WO2008130327A1

WO2008130327A1 - Biointerfaces for biomolecule detection

Info

Publication number: WO2008130327A1
Application number: PCT/SG2007/000108
Authority: WO
Inventors: Hsiao-Hua Yu; Jackie Y. Ying; Shyh-Chyang Luo; Hong Xie; Naiyan Chen
Original assignee: Agency for Science Technology and Research Singapore
Current assignee: Agency for Science Technology and Research Singapore
Priority date: 2007-04-20
Filing date: 2007-04-20
Publication date: 2008-10-30
Anticipated expiration: 2009-10-20

Abstract

Provided herein is a sensor surface for use in detecting a biological molecule. The sensor surface comprises an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film. The capture agent is able to selectively bind to the biological molecule. Also provided is a sensor system for detecting a biological molecule. The sensor system comprises a sensor surface as described above and a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding.

Description

Biointerfaces for Biomolecule Detection Technical Field

The present invention relates to sensors using functionalized electrically conductive films for the fluorescent or electrochemical detection of biological molecules, such as oligo- or polynucleotides. The films may be used in high throughput applications, such as in biosensory devices and biochip sensors.

Background of the Invention

There is an ongoing need for the ability to detect the presence of specific biologically-derived molecules, such as peptides or oligo- and polynucleotides, from amongst complex mixtures of these molecules. The completion of the human genome project, for example, provides new dimensions for clinical diagnostics and drug development based on the understanding of specific DNA sequences. As a result, there is a growing need for bio-sensing platforms with uses ranging from genotyping to molecular diagnostics. Northern blotting, ribonuclease protection, and reverse transcription- polymerase chain reaction (RT-PCR) are the most commonly used methods in gene expression analysis. However, each of these methods suffers from insufficient sensitivity to false signalling from amplified contaminants. Recently, there is a growing interest in developing electrochemical DNA biosensors due to their cost-effectiveness, the efficiency with which they may be produced using modern semiconductor fabrication processes, the high sensitivity they may provide upon signal amplification through electrocatalysis, the rapid and direct detection they can provide even in the presence of light-absorbing chemicals, and the ease of manufacturing of portable, robust, low-cost and easy-to-handle detection instrumentation suitable for field tests and home-care usage.

A key component for an electrochemical biosensor is efficient and convenient construction of conductive biointerfaces. The use of a self-assembled monolayer (SAM) is a popular method for biointerface construction due to its simplicity, chemical availability, and the ability to produce thin, uniform films. The present SAM methods may be hampered by the limited selection of grafted electrode surfaces, the technical requirement of an extremely clean electrode surface, which controls the smoothness and complete surface coverage of SAM, the extended time required for immobilization, the poorly spatial control of bioconjugation when in contact with a probe solution, and the difficulty in large-scale manufacturing. For these and other reasons, biosensor platforms based on SAM have produced inconsistent results and have been difficult to mass- produce. It is the object of the present invention to substantially overcome or at least ameliorate one or more of the above disadvantages. Provided herein is the use of functionalized polymeric, e.g. polyethylenedioxythiophene (PEDOT), biointerfaces for luminescent and electrochemical detection of biological molecules, such as oligonucleotides, polynucleotides, peptides and polypeptides.

Summary of the Invention

In a first aspect of the invention there is provided a sensor surface for use in detecting a biological molecule, said sensor surface comprising: i) an ultrasmooth film of electrically conductive polymer and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule. hi an embodiment, the ultrasmooth film of electrically conductive polymer has a surface root-mean-square roughness R_πns of less than or equal to about lOnm.

The polymer may comprise at least one type of monomer unit which is derived from a derivative of ethylendioxythiophene (EDOT). The derivative of EDOT may comprise an EDOT group coupled to a functional group. Suitable functional groups include a hydroxyl group, a thiol group, an amine group, a carboxylic acid group, an azide group, an N-hydroxysuccinimidoester group and a maleimido group. The capture agent may be coupled to the film by means of an amide linkage, an ester linkage, a 1,2,3-triazole linkage or a succinimido linkage or some other suitable linkage. The polymer may comprise more than one type of monomer unit (i.e. it may be a copolymer). More than one of the monomer units in the copolymer may be as described above. Other monomer units (e.g. derived from EDOT, pyrrole, thiophene, furan etc.) may also be present. In this context, the term "derived from" does not necessarily indicate that the monomer unit was actually formed from the stated compound, only that it could have been formed therefrom. In some embodiments the polymer comprises monomer units derived from

EDOT-OH and monomer units derived from C₂-EDOT-COOH.

The film may be generated by electropolymerisation. It may be generated by electrocopolymerisation. The film may be prepared by electropolymerisation in a microemulsion. The film may be on the surface of an electrode. The electrode may be an electrode used in the electropolymerisation. The electrode may be an ultrasmooth electrode. It may have a surface root-mean-square roughness R_n^ of less than or equal to about IOnm, or less than or equal to 8, 5 or 3nm.

The capture agent may be coupled to the film by means of an amide linkage, an ester linkage, a 1,2,3-triazole linkage or a succinimido linkage or some other suitable linkage. The capture agent may comprise an oligonucleotide or a polynucleotide. The capture agent may comprise a peptide, oligopeptide or polypeptide. It may comprise an antibody, an antibody fragment, a peptide based on the complemetarity determining region of an antibody, a ligand or a ligand-binding peptide. The film of the sensor surface may have a surface root-mean-square roughness R^ of less than or equal to about 5nm. It may have a thickness of less than or equal to about 100 nm.

In an embodiment there is provided a sensor surface for use in detecting a biological molecule, said sensor surface comprising: i) an ultrasmooth firm of electrically conductive copolymer, said copolymer comprising monomer units which are derived from a derivative of EDOT, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule. hi another embodiment there is provided a sensor surface for use in detecting a biological molecule, said sensor surface comprising: i) an ultrasmooth film of electrically conductive copolymer, said copolymer comprising monomer units which are derived from a derivative of EDOT, and ii) a capture agent coupled to the film by means of an amide linkage, an ester linkage, a 1,2,3-triazole linkage or a succinimido linkage, wherein the capture agent is able to selectively bind to the biological molecule.

In a second aspect of the invention there is provided a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding.

The detectable signal may be an electrical signal. It may be a fluorescent signal. It may be some other suitable signal. The strength or intensity of the signal may be dependent on, e.g. proportional to, the concentration of the biological molecule. It maybe dependent on the concentration of the biological molecule bound or coupled to the coupling agent. It may be dependent on the concentration of the biological molecule in a liquid to which the sensor surface is exposed. The label may be capable of coupling to the biological molecule. In an embodiment, the label is coupled to the sensor surface. It may be coupled to the capture agent. It may be coupled to the film. hi another embodiment the label is not coupled to the sensor surface prior to the binding of the biological molecule to the capture agent. The film of electrically conductive polymer may be on the surface of an electrode. hi an embodiment there is provided a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer on the surface of an electrode, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding. hi another embodiment there is provided a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer on the surface of an electrode, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label capable of binding to the biological molecule, and which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding. hi another embodiment there is provided a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer on the surface of an electrode, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label coupled to the sensor surface and capable of binding to the biological molecule, said label being able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding. hi another embodiment there is provided a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer on the surface of an electrode, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label capable of binding to the biological molecule, and which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding, wherein, prior to the binding of the biological molecule to the capture agent, the label is not coupled to the sensor surface.

In a third aspect of the invention there is provided a method of detecting a biological molecule comprising: i) providing a sensor surface, said sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) exposing said sensor surface to the biological molecule in the presence of a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal in response to said binding; and iii) detecting the detectable signal. The detectable signal may be an electrical signal It may be a fluorescent signal. It maybe some other type of signal.

In an embodiment, step ii comprises the steps of combining the biological molecule with the label to form a labelled sample, and exposing the sensor surface to the labelled sample. In this embodiment, the label is not coupled to the sensor surface prior to step ii. In another embodiment, the label is coupled to the sensor surface prior to step ii.

In another embodiment, step ii comprises: ii') exposing said sensor surface to the biological molecule so as to bind the biological molecule to the capture agent; and ii") exposing the bound biological molecule to the label, thereby producing a detectable signal.

In this embodiment, the label is not coupled to the sensor surface prior to step ii.

The method of the third aspect may be a method for quantifying the biological molecule. In this case step iii comprises quantifying the detectable signal. The film of electrically conductive polymer may be on the surface of an electrode. In this case the step of detecting or quantifying the detectable signal comprises detecting or quantifying an electrical signal from the electrode.

Thus in another embodiment there is provided a method of quantifying a biological molecule comprising: i) providing a sensor surface, said sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) exposing said sensor surface to the biological molecule in the presence of a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal in response to said binding; and iii) quantifying the detectable signal.

