which for example consist of a thin glass plate coated with partially transmissive coatings on the both side. The transmission spectrum of a FPI-filter has the shape of the Airy-function, i.e. it contains a series of sharp transmission peaks.

The expression 'numerical aperture' (N.A.) used herein denotes the divergence angle of the light beam. It is calculated by

NA = n sin θ,

where θ is the half of the opening angle of the beam and n is the refractive index of the propagation material (B. E. A. Saleh and M. C. Teich, Fundamentals of Photonics, John Wiley & Sons Inc, ISBN 0-471-83965-5, U.S.A. 1991).

The word 'doping' used herein denotes incorporation of receptors into the porous layer of the biosensor or on top of it either during synthesis of the material or after the material has been deposited to the surface of a grating.

2) MODES FORCARRYING OUT THE INVENTION

The present invention provides a sensor device and a method for label-free detection of various biological molecules. This sensor can be multiplexed to a microarray biochip for simultaneous detection of one or more different molecules.

In one embodiment of the invention a diffractive element, typically a grating comprising a set of equally spaced lines, like for example a binary or sinusoidal phase grating, is used to divide a light beam into several diffracted light beams. The grating is covered with a porous sensitive layer. The diffraction efficiencies of the reflected or transmitted diffraction orders can be modulated by the addition of molecules such as specific receptors of binding partners or both to the sensitive layer. The added molecules change the refractive index of the sensitive material which modifies the diffraction efficiencies of the grating. One embodiment of the invention provides a biosensor. A schematic diagram of an example of a biosensor structure is shown in Figure 1. The shown biosensor comprises a grating (103), a substrate layer (105) that supports the grating and a sensitive layer (102) in close contact with the grating but opposite side than the substrate. In addition to these, a biosensor can comprise a buffer layer (104) between the grating and the substrate, and/or a cover layer (101) in close contact with the sensitive layer but opposite side than the grating. More detailed diagram of the grating layer and the sensitive layer with used optical and geometrical parameters is shown in Figure 2.

A substrate can comprise, for example, glass, silicon, epoxy, plastics or other suitable material. Optionally, a substrate and a grating comprise a single unit in which the and the substrate are formed of the same material. In addition to that the shape of the surface of the substrate is plane, it can also be concave. In that case the grating is called concave grating.

Variety of techniques and materials can be used to fabricate the grating structures. Among these methods are: (1) wet and dry etch transfer technologies including patterning of the masking photoresist films, (2) replication technologies such as injection molding, hot embossing and UV molding including various techniques for mold fabrication, (3) non-lithographic technologies such as ink-jet printing and reel-to-reel printing, and (4) photolithographic patterning of a negative or positive lithography-tone photoimageable material with certain optical properties to a desired shape. The above described methods are illustrative only. Alternative fabrication methods can be used by those skilled in the art.

The grating can be manufactured from glass, plastic such as epoxy, acrylic, polystyrene, or sol-gel materials such tetraehylorthosilicate glass, hybrid organic- inorganic glass, silsesquioxane, organo silsesquioxanes, semiconductors such as silicon, doped silicon, GaAs, or metal. Some commercial materials that can be used to manufacture gratings include includes SU-8 (Microchemistry Inc.) and Cyclotene (Dow Chemical Inc.) Examples for fabrication of two-dimensional gratings are found in Wang. Opt. Soc. Am. 1990 8:1529-1544 and they are well known in the art. The gratings can be made out of epoxy or plastic by embossing which is a well-known manufacturing method.

A cross-sectional profile of the grating can comprise any periodically repeating shape, for example, binary (Figure 3A), sinusoidal (Figure 3B) or blazed (Figure 3C). Non-binary gratings, for example sinusoidal or blazed gratings can be produced using embossing or exposing photolithography sensitive material through gray scale masks. The profile depth (h₂) of the grating can vary from 100 nm to 100 μm, preferably from 500 nm 10 μm. The period (d) of the grating can vary from 1 μm to 100 μm. Both profile depth and period can be constant or they can vary over the grating area.

A grating can comprise of a repeating pattern of shapes including lines, squares, triangles, circles, ellipses, trapezoids, sinusoidal waves, ovals, rectangles and hexagons. For example a grating can comprise of a classical set of equal spaced lines. The repeating shape and the repeating period can be constant or they vary over the grating area. The above described grating structures are illustrative only. Alternative structures can be designed and used by those skilled in the art.

A scanning electron microscope picture of a grating that can be used in a biosensor is shown in Figure 14, where grating has been constructed on top of a glass substrate.

The grating can be coated with porous sensitive layer where biological receptors molecules can be attached. The pore volume of the sensitive material can be from 20 to 65 percent by weight, preferably 30 to 55 percent by weight. The pore- size of the material can vary in different applications of the biosensor depending on the analyte to be measured. Generally the pore size is at the area of 30-250 nm, preferably at the area of 40-200 nm, more preferably at the area of 50-150 nm. In one embodiment of the invention the grating is covered with porous hybrid aluminum oxide siloxane or hybrid organo modified tetraethoxysilane materials that have a water contact angle of 50° to 49° or 42° to 40° respectively. Biological receptor molecules can be incorporated into the porous layer. Said receptors are specific for a desired analyte. Such receptors can be, for example, a protein, polypeptide, polyclonal or monoclonal antibody, single chain antibody (scFv), a fragment of an antibody (Fab), antigen, small organic molecule, nucleic acid, steroid hormone, pharmacological molecule, lipid, cDNA probe, virus, viral capsule in part or in whole, bacteria, bacterial capsule or surface antigen in part or in whole, cell or biological sample. A biological sample can be for example blood, plasma, serum, gastrointestinal secretions, tissue homogenates, tumor homogenates, synovial fluid, feces, saliva sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, or prostatic fluid. In one embodiment of the invention, the receptor molecules are incorporated into the porous material. The biological molecules can be mixed into the material solution when it is in liquid form. After the biological molecules have been inserted into the material the processing of the material will be performed in substantially neutral pH (e.g. pH 6.5 to pH 7.5) conditions and in low temperature (less than 50° C) to preserve functionality of the biological binding molecules.

In another embodiment of the invention, the receptor molecules are bound to the porous sensitive layer after it has been applied on top of the grating. There are several methods to immobilize the receptors to a surface, well known in the prior art. The immobilization of the molecules can be carried out through covalent or non- covalent interactions with the host material. Some examples are shown in Figure 11. Amine activation, epoxy activation, carboxyl activation, hydroxyl activation, aldehyde activation, nickel activation, hydrazide coupling coupling of hydrophogic groups, the use of self-assembled monolayers, alkylsiloxane monolayers on hydroxylated surfaces, alkyl-thiol/dialkyldisulfide monolayers on noble metals, alkyl monolayer formation on oxide-free passivated silicon and avidin- or streptavidin- biotin coupling (Current Protocols in Protein Science, 2001; Khilko et al., 1995; Stein and Gerisch, 1996 ; O'Shannessay et al., 1992); Sigal et al., 1996; U.S. Pat. No. 5,405,766; PCT Publication WO 96/38762; U.S. Pat. No. 5,412,087; U.S. Pat. No. 5,688,642; U.S. Pat. No. 4,690,715; U.S. Pat. No. 5,620,850; Wagner et al., 1996; Linford et al. 1995; Wagner et al. 1997, U.S. Pat. No. 5,429,708) are examples of chemical binding methods that can be applied to the biosensor. Generally the immobilization of receptors can be performed similarly as immobilization of these molecules to glass.

