WO2003025547A1 - Method and device for screening analytes using surface plasmon resonance - Google Patents

Abstract

The present invention relates to a method and device for screening analytes using surface plasmon resonance. In the present invention, a disposable flow cell unit to be adapted to a miniaturized SPR sensor is provided. The flow cell incorporates reaction chambers, valves, distributing connectors, and reagents to perform a rapid one-step quantitative assay in a sample solution (e.g. whole blood, serum, saliva, urine, or other bodily fluids). Also, the flow cell could be adapted to biosensors for nucleic acid and small molecule detection. Possible applications include food quality control, environmental screening, and point-of-care test devices. The SPR sensor preferably has a plurality of sensing surfaces capable of performing independent assays.

Description

METHOD AND DEVICE FOR SCREENING ANALYTES USING SURFACE PLASMON RESONANCE

CROSS REFERENCE TO RELATED APPLICATION This application is based on application Nos. 10-2001-58777, 10-2002-

30285 and 10-2002-56131 filed in the Korean Industrial Property Office on September 21 , 2001 , on May 30, 2002 and on September 16, 2002, the contents of which are incorporated hereinto by reference. FIELD OF THE INVENTION The present invention relates to a method and device for screening analytes in a sample. The device comprises disposable integrated flow cells for performing assays. More specifically, the integrated flow cells are to be adapted in a surface plasmon resonance sensor, which provides a cost-effective and rapid method for screening biohazards or diagnosis of diseases. BACKGROUND OF THE INVENTION

Optical sensors for sensitive detection of proteins, nucleic acids, or small molecules are widely being developed. Particularly, optical sensors based on evanescent waves occurring with total internal reflectance, for example with surface plasmon resonance (SPR), are of practical importance since they provide highly sensitive measurement methods without the labeling of target molecules (See

Schuch P., Ann. Rev. Biophys. Biomol. Struct 1997, 26, 541 , Malmqvist M. Nature 1993, 361 , 186, Corr M. et al., Science, 1994, 265, 946).

Commercialized SPR sensors have shown higher or comparable sensitivity compared to prior art immunoassay methods, such as enzyme-linked immunosorbent assays (ELISA) and radio-immunoassays (RIA). Even higher sensitivities could be achieved using various signal-amplifying agents (See Lin He et al., J. Am. Chem. Soc. 2000, 122, 9071 -9077).

A major advantage of SPR sensors is their capability of real-time monitoring of molecular interactions. This allows the monitoring of binding rates between biomolecules (e.g. antigen-antibody, DNA-DNA, RNA-DNA), and thus provides a method for rapid detection and quantitation of biomolecules.

Conventional methods including ELISA and RIA require hours to days of turnaround time and tedious processing steps, and also labeled probe molecules for detection.

For rapid and sensitive monitoring of biomolecular interactions, facilitating the diffusion of analytes to a sensing surface is required. Most commercial SPR sensors (e.g. from Biacore AB, IBIS instruments, Nippon Laser Electronics, etc.) apply a flow cell and increase the diffusion rate by flowing a sample solution containing the analyte on the sensor surface. Other commercial SPR sensors (e.g. from lAsys system) apply a stirred cuvette system to increase the diffusion rate.

Biacore AB has applied microfluidics for precise flow control to increase sensitivity and reproducibility of the device. This feature enables simultaneous detection of several analytes using 100 nl -100 μl of sample volume. Other laboratory-based SPR sensors applying microfabrication technologies to sample volumes smaller than 1 μl have also been developed (Lee, H.J. et al., Anal. Chem.2001 , 73, 5525-5531 , Furuki, M. et al., Sens. Actuators B, 2001 , 79, 63-69.). The devices mentioned above are currently being applied to various fields of research, including kinetic analysis of biomolecule interaction, mass transport, conformational change, etc. However, regardless of the advantages of SPR sensors mentioned above, devices intended for screening or diagnostic purposes are unaffordable because of the high cost and large size of currently commercialized devices. In particular, current devices are not suitable for field use such as for environmental monitoring or point-of-care testing.

Recently commercialized portable miniaturized SPR sensors (e.g. Spreeta™from Texas Instruments) have high potentials in environmental monitoring and point-of-care testing. If one-step assays could be developed on a miniaturized sensor platform, it would provide a cost-effective and rapid method for applications such as for biohazard detection and diagnosis of various diseases.

To achieve this goal, the aforementioned fluidic elements for effective mass transport are essential, and it would be advantageous to have a disposable micro-channel flow cell incorporated with various microfluidic elements such as valves, reaction chambers, and reagents needed for the assay. Also, by modifying the internal structures of the flow cell, the system could be readily adapted to immunosensors performing various formats of immunoassays (e.g. competitive assays or multistep assays) and nucleic acid biosensors.

Alpha-fetoprotein (AFP) is a biomarker for hepatocellular carcinoma and yolk sac tumors, and an abnormally high level of AFP is detected in blood of patients with these diseases (Chen, D.S. et al., Cancer 1997, 40, 779-783, Knight, G.J, Methods in clinical chemistry, Chapter 62, Pesce, A.J., Kaplan, LA. eds. Mosby-Year Book Inc., 459-465 (1987)).

Human chorionic gonadotropin (hCG) is a hormone secreted from trophoblast tissue, and it exists in many forms in vivo such as non-nicked hCG

(hCG), nicked hCG (hCGn), non-nicked free beta subunit (hCGβ), free beta core factor (hCGβ cf), and free alpha subunit (hCGα) (Alfthan, H. and Stenman, U. H., Mol. Cell Endoc n. 1996, 125, 107-120) .

In maternal blood where the fetus has Down syndrome, the concentration of AFP decreases and concentrations of hCG and hCGβ increase (Erkatz, I.R. et al., Am. J. Obest. Gynecol. 1984, 148, 886-894, Bogart, M.H. et al., Prenat. Diagn.,

1987, 7, 623-630, Spencer, K. et al., Prenat. Diagn. , 1997, 17, 31 -37).

Standard test methods for Down syndrome in fetuses comprise quantitation of AFP and hCG (or hCGβ) (known as 'double test'), and also quantitation of unconjugated estriol (uE3) ('triple test') in maternal blood. Also, in pregnancies with fetuses that have trisomy 18, the concentration of AFP and uE3 in maternal blood decreases (Canik, J.A. et al., Prenat. Diagn. 1990, 10, 546-548).

All of the above proteins are representative biomarkers suitable as analytes in an immunoassay applying the device disclosed in this patent. Other biomarkers comprising proteins, nucleic acids, and small molecules (e.g. polysaccharides) could also be screened using the device and method disclosed herein.

SUMMARY OF THE INVENTION In the present invention, a disposable flow cell unit to be adapted to a miniaturized SPR sensor is provided. The flow cell incorporates reaction chambers, valves, distributing connectors, and reagents to perform a rapid one-step quantitative assay in a sample solution (e.g. whole blood, serum, saliva, urine, or other bodily fluids). Also, the flow cell could be adapted to biosensors for nucleic acid and small molecule detection. Possible applications include food quality control, environmental screening, and point-of-care test devices. The SPR sensor preferably has a plurality of sensing surfaces capable of performing independent assays.