In a fourth aspect of the invention there is provided the use of a sensor surface comprising: i) an ultrasmooth film of electrically conductive polymer, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; in the detection of a biological molecule.

In a fifth aspect of the invention there is provided a process for making a sensor surface for detecting a biological molecule, said process comprising: i) providing an ultrasmooth film of electrically conductive polymer; and ii) coupling a capture agent to said film.

The polymer may comprise at least one type of monomer unit which is derived from a derivative of ethylendioxythiophene (EDOT). In an embodiment, step i) comprises: a) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the microemulsion; and b) electropolymerizing the one or more monomers in the microemulsion to form the film.

The one or monomers may be insoluble or only sparingly soluble in water. They may comprise EDOT coupled to a functional group. The functional group may be for example a hydroxyl group, a thiol group, an amine group, a carboxylic acid group, an azide group, an N-hydroxysuccinimidoester group or a maleimido group. The microemulsion may be an oil-in-water (O/W) microemulsion. The dispersed phase may be a hydrophobic phase. The continuous phase may be an aqueous phase. The continuous phase may have a pH between about 2 and about 6. The dispersed phase may additionally comprise one or more comonomers. The comonomers may be capable of copolymerising with the one or more monomers in order to produce an electrically conductive copolymer. Step a) may comprise forming a microemulsion from the one or more monomers, and, if present, the one or more comonomers, together with a surfactant, an acid and an electrolyte. The surfactant may be present below its critical micelle concentration.

Step b) may comprise applying an electrical potential between two electrodes, each of which is at least partially immersed in the microemulsion, whereby the film of electrically conductive polymer forms on at least one of said electrodes. Step b) may applying a cyclic potential to the microemulsion. It may comprise potentiostatic polymerisation. It may comprise galvanostatic electropolymerisation. The electrode on which the electrically conductive polymer forms may be an ultrasmooth electrode. Step ii of the process of the fifth aspect may comprise reacting the capture agent with a functional group on the surface of the film.

In an embodiment there is provided a process for making a sensor surface for detecting a biological molecule, said process comprising: i) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the microemulsion; and electropolymerizing the one or more monomers in the microemulsion to form an ultrasmooth film of electrically conductive polymer; and ii) coupling a capture agent to said film. hi another embodiment there is provided a process for making a sensor surface for detecting a biological molecule, said process comprising: i) forming a microemulsion having a continuous phase and a dispersed phase from one or more monomers together with a surfactant, an acid and an electrolyte, said surfactant being present below its critical micelle concentration, wherein the one or more monomers are located in the dispersed phase of the microemulsion; and electropolymerizing the one or more monomers in the microemulsion to form an ultrasmooth film of electrically conductive polymer; and ii) coupling a capture agent to said film.

In another embodiment there is provided a process for making a sensor surface for detecting a biological molecule, said process comprising: i) forming a microemulsion having a continuous phase and a dispersed phase from one or more monomers together with a surfactant, an acid and an electrolyte, said surfactant being present below its critical micelle concentration, wherein the one or more monomers are located in the dispersed phase of the microemulsion; and electropolymerizing the one or more monomers in the microemulsion to form an ultrasmooth film of electrically conductive polymer; and ii) reacting a capture agent with a functional group on the surface of the film to couple said capture agent to said film.

In a sixth aspect of the invention there is provided a process for making a sensor surface for detecting a biological molecule, said surface comprising: i) an ultrasmooth film of electrically conductive polymer, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; said process comprising: a) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the microemulsion; and b) electropolymerizing the one or more monomers in the microemulsion to form the sensor surface; wherein at least one of the monomers comprises EDOT coupled to the capture agent. Step b) may comprise applying an electrical potential between two electrodes, each of which is at least partially immersed in the microemulsion, whereby the film of electrically conductive polymer forms on at least one of said electrodes. Step b) may applying a cyclic potential to the microemulsion. It may comprise potentiostatic polymerisation. It may comprise galvanostatic electropolymerisation. In a seventh aspect of the invention there is provided a process for making a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding; said process comprising: a) providing the sensor surface; and b) coupling the label to the sensor surface.

Abbreviations

BSA Bovine Serum Albumin

CP capture probe

Cy3 Indocarbocyanine C₂-EDOT-COOH carboxylic acid functionalized EDOT DI de-ionised EDC 1 -ethyl-3-[3-dimeihylaminopropyl] carbodiimide hydrochloride

EDOT ethylenedioxythiophene

EDOT-OH hydroxymethyl fonctionalized EDOT

GOD-A glucose oxidase-avidin D

HRP horse radish peroxidase

ITO indium tin oxide

PBS phosphate buffered saline

PEDOT polyethylenedioxythiophene

PVA-Os poly(vinylimidazole)-polymer(acrylamide) copolymer partially imidazole-complexed with

Os(4,4'-dimethyl-2,2'-bipyridine)₂Cl^+/2+

RT-PCR reverse transcription polymerase chain reaction

SAM self-assembled monolayer

SDS sodium dodecyl sulfate sulfo-NHS N-hydroxysulfosuccinimide

TBO toluidine blue

Brief Description of the Figures

Figure 1 provides structures for the C₂-EDOT-COOH and EDOT monomers.

Figure 2 provides a diagrammatic representation of lipid-bilayer-like microstructures which are believed to form near the electrode surface during electropolymerisation and which are considered to be responsible for the high density of surface -COOH groups observed in PEDOT copolymer surfaces.

Figure 3 provides a diagrammatic representation of two fluorescence-based detection schemes, fluorescence quenching and fluorescence labelling, for a sensor which utilises PEDOT surfaces.

Figure 4 provides a diagrammatic representation of an exemplary electrochemical detection system for a sensor which utilises a biotin labelled target, and a glucose oxidase-avidin D (GOD-A) and redox polymer to produce the detectable electrochemical signal.

Figure 5 provides the results of toluidine blue titration experiments to examine the surface density of COOH groups on a PEDOT film, (a) UV- visible spectra of TBO dye in acetic acid solution after dissociation from copolymer films polymerized from monomer mixtures of EDOT and C₂-EDOT-COOH. (From top to bottom: 100, 90, 80, 70, 60, 50, 40, 30, 20 and 10 mol% of C₂-EDOT-COOH). (b) Calculated surface density of COOH groups with increasing mol% OfC₂-EDOT-COOH in the monomer mixture.

Figure 6 provides the results of fluorescence detection of DNA on poly(EDOT- OH)-Co-PoIy(C₂-EDOT-COOH) films (consisting of the mol% of C₂-EDOT-COOH monomer specified) on ITO substrates, (a) Fluorescence intensity in the presence of 1 μM of complementary target (black) compared to the original films (grey) in fluorescence quenching experiments, (b) Fluorescence intensity in the presence of 1 μM of Cy3- labeled complementary target (grey) and non-complementary target (black) in the fluorescence labelling experiments.

Figure 7 provides details of the amperometric response of poly(EDOT-OH)-co- ρoly(C₂-EDOT-COOH) films (consisting of the mol% of C₂-EDOT-COOH monomer listed) on Au (grey) and Pt (black) substrates. An electrochemical detection protocol shown in Figure 4 was used with 1 μM of complementary target, and GOD-A concentrations of (a) 5 μg/mL and (b) 50 μg/mL. The control experiment was conducted on films containing 0.2 mol% of C₂-EDOT-COOH, and hybridized with non- complementary target.

Figure 8 provides a diagrammatic representation of an exemplary electrochemical detection system for a sensor which utilises an antibody as a capture agent, and an HRP- labelled second antibody and redox polymer in the presence OfH₂O₂ as a label.

Figure 9 provides results of the comparison between amperometric responses of a sensor coated with 500 ng/ml IgG and a sensor coated with BSA (control) after reacting with goat anti-Rat IgG-HRP conjugate and applying redox polymer overcoating.

Detailed Description hi the context of this specification, the term "comprising" will be understood to imply the inclusion of a stated step or element or integer or group of stems or elements or integers, but not the exclusion of any other step or element or integer or group thereof. Thus, in the context of this specification, comprising means "including, but not necessarily solely including". Furthermore, variations of the word "comprising", such as "comprise" and "comprises", have correspondingly varied meanings.