Furthermore, different adaptations of the porous material can be used in the binding process. The porous material can be tailored to include number of different functional groups including but not limited to amine, aldehyde, carboxyl, hydroxyl, epoxy groups.

In one embodiment the sensitive layer comprises material, which can be obtained by hydrolyzing a first silane having the general formula I

X _4-a-SιR ι,R _CR d I

with a second component having the general formula II

X _-eMR _fR _gR j, II

Optionally together with a third compound having the general formula III

X³ _4-i-M'R⁷ _jR⁸ _kR⁹, III

In the above formulas, the groups "X" (i.e. the reactive groups X¹, X² and X³) are groups, which are cleaved off by the hydrolysis or condensation reaction. They are independently selected from hydroxyl, alkoxy, acyloxy, nitrade, carboxyl, acetylacetonate and halogen. It is possible to use silanes, metals or metalloids wherein the X , X and X are different or identical. By using different leaving groups, certain important advantages can be obtained, as will be explained below. In the preferred hydrolysable groups, X , X and X stand for halogen, preferably chlorine or bromine, or an alkoxy group, such as methoxy, ethoxy or propoxy. If X¹, X² and X³ groups are condensable it is preferred that they are hydroxyl groups. M and M' stand for metal or metalloid groups.

In order to provide an organically crosslinked material, there are reactive unsaturated or oxane groups present in at least one of the silane reactants. There can be such reactive groups present in two or more of the reactant groups. Thus, silanes of formulas I and III can contain unsaturated groups bonded to the silicon atom in addition to the aryl or alkyl groups, respectively, also present therein. The unsaturated groups contain double- or triple bonds (-C=C- or -C≡C-) or oxanes. Such groups are represented by alkenyl, alkynyl, epoxy groups.

In particular, alkenyl groups are preferred because they provide high reactivity combined with reasonable stability. The "alkenyl" has preferably the following meanings in the definitions of substituents R¹ to R³, R⁵, R⁶, R⁸ and R⁹: linear or branched alkenyl group containing 2 to 18, preferably 2 to 14, and in particular 2 to 12 carbon atoms, the ethylenic double bond being located at the position 2 or higher, the branched alkenyl containing a CI to C6 alkyl, alkenyl or alkynyl group, which optionally is per-fluorinated or partially fluorinated, at alpha or beta positions of the hydrocarbon chain. Particularly preferred alkenyl groups are vinyl and allyl.

Substituents R² to R⁶, R⁸ and R⁹ can stand for aryl, which means for a mono-, bi-, or multicyclic aromatic carbocyclic group, which optionally is substituted with C] to C₆ alkyl groups or halogens. The aryl group is preferably phenyi, which optionally bears 1 to 5 substituents selected from halogen alkyl or alkenyl on the ring, or naphthyl, which optionally bear 1 to 11 substituents selected from halogen alkyl or alkenyl on the ring structure, the substituents being optionally fluorinated (including per-fluorinated or partially fluorinated)

Substituents R , R , R to R stand for hydrogen, an alkyl group, including linear or branched alkyl groups containing 1 to 18, preferably 1 to 14, and in particular 1 to 12 carbon atoms, the branched alkyl containing a to C₆ alkyl, alkenyl or alkynyl group, which optionally is per-fluorinated.

In particular, the alkyl group is a lower alkyl containing 1 to 6 carbon atoms, which optionally bears 1 to 3 substituents selected from methyl and halogen. Methyl, ethyl, n-propyl, i-propyl, n-butyl, i-butyl and t-butyl are particularly preferred.

Generally the materials of present invention are produced by the steps of a) hydrolyzing and/or condensing the above mentioned silanes and other precursors to produce a hybrid siloxane material; b) doping the material with biological agent; c) depositing the material in the form of a thin layer; and d) curing the thin layer to form a film or layer. Or alternatively,

The present materials are produced by the steps of

e) hydrolyzing and/or condensing the above mentioned silanes and/or other precursors to produce a siloxane material; f) depositing the material in the form of a thin layer; curing the thin layer to form a porous film or layer.; and g) doping the material with biological agent.

Typically, the method comprises hydrolyzing the first, second and optionally third compounds in a liquid medium formed by a first solvent to form a hydrolyzed product comprising a hybrid siloxane material; depositing the hydrolyzed, condensed or partially condensed product on the substrate as a thin layer; and curing the thin layer to form a thin film having a thickness of 0.01 to 1000 μm. The various solvents, which can be used in the methods according to the invention, are water and various organic solvents. However, preferably the solvent is water when the present materials are produced through steps a) to d) as described above.

The hydrolyzed product comprising a siloxane material can be recovered and mixed with a second solvent to form a solution, which is applied on a substrate. The second solvent is removed to deposit the hydrolyzed product on the substrate as a thin layer, and then the thin layer to form a thin film having a thickness of 0.01 to 1000 μm.

The above hydrolysis steps of the first, second and third silicon compounds to form a hydrolyzed product and the step of curing the hydrolyzed product are all performed at a temperature of 0 to 500 °C, preferably less than 60 °C.

The hydrolysable group can be alkoxy, halogen, acyloxy, deuteroxyl, carboxyl, nitride or amine. The condensable groups can be for example, hydroxyl, alkoxy or halogen. The hybrid siloxanes are formed by hydrolyzing and condensating metal or metalloid compounds that contain one or more reacting group so that final material contains at least the Si-O-Si group. A scanning electron microscope picture of porous hybrid siloxane composite material that forms the sensitive layer on top of a grating structure is shown in Figure 15.

One embodiment of the invention provides a biochip. A biochip comprises one or more biosensors produced to the same substrate layer. The physical size of a biochip can vary e.g. from 1 mm x 1 mm to 50 cm x 50 cm, preferably from 5 mm x

5 mm to 100 mm x 100 mm. A biochip can have arbitrary shape. It can be for example square, rectangular, triangle, hexagonal or circle. The biosensors on a biochip can be different or identical in their size and structure. The biosensors can be miniaturized according to the rule that the smallest diameter of a grating is limited to be at least few times the period of the grating. So, a biosensor can be about 25 μm to about 1 cm in diameter, preferably from 100 μm to 3 mm. A biochip can comprise e.g. 1, 10, 100, 1000, 10000, 100000 or more than 1000000 biosensors, typically from 1 to 1000 biosensors. Such a biochip is called a microarray because one or more biosensors are typically laid out in a regular grid pattern. However, a microarray of the invention can comprise one or more biosensors laid out in any type of regular or irregular pattern. Figure 4 presents a schematic diagram of an example of a biochip, which comprises of 36 1 mm x 1 mm biosensors on a 10 mm x 10 mm substrate.