According to this invention, the flow cell is formed of layers, each comprising different functional parts needed to perform an assay. The layers include: (a) a fluid inlet layer comprising at least one inlet port in fluid contact with an external fluid transporting means;

(b) a sample-introducing layer sealed under the fluid inlet layer, comprising: i) a sample chamber having an inlet to receive a sample, and an outlet; and ii) an inlet port in fluid contact with the inlet port of the fluid inlet layer, and also in fluid contact with the inlet of the sample chamber; (c) a reaction chamber layer sealed under the sample-introducing layer, comprising: i) an inlet port in fluid contact with the outlet of the sample-introducing layer; ii) at least one reaction chamber having a precisely defined volume in fluid contact with the inlet port; iii) a first channel in fluid contact with the inlet port connected to a waste chamber through which excess sample flows out; and iv)a second channel narrower than the first channel in fluid contact with the reaction chamber and connected to an outlet through which reacted sample flows out;

(d) a fluid distributing layer sealed under the reaction chamber layer to direct the sample solution and other aqueous solutions into desired positions with a desired sequential order, comprising: i) at least one inlet port in fluid contact with the outlet of the reaction chamber layer; ii) at least one distributing port in fluid contact with the inlet port through a narrow channel; and iii) a waste chamber having an inlet through which waste sample flows in;

(e) a sealing layer comprising: i) a top face sealed onto the fluid distributing layer; ii) a bottom face sealed onto the sensing surface; and iii) a narrow flow path having an inlet in fluid contact with the distributing port of the fluid distributing layer and a outlet in fluid contact with the inlet of the waste chamber, through which reacted sample flows.

The term 'sensing surface' means an interface where two mediums of different refractive indexes meet where preferably a thin metal film is located to produce surface plasmons. Preferably, the metal film is sputtered on the side of a transparent substrate which is the higher refractive index medium. Also preferably, a thin layer of biomolecules having a specific affinity to the target analyte could be placed on the sensing surface.

Each the layer is preferably formed in a separated substrate and stacked upon one another by forming irreversible or reversible bonding between layers. In other embodiments, however, one layer may incorporate multiple functional parts mentioned above (e.g. sample receivers, reaction chambers, mixers, etc.). To measure the temperature inside the flow cell, a thermocouple may be incorporated therein. The thermocouple should be incorporated near the sensing surface and be preferably incorporated in one wall of the sealing surface where the surface is in direct contact with the fluid. In other preferred embodiments, a temperature control means may also be incorporated with the thermocouple. Temperature control means include peltier devices, metal film heaters, radiation heaters, etc. Temperature control means preferably act as a thermostat and retain the temperature set for each assay condition. In yet other embodiments, the temperature control means may be used for thermocycling reactions such as polymerase chain reactions (PCR). In this embodiment, the temperature controller is adjacent to the reaction chamber layer, and a plurality of temperature controlling means may be adapted to prevent evaporation or unwanted movement of the sample from the chamber to other layers. In yet other embodiments, the temperature control means may be used for facilitating reactions occurring at elevated temperatures. Examples of these reactions include cell lysis, denaturation of biomolecules, and melting of certain reagents.

In a preferred embodiment, the the reaction chambers contain reagents needed for the assay to be performed. In one embodiment where immunoassay is performed, the reagents may include: antibodies or antigens, antibodies or antigens conjugated with organic or inorganic substances, salts, surfactants, etc. In another embodiment where nucleic acid is to be analyzed, the reagents may include: primers, enzymes (e.g. polymerase, ligase, DNAse, RNAse, etc.), salts, surfactants, etc. The reagents are preferably dried on one of more surfaces of the reagent chamber. In other embodiments, the reagents may be incorporated in but not covalently bonded to a polymeric matrix such as poly(ethylene glycol), poly(lactic acid), poly(glycolic acid), poly(ε-caprolactone), polydioxanone, poly(maleic acid), polyanhydrides, and polysaccharides. Additional to the sealing layer, the flow cell may further incorporate a metal- sputtered transparent substrate bonded to a transparent elastomeric substrate. In this embodiment, the flow cell is connected to the external light source through the reversible bonding between the transparent elastomeric substrate and another transparent substrate, which is a component of the SPR sensor. In this manner, the flow cell also integrates the sensing surface, which may be functionalized with various biomolecules depending on the assay to be performed. The provided flow cell could be adapted to miniaturized SPR sensors such as disclosed in US patent 5912456, entitled 'Integrally formed surface plasmon resonance sensor'.

Also, another aspect of this invention is to provide a method to screen analytes in a sample solution. Following is a method disclosed herein for quantifying the analytes:

(a) forming a self assembled monolayer on the sensing surface by delivering a chemical species capable of binding on the sensing surface;

(b) immobilizing an antibody to the analyte on the self assembled monolayer; (c) contacting a sample solution expected to contain the analyte on the antibody immobilized on the self assembled monolayer;

(d) emitting light to the sensing surface and measure the change of refractive index; and

(e) quantifying the amount of analyte in the sample by comparing the change of refractive index with the change of refractive index of standard analyte solution.

Detailed description of the device and method is shown below.

BRIEF DESCRIPTION OF THE DRAWINGS Fig. 1 is a schematic illustration showing representative layers stacked to make a flow cell. Figs. 2A and 2B show a fluid inlet layer having fluid inlet ports.

Fig. 3 shows a sample-introducing layer having a sample inlet port and a sample chamber connected with three air vents on each side.

Fig. 4A shows a reaction chamber layer with one reaction chamber connected to an inlet port and a waste channel, while Fig. 4B shows a reaction chamber layer with one reaction chamber.

Fig. 4C illustrates a micropattern which is formed at the inlet of a reaction chamber in Fig. 4B.

Figs. 5A and 5B illustrates an embodiment of the first fluid distributing layer. Figs. 6A through 6C illustrate an embodiment of the second fluid distributing layer.

Fig. 7A illustrates an air bubble filter formed in part "B" of Fig. 6A, and Figs. 7B and 7C show a side view of a distributing port of the fluid distributing layer. Figs. 8A and 8B show sealing layers with three and one fluid paths, respectively. Fig. 9 shows the change of a refractive index while an antibody is flowed on a sensing surface modified with a SAM; Fig. 10 shows the change of a refractive index while different concentrations of hCG are flowed on an anti-hCG immobilized sensing surface;

Fig. 1 1 shows the change of a refractive index while serum with or without hCG is flowed on an anti-hCG immobilized sensing surface; Fig. 12 shows the change of a refractive index while different concentrations of AFP are flowed on an anti-AFP immobilized sensing surface; and Fig. 13 shows the amplification of a change of a refractive index with polyclonal antibodies.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS Preferably all the layers of the flow cell are formed within a polymeric substrate. A non-limiting list of polymers that can be used includes: silicones such as poly(dimethylsiloxane) (PDMS), acrylic polymers such as poly(methylmethacrylate) (PMMA), fluorine-containing polymers such as poly(tetrafluoroethylene) and perfluoroalkoxyalkane (PFA), polycarbonate (PC), polyesters, polyurethane (PU), polyamides, polyimides, epoxy resins, and blends or copolymers of two or more monomer units or modified polymers listed above. Also, inorganic substrates such as glass, silicon, or metallic substrates could be used.