Throughout this specification, reference to "a" or "one" element does not exclude the plural, unless context determines otherwise. For instance, reference to "a monomer" should not be read as excluding the possibility of multiple monomers. Provided herein is a sensor system for detecting a biological molecule. The present inventors have developed an improved strategy based on a conductive polymer such as

PEDOT (poly-3,4-ethylenedioxythiophene) and its derivatives for the construction of sensor surfaces which utilize a thin-film biointerface. These films may offer the following advantages:

(1) when electropolymerized from a microemulsion the films may be less than about 100 nm in thickness;

(2) the films may be ultrasmooth. An ultrasmooth surface may have a surface roughness (Rims) of less than about lOnm; (3) the films may be polymerized, or copolymerized under uniform conditions with various different functional groups and consequently are capable of a wide variety of different bioconjugation pathways;

(4) the films are conductive, allowing their use in circuits involving biological material;

(5) the deposition of the film may be spatially and/or temporally controlled by using, for example, electropolymerisation techniques;

(6) the films may be synthesized relatively rapidly when compared to other SAMs;

(7) the synthesis of the films is amenable to large-scale manufacturing; and

(8) the films exhibit very low intrinsic cytotoxicity.

The films which are produced may be "ultrasmooth", that is they have a surface morphology sufficiently smooth to reduce non-specific binding of biomolecules which results from interactions between the biomolecule and surface morphology, such as pits, indentations or peaks in the film. The ultrasmooth surface may provide a reduced nonspecific binding of biological molecules due to the presence of surface traps such as pits or indentations when compared to the polypyrrole surfaces described in Ho-Pun-Cheung et al., (2006) Lab Invest. 86:304-313. The ultrasmooth surface described herein may be at least as smooth as the SAM surfaces currently described in the art. hi one embodiment ultrasmooth surface may have a surface roughness (R_πns) of less than about lOnm.

The sensor systems are suitable for use in the fluorescent or electrochemical detection of a biological molecule. The sensor surface may advantageously be part of an array or micro array which is able to simultaneously detect a plurality of individual biological molecules. The electrically conductive films described herein lend themselves to such applications because of the short time required for their synthesis, and their ability to be produced in a spatially and temporally controlled manner, for example by electropolymerisation. The inventors have observed that the deposition of thin film polymers produced from a monomer microemulsion by electropolymerisation localises precisely to the conducting substrate. Accordingly, the electrically conductive films described herein may be deposited in patterns on a substrate to create a microarray in which the film is present in one or more defined regions of the substrate by controlling the location of the electropolymerisation. In the case of microarray chip, the control of the location of electropolymerisation is readily achieved by the selective application of the electropolymerisation voltage to the different electrode surfaces on the chip.

Where it is desirable for a microarray to present a plurality of sensor surfaces having a variety of different capture agents, one may selectively produce the electrically conductive film in a defined region of the microarray by only applying the electropolymerisation voltage to that region. The desired capture agent may then be selectively applied to the region of the microarray which is coated with the electropolymerized film, for example by immersing the microarray in a solution containing the capture agent, in which case only those areas of the microarray having the film will be capable of binding the capture agent, or by selectively providing the capture agent in defined positions, for example by microdroplet printing techniques. If required, exposed surfaces of the film may be blocked with a blocking agent to prevent the coupling of any further capture agent to the film. Alternatively, the activation of functionalizations on the film surface to couple to the capture agent may be made selectively, for example by the use of microdroplets of activating agents, so that only the selectively activated regions of the film is able to couple to the capture agent. Further rounds of selective electropolymerisation and/or selective coupling of specific capture agents may then be carried out to produce a microarray which provides a plurality of sensor surfaces having a variety of different capture agents. A further exemplary process for the selective electropolymerization of a polymer biointerface (but using a polypyrrol polymer) is described for example, in Bidan et al. (1999) Synthetic Metals 102:1363- 1365, the entire contents of which are incorporated herein by reference.

The sensor systems described herein have a wide range of applications for which sensors for biological molecules, such as arrays or micro-arrays have previously been used. A person of skill in the art will recognise that the applications for sensors for detecting a biological molecule are not limited to those exemplified in the present application. It will be understood that the detection of a particular biological molecule for the sensor systems described herein is limited only by the requirement for a specific binding partner to act as the capture agent and by the requirement for a label which detects the binding of the capture agent to the biological molecule. Where the biological molecule comprises an oligo- or polynucleotide, the capture agent may be a complementary oligo- or polynucleotide which is capable of hybridizing to at least a portion of the biological molecule, preferably under stringent hybridization conditions. By "stringent hybridization conditions" is intended overnight incubation at 42° C. in a solution comprising: 50% formamide, 5x SSC (750 mM NaCl, 75mM trisodium citrate), 50 mM sodium phosphate (pH 7.6), 5x Denhardt's solution, 10% dextran sulfate, and 20 μg/ml denatured, sheared salmon sperm DNA, followed by rinsing in O.lx SSC at about 65° C. By a polynucleotide which hybridizes to a "portion" of a polynucleotide is intended a polynucleotide (either DNA or RNA) hybridizing to at least about 15 nucleotides (nt), and more preferably at least about 20 nt, still more preferably at least about 30 nt, and even more preferably about 30 to 70 nt of the reference polynucleotide. Optimization of DNA microarrays using tethered oligonucleotides is described for example in Vainrub and Pettitt (2003) J. Am. Chem. Soc. 125:7798-7799, the entire contents of which is incorporated herein by reference. DNA microarrays are examples of exquisite selectivity and sensitivity of detection. The assembly, optimization and use of "DNA-chip" microarray to detect single nucleotide polymorphisms is, for example, described in Ho-Pun-Cheung et al. (2006) Lab. Invest. 86:304-313, the entire contents of which is incorporated herein by reference.

Where the biological molecule comprises an antigenic molecule or a molecule against which specific antibodies may be raised, the capture agent may be an antibody or an antigen-binding antibody fragment. The term "antibody" as used herein includes IgG (including IgGl, IgG2, IgG3, and IgG4), IgA (including IgAl and IgA2), IgD, IgE, or IgM₅ and IgY, and is meant to include whole antibodies, including single-chain whole antibodies, and antigen-binding fragments thereof. Antigen-binding antibody fragments include, but are not limited to, Fab, Fab' and F(ab')₂, Fd, single-chain Fvs (scFv), single- chain antibodies, disulfide-linked Fvs (sdFv) and fragments comprising either a VL or VH domain. The antibodies may be from any animal origin. Antigen-binding antibody fragments, including single-chain antibodies, may comprise the variable region(s) alone or in combination with the entire or partial of the following: hinge region, CHl, CH2, and CH3 domains. Also included are any combinations of variable region(s) and hinge region, CHl, CH2, and CH3 domains. Antibodies may be monoclonal, polyclonal, chimeric, humanized, and human monoclonal and polyclonal antibodies which specifically bind the biological molecule.

Conversely, where the biological molecule is an antibody, the capture agent may be an antigen or an anti-idiotype antibody. Where the biological molecule comprises a polysaccharide, the capture agent may be a lectin or a polysaccharide-binding fragment thereof.

The electrically conductive polymer film of the sensor surface comprises a polymer, and may optionally comprise other components, e.g. included salts, swelling solvents etc. The polymer may comprise monomer units which are derived from EDOT or derivatives thereof. It may be a homopolymer or a copolymer.

The ultrasmooth film may have a surface root-mean-square roughness R_11nS of less than or equal to about 1 Onm, or less than about 5nm. The R_ms may be between about 1 and 10 ran, 1 and 9 ran, 1 and 8 ran, 1 and 7 ran, 1 and 6 ran, 1 and 5 ran, 1 and 4 nm, 1 and 3 nm or 1 and 2 nm. The R may be between about 2 and 10 nm, 2 and 4 nm, 2 and 5 ran, 2 and 7 nm, or 2 and 9 nm. The R_11nS may be between about 3 and 10 nm, 4 and 10 nm, 5 and 10 nm, 6 and 10 nm, 7 and 10 nm, 8 and 10 nm, 9 and 10 nm. The R_ππs may be between about 3 and 4 nm, 3 and 5 ran, 3 and 7 nm, or 3 and 10 nm. The R_11nS may be between about 4 and 5 nm, 4 and 7 nm, or 4 and 10 nm. The R_ms may be less than about 10, 9, 8, 7, 6, 5, 4, 3, 2 or Iran. The R_n^ may be about, 1, 2, 3, 4, 5, 6, 7, 8, 9 or lOnm. The smoothness of the film may provide increased accuracy relative to a less smooth film. It may provide increased sensitivity relative to a less smooth film. It may provide increased selectivity relative to a less smooth film. It may provide reduced non-specific binding relative to a less smooth film. The film may have a water contact angle of less than or equal to 90°. The film may have a water contact angle of less than or equal to 85°. The film may have a water contact angle in any one of the ranges of between 90° and 30°, 87° and 30°, 85° and 30°, 84° and 30°, 83^° and 32°, 81° and 35°, 81° and 38°, 81° and 39°, or 81° and 40°. The film may have a contact angle of any one of the following values: 30^°, 31°, 32°, 33^°, 34°, 35°, 36^°, 37°, 38°, 39°, 40°, 41°, 42°, 43°, 44°, 45°, 46°, 47^°, 48°, 49°, 50°, 51°, 52°, 53°, 54° 55°, 56°, 57^°, 58^°, 59°, 60°, 61°, 62°, 63°, 64°, 65°, 66°, 67°, 68°, 69°, 70°, 71°, 72°, 73°, 74°, 75°, 76°, 77°, 78°, 79°, 80°, 81°, 82°, 83°, 84° or 85°. The film may have hydrophobic groups on the surface. The hydrophobic groups may be trialkylsilyl or dialkylsilanediyl groups, e.g. trimethylsilyl, triphenylsilyl, dimethylsilanediyl or diphenylsilanediyl groups or some other suitable hydrophobic groups. The hydrophobic groups may be introduced on the surface following coupling of the surface to the capture molecule or prior thereto, or may be introduced at the same time. The film may have no hydrophobic groups. It may be hydrophilic.