Figure 5 presents a schematic diagram of an example of a biochip, where biosensors have been arranged in circular pattern.

A photograph of a biochip that comprises multiple biosensors is shown in Figure 16. The biochip is constructed on glass substrate.

Each biosensor on a same biochip can be sensitive to different analyte by choosing the receptor that is present in the sensitive layer of the sensor, i.e. each biosensor on a same substrate can simultaneously measure different molecules.

In one embodiment of the invention, one or more biosensors on a biochip can be used as a reference by not including or immobilizing specific receptors in the reference sensors sensitive layer, that are included in the sensitive layer of another otherwise identical biosensor. In one embodiment of the invention, reference sensors can be used to detect unspecific binding in a sample. In one embodiment of the invention, the measurements from reference channel can be compared to measurements from another, otherwise identical biosensor.

Optionally, all biosensors on a biochip can comprise a single grating, i.e., a biochip can consist only one grating, which contains many distinct locations, each with a different receptor or with a different amount of a specific receptor.

One example of a biochip of the invention is a circular biochip on which all biosensors are arranged in circles with certain radii. The circular biochip can be rotated which is especially good for rapid measurement of many biosensors.

In one embodiment of the invention, a biosensor can be attached to a tip of an optical fiber and the optical fiber can act as a light source as shown in Figure 12).

In one embodiment of the invention, multiple biosensors can be coupled to a single waveguide, which acts as a light source to all of the biosensors attached to the guide as shown in Figure 13).

In yet another embodiment of the invention, the waveguide can be constructed directly on top of the waveguide (Figure 13). A guide layer (1301) is applied on top of a substrate (1302), The guide layer is exposed to UV light through a waveguide mask (1303) and chemically developed. Thereafter, a planarization layer (1304) is applied on top of the waveguide. A grating is developed into the planarization layer, directly on top of the waveguide, by exposing the layer to UV light through a grating mask (1305). Sensitive layer is spinned on top of the grating structure to complete the biosensor. The sensitive material can include the biological receptors, or they can be applied after the device has been completely processed. An additional array layer

(1306) can be constructed on top of the planarization layer to direct the liquid flow and handling. When a biological sample is analyzed, the biosensor is put in contact with the sample. For example, the sample can be applied onto the sensor or the sensor can be placed into the sample. The receptor molecules that are bound to the sensitive layer interact with the analytes that might be present in the sample.

In one embodiment of the invention, the analytes specifically bind to the receptors that are attached to the sensitive layer. Figure 6 shows an example of a situation where a sample has been applied on tope of the sensitive layer (601) and analytes (603) that are present in the sample have bound to antibodies (604) in the sensitive layer.

In one embodiment of the invention, specific DNA or RNA sequences or molecules (703) can be identified from a sample, when DNA probes with complementary nucleotide sequences (704) act as a receptor in the sensitive layer as shown in Figure 7. When a molecule binds to the receptors present in the sensitive layer, the binding will have an effect on the optical properties of the said layer, and therefore the a biosensor can be used to detect wide variety of different biological, pharmaceutical, organic and inorganic molecules, given that a proper receptor is for the desired analyte can be used.

In one embodiment of the invention, the effect of the analyte binding can be amplified by further attaching more molecules to the complex that has been bound to the specific receptors in the sensitive layer. Figure 8 shows how the measurement of biotinylated analyte (804) can be amplified using streptavidin coated nanoparticle (803). The biotinylated analyte can be bound to the nanoparticle either before the sample has been applied to the biosensor, or thereafter.

In one embodiment of the invention, a biosensor that can be easily tailored, can be constructed by doping avidin or streptavidin (904) into the sensitive layer (901) as shown in Figure 9. Before binding the receptors into the sensitive layer of said biosensor, the receptors can be linked with biotin by methods that are well known in the art. The biotinylated receptors (903) can then be bound to the avidin/streptavidin that is bound to the sensitive layer of the biosensor.

In another embodiment of the invention, enzymes that are present in a sample can react with their substrates and/or cleaving targets that are present in the sensitive layer as shown in Figure 10. The analytes (1003) in the sample interact with the receptors (1005) that are bound to the sensitive layer (1001). The result is the cleavage of the receptor (1005) in to two parts, of which only one is bound to the sensitive layer (1006). The analyte and the cleaved part of the receptor can be washed away from the sensitive layer and thereafter the change in the optical properties of the sensitive layer can be measured to determine the activity of the analyte towards the said receptor.

A sample, biological sample or other sample, can be put in contact directly or it can be pre-treated with variety of ways, well known to those skilled in the art.

In one embodiment of the invention, a sample can be applied on to a biochip so that a single sample is in contact with several biosensors simultaneously as shown in Figure 38. The biosensors that become in contact with the sample can be each specific to their own analytes, or alternatively one or more may measure the same analyte. Furthermore, one ore more biosensors can be constructed so that it does not contain any receptors. Therefore these sensors act as a reference sensor for measurement of unspecific binding.

In another embodiment of the invention, any number of samples can be applied to a biochip so that each sample comes in contact only one biosensors as shown in Figure 39. The biosensors can all measure the same or different analyte. Any variation with the number of biosensors that a sample comes in contact with is also possible.

In one embodiment of the invention, a sample can be guided to a biosensor by using microfluidistic channels as is shown in Figure 40.

A biochip can also comprise of biosensors that all or some are specific towards a single analyte but have different receptors.

The sample can be washed away for example with water, aqueous solution, like buffer or with any suitable solvent. When the properties of the sensitive layer change also the optical properties of the grating change. To measure these changes, a detection system is required.

A detection system measures the intensities of certain diffracted light beams.

These intensities, i.e. diffraction efficiencies, change as the optical properties of the grating change. The sensitivity of the detection system can be optimized by properly choosing the grating geometry, materials, measured diffraction orders and detection system configuration, and can be done by a person skilled in the art.

A detection system of the innovation can rely on the following principle. One or more light beams containing one or more wavelengths, one or more polarization, and possibly one or more different income angles are collimated or focused onto the same spot on the grating. Some of the intensities of the diffracted (reflected and/or transmitted) beams are detected by using one or more detectors. The changes of the refractive index of the sensitive layer of the biosensor can be calculated from the changes in the measured intensities. There are many possible variations about how the change in the refractive index is obtained from the intensities of the diffracted beams, and a skilled professional can easily find suitable solution for each application.

