WO2009122159A2 - Biosensor for analyte detection - Google Patents

Abstract

This invention relates to biosensors for detecting analyte molecules, such as nucleic acids, in a sample solution. The biosensors comprise a capture electrode which has probe molecules immobilised at its surface which specifically bind to analyte molecules. The probe molecules are separated on the surface of the capture electrode by spacer molecules and the proportion of probe molecules to spacer molecules is less than 10%. Changes in the electrical properties of the capture electrode surface are indicative of the binding of the immobilised probe molecules to analyte molecules. Biosensors as described may be useful for example, in diagnostic, genomic and forensic applications.

Description

Biosensor for Analyte Detection

This invention relates to biosensors for the detection of analyte molecules, such as nucleic acids, in a sample solution. This may be useful for example, in diagnostic, genomic and forensic applications.

Low-cost portable DNA biochips are in high demand in many fields, including medical diagnostics, genomics and forensics. Label-free electrochemical detection based upon the intrinsic charge of the DNA phosphate backbone provides a range of promising techniques to fulfil this demand. Of these techniques, Electrochemical Impedance Spectroscopy (EIS) is very effective for the characterization of bio- functionalized electrodes and the detection of DNA hybridization. Hybridization of target DNA to immobilized probe DNA results in an increased negative charge at the sensor surface. This results in an increased electrostatic barrier for negatively charged redox molecules in solution, causing a change in charge transfer resistance R_ct that can be detected using EIS (Bardea et al., 1999).

The signal, corresponding to the fractional change in R_ct obtained upon hybridization, is often relatively small and requires amplification. This can be achieved by post-affinity attachment of charged and/or bulky particles to form secondary superstructures (Bardea et al., 1999; Patolsky et al., 2000), or the attachment of enzymes that catalyse the formation of an insoluble product, which precipitates on the sensing surface to further increase R_ct (Patolsky et al., 1999). For example, Patolsky et al. (2000) report a 50% increase in R_ct upon hybridization of 5 μM 27-mer target DNA with 14-mer probe DNA. Secondary binding of a biotinylated-oligonucleotide, complementary to the residual sequence of the target DNA left unhybridized, allowed the formation of a superstructure of avidin and biotinylated liposomes. This amplified the charge transfer resistance change to 400%. However, these additional amplification steps increase the complexity and cost of the detection system. Thiol-modified peptide nucleic acid (PNA) probes immobilized onto gold and backfilled with mercaptohexanol (MCH) to form a mixed self- assembled monolayer (SAM) results in a 400% increase in charge transfer resistance upon hybridization with 0.5μM target DNA, without any additional amplification steps (Liu et al. 2005).

The present inventors have developed biosensors with improved analyte sensitivity by controlling the density of the probe which is immobilised on the electrode. This amplifies the changes in the electrical properties of the electrode that occur when probe binds to analyte on the electrode surface and allows sensitive analyte detection without additional amplification steps.

An aspect of the invention provides a biosensor for detecting an analyte comprising; a capture electrode having probe molecules and spacer molecules immobilised at the surface thereof, wherein the probe molecules are separated on the surface by the spacer molecules and the proportion of probe molecules to spacer molecules is less than 10%, and; wherein the probe molecules specifically bind to the analyte to form a probe : analyte complex, the capture electrode being connectable to a detector which measures changes in the electrical properties of the capture electrode surface indicative of the formation of the probe: analyte complex.

Other aspects of the invention relate to methods of detecting analyte using a biosensor. For example, a method of detecting an analyte in accordance with the invention may comprise; providing a capture electrode having probe molecules and spacer molecules at the surface thereof, wherein the probe molecules are separated on the surface by spacer molecules and the proportion of probe molecules to spacer molecules is less than 10%; contacting the capture electrode with a sample solution such that analyte in the solution specifically binds to the probe molecules to form probe : analyte complexes at the electrode surface, and; measuring the changes in the electrical properties of the capture electrode indicative of the formation of probe : analyte complex at the electrode surface.