The film may have a thickness of less than about lOOnm, or less than about 90, 80, 70, 60 or 50nm. It may have a thickness of between about 20 and 100, 30 and 100, 40 and 100, 50 and 100, 60 and 100, 70 and 100, 80 and 100, 90 and 100, 20 and 80, 20 and 50, 20 and 30 or 50 and 80nm. It may have a thickness of e.g. about 20, 30, 40, 50, 60, 70, 80, 90 or lOOnm. The conductivity of the film may be dependent on the concentration of a biological molecule in a medium in which the film is immersed or partially immersed, or which is placed on the surface of the film.

The flexibility of the coupling strategy for the film allows the sensor to be constructed to allow the immobilization of one or more of a wide variety of capture agents, and hence allows the sensor to potentially interact with a wide variety of biological molecules. The electrically conductive film may comprise functional groups which associate with, or bind to, the capture agent to couple the capture agent to the film. For example, where the capture agent comprises an oligonucleotide, polynucleotide, or a peptide or polypeptide, COOH functionality provided by C₂-EDOT-COOH groups present on the film surface may be used to couple the capture agent to the film. The COOH groups may be activated by exposure to sulfo-NHS (N-hydroxysulfosuccinimide) and EDC (1-ethyl- 3-[3-dimethylaminopropyl]carbodiimide hydrochloride). It will be understood, however, that numerous other functional groups and activation strategies commonly employed in the art to specifically couple biological molecules to polymer surfaces may be employed. Other coupling reactions that are well known in the art may also be used. These may comprise any of the well known methods of introducing chemical groups into a molecule or onto a functionalised solid. These include nucleophilic substitution of a benzyl halide (e.g. chloride or bromide) group, "click" chemistry etc. Suitable click chemistry may include for example cycloaddition reactions, such as the Huisgen 1,3 -dipolar cycloaddition, Cu(I) catalyzed azide-acetylene cycloaddition, Diels-Alder reaction, nucleophilic substitution to small strained rings (e.g. epoxy and aziridine rings), formation of ureas and amides and addition reactions to double bonds, e.g. epoxidation, dihydroxylation.

Alternatively or in addition, the capture agent may be incorporated into the solution of monomer units prior to polymerization, so that upon polymerization of the film capture agent is exposed on the surface of the film. The capture agent may be conjugated directly or via a linking group to a monomer, or it may be dispersed in solution with the monomers.

In the context of the present specification, the term "coupled" and related terms (such as coupling etc.) should be taken to encompass immobilised, bonded to or other similar forms of coupling. The bonding may be covalent, ionic or may be a conjugation or a complexation, or may be a combination of any two or more of these. The coupling may comprise chemical coupling. It may comprise physical coupling. It may comprise a combination of physical and chemical coupling. The coupling may be any suitable coupling such that the capture agent is not readily removed from the film under normal conditions of use.

The onset potential for conductivity may be dependent on the concentration of the substance in the medium in which the film is immersed or partially immersed. In particular embodiments, the onset potential is between about -1 and about 0 V against saturated calomel electrodes (SCE), and the film's conductivity is between about 0.1 to about 1000 S/cm. The onset potential may be between about -1 and -0.5, -0.5 and 0 or - 0.8 and -0.3, e.g. about -1, -0.9, -0.8, -0.7, -0.6, -0.5, -0.4, -0.3, -0.2, -0.1 or OV against SCE. The conductivity may be between about 0.1 and 100, 0.1 and 10, 0.1 and 1, 1 and 1000, 10 and 1000, 100 and 1000, 1 and 100, 1 and 10 or 10 and 100 S/cm, e.g. about 0.1, 0.2, 0.5, 1, 2, 5, 10, 20, 50, 100, 200, 300, 400, 500, 600, 700, 800, 900 or lOOOS/cm. The polymer of the electrically conductive film may additionally comprise other monomer units that are copolymerised with the monomer units which are derivatives of ethylendioxythiophene. These other monomer units (comonomer units) may be ethylendioxythiophene units, substituted ethylenedioxythiophene units, thiophene units, substituted thiophene units, pyrrole units, substituted pyrrole units, furan units, substituted furan units, or other comonomer units that, when incorporated with the derivatives of ethylendioxythiophene into the polymer, allow the polymer to conduct a current. The solubility in water of the comonomers from which these units are derived should be sufficiently low that the process of the present invention provides a film having R_πnS of less than about lOnm. The roughness of polymer film is strongly dependent on the substrates, however, the increase Of R_11nS roughness after coating of polymer thin films is commonly less than 10 nm, optionally less than about 5 or 2 nm. The substrate (e.g. an electrode on which the film is formed) may be ultrasmooth. It may be sufficiently smooth that the electrically conductive film is ultrasmooth as described elsewhere herein. The substrate may have a surface root-mean-square roughness R_πn_s of less than or equal to about lOnm, or less than about 8, 5 or 3nm. The R_nJ13 maybe between about 1 and 10 nm, 1 and 9 nm, 1 and 8 nm, 1 and 7 nm, 1 and 6 nm, 1 and 5 nm, 1 and 4 nm, 1 and 3 nm or 1 and 2 nm. The R_πns may be between about 2 and 10 nm, 2 and 4 nm, 2 and 5 nm, 2 and 7 nm, or 2 and 9 nm. The R_πns may be between about 5 and 8nm, 2 and 8nm, 3 and 10 nm, 4 and 10 nm, 5 and 10 nm, 6 and 10 nm, 7 and 10 nm, 8 and 10 nm, 9 and 10 nm. The R_πns may be between about 3 and 4 nm, 3 and 5 nm, 3 and 7 nm, or 3 and 10 nm. The Rπns may be between about 4 and 5 nm, 4 and 7 nrn, or 4 and 10 nm. The R_πns may be less than about 10, 9, 8, 7, 6, 5, 4, 3, 2 or lnm. The R_ms may be about, 1, 2, 3, 4, 5, 6, 7, 8, 9 or IOnm.

The derivatives of ethylendioxythiophene, prior to coupling of the capture agent, may comprise an ethylendioxythiophene group coupled to a functional group. The functional group may be for example a hydroxyl group, a thiol group, an amine group, a carboxylic acid group, an azide group, an N-hydroxysuccinimidoester group or a maleimido group. The functional group may be a group that is capable of reacting with the capture agent in order to couple the capture agent to the ethylenedioxythiophene group. The functional group may be coupled to the ethylenedioxythiophene group through a linker group, for example a hydrocarbon chain, an ether containing chain, an polyether chain, an ester containing chain or a chain containing both an ether and an ester. The hydrocarbon chain may have between 1 and 12 carbon atoms, and may be straight chain, branched, cyclic or a combination of these. The linker chain may have a total of between about 1 and 20 atoms in the main chain linking the ethylenedioxythiophene group to the functional group. Alternatively there may be no linker group.

The sensor surface may form part of a sensor system. A suitable sensor system comprises a sensor surface as described above and a label. The label is capable of detecting the binding of the biological molecule (i.e. an analyte) to the capture agent of the sensor surface. The detection generates a detectable signal in the sensor system, for example an electrical (e.g. electrochemical) or fluorescent signal, which then may be detected in order to indicate the presence of the biological molecule, hi some embodiments, the intensity of the signal depends on the concentration of the biological molecule. Thus in these embodiments the sensor system is capable of providing a signal which may be used to indicate the concentration of the biological molecule. The signal may be directly proportional to the concentration of the biological molecule. It may be indirectly proportional thereto, hi may be dependent in some other manner thereon. Since the sensor surface is capable of selectively binding to the biological molecule (as the capture molecule coupled to the sensor surface is capable of selectively binding to the biological molecule), the signal is capable of detecting the presence, optionally the concentration, of the biological molecule in the presence of other species, optionally in the presence of other biological species.