For easier treating of the subject, let us divide the illumination light into "channels" so that each channel contains only one linearly polarized monochromatic beam. So every incoming photon belongs to some channel with specified income angle, θ, polarization, p, and wavelength, λ. Every channel divides in the grating and is detected by one or more detectors. The detected intensity of the &'th diffraction order beam of the channel (λ, p,θ) can be written in the following way

l(k, λ,p,θ,t) =

L_x (λ, p, θ)

L₃ (k, λ, p, θ, t) ^■ L₄ (k, λ, p, θ) ^■ S(k, λ,p,θ, t) where λ = wavelength, θ = income angle, p = polarization, t = time, P = intensity of the source, Lj = the intensity loss factor from the source to the biochip, L₂ = the intensity loss factor from the biochip border to the grating layer, D = diffraction efficiency, L₃ = the intensity loss factor from the grating layer to the biochip border, L₄ = the intensity loss factor from the biochip border to the detector. S = the sensitivity of the detector which detects the k't order beam. In addition to the channel parameters (λ,p,θ), these variables are functions of the order number k and the measurement time t. P, Lj and L₂ are the same for all orders because the orders have common path before the grating layer. Lj and L₄ are time-independent because they are properties of the detection device whereas L₂ and L₃ are properties of the biochip which can change with time.

In principle the intensities of two or more diffraction orders are measured (the set of these diffraction orders is denoted with K) at the same time. Then the sum of a subset of the measured diffraction orders divided by the sum of another subset of the diffraction orders as

G{λ,_P,θ,t) =

where K_λ ,K₂ e K , is calculated. This simplifies to the form

i.e. G is insensitive to the intensity noise of the source. In addition to that, it is insensitive also to the changes in L₂ which can happen for example when contamination or overly solution layer remains on the biochip between the source and the grating layer when the analyte is applied on it.

In most of the cases it can be supposed that L₃ is time independent. When measuring with the best sensitivity the following has to be considered: when measuring a transmission order and the light beam is coming through the grating layer to the substrate layer, it can be supposed that L₃ is really time independent. That is because it can be supposed that the loss factor through the substrate layer is time independent. When measuring the reflection orders one has to ensure that no overly solution layer remains on the biochip because that would cause time dependence to L₃. Most of the cases it can also be supposes that S is time independent. When measuring with the best sensitivity the effect of the time dependence of S can be minimized by careful device design and calibration. After all G simplifies to

∑ D(k, λ, p, θ, t) ■ L₃ (k, λ, p, θ) ■ L₄ {k, λ, p, θ) ■ S{k, λ, p, θ) ■ a{k, λ,p,θ) K_X

G(λ,p,θ,t) =

∑D(k,λ,p,θ,t)-L₃(k,λ,p,θ)- L₄(k,λ,p,θ)- S(k,λ,p,θ a{k,λ,p,θy k<=K,

where a is a known calibration function related to S. Now, the only changes of G as a function of time are caused by the changes in diffraction efficiency D. It was supposed that the intensities of all the measured diffraction orders were measured at the same time. It is also possible to measure some or all of them in a series in a short time period during which the intensity of the source and the sensitivity of the detector are constant. In order to measure the reflection orders the reflectance must be high enough. In that case the refractive index difference between the grating and sensitive material should preferably be high, more than 0.1. Another possibility is to use reflective substrate, for example silicon, with the grating and sensitive material which both have refractive index around 1.5. Still another possibility is to use metal substrate or metal grating. With abovementioned measurement principle one or more channels can be used to illuminate the biosensor. When using only one channel, i.e. only one linearly polarized monochromatic beam only one G - number will be monitored. In that case one has to had fully known and calibrated biosensor and detection system in order to be able to translate the changes in G to the changes in the refractive index of the sensitive material. This is practical in simple low cost applications where the sensitivity need not be optimized. However, for a device of high sensitivity it can be advantageous to use several channels by using several wavelengths for example. By using several channels it is possible to obtain more information about the biosensor than by using only one channel. Thus it is not necessary to know for example the exact shape of the grating beforehand if the refractive index difference between the grating layer and the sensitive layer is known. On the other hand, if the geometrical shape of the grating is known, it is not necessary to know the refractive index of the grating layer beforehand. Also by using several channels it is possible to see if the biosensor is broken before or during the measurement and so it is possible to know if the measurement result is right or not. Another important advantage of using several channels is that it is possible to average measurement results over the channels so that the measurement result will be better. These advantages are clear to a skilled professional and so the exact mathematical and physical treatment of the subject is not necessary in this context.

If power distribution between the polarization states is known, as usually is especially with bulbs and LEDs, it is possible to treat a monochromatic beam which includes both polarization states as one channel.

One embodiment of the detection principle is to use phase grating, for example a binary phase grating or a sinusoidal phase grating, the phase difference of which changes as a function of refractive index of the sensitive material. A phase grating also works as a weak amplitude grating the efficiency of which changes as a function of refractive index of the sensitive material.

In order to optimize the detection system it is important to solve the diffraction properties of the grating. The diffraction efficiencies of gratings can be solved analytically when the geometry of the grating is simple. In many cases however the analytical solution is too complex in comparison to relatively simple numerical modeling. Numerical modeling of diffraction gratings is possible for example by using GSOLVER (Grafting Solver Development Company, Allen, Texas, USA) software. GSOLVER utilizes a full 3-dimensional vector code using hybrid Rigorous Coupled Wave Analysis and Modal analysis for solving diffraction efficiencies of arbitrary grating structures for plane wave illumination.

One embodiment of the detection principle is to measure the ratio of the intensities of the first transmission order and the zeroeth transmission order.

In addition to abovementioned preferred measurement principles a person skilled in the art can construct a wide variety of other mathematical functions which have the intensities of the diffracted light as parameters and which can be used to measure the change of the refractive index of the sensitive material. The diffractive element can have both phase and amplitude grating properties and they both can be taken in account.

Detection systems working with abovementioned principles can be constructed in many various ways depending on the application. Detection systems can vary in sensitivity, complexity and cost according to what is needed. A very sensitive device could use several wavelengths, several income angle and measure several diffraction orders. The wavelength of the light can vary from visible to near infrared (VIS - NIR), preferable from 400 nm to 780 nm. The income angle can vary from 0 degrees to 90 degrees. On the other hand the most simple device could measure the zeroeth and the first order intensities by using only one beam with zero degrees income angle and one wavelength. The following detection system configurations are only examples as a person skilled in the art can easily construct many different variations which rely on the same measurement principles. A detection system can comprise (See Figure 37) sensor chip, a light source providing light, light detector(s) and optionally transmitting optics directing light from the source to the sensor chip, collecting optics which gathers light from the sensor chip to the detector(s), signal processing unit and possibly also a modulation unit.

The transmitting optics are used to direct light from the source unit to the sensor chip. The transmission optics can be constructed in various ways depending on the application. Typically transmitting optics comprises a collimating or focusing unit, which shoots the beam through the biosensor. In many cases transmitting optics comprises also polarization filters and/or spectral filters.