Preferably soft lithographic techniques described by Younan Xia and George M. Whitesides (Soft Lithography, Angew. Chem. Int. Ed. 1998, 37, 550 - 575) are adapted to mold silicones such as PDMS. Also phenyl substituents of

PDMS (e.g. poly(diphenylsiloxanes), poly(phenylmethylsiloxanes), and copolymers thereof with PDMS) and other modified silicones such as mercaptosilicones could be fabricated by this technology. Also, injection molding or milling processes could be adapted to manufacture the one or more layers. The layers are in fluid contact with each other through ports described below. A hole makes the connection between the ports, and the hole preferably has a diameter between 10-1000 μm, and most preferably between 200-500 μm. The layers of the flow cell are described in detail below.

Fluid inlet layer

Figs. 2A and 2B show the bottom side of the fluid inlet layer 1 . In Fig. 2B, an inlet port 1 1 is connected to an external fluid transport means, and in Fig 2A, two inlet ports 1 1 and 12 are connected to an external fluid transport means through tubing. The tubing is made of polymeric materials (e.g. poly(tetrafluoroethylene) (PTFE), poly(etheretherketone) (PEEK), perfluoroalkoxyalkane (PFA), and silicone) and typically has an inner diameter of 500 μm. Various aqueous solutions containing phosphate, citrate, formate, acetate, carbonate, tris (tris(hydroxymethyl)aminomethane), HEPES (N-2-Hydroxyethylpiperazine-N'-2- ethanesulfonic acid), borate, and ammonia are transported to the inlet ports 11 and 12. Preferably, the aqueous solutions also contain surfactants such as Tween-20, Tween-80, Triton X-100, Triton X-114, octyl glucoside, 3-

[(Cholamidopropyl)dimethyl-ammonio]-1-propanesulfonate (CHAPS), and sodium dodecyl sulfate (SDS). Additionally, to inhibit adsorption of biomolecules onto substrate surfaces, adhesion inhibitors such as bovine serum albumin (BSA) may be added to the aqueous solution. The aqueous solution is transported into the inlet ports 11 and 12.

To transport fluid in the flow cell, the fluid inlet layer of the flow cell is connected to an external pump system. Preferably, peristaltic pumps, syringe pumps, diaphragm pumps, vacuum pumps, or piezoelectric pumps could be adapted, and most preferably a peristaltic pump is used. In this embodiment, where a peristaltic pump is used, a dampening chamber may be incorporated to the fluid inlet layer to reduce the pulsations produced from the pump. In other embodiments where two or more fluids are separately transported or one fluid is transported at a different rate through different channels, a plurality of pumps could be adapted.

The aqueous solutions are mainly used to drive the sample through the flow cell, and they may also be used to clean the sensing surface or to dilute the sample solution.

Alternatively, air instead of aqueous solutions could be used to drive the sample solution. By using air, unexpected mixing of the aqueous solutions and sample solutions inside channels or chambers could be prevented. In this embodiment, all air vents described below are non-necessary unless particularly mentioned.

In another embodiment, a plug of air followed by the aqueous solution may be used to drive the sample solution. The air plug is adapted to prevent mixing of the sample with the aqueous solution. In this embodiment, an air bubble filter described below may be adapted to prevent^' air bubbles from flowing into the sensing surface.

As shown in Fig. 2B, a fluid inlet layer 1 ' in another embodiment may have one fluid inlet port 11 connected to external fluid transporting means.

Sample-introducing layer

Fig. 3 describes the sample-introducing layer 2. The sample chamber 23 has an inlet 22 and an outlet 24 at respective ends thereof. Also, an inlet port 21 in fluid contact with the fluid inlet port 11 in the fluid inlet layer 1 , is formed at the outer side of the inlet 22.

Additionally, a flow-through hole 26 may be formed in fluid contact with the fluid inlet port 12 of the fluid inlet layer 1 , through which the aqueous solution or air may flow.

The sample is introduced into the inlet 22 by means such as a syringe or a capillary tube, and preferably with a pipette. Preferably, a predefined volume of sample in the range of between 1 - 500 μL is introduced, and most preferably, 20 - 200 μL of the sample is introduced. The inlet 22 has a diameter of 1.5 mm, and it can be larger of smaller depending on the introducing means.

Pretreatment of the sample may be done before introducing it into the inlet

22. Mixing the sample with other aqueous solutions containing reagents such as a detergent, an antibody, etc. could be done to reduce non-specific binding and to amplify the signal. Also, a separation process such as centrifugation or filtration could be done. However, the sample is preferably introduced without the above- mentioned pretreatment steps.

After the sample is introduced, the fluid inlet layer 1 is reversibly sealed onto the sample-introducing layer 2. Both layers are aligned carefully to make sure that port 21 is in fluid contact with port 11 in Fig. 2. Thus the fluid transported into the fluid inlet layer 1 is in connection with the sample, and it drives the sample into other parts of the flow cell.

The sample-introducing layer 2 may be composed of an elastomeric substrate (e.g. PDMS, PU), or the upper side of the sample-introducing layer 2 may comprise an additional layer, not shown in Fig. 3, composed of the elastomeric substrate. ι

The sample introduced into the inlet 22 is filled into the sample chamber

23. 25 indicates an air vent. There are six air vents, on both sides of the sample chamber 23, and one on each side of the ports 21 and 24. The air vents prevent the ^• formation of bubbles and enable complete filling of the sample chamber. Typically, the air vents have a channel width of 1 -20 μms, and they may have a channel width smaller than 1 μm. Channels thinner than 1 μm, however, may be difficult to fabricate. When the channel width exceeds 20 μms, the sample solution may flow out through the air vent. Additionally, the sample-introducing layer may comprise a filter to remove unwanted particles in the sample fluid. The filter may be composed of polymeric porous materials such as polysulfone, poly(vinylidene difluoride), nylon, cellulose acetate, and polyacrylate. Also, glass fibers such as borosilicate glass fibers may be used.

Reaction chamber layer

Fig. 4A and 4B show reaction chamber layers according to this invention. Fig. 4C shows a magnified view of part "A" of Fig. 4B.

According to Fig. 4A, the reaction chamber layer comprises at least one reaction chamber (32a, 32b, 32c) in which the sample could react. Reaction chambers 32a-c are connected to an inlet 31 which is in fluid contact with the outlet

24 of the sample chamber 23. At the other side of reaction chambers 32a-c, outlets 33a-c connected through a narrow channel are formed.

Also, at the inlet 31 of the reaction chambers 32a-c, another channel 36 is formed through which excess sample flows out after filling the reaction chambers 32a-c.

As shown in Fig. 4B, in other embodiments only one reaction chamber 32 could be formed in the reaction chamber layer 3'.

Additionally, a flow-through hole 39 may be formed in fluid contact with the fluid inlet port 12 of the fluid inlet layer 1 , and also with the flow-through hole 26 in the sample-introducing layer, through which the aqueous solution or air may flow.