The label may be coupled to the sensor surface. Prior to the coupling of the biological molecule to the sensor surface (i.e. to the capture molecule) there may be no coupling of the label to the sensor surface. Thus there may be three possibilities for detection:

1) the label is coupled to the sensor surface prior to the coupling of the biological molecule to the sensor surface. The label may be coupled to the capture molecule, or it may be coupled to some other location of the surface. It may be coupled to the polymer. It may be coupled through a tether group. The tether group may be sufficiently flexible, of suitable hydrophilicity and of sufficient length to permit coupling of the label to the biological molecule once the capture agent is coupled to the biological molecule. Exposure of the surface to biological molecule leads to both biological molecule and label being coupled to the surface and thereby gives rise to a detectable signal.

2) the label is not coupled to the sensor surface prior to the coupling of the biological molecule to the sensor surface. Exposure of the sensor surface to the biological molecule leads to binding of the biological molecule to the surface. Exposure of the bound surface to the label leads to both biological molecule and label being coupled to the surface and thereby gives rise to a detectable signal.

3) the label is not coupled to the sensor surface prior to the coupling of the biological molecule to the sensor surface. Exposure of the biological molecule to the label gives rise to a labelled biological molecule. Exposure of the sensor surface to the labelled biological molecule leads to both biological molecule and label being coupled to the surface and thereby gives rise to a detectable signal.

In the above options, the biological molecule may be in solution, in suspension, in emulsion, in microemulsion or in dispersion. It may dissolved, emulsified, microemulsified or dispersed in a medium. The medium may comprise suitable salts, buffers etc. The medium may also comprise other species, e.g. other biological species. The detection or quantitation of the biological molecule may be selective for the biological molecule in the presence of the other species.

The smallest detectable concentration of biological molecule may be less than 100 fM, and in certain embodiments less 50 fM, 25 fJVI, or 10 fM. The concentration of label may be at least as great as the concentration of biological molecule on a molar basis. There may be at least 1.1 times as much label as biological molecule on a molar basis or at least about 1.2, 1.3, 1.4, 1.5, 2, 2.5, 3, 3.5, 4, 4.5 or 5 times as much. The sensor surface may be produced by a process comprising: i) providing an ultrasmooth film of electrically conductive polymer, e.g. a film having a surface root-mean-square roughness R_πns of less than or equal to about lOnm; and ii) coupling a capture agent to said film. Step i) may comprise the steps of: a) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the

5 microemulsion; and b) electropolymerizing the one or more monomers in the microemulsion to form the film.

There are two factors contributing to the formation of a smooth surface, as described herein. The monomer and any comonomers present should have low solubilityo in water. They are therefore mainly confined in the micelle and lipid bilayer formed by surfactants on substrates. The presence of a lipid bilayer offers a template for the creation of smooth surfaces on polymer thin films and the use of lipid micelles allows control in film growing.

The dispersed phase may additionally comprise one or more comonomers. It mays comprise 2, 3, 4, 5 or more than 5 comonomers. The comonomers may be capable of copolymerizing, e.g. electrocopolymerizing, with the derivative of ethylendioxythiophene in order to produce an electrically conductive copolymer. The comonomers may each be selected from the group consisting of ethylendioxythiophene, substituted ethylenedioxythiophene, thiophene, substituted thiophene, pyrrole, substituted pyrrole,0 furan, substituted furan, or other comonomers that, when copolymerised with the derivatives of ethylendioxythiophene, allow the resulting copolymer to conduct a current. The solubility in water of the comonomers should be sufficiently low that the process of the present invention provides a film having R^ of less than about 1 Onm. In particular embodiments the solubility of the comonomers in water is less than 3 mM, less than 2S mM or less than ImM. The solubility of each monomer separately, or of the monomers and comonomers in combination, may be between about 0 and about 3mM, or about 0 to 2, 0 to 1, 0 to 0.5, 0.5 to 3, 0.5 to 2, 1 to 3 or 1 to 2mM, e.g. about 0.5, 1, 1.5, 2, 2.5 or 3mM.

Step a) may comprise forming a microemulsion from the one or more monomers0 (each of which may be a derivative of ethylendioxythiophene), and, if present, the one or more comonomers, together with a surfactant, an acid and an electrolyte. The one or more monomers may be present in the microemulsion in a concentration of between about 0.001 and 0.1M, or between about 0.005 and 0.1, 0.01 and 0.1, 0.05 and 0.1, 0.001 and 0.05, 0.001 and 0.01, 0.005 and 0.05 or 0.005 and 0.2. The monomers may be present in5 the microemulsion in a concentration (either individually or in combination) of e.g. about 0.001, 0.002, 0.003, 0.004, 0.005, 0.006, 0.007, 0.008, 0.009, 0.01, 0.015, 0.02, 0.025, 0.03, 0.035, 0.04, 0.045, 0.05, 0.06, 0.07, 0.08, 0.09 or 0.1M. The above concentrations may also apply to each comonomer present. They may apply to the total concentration of monomer plus all comonomers present. The acid may be a mineral acid, for example hydrochloric, sufuric, nitric, phosphoric, hydrobromic or some other acid. It may be a strong acid. It may be in sufficient concentration that the pH of the microemulsion is between about 1 and about 6 or between about 2 and about 6, or between about 2 and 4, 2 and 3, 2 and 6, 4 and 6, or 3 and 5. It may be in sufficient concentration that the pH of the microemulsion is e.g. about 1, 1.5, 2, 2.5, 3, 3.5, 4, 4.5, 5, 5.5 or 6. The pH of the microemulsion should be within a range that does not degrade or denature any of the monomers or comonomers.

In particular, some monomers usable in the invention may comprise biomolecule fragments, e.g. protein fragments, and the pH should in that case be such that it does not hydrolyse or denature the protein fragments. The biomolecule fragments may comprise the capture agent. They may comprise the capture agent coupled to the label, hi those cases the pH should be such that it does not hydrolyse or denature the capture agent, and in the later case also the label.

The surfactant may be any suitable surfactant capable of forming a microemulsion with the monomer and, if present, the comonomers. It may be an ionic surfactant. It may be a non-ionic surfactant. It may be a zwitterionic surfactant. It may be an anionic surfactant. It may be a cationic surfactant. It may be, for example, sodium dodecyl sulfate or sodium 4-dodecylbenzene sulfate (SDBS). In some embodiments, the surfactant is present below its critical micelle concentration. The concentration of the surfactant in the microemulsion may be between about 0.01 and 0.1 M, or between about 0.01 and 0.05, 0.05 and 0.1 or 0.03 and 0.07M. The concentration of the surfactant in the microemulsion may be e.g. about 0.01, 0.02, 0.03, 0.04, 0.05, 0.06, 0.07, 0.08, 0.09 or 0.1M. It may be sufficient concentration to form a microemulsion with the monomer. It may be sufficiently low concentration that the process of the invention provides a polymeric film having R_11nS of less than about IOnm. The microemulsion also comprises an electrolyte. The electrolyte may be an ionic salt. It may be a water soluble ionic salt. It may be for example lithium perchlorate. It may be present in the microemulsion, particularly in the aqueous phase of the microemulsion, in a concentration sufficient to allow electropolymerisation of the monomer and, if present, the comonomer. The concentration of the electrolyte may be between about 0.01 and IM, or between about 0.05 and 1, 0.1 and 1, 0.5 and 1, 0.05 and 0.5 or 0.08 and 0.12M. The concentration of the electrolyte may be e.g. about 0.01, 0.02, 0.03, 0.04, 0.05, 0.06, 0.07, 0.08, 0.09, 0.1, 0.11, 0.12, 0.13, 0.14, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 0.6, 0.7, 0.8, 0.9 or IM.

In forming the microemulsion, the components thereof may be mixed in any convenient order. Once mixed, they may be agitated in order to form a microemulsion, although in general vigorous agitation is not necessary.

Step b) of the process may comprise applying a cyclic potential to the microemulsion. The upper voltage of the cycle may be between about 0.5 and 1.5V, or between 0.5 and 1 or 1 and 1.5 or 0.8 and 1.2V, and maybe about 0.5, 0.6, 0.7, 0.8, 0.9, 1, 1.1, 1.2, 1.3, 1.4 or 1.5V. The lower voltage of the cycle may be between about -0.2 and about -IV or between about -0.2 and -0.5 or -0.5 and -1, or -0.5 and -0.8V, e.g. about -0.2, -0.3, -0.4, -0.5, -0.6, -0.7, -0.8, -0.9 or -IV. The voltage gap between upper and lower voltage may be between about 1.5 and 2V, e.g. about 1.5, 1.6, 1.7, 1.8, 1.9 or 2V. The scan rate may be between about 10 and about 500mV/s, or between about 50 and 500, 100 and 500, 200 and 500, 10 and 200, 10 and 100, 10 and 50, 10 and 20, 50 and 200 or 80 and 120mV/s. The voltage gap between upper and lower voltage may be e.g. about 10, 20, 30, 40, 50, 60, 70, 80, 90, 100, 110, 120, 130, 140, 150, 200, 250, 300, 350, 400, 450 or 500mV/s.