The collecting optics is used to collect light from the sensor chip to the detector(s). In some cases there is no need for the collecting optics. In other cases it comprises a focusing optics, which focuses the diffracted light to the one spot to the detector.

The collimating and focusing unit collimates or focuses the beam before it is directed to the biosensor. This is done for the following reasons: the beam diameter must be correct on the biosensor, the numerical aperture of the beam must be small enough so that different diffraction orders and channels do not mix together, and, the beam diameter on the detector must be small enough so that the whole spot fits in the detector area. Focusing can be done by using lenses, mirrors or diffractive optical elements.

In the following, the preferred embodiments of the detection system are described in more detail. Depending on the chosen configuration and measurement system, we might measure by using one or two polarizations. Almost every embodiments below, it is possible to choose whether or not to use polarizers and that is why polarizers are not drawn in the figures. Polarizers may be placed for example just after the collimating or focusing unit.

One embodiment of the detection system uses a monochromatic source. Monochromatic beam is collimated or focused to have small numerical aperture. The beam goes through the biosensor and divides into several transmission order beams by diffraction. By using one or more detectors two values are measured: the intensity of the zeroeth order (I₀) and the sum of the intensities of other orders (I_s). Is can contain only the intensity of the +l:st order or I_s can contain the sum of the intensities of the +l:st and — l :st orders, or optionally also the intensities of several other orders. When the ratio I_s 11₀ is calculated, the change of the refractive index of the sensitive material is obtained when the original refractive index difference between the grating material and the sensitive material is known. It is supposed that the shape of the grating is well known too.

Figure 18 illustrates one embodiment of the detection system. A beam from a wavelength stabilized and pulse modulated laser is directed through a collimating or focusing optics (1801,1802) to the biosensor (1803). The intensities of the zeroeth and the first transmission orders are detected by using two separate detectors (Det_l5 Det₂) the both of which are connected into the same metal piece for temperature stabilization. For wavelength stabilization a part of the beam can be directed by using a beam splitter (1804) into a simple spectrometer which can comprise for example a blazed reflection grating (1805) with a CCD row detector.

Figure 19 presents one embodiment of the collecting optics where several transmission order beams from the biosensor (1901) are collected to the one detector (Det by using a focusing lens (1902).

Figure 20A presents another embodiment of the collecting optics where the zeroeth order beam is deflected to the separate detector (Det₂) by using a small mirror (2001).

Figure 20B shows another embodiment of the collecting optics, where several transmission order beams from the biosensor are collected to only one detector (Det by using a focusing lens. By using a spatial modulator (2002) like optical chopper, or LCD or DMD modulator, several diffraction orders can be measured in a short period of time. Using this solution the pulse modulation of the source is not needed.

In another embodiment of the detection system a beam from a wavelength stabilized and pulse modulated laser is directed through a collimating or focusing optics to the biosensor. The diffraction orders are detected by using a CCD-row or CCD-matrix detector as shown in Figure 21. By this embodiment all wanted diffraction orders can be detected at the same time. In addition to that the CCD- detector controls the wavelength of the light so additional spectrometer is not needed.

Instead of the laser, LED's or bulbs with filters can be used as a light source in the abovementioned embodiments. Figure 22 illustrates one embodiment in which a light from a LED or bulb source (2201) is focused into a pinhole (2202) by using a mirror (2203) and lens (2204). Needed filters (2205) are placed after the collimating or focusing optics (2206).

Figure 23 illustrates an embodiment in which a LED or bulb source (2301) is located inside an integrating sphere (2302), which contains a pinhole for light output (2303). When filters are used, the spectrometer is not needed for wavelength control.

When using a bulb source a spatial modulator is needed for light modulation. That is because bulbs can not be pulsed as lasers or LEDs. Figures 24, 25 and 26 illustrate embodiments where light is guided through optical fiber or optical waveguide from source to collimating or focusing optics. In Figure 24 the optical fiber (2401) collects light from the integrating sphere (2402) with a light source (2403) and guides it to the collimating or focusing optics (2404). In Figure 25 the light from a fiber pigtailed laser (2501) is divided by a fiber coupler (2502) to the collimating or focusing optics (2503) and to the spectrometer (2504). Figure 26 illustrates an embodiment where source (2601) is integrated into the substrate (2602), light couples from the source to the waveguide (2603), goes through Bragg grating filters (2604), is modulated by an electro-optical modulator (2605) and arrives to the collimating or focusing optics (2606).

Another embodiment of the detection system uses a source with a continuous spectral band. The spectral band can be from 10 nm to 1000 nm wide. The light is collimated or focused to have small numerical aperture. The beam goes through the biosensor and divides into several transmission order beams by diffraction. By using one or more detectors and the light modulator the ratio of the intensity of the zeroeth order and the sum of the intensities of some other orders are detected separately for each wavelength. When this ratio is known in the whole spectral band, the change in refractive index of the sensitive material can be calculated. The benefit for using several wavelengths is that the biosensor need not to be fully calibrated before measurement. For example if we know the original refractive index difference between and the grating material and the sensitive material, we can use the information obtained by using several wavelengths to calculate the refractive index change without previous information about the geometry of the grating. On the other hand, if we know the geometry of the grating, we can calculate the refractive index change without information about the original refractive index difference, for example. If the biosensor is well calibrated already we, by using wideband source, have possibility to make measurements with a better accuracy.

One embodiment of the detection system using a wide band source is shown in Figure 27. The light from a bulb (2701) is focused by a mirror (2702) and lens (2703) to a pinhole (2704). The light from the pinhole goes through a collimating or focusing (2705) unit after which it is filtered (2706) into the wanted spectral band. The beam is then modulated by a modulator, which can be for example a chopper or LCD (2707). The beam goes through the biosensor (2708) and divides into several diffraction order beams. Apart from the zeroeth order beam all other orders have spread into their spectra in horizontal direction. The spectrum of the zeroeth order is spread in vertical direction by using another grating (2709). These spectra are detected by using one large CCD-matrix or three or more smaller CCD-matrices or row detectors (2710).

As monochromatic sources, the wideband bulb sources can also be connected to an integrating sphere. Similarly, waveguides or fibers can be used for guiding light and fiber or waveguide components can be used for filtering, modulating or dividing light. Figure 28 shows an embodiment of the invention, where an integrating sphere (2801) is used with a wideband bulb source (2802). In this case the zeroeth order beam is spread into its spectrum by using a prism-grating-prism (2803) component so that zeroeth order spectrum can be detected by using the same row detector than for the +1 'st and -1 'st order spectra (2804).

One embodiment of the detection system uses wideband source with optical filters which filter out certain known bands from the spectrum of the source. These known spectral features can be used to calibrate the wavelength on the CCD-matrix or row detector. One embodiment of the detection system uses wavelength tunable laser or a bulb with a scanning filter as a light source to get the same measurement result than with the wideband source. When using scanning light sources, the possible optical configurations are similar with the ones with monochromatic light sources.