As in the sample-introducing layer 2, air vents 38 may be formed in the reaction chamber layer 3, 3'. A plurality of air vents may be formed on each side of the reaction chambers 32, 32a, 32b, and 32c, which prevent the entrapment of air bubbles inside the reaction chambers. Various reagents can be dried in the reaction chambers 32, 32a, 32b, and

32c. The reagents include antibodies or antigens, antibodies or antigens conjugated with organic or inorganic substances, primers, enzymes (e.g. polymerase, ligase, DNAse, or RNAse), salts, and surfactants. Preferably, antibodies or antigens conjugated with metal nanoparticles or polymeric nanoparticles are used. The reagents may be incorporated in a polymeric matrix such as polyethylene glycol), poly (lactic acid), poly(glycolic acid), poly(ε-caprolactone), polydioxanone, poly(maleic acid), polyanhydrides, and polysaccharides. In another embodiment, the reagents may be adsorbed onto polymeric porous materials and then applied to the reaction chambers 32 or 32a, 32b, and 32c. The porous materials may be composed of polyolefins, polyacrylates, or polycarbonates. When a sample is driven into the reaction chambers 32 or 32a, 32b, and 32c, reagents dried in the reaction chambers 32 or 32a, 32b, and 32c are reconstituted and mixed with the sample.

To facilitate mixing, the surface of the reaction chambers 32 or 32a, 32b, and 32c may have micro-scale patterns as shown in Fig. 4C. Fig. 4C shows a magnified view of part "A" of the reaction chamber 32a of Fig. 4A. 301 indicates a micro-scale pattern where the surface is 20 μm higher that the adjacent surface 302. Alternatively, the region 301 could be rendered more hydrophobic than the region 302. Micro-scale patterns can facilitate mixing in microchannels through mechanisms described in Stroock et al., "Chaotic mixers in microchannels", Science 2002, 295, 647-651 . Additionally, temperature controlling and sensing means (not shown) may be in contact with the reaction chambers 32 or 32a, 32b, and 32c. Preferably, the temperature controlling means are used as a thermostat. In other embodiments, by elevating the temperature in the reaction chambers, mixing of reagents with the sample, cell lysis, denaturation of biomolecules, and dissolution could also be facilitated.

Using the temperature-controlling means, a thermocycling reaction such as polymerase chain reaction (PCR) may also be performed. Other nucleic acid amplification reactions such as transcription mediated amplification (TMA), nucleic acid sequence-based amplification (NASBA), strand displacement amplification (SDA), and ligase chain reaction (LCR) could also be performed.

A precise volume of sample should react with reagents in the reaction chamber 32 for reproducible and accurate quantitation of the target analyte. In Figs. 4A and 4B, a pressure barrier 35 temporarily stops the sample from flowing into ports 33, 33a, 33b, and 33c. The pressure barrier 35 has a smaller width than channels 34 and 36, and it preferably has a width of 10-100 μm. After completely filling the reaction chamber 32 or 32a, 32b, and 32c, excess sample flows into a waste chamber 59 (see Figs. 6A-C) through a port 37, which is in fluid contact with waste chamber 59.

The three reaction chambers 32a, 32b, and 32c in Fig. 4A can be filled with identical components of reagents or each can be filled with different reagents, depending on the assay format. By filling the reaction chambers 32a, 32b, and 32c with different reagents, multi-analyte detection can be performed in one assay.

The reagents may be applied on the reaction chambers 32 or 32a, 32b, and 32c by using an ink-jet printer, a microarrayer, or a screen-printer, and also by using contact printing methods. A contact printing method is described in US patent

No.5,776,748 entitled 'Method of formation of microstamped patterns on plates for adhesion of cells and other biological materials, devices and uses therefor'.

In other embodiments, multiple reaction chamber layers may be adapted to perform multi-step reaction processes for an assay.

First fluid distributing layer

Figs. 5A and B show embodiments of a fluid distributing layer. In Fig. 5A, a small fluid chamber 45 is connected to three outlets 41 a, 41 b, and 41 c. An inlet 42 is in fluid contact with the flow-through hole 39 of the reaction chamber layer 3, and it is connected to the fluid chamber 45 through a channel. Flow-through channels 46a, 46b, and 46c are in fluid contact with the outlets 33a, 33b, and 33c of the reaction chamber layer 3. Also, another flow- through channel 43 is formed in fluid contact with the outlet 37 of the reaction chamber layer 3.

In Fig. 5B, another embodiment of a first fluid distributing layer 4' wherein the fluid chamber is omitted, has an outlet 41 and a flow-through channel 43 in fluid contact with the outlet 37 of the reaction chamber layer 3.

Fluid delivered into the inlet port 12 of the fluid inlet layer 1 is distributed to outlets 41 , 41 a, 41 b, and 41 c. Preferably, aqueous solutions containing surfactants flowed in this manner wash the sensor surface and initialize the sensor. The port 43 is connected to the waste chamber 59 in Fig. 6. The sample reacted in the reaction chamber layer 3 or 3' flows through channels 46, 46a, 46b, and 46c to the fluid distributing layer 4 or 4', and is delivered to fluid distributing layer 5, 5', or 5" described below.

As in the reaction chamber layer 3, 3', air vents 44 may be formed in a fluid distributing layer 4, A'. A plurality of air vents may be formed on each side of the fluid chamber 45 and the inlet 42 and outlet 41 , 41 a, 41 b, and 41 c, which prevent the entrapment of air bubbles inside.

In yet other embodiments, the fluid chamber 45 in Fig. 5B can function as a dampening chamber to reduce pulsations in flow produced by the external fluid transport means. Without the air vent 44 shown in Fig. 5A, air is trapped in the fluid chamber 45 and effectively absorbs pulsations in flow. This embodiment is useful when a peristaltic pump is used as an external fluid transport means.

Second fluid distributing layer Fig. 6A through 6C shows an embodiment of the second fluid distributing layer. According to Fig. 6A, inlets 51a, 51 b, and 51 c are in fluid contact with outlets 33a, 33b, and 33c of the reaction chamber layer 3, and they are connected to the distributing ports 52a, 52b, and 52c through a narrow channel. Between each inlet 51 a, 51 b, and 51 c and distributing port 52a, 52b, and 52c, other inlets 53a, 53b, and 53c, in fluid contact with the outlet 41 a, 41 b, and 41 c of the fluid distributing layer 4, are formed.

The distributing ports 52a, 52b, and 52c are in fluid contact with fluid paths

61 a, 61 b, and 61 c of a sealing layer 6, through which the sample flows. The inlets

53a, 53b, and 53c are in fluid contact with the outlets 41a, 41 b, and 41c of the fluid distributing layer 4, and the fluid in the fluid chamber 45 flows therein. Preferably, a solution containing a surfactant flows through the inlets 53a, 53b, and 53c.

A channel 57 connecting inlets 51 a, 51 b, and 51 c and 53a, 53b, and 53c preferably has a smaller width than the channel 58 which connects inlets 53a, 53b, and 53c and distributing ports 52a, 52b, and 52c to prevent back-flow (i.e. flow from port 53a, 53b, and 53c to 51 a, 51 b, and 51 c) of solution.