Alternatively the process of forming the electrically conductive film may comprise potentiostatic electropolymerisation. The applied voltage in this case may be between about 0.8 to about 1.5 V, and usually between about 0.9 to about 1.3 V. It may be about 0.8 to 1, 1 to 1.5, 0.9 to 1.1 or 1.1 to about 3, e.g. about 0.8, 0.9, 1, 1.1, 1.2, 1.3, 1.4 or 1.5V. As the polymerisation of the electrically conductive film may depend on the application of a voltage, the deposition of polymerised film advantageously may be temporally and spatially controlled by controlling the timing and period of electropolymerisation and the surface over which a voltage is applied.

In another option the process of forming the electrically conductive film may comprise galvanostatic electropolymerisation. Suitable currents for galvanostatic electropolymerisation may be between about 5 and about 20 niA/cm², and in one embodiment 10 mA/cm². The current may be between about 5 and 10, 10 and 20, 10 and 15, 15 and 20 or 8 and 15mA/cm², e.g. about 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19 or 20 mA/cm².

The derivative of ethylendioxythiophene may comprise an ethylendioxythiophene group coupled to a functional group. The functional group may be a hydroxyl group, a thiol group, an amine group, a carboxylic acid group, an azide group, an N- hydroxysuccinimidoester group or a maleimido group, or may be some other group capable of being coupled to a biomolecule. The derivative of ethylendioxythiophene may comprise a biomolecule residue. In this case, it may be coupled to the ethylendioxythiophene group by reaction with one of the above stated groups. Step ii) may comprise exposing the film of polymer to the capture agent so that the capture agent can bind to, e.g. react with, a functional group on the polymer. Suitable chemistry, e.g. click chemistry, for this binding has been discussed elsewhere. The exposing may comprise at least partially immersing the film in a solution, suspension, emulsion, microemulsion or dispersion of the capture agent in a carrier. The carrier may be an aqueous carrier. It may comprise suitable salts, buffers etc.

In some embodiments, as noted earlier, the label is coupled to the film. In this case, the process may comprise coupling the label to the film. The coupling of the label to the film may comprise coupling the label to the capture agent. It may comprise coupling the label to the polymer. This may be conducted before, during or after coupling of the capture agent to the polymer. The coupling of the label to the film may use similar chemistry to that described earlier for coupling the capture agent to the polymer or it may use different chemistry. It may use chemistry that is well known to those skilled in this art.

The sensor surfaces described herein allow the use of a variety of known techniques to detect binding of the biological molecule to the binding agent. The use of electrochemical detection techniques or fluorescent techniques are particularly contemplated.

The binding of the biological molecule to the capture agent is detected by a label. The label may be attached to or comprised within the binding agent. Alternatively or in addition, the label may be provided after the biological molecule is bound to the binding agent. Where the label is provided after the biological molecule is bound to the capture agent, the label may recognise the immobilised biological molecule, such as in a "sandwich" technique. Alternatively, or in addition, the label may be coupled to the biological molecule prior to its binding to the capture agent. This may be the case where, for example, all of the biological molecules within a sample are labelled prior to screening them with the sensor.

The use of a "molecular beacon", for example, as a label to detect binding of an oligo- or polynucleotide biological molecule is contemplated. Molecular beacons comprise an oligonucleotide sequence, which is able to specifically hybridize with a complementary oligo- or polynucleotide, together with a fluorescent label which is quenched when the molecular beacon is not bound to a complementary oligo or polynucleotide, but which becomes unquenched when the molecular beacon is hybridized (Bonnet et al. (1999) Proc Natl Acad. Sd USA 96: 6171-6176, the entire contents of which is incorporated herein by reference). The detection of binding of the biological molecule by the label may produce an increase in a signal, for example where binding of a fluorescently tagged label agent produces an increase in fluorescence. The detection agent may produce a decrease in a signal, for example where binding of the biological molecule to a fluorescently labelled binding agent quenches or otherwise inhibits the fluorescence of the fluorescent label. Electrochemical techniques rely on may provide a particularly sensitive approach to identify the presence of target biomolecule binding. Numerous electrochemical detection techniques are described, for example, in the reviews in Drumrnond et ah, (2003) Nature Biotechnology 21: 1192-1199, Wang (2000) Nucleic Acids Res 28: 3011-3016 and Christopoulos, (1999) Anal. Chem 71: 425R-438R, the entire contents of which are incorporated herein by reference. hi a particularly preferred embodiment, an electrochemical technique utilizing cationic redox polymers containing osmium- bipyridine complexes interacting with anionic enzymes is used. This technique is described in Xie et al (2004) Anal. Chem. 76: 1611-1617, the entire content of which is incorporated herein by cross reference. Where electrochemical detection techniques are used, it may be advantageous for the film to be present on the surface of an electrode, so that changes in current flow may be conveniently detected.

Examples Example 1. Preparation of PEDOT films

Ethylenedioxythiophene (EDOT, Sigma-Aldrich), lithium perchlorate (LiClO₄, Fluka), sodium dodecyl sulfate (SDS, Alfa Aesar), D-(+)-glucose (Sigma) were used as received. Hydroxymethyl-functionalized EDOT (EDOT-OH) was synthesized according to the procedure described in Lima et al. (1998) Synth. Met. 93: 33. Carboxylic acid- functionalized EDOT (C₂-EDOT-COOH) was synthesized as described in International patent application No PCT/SG2006/000282, the entire contents of which is incorporated herein by reference. A phosphate-buffered saline (PBS) consisting of 137 mM of NaCl, 2.7 mM of KCl, and 10 mM of phosphate buffer was used as supporting electrolyte solution. The redox polymer used in this study was poly(vinylimidazole)- polymer(acrylamide) copolymer partially imidazole-complexed with Os(4,4'-dimethyl- 2,2'-bipyridine)₂Cl^+/2+ (PVA-Os), which was synthesized as described before (Gao, et ah; (2002) Angew. Chem. Int. Ed. 41: 810.). Glucose oxidase-avidin D (GOD-A, Vector Laboratories) was diluted in PBS by 100 and 1000 times in volume to form 50 μg/mL and 5 μg/mL solutions. The probe and target oligonucleotide sequences in Table 1 were custom-prepared by 1st Base, Inc. Indium tin oxide (ITO) coated glass (Delta- Technologies, Ltd.) was cleaned by standard procedure prior to use. Au and Pt disk working electrodes (CHI Instrument) were polished by Polishing Kits (PK-4, Bioanalytical Systems, Inc.) with 0.05-μm alumina (Gamma Micropolish, Buehler) before use. The acid-catalyzed microemulsion electropolymerization allows for the formation of carboxylic acid-functionalized, thin, ultrasmooth, non-cytotoxic PEDOT films. PEDOT films from the EDOT monomers Hydroxymethyl-functionalized EDOT (EDOT- OH) and Carboxylic acid-functionalized EDOT (C₂-EDOT-COOH) (10 mM respectively) were electropolymerized on Au, Pt and ITO electrodes in 2cm x 1 cm areas from 10 mM of EDOT aqueous solutions containing 0.1 M Of LiClO₄ as supporting electrolyte, in the presence of 1 mM of HCl and 0.05 M of SDS, by applying cyclic potentials (-0.6 to 1.1 V vs. Ag/ AgCl at a scan rate of 100 mV/s) or by potentiostatic methods in which the films were deposited at a constant potential (1.0 V) with the electropolymerization cut off when the charge exceeded 0.15 mC. The potentiostatic method was used where it was desirable to retain the consistency of the amount of polymer deposited on the electrode surface.

The surface morphology of polymer films was observed with field emission scanning electron microscopy (FESEM) and atomic force microscopy (AFM). FESEM was carried out with JEOL JSM-7400 at a vacuum of 10^"8 torr and an accelerating voltage of 10 kV. AFM was performed in the tapping mode at room temperature in air with BioScope™ Digital Instruments. R_ms was obtained for a scan range of 20 μm x 20 μm. These methods yielded ultrasmooth polymer films with a Rπ_ns of <5nm. The surface density of carboxylic acid groups (COOH) on polymer films was determined by toluidine blue (TBO) staining using the method described in Uchida et al. (1993) Langmuir, 9: 1121. Polymer films were immersed in 0.5 mM of TBO solution at pH 10. After removing non-complexed dye with water, the dye on polymer films was desorbed in a 50% acetic acid solution, and the dye concentration was determined by measuring the absorbance at 633 ran (Agilent 8453), with the absorbance signifying the presence of COOH on the PEDOT film surface. Density was then obtained from a calibration curve of the absorbance intensity vs. dye concentration. Assuming that TBO was bound to surface COOH at 1:1 molar ratio, the density of surface functional carboxylic acid group was quantified as 3.5x10 /cm based on the absorbance of redissolved TBO. Using this method, the surface COOH group of films electropolymerized in CH₃CN was found to have a surface density of 2.9^χlO¹⁵/cm². The higher density of surface COOH groups in the film synthesized by microemulsion electropolymerization was attributed to the lipid bilayer microstructures near the electrode surface (as illustrated in Figure 2). Without wishing to be bound by any proposed mechanism, it is suggested that such microstructure allows hydrophilic COOH groups to stay closer to the material surface during electropolymerization, thereby increasing the availability of these functional groups compared to that in the films randomly grown from organic medium, although the latter films have a higher surface area due to their greater roughness.