Figure 29. illustrates one embodiment of the source with a scanning filter. The light from a bulb (2901) is focused to the pinhole (2902) by using a mirror (2903) and a lens (2904). The light from the pinhole is guided through a collimating or focusing unit (2905) to the blazed reflection grating (2906) which spreads the beam to its spectrum. The spectrum is modulated with a spatial modulator (2907) like LCD and collimated again with a pair of lenses (2908). Another embodiment of the detection system uses a source with several (two or more) narrow spectral bands. The advantage of this source is the same than with the case of the wideband source, i.e. we have possibility to better accuracy and the biosensor need not to be fully calibrated before measurement.

One embodiment of the detection system with a source with several narrow spectral peaks is presented in Figure 30 The source comprises of an integrating sphere (3001) with several LEDs (3002) and a pinhole (3003) for light output. LED's are electrically modulated so that only one wavelength band is active at each time moment. The light from the pinhole is collimated or focused (3004) and guided through the biosensor (3005). The beam divides into several diffraction orders which are detected by a CCD - matrix or a row detector (3006). At the same time, the wavelength stability of the spectral bands can be monitored by detecting the position of the diffraction orders at the detector. This wavelength information can be used to control the LED's temperature and current if needed. In addition to set of LEDs, the set of narrow wavelength bands can be produced by coupling light from lasers or filtered bulbs to the integrating sphere.

One embodiment of the detection system, which is presented in Figure 31 uses white light source (3101) with Fabry-Perot interference filter (3102) to produce a series of narrow wavelength peaks.

Instead of CCD or row detectors, conventional silicon photodiodes can be used also when a series of narrow spectral bands are used as is illustrated in Figure 32. The optical configuration is the same that the ones with monochromatic light source apart from the light source. Light source, which comprises an integrating sphere (3201) with LEDs (3202) and a pinhole (3203), is modulated so that only one spectral peak is activated at each time moment. Another suitable light source for this kind of serial measurement is also a light bulb modulated by a filter wheel.

One embodiment of the detection system uses several beams each of which illustrate the same spot on the biosensor but with different income angles. Figure 33 presents one possibility where collimated monochromatic beam coming from the source and focusing units, is divided vertically by using an additional grating (3301) into several beams with different propagation directions. These beams are focused into the one spot on the biosensor with vertical line groove pattern (3302) by using a lens (3303). Beams divide again but in horizontal direction. The following beam matrix can be detected for example by a CCD-matrix (3304). This solution can also be applied together with a beam including several wavelengths at the same time.

These abovementioned detection systems were examples of preferred configurations. A skilled professional can vary already mentioned optical configurations and construct new optical configurations, which rely on the abovementioned measurement principles. For example, it is well known that in most of the cases lenses can be replaced by mirrors or diffractive optical elements. Also bulk optical components can be replaced in many cases by fiber optical components or integrated optics.

EXAMPLE 1

Fabrication of biosensor

Biosensor fabrication begins with an untreated flat glass substrate (4" x 4"). The glass substrate is first cleaned by using acetone, isopropanol and methanol baths in ultrasonic cleaner. After this the substrate surface is treated in a plasma etcher for 5 min at 300W (0₂-gas). The glass substrate is coated with a negative lithography- tone photoimageable material by using a spin coater. After this the deposited film is exposed through a photomask in contact exposure-mode using a maskaligner (Karl- Suss, MA-6, UV-400 optics). The used photomask is specifically designed to produce the desired grating structure. In this case the photomask was designed to produce grating structures with period of 6 μm. A 5" x 5" photomask was used that had 25 similiar gratings with dimension of 0.5 mm x 0.5 mm. The exposure step is followed by a development step where the unexposed areas of the film dissolve into the used developer solvent. As a final step the produced grating structures are fully solidified by baking at elevated temperatures. In this example the gratings were baked at 200°C for 3 hours. The used negative lithography-tone photoimageable material can be choosed to have a low refractive index (e.g. 1.47 at 632.8 nm) or a high refractive index (e.g. 1.60 at 632.8 nm). The choice of the used grating material (with certain refractive index) has to be done based on the selected sensitive material layer (with certain refractive index) to be able to achieve the optimal sensitivity of the biosensor.

In this example the sensitive material layer was selected to be a hybrid aluminum oxide siloxane material with refractive index of 1.51 at 632.8 nm. The sensitive material layer is deposited on top of the gratings by using a spin coater. Finally the sensitive material layer is solidified by baking the sensor construct at elevated temperatures. In this example the samples were baked at 50°C for 15 hours.

EXAMPLE 2

Incorporation of Biological Receptors into the Sensitive Layer

The biological receptor molecules that specifically bind to the desired analyte can be incorporated to the sensitive layer of the sensor device during the synthesis of the sensitive layer material.

Purified commercially available antibodies (mouse anti-human leptin, part 840279 in Human Leptin DuoSet ELISA Development Kit DY398, purchased from R&D Systems Inc. Minneapolis, MN, USA) were diluted in PBS into concentration of 4.0 μg/ml. 10 μl of the diluted antibody mixture (40 ng of anti-leptin antibodies) was mixed with 10 g of a soluble hybrid aluminum oxide siloxane material and mixed 10 minutes in ultrasonic bath for 10 minutes. The prepared hybrid material doped with anti-human leptin antibodies were deposited on top of the fabricated grating structures by using a spin coater as described in Example 1. After this the sensor construct was baked at 50°C for 15 hours.

EXAMPLE 3

Manufacturing of streptavidin doped biosensors

Streptavidin is a bacterial protein that has great affinity and specificity for biotin. Receptor proteins can be easily biotinylated by using methods well known in the art. Biosensors that have been doped with streptavidin can be easily modified to detect desired analytes by binding biotinylated receptors to the streptavidin that has been incorporated to the surface of the biosensor. Purified commercially available streptavidin (AS-5000, purchased from R&D

Systems Inc. Minneapolis, MN, USA) was diluted with H₂0 into concentration of 100 μg/ml. 100 μl of diluted streptavidin mixture (10 μg of streptavidin) was mixed with 10 g of a hybrid aluminum oxide siloxane material using an ultrasonic bath for 10 minutes. The prepared hybrid material doped with streptavidin was deposited on top of the fabricated grating structures by using a spin coater as described in Example 1. After this the sensor construct was baked at 50°C for 15 hours.

EXAMPLE 4

Immobilization of biological receptors into the sensor chip

Biological receptors that bind to the desired analyte can also be immobilized to the sensitive layer after the deposition and curing of the sensitive material layer on top of the gratings. Receptor molecules can be immobilized to a surface through variety of surface chemistries. Avidin/streptavidin - biotin coupling offers a precise method to attach biotinylated receptor molecules into biosensor which has been doped with streptavidin (Example 3).