As in the first fluid distributing layer 4 and A', air vents may be formed in the second fluid distributing layer 5, 5', 5". A plurality of air vents may be formed on each side of inlets 51 a, 51 b, and 51 c, inlets 53a, 53b, and 53c, distributing ports

52a, 52b, and 52c, and waste chamber 59, and they prevent the entrapment of air bubbles inside.

The fluid distributing layer described above is an embodiment wherein the sample is delivered to three sensing surfaces simultaneously, while in Fig. 6B, the sample is delivered to each of the sensing surfaces sequentially. The second fluid distributing layer 5' is connected to the reaction chamber layer 3' in Fig. 4B and the fluid distributing layer 4' in Fig. 5B, and a sample reacted in, one reaction chamber is delivered sequentially to three sensing surfaces. This may be embodied through the fluid distributing layer 5, but the fluid distributing layer 5' may be easier to fabricate.

The second fluid distributing layer 5" is connected to the reaction chamber layer 3' in Fig. 4B and the fluid distributing layer A'. This layer is adapted into SPR sensors with one sensing surface. '^■

An important feature of the fluid distributing layer 5 is the air bubble filter incorporated at ports 52a, 52b, and 52c, and it is shown in detail in Fig. 7A.

Fig. 7A shows a magnified version of part "B" in Fig. 6A. Figs. 7B and 7C shows a side view of ports 52a, 52b, and 52c. In Fig. 7A, region 71 is a hydrophobic zone to trap air bubbles and drain into air vents 73. Preferably, region 72 is relatively more hydrophilic than region 71 , and the air bubble-filtered sample or solution flows into region 72.

In one embodiment, the hydrophobic zone 71 can be formed by methods described in Fig. 7B and Fig. 7C.

In Fig. 7B, the fluid distributing layer 5 is composed of hydrophobic material such as PDMS, PFA, and PTFE, and the hydrophobic zone 71 is formed by narrowing the fluid path.

In Fig. 7C, the hydrophobic zone 71 is formed by patterning hydrophobic materials on the surface.

In other embodiments, the air bubble filter composed of a polymeric matrix may be adapted to the ports 52a, 52b, and 52c. Preferably, a porous polymeric matrix composed of polymers such as polysulfone, poly(vinylidene difluoride), nylon, cellulose acetate, and polyacrylate is adapted.

Sealing layer Figs. 8A and 8B show the sealing layer according to this invention.

In Fig. 8A, sealing layer 6 has three flow paths 61a, 61 b, and 61c that are in fluid contact with the distributing ports 52a, 52b, and 52c in the second fluid distributing layer 5. The inlets of the flow paths 61a, 61 b, and 61c are in fluid contact with the distributing ports 52a, 52b, and 52c, and the outlets are connected to the inlets 56a, 56b, and 56c of the waste chamber 59.

The reacted sample flows from the inlets 52a, 52b, and 52c to the flow paths 61a, 61b, and 61c, and is collected at the waste chamber 59.

As shown is Fig. 8A, the sealing layer 6 having three flow paths could be sealed onto an SPR sensor with at least three sensing surfaces. Fig. 8B shows a fluid distributing layer 6' to be adapted to the sensor with only one sensing surface.

The flow paths 61 and 61 a, 61 b, and 61 c typically have a channel width of between 10-500 μm, and preferably between 100-300 μm. The distance between the flow paths is 50-1000 μm, and preferably between 100-300 μm.

Preferably, temperature sensing means (not shown in Figs. 8A and B) are incorporated into at least one flow path 61 or 61a, 61 b, and 61 c in the sealing layer

6 or 6'. The temperature sensing means include a thermocouple, a thermister, and a resistance temperature detector (RTD). Preferably, a thermocouple is incorporated in one wall of a flow path.

Flow cell further incorporating sensing surfaces

Additionally, the flow cell may further incorporate a metal-sputtered transparent substrate. The transparent substrate is preferably sealed onto the sealing layer 6 or 6' by reversible bonding (physical adsorption), and another elastomeric transparent substrate, preferably composed of phenyl substituents of PDMS (e.g. poly(diphenylsiloxanes), poly(phenylmethylsiloxanes), and copolymers thereof with PDMS) is additionally bonded to the metal-sputtered transparent substrate.

In this embodiment, the SPR sensor only provides an external light source and a transparent medium (e.g. glass, epoxy resin, etc.) through which light is emitted. The transparent elastomeric substrate is reversibly bonded (via physical adsorption) onto the transparent medium of the SPR sensor.

According to this embodiment, by replacing the flow cell, the sensing surface could also be replaced. This feature provides a more cost effective means for sensing and can give more reproducible data since variations coming from replacement of the external light source could be eliminated. By using the integrated flow cell disclosed herein, each functional part is formed in each layer enabling the analysis of various samples by adapting different layers as needed. Each layer may be easily modified as the assay condition requires.

By designing each component of the flow cell, the assay condition is determined and commercialized tests which require the same experimental condition every time could be preformed efficiency with good reproducibility.

Also, this flow cell makes the screening system portable and requires an analysis time of less than 30 min which may be useful in emergency situations.

Additionally, polymeric components have low costs, thus the flow cell could be disposable. Hazardous analytes could be easily abandoned.

Method of screening analyte

When light is emitted on an interface where two mediums with different refractive indexes meet, one portion of the light transverses the interface and the other portion is reflected from the interface. When light reaches the interface at a certain incident angle (determined by the ratio of the refractive index of two mediums) from the medium with a higher refractive index to the lower refractive index medium, the total amount of light is reflected from the interface. As the light is reflected, an evanescent wave is produced at the interface which penetrates a short distance (about 200 nm) into the medium.

When a thin metal layer is placed at the interface, and the incident light is p-polarized, free electrons of the metal can absorb the evanescent wave and produce surface plasmon. This phenomenon, termed surface plasmon resonance (SPR), causes the intensity of reflected light to be reduced and the amount of absorbed light is dependent on the angle of incidence. The incident angle where a minimum intensity of reflected light can be observed is often called the SPR angle.

The SPR angle changes with the refractive index near the sensing surface.

If a receptor having a specific affinity to the target analyte is immobilized on the sensing surface, the target analyte near the sensing surface will bind on the receptor and cause a change in refractive index near the sensing surface. Since this change of refractive index corresponds to the change of the SPR angle, monitoring the change of SPR angle will give information on the amount of analyte in a sample solution.

The analyte and receptor respectively may be a protein, a chemical, DNA or RNA and it is preferably a protein. The sensing surface is a metal selected from the group consisting of gold, copper, silver, aluminum, chromium, germanium, an alloy thereof, and an oxide thereof.

Also, a protein having a specific affinity to the target analyte can be immobilized on the sensing surface. The protein may be monoclonal antibody, a polyclonal antibody, or an antigen.