Example 2. Immobilization of capture probe oligonucleotides and detection of hybridization

Carboxylic acid groups on polymer films produced in Example 1 were activated by N-hydroxysulfosuccinimide (sulfo-NHS, Pierce) and l-ethyl-3-[3-dimethylaminopropyl] carbodiimide hydrochloride (EDC, Pierce) coupling, using techniques described in Navarro et al, (2005) Tetrahedron 61: 3947. After 15-20 min of incubation at room temperature, excess sulfo-NHS and EDC were removed by rinsing with deionized (DI) water. The activated surfaces were then coupled with 3'-amino-functionalised oligonucleotides as shown in Table 1.

The whole assembly process combining polymerisation, activation and immobilization required less than 5 h. Compared to the 18 to 24 h of assembling time for similar devices based on SAM, the procedures described herein are more suitable for the efficient manufacturing or array applications.

The following oligonucleotides were used in this example: Table 1. Oligonucleotide Sequences

Function 5'— >3' sequence capture probe TTTGAGTCTGTTGCTTGGAAAAAA(CH₂)₆NH₂ (SEQ ID NO: l)^a complementary CCAAGCAACAGACTCAAA (SEQ ID NO:2) ^b non-complementary CAAGCACTTGCTGACCAAGCAAC (SEQ ID NO:3)^b aCapture probe was 5 '-labeled with Cy3 or unlabeled for different detection schemes. ^bTargets were 3 '-labeled with Cy3, biotin, or unlabeled for different detection schemes.

Amine-modified capture probe (CP) oligonucleotides (SEQ ID NO:1) were immobilized onto electrodes by immersing the electrodes in PBS containing 1 μM of CP. After adsorption, the electrodes were rinsed with PBS and DI water to remove non- specifically adsorbed materials, and then blown dry with nitrogen.

The hybridisation of target oligonucleotides (SEQ ID NOS :2 and 3) was performed in a moisture-saturated chamber maintained at room temperature. CP -immobilized electrodes were immersed in PBS with 1 μM of either of the target oligonucleotides overnight. Cy3-labeled and biotin-labelled target oligonucleotides were used for fluorescence detection and electrochemical detection, respectively. Fluorescence measurement was performed with a fluorescence microscope (BX-51, Olympus) with a CCD camera (DP70, Olympus). Images were then analyzed with Image-Pro 3D Suite software (MediaCybernetics) to calculate the fluorescence intensity.

For electrochemical detection, the method described in Xie (Xie et al (2004) Anal. Chem. 76: 1611-1617) was used. After rinsing thoroughly with PBS and DI water, electrodes of biotin-labelled target oligonucleotides were immersed in GOD-A solutions. GOD-avidin was attached to the biotinylated target oligonucleotides via the biotin-avidin interaction. The glucose electro-oxidation current was measured amperometrically in PBS containing 40 mM of glucose with a potential of 0.4 V (vs. Ag/ AgCl) in a Faradic cage. .

A primer design targeting the viral neuraminidase gene Nl was adapted for avian flu detection as the CP on the polymer surface. Initial tests upon hybridization of Cy3 fluorescent dye labelled complementary target showed almost no fluorescence when the films from homopolymer poly(C₂-EDOT-COOH) were employed as biointerface for activation and CP immobilization. This was attributed to the high CP density on the polymer surface, which interfered with efficient hybridization with the target sequence. Theoretical calculations and previous reports suggested that the efficiency of DNA hybridization to surface-bound probe was maximized when the probe density was 10¹² to

1¹X 0

10 /cm . This value was 23 orders of magnitude lower than that obtained with the PoIy(C₂-EDOT-COOH) film. To reduce the COOH surface density, copolymers of C₂- EDOT-COOH and EDOT were synthesized from a mixture of monomers, as COOH density is controlled by the percentage of C₂-EDOT-COOH in the monomer mixture. A linear relationship (R = 0.99179) between the TBO absorbance at 633 nm and the molar percentage of C₂-EDOT-COOH in the monomer mixture was observed (see Figure 5), which confirmed these calculations.

Fluorescent detection techniques remain a popular approach for DNA detection due to their simplicity and the ready availability of appropriate instrumentation. Although a strategy to construct fluorescent DNA detection systems based on electropolymerized polypyrrole surfaces has been attempted, the films which were used were not uniform. The rough surface which was produced may have affected the hybridization efficiency and detection limit if a long DNA sequence from sample tissues was targeted.

In contrast, the uniform PEDOT film which is described herein increases the accessibility of each capture probe, allowing efficient target hybridization and lowering the detection limit. Two fluorescence-based detection schemes were employed, fluorescence quenching and fluorescence labelling (see Figure 2) to demonstrate the versatility of this platform. A copolymer was electropolymerized from a mixture of EDOT-OH and C₂-EDOT-COOH monomers containing 0.1-1% C₂-EDOT-COOH, providing and optimized surface COOH density for detection. The fluorescent response with PEDOTs was examined. Initially, EDOT was employed to modulate the monomer composition, but there was undesired level of attachment of polymer films to the substrate surface. Therefore, the more hydrophilic EDOT-OH was subsequently used in place of EDOT for the copolymerization. The resulting poly(EDOT-OH)-co-ρoly(C₂-EDOT-COOH) system greatly enhanced the interfacial adhesion between the polymer film and the substrate (Au, Pt or ITO).

In the fluorescence quenching scheme, the polymer films were first deposited on an ITO substrate. CPs with amino group functionalized the 3'-ends and Cy3 label at the 5'- ends were immobilized after the surface COOH groups were activated by NHS/EDC coupling. Fluorescence micrographs of these films showed no fluorescence after the immobilization of dye-labelled probes due to the strong energy transfer between the dye molecules and PEDOT. PEDOT films, in oxidized and conductive form, displayed broad absorption wave in the visible to near-infrared range. This broad absorption peak overlapped with the emission spectrum of Cy3. Before complementary target hybridization, single-stranded DNA (ssDNA) CPs behaved like flexible polymers and laid on the surface of the electrodes with the 5 '-end close to the surface, prompting efficient energy transfer between Cy3 and PEDOT films.

Upon hybridization of the complementary target, the double-stranded DNA (dsDNA) became more rod-like, and the 5'-end was pulled away from the film surface, reducing the efficiency of energy transfer. As a result, an increased fluorescent intensity from Cy3 emission was observed (see Figure 6(a)). This phenomenon could be extended to applications similar to the "molecular beacon" method for DNA detection.

After hybridization with Cy3 -labeled targets on the CP-grafted biointerface, much greater fluorescence intensity was obtained with complementary targets, compared to non-complementary targets (Figure 6(b)), regardless of the surface density of CPs. Copolymers with 5% and 0.1% C₂-EDOT-COOH displayed < 60% fluorescence intensity after complementary target hybridization in both quenching and labelling schemes, compared to copolymers with 1% and 0.2% C₂-EDOT-COOH.

In addition to fluorescence detection, the conductive biointerfaces were also suitable for electrochemical detection. A SAM-based detection protocol based on the method of Xie et al. (Xie et al (2004) Anal. Chem. 76: 1611-1617) was developed, as illustrated diagrammatically in Figure 4. After the biotin-labelled target oligonucleotide was hybridized onto the CP-biointerface, GOD-A and PVA-Os were subsequently introduced by providing an aliquot on the surface of the PEDOT films or by soaking the electrode in the solution. The amperometric signal was calculated from the difference in current output between soaking the electrodes in PBS buffer, and soaking the electrodes in PBS buffer containing 40 mM of glucose.