Purified commercially available biotinylated antibodies (mouse anti-human leptin, from Human Leptin DuoSet ELISA Development Kit DY398, purchased from R&D Systems Inc. Minneapolis, MN, USA) were diluted in PBS into concentration of 1.0 μg/ml. 10 μl volume of the diluted antibody mixture (10 ng of biotinylated anti-leptin antibodies) was applied to the surface of biosensor. After an immobilization period of 60 minutes in 37 °C the sensor was rinsed extensively with PBS and air-dried.

EXAMPLE 5

Optical modeling of biosensor

In this example biochips comprised of a glass substrate, a grating layer and a sensitive layer. The grating was a sinusoidal phase grating with a period of 6 μm. In this simple demonstration we wanted to design a detection system which uses 632.8 nm laser as a source, illuminates the biosensor with zero degrees income angle and uses both TE and TM polarized light. The measurement will comprise of the measurements of the zeroeth and the first order intensities before and after applying the analyte to the biosensor. After the measurements, the change in the refractive index of the sensitive layer could be calculated by using the information obtained in the simulations. These simulations can many times be replaced by measurements, but typically simulations can be done faster. Measurements have to be done instead of simulations when good accuracy is needed, at least simulations have to be verified by measurements partially.

At first, the numerical modeling of the grating was used to optimize the modulation depth of the grating before manufacturing. The modeling was made by abovementioned GSOLVER software. The refractive index of the grating layer was 1.600 and 1.514 for the sensitive layer for source wavelength. Figure 34 shows the simulated diffraction efficiencies for zeroth (TO) and first (TI) transmission orders for TE and TM polarized light as a function of grating modulation depth. From this we can see that a good sensitivity can be obtained when the grating depth is more than 1.5 μm, for example 2.5 μm. On the other hand, greater depths than 3 μm are difficult to manufacture with a good quality when the period of the grating is 6 μm. From the figure we can also see that the diffraction is not significantly polarization dependent with zero degrees income angle.

If the manufacturing process preserves the shape of the grating constant and if we know accurately the refractive indexes it will be enough to measure the grating only after applying the analyte. In that case we can use simulation to approximate the change in the refractive index of the sensitive material from the measured diffraction efficiencies: Figure 35 shows the simulated diffraction efficiencies as a function of the change of the refractive index («_/) of the sensitive layer. In this simulation the grating modulation depth was 2.5 μm. Figure 36 shows the change of «_/ as a function of the ratio of the intensities of the first and zeroeth diffraction orders. This graph can be used to convert the measured intensities into refractive index change.

Preferably we will measure the ratio before and after the analyte has been applied to the biosensor. Supposing that we know the shape of the biosensor, from the first measurement we can calculate the exact height of the biosensor. This result can be used when calculating the refractive index change from the second measurement.

EXAMPLE 6

Detection of leptin using biosensor

An assay was performed to demonstrate the capability to detect biological molecules label free. The biosensor was fabricated by mixing rat monoclonal anti- leptin antibody (mouse anti-human leptin, part 840279 in Human Leptin DuoSet ELISA Development Kit DY398, purchased from R&D Systems Inc. Minneapolis, MN, USA) into a hybrid aluminum oxide siloxane material using an ultrasonic bath for 10 minutes. The prepared hybrid material doped with mouse anti-human leptin antibodies was deposited on top of the fabricated grating structures by using a spin coater as described in Example 1. After this the sensor construct was baked at 50°C for 15 hours. In order to prevent non-specific binding, the biosensor surface was exposed to

1% solution of Bovine Serum Albumin (BSA) in PBS for 30 minutes and rinsed extensively. Following blocking, 10 μl drops of recombinant human leptin hormone (human leptin, in Human Leptin DuoSet ELISA Development Kit DY398, purchased from R&D Systems Inc. Minneapolis, MN, USA) in 1 μg/1 concentration was applied to the sensor surface. The droplets were allowed to incubate for 5 minutes before thorough washing of the unbound material with Dl water.

The binding of the analyte was then measured using a laser beam and measuring the first zeroeth-order ratios of transmitted light. A sensor constructs with no specific receptors incorporated into the sensitive layer were used as controls. The results of the binding assay are shown in Figure 17. Alternative process method in all previously mentioned manufacturing processes may be accomplished by using so called reel-to-reel (also called roll-to- roll) processing, wherein the view of achieving of the objectives stated above the method for manufacturing optical elements is mainly characterized in that the method comprises the steps of supplying a printing cylinder with printing elements for forming optical structures, applying optical material on the printing cylinder and creating the optical structures on the substrate material web or substrate material sheets.

According to the method optical elements are produced in a printing system in which the optical element is transferred from the printing cylinder to a suitable substrate material. In an extreme case the whole substrate material can be covered with optical material. The substrate material is paper or plastic or other passive or active optical/electrical material, such as a semiconductor material. The substrate material is in a form of a web or separate sheets of suitable size. Optical material is a material system that can be handled and delivered in the liquid format to its final location, e.g. substrate, in which it forms to a stable or metastable phase, i.e. a solid or metasolid form. After taking the stable or metastable phase the material presents optical properties, which can be for example transparency or selective transparency, reflectivity, diffraction, light emission, polarization selectivity, modulation or phase modulation. The optical material can be e.g. an organic polymer that is dissolved in an appropriate solvent such as organic solvent or water. The material can also be a suspension of solid particles in a liquid carrier. In both cases the material forms stable or metastable form when the solvent or the carrier is removed. However, the invention is not restricted to these materials. The manufacturing of the optical elements according to the method comprises the primary printing step in which the optical element is formed on the substrate surface using the primary printing method. The primary printing system is preferably a gravure printing system, a gravure offset printing system, a flexographic printing system, an offset lithographic system, electrophotographic printing system, or a combination of these.

Gravure printing includes direct gravure printing, in which the printable pattern is transferred from the printing cylinder to the printing surface, gravure offset printing in which the printable pattern is transferred from the printing cylinder to a second cylinder and from it to the printing surface, and intaglio printing. In intaglio printing process viscous inks are used which allow the printing patterns of larger uniform areas. After the primary printing phase the printed optical element is optionally treated with an additional printing method(s). In the additional printing phase devices for digital printing, hot stamping, silk, screen printing and/or photolithographic printing may be applied.

The method makes it possible to produce high quality optical elements at a cost which is a remarkably lower than whoa using conventional methods. This is preferably achieved by manufacturing a printing cylinder provided with surface structures to form optical elements on a substrate material. The printed optical component is formed by using a liquid form optical material that is suitable for printing systems, and is, if needed, cured with suitable method. The printed optical elements can further be laminated, covered or printed with additional optical layers. The curing method can be such as thermal curing or UV curing.

Advantages of using the gravure printing method are deep enough, structures achievable in gravure printing, high quality of the transfer of the printing pattern, high throughput, and low price compared to the conventional methods of producing optical elements. The gravure printing method can also be easily integrated to other process parts such as lamination, coating or embossing.