For amplifying the signal, a signal amplifier can be used. The signal amplifier can be selected from the group consisting of an antibody, an antigen, an antibody conjugated with an organic or inorganic substance, and an antigen conjugated with an organic or inorganic substance. The antibody can be monoclonal or polyclonal and a concentration of the antibodies is from 5 to 100 βg/xxixl, and preferably from 5 to 20 μg/xxxt. Additionally, a second antibody, also having a specific affinity for the analyte could be contacted to the previously bound antibody-analyte complex to amplify the signal by causing an additional change in the refractive index The organic or inorganic substances are preferably spherical and have a diameter of 1 -200 nm, and more preferably a diameter of 1 - 50 nm. The inorganic substances include a metal such as gold, silver, copper, aluminum, or alloys thereof, and also oxides of metals such as alumina, titanium dioxide, and other inorganic substances such as glass, silicon, and silica. The organic substances include synthetic polymers such as polyethylene, polystyrene, polyethylene glycol, polyurethane, polyacrylic acid, polymethyl methacrylate, or their copolymers, as well as natural macromolecules such as horseradish peroxidase (HRP).

Preferably, portable miniaturized SPR sensors are used to measure the change of the refractive index. The SPR sensors preferably comprise a sensing surface, a light emitting diode (LED) light source, a mirror, and a photodiode array. The LED has a wavelength of 800-850 nm, which could be adjusted. The LED and the photodiode array are incorporated in a single chip having a housing, and the signal from the chip is processed by a digital signal processor (DSP). An embodiment of the miniaturized SPR sensor is disclosed in US patent 5912456.

In one embodiment, the analytes comprise protein biomarkers, and they are preferably proteins of which their concentration in bodily fluids changes according to specific diseases. The analytes can be alpha-fetoprotein (AFP), human chorionic gonadotropin (hCG), and unconjugated estriol (uE3).

Concentrations of AFP in blood according to various diseases are shown in Table 1 (From I.T. Seo, Clinical Nuclear Medicine Methods, Korea Medicine).

Table 1.

Table 2 shows symptoms resulting in an increase of hCG in various forms in blood (See Alfthan, H. and Stenman, U. Pathophysical importance of various molecular forms of human choriogonadotropin, Mol. Cell. Endocrinol. 1996, 125, 107-120).

Table 2.

In healthy males and females, concentrations of hCG and hCGβ in serum is as shown in Table 3 (See Alfthan, H. et al., Clin. Chem., 1992, 38, 1981-1987). In healthy females, concentration of hCG in blood is as shown in Table 4 (From I.T. Seo, Clinical Nuclear Medicine Methods, Korea Medicine). Table 3.

Estriol is a well-known estrogen hormone synthesized at the placenta and found in the blood during pregnancy. Unconjugated estriol (uE3) is one form of estriol commonly found in blood. uE3 of the fetus that flows into maternal blood usually exists for only 1 hr and is then degraded. In normal cases, the concentration of uE3 increases during pregnancy. However, in pregnancies where the fetus has Down syndrome, the level of uE3 is lower (Canick, J.A. et al. J, Obstet. Gynecol. 1988, 95, 330 - 333). In normal pregnancies, the concentration of uE3 in maternal blood is as shown in Table 5 (See Katagiri, H. et al., Amer. J. Obstet Gynecol., 1976, 124, 272 - 280; Buster, J.F. et al., Amer. J. Obstet. Gynecol., 1976, 125, 672).

Table 5

Therefore, by quantifying AFP, hCG, and uE3 in bodily fluids, diagnosis of certain diseases would be possible.

In an embodiment, biomolecules having a specific affinity to a target analyte (e.g. an antibody) were immobilized on the sensing surface following steps comprising forming a self-assembled monolayer (SAM) on the sensing surface by reacting chemical species capable of binding on the sensing surface, and immobilizing the biomolecule on the SAM.

The SAM allows a dense layer of the biomolecules to form on the sensing surface, and in other embodiments, three-dimensional structures could further increase the binding of the biomolecules.

The chemical comprises a moiety capable of forming a non-covalent bond with a positively charged group of a molecule and a moiety capable of binding on the sensing surface. Chemicals capable of binding on the sensing surface includes trichlorosilanes, fatty acids, or compounds having at least one thiol group (See Flink, S. et al. Sensor functionalities in Self-Assembled Monolayers, Adv. Mat,

2000, 12, 1315 - 1328). Additionally, these compounds preferably have other functional groups such as carboimide groups, acetic acid groups, or calixarene derivatives.

The calixarene derivatives can form the SAM though thiol groups. A calixarene-crown derivative specifically recognizes amine groups of biomolecules and forms a stable non-covalent bond via a π-π interaction and hydrogen bonding. Thus, a calixarene-crown derivative enables immobilization of biomolecules onto the SAM without further processes (i.e. chemical modifications). A detailed description of the compound is disclosed in EP 1110964, entitled 'Calixcrown derivatives for immobilizing proteins'.

SAMs could also be formed on the sensing surface by reacting thiol- containing compounds such as mercaptohexadecanoic acid, mercaptoundecanoic acid, and 1 ,6-mercaptohexadecylamine. The calixarene-crown derivatives include calix[4]arene-biscrown-4 shown in formula 1 , calix[4]arene-crown-5 shown in formula 2, and calix[4]arene-crown-6 shown in formula 3.

(Formula 1)

wherein the R1 , R2, R3, and R4 independently represent thiol groups such as -CH₂SH, respectively, or a pair of each side chain can form a disulfide bond, i.e. -CH₂-S-S-CH₂, respectively; or R1 , R2, R3, and R4 independently represent -CH₂CI, -CH₂CN, -C H₂CHO, CH₂N H₂, or -CH₂COOH. (Formula 2)

wherein n is 1 and R1 , R2, R3, and R4 independently represent -CH₂SH, - CH₂CI, -CH₂CN, -CH₂CHO, CH₂NH₂, or -CH₂COOH; or

R1 and R3 represent -CH₂SH, -CH₂CI, -CH₂CN, -CH₂CHO, CH₂NH₂, or - CH₂COOH, R2 and R4 represent H, and R5 and R6 represent H, methyl, ethyl, propyl, isopropyl, or isobutyl, respectively.

(Formula 3)

wherein n is 2 and R1 , R2, R3, and R4 independently represent -CH₂SH, - CH₂CI, -CH₂CN, -CH₂CHO, CH₂NH₂, or -CH₂COOH; or

R1 and R3 represent -CH₂SH, -CH₂CI, -CH₂CN, -CH₂CHO, CH₂NH₂, or - CH₂COOH, R2 and R4 represent H, and R5 and R6 represent H, methyl, ethyl, propyl, isopropyl, or isobutyl, respectively.

Various immunoassay methods known in the prior art could be adapted to methods disclosed herein. A signal from an optical sensor is converted into a refractive index, and a change of the refractive index is compared to a standard change of a refractive index at a known concentration of an analyte to quantify the analyte in a sample solution.

In one embodiment, the present invention provides a method for quantifying a biomarker for Down syndrome, comprising: (a) forming a self- assembled monolayer on sensing surface with a chemical species capable of binding on the sensing surface (b) immobilizing an antibody for an analyte on the self-assembled monolayer; (c) contacting a sample solution expected to contain the analyte on the antibody immobilized on the self-assembled monolayer; (d) emitting light to the sensing surface and measuring the change of refractive index; and (e) quantifying the amount of analyte in the sample by comparing the change of refractive index with the change of refractive index of a standard analyte. The method further comprises mixing the sample solution with a signal amplifier before the (c) step, or reacting a signal amplifier on the SAM after the (c) step.