Experiments were first conducted on films with 0.2% C₂-EDOT-COOH coated on Au substrate with a lower GOD-A concentration (5 μg/mL) in the detection protocol. A much greater amperometric signal (84 nA) with complementary target hybridization was obtained, compared to the signal (25 nA) from control experiment using non- complementary targets (Figure 6). An even greater difference in signal was attained between complementary (128 nA) and non-complementary (24 nA) targets when Pt substrate was employed. This was attributed to the better work function matching between Pt and PEDOT. When a high GOD-A concentration was used, enhanced signals were achieved in both complementary targets (151 nA and 222 nA for Au and Pt substrates, respectively) and non-complementary targets (50 nA and 55 nA for Au and Pt substrates, respectively). Comparing these results, a low GOD-A concentration was preferred as the low control signal allowed for improved detection limits. The comparison between signal outputs from different surface probe densities also agreed with the theoretical prediction and results from fluorescent detection.

PEDOT films containing 0.2% C₂-EDOT-COOH consistently yielded the largest signal output. This study demonstrated the feasibility of applying PEDOT biointerfaces for fluorescent and electrochemical DNA detection. Compared to SAM or other conducting polymer-based methods for constructing DNA-grafted conducting surfaces, PEDOT biointerfaces are advantageous because they are uniform, compositionally tunable, and allow for efficient manufacturing. By varying the molar fraction of C₂-EDOT-COOH in the monomer mixture, it is possible to synthesize PEDOT biointerfaces with different CP densities.

Example 3. Immobilization of capture probe antibodies and detection of antibody binding The electrochemical detection of a "sandwich-type" antibody reaction was examined. A schematic illustration of this electrochemical antibody detection system is shown in Figure 8.

Films were generated as described in Example 2. C₂-EDOT-COOH was activated by NHS/EDC coupling and goat anti-rat IgG (10 μg/ml, Rat IgG Elisa Quantitation Kit, BETHYL Laboratories, Inc.) was coupled to the activated polymer thin films for one hour. No functionalisation of the immunoglobulin was required. After washing the films with washing buffer (5OmM tris, 0.14 M NaCl, 0.05% Tween 20), the films were exposed to a solution containing an excess of bovine serum albumin to block any unoccupied functionalisation sites on the activated film. After rinsed with washing buffer, target rat immunoglobulin (500 ng/ml) (Rat

Reference Serum, Rat IgG Elisa Quantitation Kit, BETHYL Laboratories, Inc.) was incubated with the films for one hour. After another washing, the films were exposed to a labelled antibody (goat anti-rat Ig conjugated to HRP) (20 ng/ml, Rat IgG Elisa Quantitation Kit, BETHYL Laboratories, Inc.) and incubated for one hour. After a further rinse, the electrochemical detection of bound labelled antibody using the redox polymer as described in Example 2 was carried out.

Electrochemical characterization was carried out with a gold electrode with Ag/AgCl electrode as reference electrode and platinum wire as counter electrode. The amperometric signal (at potential = 0.15 V) was calculated from the difference in current output between soaking the electrodes in PBS buffer, and soaking the electrodes in PBS buffer containing 5 mM H₂O₂. An example of the results of these experiments is presented in Figure 9. The presence of HRP-labelled antibody, indicating the binding of rat immunoglobulin to the capture antibody immobilized to the film of the sensor, was readily detectable electrochemically using this system.

Claims

Claims:

1. A sensor surface for use in detecting a biological molecule, said sensor surface comprising: i) an ultrasmooth film of electrically conductive polymer, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule.

2. The sensor surface of claim 1 wherein the polymer comprises at least one type of monomer unit which is derived from a derivative of ethylendioxythiophene (EDOT).

3. The sensor surface of claim 2 wherein the derivative of EDOT comprises an EDOT group coupled to a functional group selected from the group consisting of a hydroxyl group, a thiol group, an amine group, a carboxylic acid group, an azide group, an N-hydroxysuccinimidoester group and a maleimido group.

4. The sensor surface of any one of claims 1 to 3 wherein the polymer is a copolymer.

5. The sensor surface of any one of claims 1 to 4 wherein the polymer comprises monomer units derived from EDOT-OH and C₂-EDOT-COOH.

6. The sensor surface of any one of claims 1 to 5 wherein the film is prepared by electropolymerisation in a microemulsion.

7. The sensor surface of any one of claims 1 to 6 wherein the film is on the surface of an electrode.

8. The sensor surface of any one of claims 1 to 7 wherein the capture agent is coupled to the film by means of an amide linkage, an ester linkage, a 1,2,3-triazole linkage or a succinimido linkage.

9. The sensor surface of any one of claims 1 to 8 wherein the capture agent comprises an oligonucleotide or a polynucleotide.

10. The sensor surface of any one of claims 1 to 8 wherein the capture agent comprises an antibody or an antigen-binding antibody fragment.

11. The sensor surface of any one of claims 1 to 10 wherein the film has a surface root- mean-square roughness R_nH8 of less than or equal to about lOnm.

12. The sensor surface of any one of claims 1 to 11 wherein the film has a thickness of less than or equal to about 100 rnn.

13. A sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding.

14. The sensor system of claim 13 wherein the detectable signal is an electrical signal or a fluorescent signal.

15. The sensor system according to claim 13 or claim 14 wherein the label is capable of coupling to the biological molecule.

16. The sensor system of any one of claims 13 to 15 wherein the label is coupled to the sensor surface.

17. The sensor system of claim 16 wherein the label is coupled to the capture agent.

18. The sensor system of claim any one of claims 13 to 17 wherein the film of electrically conductive polymer is on the surface of an electrode.

19. A method of detecting a biological molecule comprising: i) providing a sensor surface, said sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) exposing said sensor surface to the biological molecule in the presence of a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal in response to said binding; and iii) detecting the detectable signal.

20. The method of claim 19 wherein the detectable signal is an electrical signal or a fluorescent signal.

21. The method of claim 19 or claim 20 wherein step ii comprises combining the biological molecule with the label to form a labelled sample, and exposing the sensor surface to the labelled sample.

22. The method of claim 19 wherein the label is coupled to the sensor surface.

23. The method of claim 19 wherein step ii comprises: ii') exposing said sensor surface to the biological molecule so as to bind the biological molecule to the capture agent; and ii") exposing the bound biological molecule to the label, thereby producing a detectable signal.

24. The method of any one of claims 19 to 23 wherein the method is a method for quantifying the biological molecule, and step iii comprises quantifying the detectable signal.

25. The method of any one of claims 19 to 24 wherein the film of electrically conductive polymer is on the surface of an electrode, and the step of detecting or quantifying the detectable signal comprises detecting or quantifying an electrical signal from the electrode.

26. Use of a sensor surface comprising: i) an ultrasmooth film of electrically conductive polymer, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; in the detection of a biological molecule.

27. A process for making a sensor surface for detecting a biological molecule, said process comprising: i) providing an ultrasmooth film of electrically conductive polymer; and ii) coupling a capture agent to said film.

28. The process of claim 27 wherein the polymer comprises at least one type of monomer unit which is derived from a derivative of ethylendioxythiophene (EDOT).

29. The process of claim 27 or claim 28 wherein step i) comprises: a) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the microemulsion; and b) electropolymerizing the one or more monomers in the microemulsion to form the film.

30. The process of claim 29 wherein the one or monomers are insoluble or only sparingly soluble in water.

31. The process of claim 29 or claim 30 wherein the dispersed phase additionally comprises one or more comonomers capable of copolymerising with the one or more monomers in order to produce an electrically conductive copolymer.

32. The process of any one of claims 29 to 31 wherein step a) comprises forming a microemulsion from the one or more monomers, and, if present, the one or more comonomers, together with a surfactant, an acid and an electrolyte.

33. The process of claim 32 wherein the surfactant is present below its critical micelle concentration.

34. The process of any one of claims 29 to 33 wherein step b) comprises applying an electrical potential between two electrodes, each of which is at least partially immersed in the microemulsion, whereby the film of electrically conductive polymer forms on at least one of said electrodes.

35. The process of any one of claims 29 to 34 wherein step b) comprises applying a cyclic potential to the microemulsion.

36. The process of any one of claims 27 to 35 wherein step ii comprises reacting the capture agent with a functional group on the surface of the film.

37. A process for making a sensor surface for detecting a biological molecule, said surface comprising: i) an ultrasmooth film of electrically conductive polymer, and ii) a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; said process comprising: a) providing an acidic microemulsion having a continuous phase and a dispersed phase, wherein one or more monomers are located in the dispersed phase of the microemulsion; and b) electropolymerizing the one or more monomers in the microemulsion to form the sensor surface; wherein at least one of the monomers comprises EDOT coupled to the capture agent.

38. A process for making a sensor system for detecting a biological molecule, said sensor system comprising: i) a sensor surface comprising an ultrasmooth film of electrically conductive polymer, and a capture agent coupled to the film, wherein the capture agent is able to selectively bind to the biological molecule; and ii) a label which is able to detect binding of the biological molecule to the capture agent, thereby producing a detectable signal from the sensor system in response to said binding; said process comprising: a) providing the sensor surface; and b) coupling the label to the sensor surface.