In additional step the printed optical element is further provided with additional layers to form the desired optical and biochemical coatings. Suitable methods for additional treatment of the optical elements are hot stamping, photolithographic printing method and silk screen printing. In hot stamping printing method, ink coated on a film transfers by heat and pressure to a web. The raised parts of the profile contact the film, and the resulting heat flow causes liquidifϊcation of the ink. In silk screen printing method the printing plate is replaced by a stencil having different porosity in the printing and nonprinting areas. Ink is pressed through the stencil to the paper or other substrate positioned below the stencil. According to the process method the additional step for manufacturing optical elements may also include using a stamping unit in which an area consisting of an optical layer is printed on the web and then an optical pattern is stamped on this area. In the method for producing optical elements a printing cylinder is prepared containing printing elements of the form of optical elements. The optical elements are created on a substrate surface running as a web through the printing system. According to the method the printing elements in the printing cylinder are preferably of the form of lines or other three-dimensional structures instead of point structures of the prior art printing cylinders.

According to the process method the additional step for manufacturing may include coating of biochemically sensitive coating on a previously prepared optical components. The biochemically sensitive coatings are created on a substrate surface running as a web through the printing system.

Claims

1. A device for detecting the presence and/or quantity of an analyte of interest in a sample, which device comprises an optical grating and a sensitive layer opposite to the grating and in close contact with it characterized in that the sensitive layer is substantially porous and capable of binding specifically the analyte of interest throughout the layer structure and the binding of the analyte of interest to the sensitive layer causes a measurable change in the optical properties of the sensitive layer.

2. The device according to claim 1, characterized in that it comprises detection means for measuring the change in the optical properties of the sensitive layer.

3. The device according to claim 1, characterized in that the sensitive layer comprises porous hybrid aluminum oxide siloxane or hybrid organo siloxane material or combination of both.

4. The device according to claim 1, characterized in that the sensitive layer comprises aluminum oxide, silicon oxide or organic polymer ingredients or mixtures thereof.

5. The device according to claim 1, characterized in that the sensitive layer contains receptor molecules capable of specifically binding the analyte of interest.

6. The device according to claim 1, characterized in that the sensitive layer contains receptor molecules that are specific targets of an analyte of interest.

7. The device according to claim 5, characterized in that the receptor molecules are incorporated in the sensitive layer or they are located on top of the sensitive layer.

8. The device according to claim 6, characterized in that the receptor molecules are incorporated in the sensitive layer or they are located on top of the sensitive layer.

9. The device according to claim 5, characterized in that the receptor molecules are biological molecules, such as proteins, peptides, polyclonal or monoclonal antibodies, single chain antibodies (scFv), antibody binding fragments (Fab), antigens, DNA or RNA molecules (nucleic acids), ligands, lipids or carbohydrate molecules or like.

10. The device according to claim 5, characterized in that the receptor molecules are pharmacological molecules.

11. The device according to claim 5, characterized in that the receptor molecules are aggregates bigger than a single molecule, such as cell organelles, viruses, bacteria or cell either in part or in whole.

12. The device according to claim 5, characterized in that the receptor molecules are inorganic or organic molecules.

13. The device according to claim 1, characterized in that the grating is a transmission or reflection grating.

14. The device according to claim 1, characterized in that the grating is integrated with an optical waveguide.

15. The device according to claim 1, characterized in that the grating is integrated with an optical fiber.

16. The device according to claim 1, characterized in that it contains one or more units of said gratings with said sensitive layers in a form of a biochip.

17. The device according to claim 16, characterized in that biochip contains one grating with several sensitive layer locations.

18. The device according to claim 16, characterized in that multiple detections of analytes can be run simultaneously or separately on the microarray plate by detecting the changes in the optical properties of the sensitive layers by using multiple light sources.

19. The device according to claim 16, characterized in that multiple detections of analytes can be run simultaneously or separately on the microarray plate by detecting the changes in the optical properties of the sensitive layers by using a single light source which is optically divided.

20. The device according to claim 16, characterized in that one or more biosensors in a biochip do not have receptors incorporated into the sensitive layer.

21. The device according to claim 20, characterized in that biosensors with no receptors towards specific analyte function as a reference channel to measure unspecific binding.

22. The device according to claim 1, characterized in that the grating is integrated with a heater and/or a cooler, and/or temperature stabilizer.

23. The device according to claim 1, characterized in that the sample is delivered to the sensitive layer through microfluidistic channel.

24. The device according to claim 23, characterized in that the microfluidistic channel is integrated into the biochip, which contains multiple biosensors.

25. A method for detecting the presence and/or quantity of an analyte of interest in a sample, using an optical grating and a sensitive layer opposite to the grating and in close contact with it characterized in that the sensitive layer is substantially porous and capable of binding specifically the analyte of interest throughout the layer structure and the binding of the analyte of interest to the sensitive layer causes a measurable change in the optical properties of the sensitive layer.

26. Method for detecting the presence and/or quantity of an analyte of interest in a sample characterized in that the sample is brought in contact with the device of claim 1 and the change in the optical properties of the sensitive layer caused by the binding of analyte of interest to the sensitive layer is measured

27. Method according to claim 25, characterized in that the amount of change in the optical properties of the sensitive layer is proportional to the amount of the analyte of interest in the sample.

28. Method according to claims 25 or 26, characterized in that grating is illuminated.

29. Method according to claim 25, characterized in that grating is illuminated with a broad band source and reflected or transmitted spectrum is detected.

30. Method according to claim 25, characterized in that the grating is illuminated with excitation source and reflected or transmitted spectrum is detected

31. Method according to claim 27, characterized in that the ratio of the detected intensities of the first transmission order and the zeroeth transmission order from the grating is measured.

32. Method according to claim 25, characterized in that the diffraction efficiencies of 0^l , ±l^st and ± 2ⁿ of diffraction orders of the grating are used for the determination of the changes in the optical properties.

33. A Device according to claim 1, characterized in that the diffraction efficiencies of 0^th, ±l^st and ± 2^nd of diffraction orders of the grating are used for the determination of the changes in the optical properties.

34. The device according to claim 1, characterized in that the sensitive layer material is optically transparent between 300 nm 1700 nm

35. The device according to claim 1, characterized in that the sensitive material contact angle with water is 50 degrees or less.

36. The device according to claim 1, characterized in that the diffraction grating is manufactured by using reel-to-reel embossing process;

Biological layer is deposited on the diffraction grating in term of gravure or flexografic printing.

37. The device according to claim 1, characterized in that the grating manufacturing and biological layer deposition are made in one integrated process line.

38. The device according to claim 36, characterized in that the diffraction grating is manufactured so in that it contains one or more units of said gratings with said sensitive layers in a form of a biochip.

39. The method according to claim 26, characterized in that the optical properties of the sensitive layer are amplified by using nanoparticles.

40. The device according to claim 5, characterized in that the receptor molecules are incorporated in the sensitive layer.

41. The device according to claim 5, characterized in that the receptor molecules are located on top of the sensitive layer.