The biomarker is preferably selected from the group consisting or alpha- fetoprotein, human chorionic gonadotropin and unconjugated estriol.

Concentrations of hCG, AFP, and uE3 in bodily fluids represent important information in diagnosing disorders such as Down syndrome. By quantifying biomarkers such as hCG, AFP, and uE3, the status of patients could be monitored.

The following examples are intended to further illustrate the present invention. However, these examples are shown only for better understanding of the present invention without limiting its scope.

Example 1 : SAM formation on the sensing surface by flowing

As an SPR sensor, a Spreeta™ sensor (Texas Instruments, Dallas) was used. The Spreeta™ sensor is a miniaturized SPR sensor which incorporates an LED, a polarizer, a mirror, a photodiode array, and a sensing surface in one housing. A signal from the sensor is processed by a digital signal processor connected to a computer.

The sensing surface was washed with 0.1 % TritonX-100 (Sigma) and 0.1 M NaOH solution by flowing for at least 10 min, and dried under nitrogen gas. Then the sensing surface was further incubated in a piranha solution (sulfuric acid: hydrogen peroxide = 1 : 1 , v/v) for at least 10 min, washed in a stream of distilled water, and dried under nitrogen gas. A flow cell provided by Texas Instruments (a different flow cell from what is disclosed herein) was assembled on the Spreeta™ sensor, and a 0.1 % buffer containing TritonX-100 (Sigma), a 0.1 M NaOH solution, and distilled water was alternately flowed at least 3 times for further washing of the sensing surface. All flow rates were 50 μL/min during the above process. Prolinker™ B (Proteogen, Seoul), which is a calix[4]arene-crown derivative, was dissolved in DMF (N,N-Dimethylformamide, Sigma) to a concentration of 1 .5 mM, and flowed on the sensing surface at a flow rate of 7 μL/min for 30 min using a syringe pump (Keon-A information technology, Co., Seoul). After flowing, the sensing surface was further incubated for 30 min and washed with DMF at 50 μL/min for 10 min. The flow cell was dissembled and the sensing surface was washed in a stream of distilled water, and then dried under nitrogen gas.

Example 2: SAM formation on the sensing surface by immersion

The sensing surface (~50 nm gold layer on borosilicate glass) was washed by the same method as described in Example 1 , then it was immersed in DMF containing 1.5 mM Prolinker™ B (Proteogen, Seoul) for 30 min.

The gold and chromium layer of the Spreeta™ sensor was removed to completely clean the borosilicate glass by dipping it in nitrohydrochloric acid (nitric

^' acid : hydrochloric acid = 1 : 3, v/v) for 30 s. The sensing surface modified with the

SAM was attached to the glass of the Spreeta™ sensor by applying an index matching liquid (Triallyl trimellitate, Sigma).

Example 3: Immobilization of antibodies

An antibody (anti-AFP or anti-hCG) was immobilized on the SAM modified sensing surface prepared in Example 1. A flow cell provided by Texas Instruments was assembled onto the sensing surface, and PBS-Triton (0.1% Triton X-100 in phosphate buffered saline, pH 7.2) was flowed over the sensing surface for 10 min. An antibody (Monoclonal Anti-AFP, Sigma; Monoclonal Anti-hCG, Calbiochem) dissolved in PBS to a concentration of 50 μg/-nX was then flowed over the sensing surface at 30 βi/πxm. The change of the refractive index during this procedure was monitored.

Fig. 9 depicts the change of the refractive index while the antibody was binding onto the SAM. The binding occurs via a π-π interaction and hydrogen bonding between amine groups of the antibody and the SAM. It can be seen from the refractive index change after 45 min that during this time most of the binding occurs. PBS-T (0.01% TritonX-100 in PBS, pH 7.2) was further flowed for 10 min to wash weakly bound antibodies. The sensing surface was dried under nitrogen gas or used immediately for the following assay.

Example 4: Quantitative measurement of hCG in buffer solution By using the anti-hCG immobilized sensing surface prepared in Example 3, a quantitative measurement of hCG was performed. On the Spreeta™ sensor, PBS-T containing hCG (Calbiochem) was flowed at 30 /- /min for 15 min, followed by a washing step with PBS-T. The concentration of hCG was set to 5, 20, 40, 80, and 160 ng/inϋ by diluting in PBS-T. The change of refractive index after flowing for 15 min is shown in Fig. 10. It can be seen that the refractive index increases as the concentration of hCG in PBS-T increases.

Example 5: Detection of hCG in serum

A blood sample obtained from a healthy male (23 years old) was centrifuged to obtain serum, and it was diluted 100-fold in PBS-T. The diluted serum was flowed by the same method as described in Example 4, and the change of the refractive index was monitored.

In another experiment, hCG was added to the diluted serum, and then flowed at 30 μJt/ \n onto the antibody immobilized sensing surface. Fig. 11 shows the difference between the change of refractive index of the two serums. The increase of refractive index when serum (no hCG added) is flowed can be considered to come about as a result of non-specific binding of serum proteins on the sensing surface, since a very low concentration of hCG (less than 0.08 ng/m£) is contained in healthy male serum. The larger change of the refractive index as shown in Fig. 11 is due to specific binding of hCG with the immobilized antibody.

This shows the capability of the sensor for quantitative measurement of hCG in serum.

Example 6: Quantitative measurement of AFP in buffer solution

AFP in PBS-T was quantified by the same method as described in Example 4.

AFP (Biodesign) was diluted in PBS-T to concentrations of 10, 20, and 80 ng , and flowed onto the sensor surface at 30 ft/min for 15 min. Fig. 12 depicts the change of the refractive index after 15 min. It can be seen that the change of refractive index increases as the concentration of AFP in PBS-T increases.

Example 7: Signal amplification using polyclonal antibodies Polyclonal antibodies may further bind on analytes bound on antibodies immobilized on a surface since polyclonal antibodies can recognize many different epitopes on the analyte.

Following the method in Example 3, polyclonal anti-BSA (anti-bovine serum albumin, Sigma) was immobilized on the SAM surface. 0.1 ng/inϋ of BSA (Sigma) was flowed onto the immobilized polyclonal anti-BSA at 30 -1/min. No detectable change of the refractive index was measured, as shown in Fig. 13.

Then, 50 μglxxi of polyclonal anti-BSA in PBS-T was flowed at 30 μt/rn^'m. Fig. 13 depicts the change of the refractive index during this procedure. Considering that almost no binding occurs when 50 μg/xxύ of polyclonal anti-BSA in PBS-T was flowed on the anti-BSA immobilized sensing surface, it can be seen from Fig. 13 that by using polyclonal anti-BSA, the change of the refractive index from 0.1 ng/mX BSA was amplified more than 10-fold.

While this invention has been described in connection with what is presently considered to be the most practical and preferred embodiment, it is to be understood that the invention is not limited to the disclosed embodiments, but, on the contrary, is intended to cover various modifications and equivalent arrangements included within the spirit and scope of the appended claims.

Claims

WHAT IS CLAIMED IS:

1. An integrated flow cell to be used in a surface plasmon resonance sensor, comprising:

(a) a fluid inlet layer comprising at least one inlet port in fluid contact with an external fluid transporting means;

(b) a sample-introducing layer sealed under the fluid inlet layer, comprising: i) a sample chamber having an inlet to receive a sample, and an outlet; and ii) an inlet port in fluid contact with the inlet port of the fluid inlet layer and also in fluid contact with the inlet of the sample chamber; (c) a reaction chamber layer sealed under the sample-introducing layer, comprising: i) an inlet port in fluid contact with the outlet of the sample-introducing layer; ii) at least one reaction chamber having a precisely defined volume in fluid contact with the inlet port; iii) a first channel in fluid contact with the inlet port connected to the waste chamber through which excess sample flows out; and iv) a second channel that is narrower than the first channel, in fluid contact with the reaction chamber and connected to an outlet, through which reacted sample flows out;

(d) a fluid distributing layer sealed under the reaction chamber layer to direct the sample solution and other aqueous solutions into desired positions with a desired sequential order, comprising: i) at least one inlet port in fluid contact with the outlet of the reaction chamber layer; ii) at least one distributing port in fluid contact with the inlet port through a narrow channel; and iii) a waste chamber having an inlet through which a waste sample flows in; and (e) a sealing layer comprising: i) a top face sealed onto the fluid distributing layer; ii) a bottom face sealed onto the sensing surface; and iii) a narrow flow path having an inlet in fluid contact with the distributing port of the fluid distributing layer and an outlet in fluid contact with the inlet of the waste chamber through which reacted sample flows.

2. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , wherein the fluid which enters the fluid inlet layer is air.

3. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising reagents dried within the reaction chamber.

4. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 3, wherein the reagents are incorporated in a polymeric matrix.

5. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 3, wherein the reagent comprises antibodies.

6. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 5, wherein the antibodies are monoclonal antibodies.

7. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 5, wherein the antibodies are polyclonal antibodies.

8. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 5, wherein the antibodies are conjugated to organic or inorganic substances.

9. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 3, wherein the reagent comprises nucleic acids.

10. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 3, wherein the reagent comprises surfactants.

11. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising temperature controlling means and temperature sensing means in contact with the reaction chamber.

12. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising a micro-patterned structure within the reaction chamber.

13. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising an air bubble filter within the fluid distributing layer.

14. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising temperature sensing means in contact with at least one of the flow paths in the sealing layer.

15. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , wherein at least one of the layers further comprises an air vent channel.

16. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 1 , further comprising a second reaction chamber layer sealed under the reaction chamber, comprising: an inlet port in fluid contact with the inlet of the reaction chamber layer; at least one reaction chamber having a precisely defined volume in fluid contact with the inlet port; a first channel in fluid contact with the inlet port connected to the waste chamber through which excess sample flows out; and a second channel narrower than the first channel in fluid contact with the reaction chamber and connected to an outlet through which reacted sample flows out;

17. An integrated flow cell to be used in a surface plasmon resonance sensor comprising:

(a) a fluid inlet layer comprising at least one inlet port in fluid contact with an external fluid transporting means;

(b) a sample-introducing layer sealed under the fluid inlet layer, comprising: i) a sample chamber having an inlet to receive a sample, and an outlet; and ii) an inlet port in fluid contact with the inlet port of the fluid inlet layer and also in fluid contact with the inlet of the sample chamber;

(c) a reaction chamber layer sealed under the sample-introducing layer, comprising: i) an inlet port in fluid contact with the outlet of the sample-introducing layer; ii) at least one reaction chamber having a precisely defined volume in fluid contact with the inlet port; iii) a first channel in fluid contact with the inlet port connected to a waste chamber, through which excess sample flows out; and iv) a second channel narrower than the first channel in fluid contact with the reaction chamber and connected to an outlet through which reacted sample flows out;

(d) a fluid distributing layer sealed under the reaction chamber layer to direct the sample solution and other aqueous solutions into desired positions with a desired sequential order, comprising: i) at least one inlet port in fluid contact with the outlet of the reaction chamber layer; ii) at least one distributing port in fluid contact with the inlet port through a narrow channel; and iii) the waste chamber having an' inlet through which the waste sample flows in; and

(e) a sealing layer comprising: i) a top face sealed onto the fluid distributing layer; ii) a bottom face sealed onto the sensing surface; and iii) a narrow flow path having an inlet in fluid contact with the distributing port of the fluid distributing layer, and a outlet in fluid contact with the inlet of the waste chamber through which reacted sample flows.

(f) a first rigid transparent substrate having a top surface whereon metal film is sputtered and sealed onto the bottom face of the sealing layer; and (g) a second elastomeric transparent substrate sealed onto the first transparent substrate.

18. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 17, wherein the elastomeric transparent substrate is composed of poly(diphenylsiloxanes) or their copolymers with poly(dimethylsiloxanes).

19. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 17, wherein the elastomeric transparent substrate is composed of poly(phenylmethylsiloxanes) or their copolymers with poly(dimethylsiloxanes).

20. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 17, wherein the elastomeric transparent substrate is composed of copolymers of poly(diphenylsiloxanes) and poly(phenylmethylsiloxanes), or copolymers of poly(diphenylsiloxanes), poly(phenylmethylsiloxanes), and poly(dimethylsiloxanes) .

21. The integrated flow cell to be used in a surface plasmon resonance sensor of claim 17, wherein the elastomeric transparent substrate is composed of mercaptosilicones.

22. A method for quantifying a biomarker for Down syndrome comprising:

(a) forming a self-assembled monolayer on a sensing surface with a chemical species capable of binding on the sensing surface;

(b) immobilizing an antibody for an analyte on the self-assembled monolayer; (c) contacting a sample solution expected to contain the analyte on the antibody immobilized on the self-assembled monolayer;

(d) emitting light to the sensing surface and measuring the change of refractive index; and

(e) quantifying the amount of analyte in the sample by comparing the change of the refractive index with the change of a refractive index of a standard analyte.

23. The method of claim 22, wherein the sensing surface is selected from the group consisting of gold, copper, silver, aluminum, chromium, germanium, an alloy thereof, and an oxide thereof.

24. The method of claim 22, wherein the chemical species comprise a moiety capable of forming a non-covalent bond with a positively charged group of a molecule and a moiety capable of binding on the sensing surface.

25. The method of claim 22, wherein the chemical species comprise a thiol group and a crown ring.

26. The method of claim 22, wherein the light is produced from a light emitting diode (LED) and has a wavelength of 800 to 850 nm.

27. The method of claim 26, wherein the LED is incorporated within a photodiode array in a single housing.

28. The method of claim 22, wherein the biomarker is alpha-fetoprotein, human chorionic gonadotropin, or unconjugated estriol.

29. The method of claim 22, wherein the antibody is a monoclonal or polyclonal antibody.

30. The method of claim 22 further comprising a step of reacting a signal amplifier on SAM after the (c) step.

31. The method of claim 22 further comprising a step of mixing the sample solution with a signal amplifier before the (c) step.

32. The method of claim 30 or 31 , wherein the signal amplifier is antibodies or antibodies conjugated with organic or inorganic substances.