The analyte may be any target molecule whose detection in a sample solution is desired. Suitable analytes include small molecules, such as vitamins and monosaccharides, and biological molecules, such as peptides, proteins, polysaccharides and nucleic acids, such as DNA and RNA.

Peptide or protein analytes may include antibodies, antigens, enzymes, substrates, hormones or cytokines.

Nucleic acid analytes may include DNA, RNA or cDNA molecules having a target nucleotide sequence. The target nucleotide sequence may be a genomic sequence or a transcribed nucleic acid sequence (e.g. mRNA or cDNA) . In some embodiments, the nucleic acid analyte may be amplified before detection as described herein, for example by PCR techniques, such as asymmetric PCR.

The capture electrode may be any conductive structure which allows the passage of current to or from the sample and/or measurement solution. Suitable capture electrodes are well known in the art and may be metal, for example a noble metal such as gold, silver, copper, platinum, palladium, ruthenium and iridium or alloys thereof. Other suitable capture electrodes may be carbon (or carbon paste) , conductive semiconductor-based compounds, conductive polymers and porous conducting materials. The probe molecule is a molecule or molecular complex which binds specifically to the analyte i.e. the probe molecule binds to the analyte and shows no significant binding to other molecules in the sample solution. For example, the probe molecule may bind to nucleic acid analyte having the target nucleotide sequence and show little or no binding to other nucleic acid molecules in the sample solution.

Suitable probe molecules may include specific binding members, such as antibodies (or antigens) , substrates and cofactors, which bind to the analyte; and nucleic acids or nucleic acid analogues which specifically hybridize to the analyte.

The probe molecule is selected such that the charge or charge distribution at the capture electrode surface changes when the probe molecule binds to the analyte. Preferably, one of the probe molecule and the probe: analyte complex is charged and the other is uncharged or the probe molecule and the probe: analyte complex have opposite charges. In other words, the analyte may have the opposite charge to the unbound probe molecule; the probe : analyte complex may be uncharged and the unbound probe molecule charged; or the probe: analyte complex may be charged and the unbound probe molecule uncharged. For example, for detection of a charged analyte in a sample solution, the probe molecule may be uncharged.

In some preferred embodiments, the probe molecule is uncharged and the analyte is a negatively charged molecule, such as a nucleic acid. Binding of the negatively charged analyte to the uncharged probe molecule leads to the formation of a negatively charged probe: analyte complex on the capture electrode surface. The increased negative charge on the electrode surface caused by the formation of a negatively charged probe : analyte alters the electrical properties of the electrode and these changes in properties may be detected, for example by potentiometric, amperometric, voltammetric or impedance techniques, such as EIS.

For example, an increased negative charge on the electrode surface may increase the R_ct of the capture electrode for negatively charged redox couples or reduce the R_ct of the capture electrode for positively charged redox couples. The change in R_ct may be detected by measuring the EIS characteristic over a wide frequency range, measuring a change in impedance at a fixed or limited range of frequencies, or by amperometric techniques such as cyclic voltammetry, chronoamperometry, square wave voltammetry or pulse voltammetry.

Suitable uncharged probe molecules which specifically bind to analyte nucleic acid, such as DNA and RNA molecules, include nucleic acid analogues.

Nucleic acid analogues suitable for use as probe molecules may comprise oligonucleotide backbones which are modified to be non- negatively charged, for example by removal or replacement of phosphate groups. A nucleic acid analogue may comprise a modified oligonucleotide backbone formed by short chain alkyl or cycloalkyl internucleoside linkages, mixed heteroatom and alkyl or cycloalkyl internucleoside linkages, or one or more short chain heteroatomic or heterocyclic internucleoside linkages. These include those having morpholino linkages (formed in part from the sugar portion of a nucleoside) ; siloxane backbones; sulfide, sulfoxide and sulfone backbones; formacetyl and thioformacetyl backbones; methylene formacetyl and thioformacetyl backbones; alkene containing backbones; sulfamate backbones; methyleneimino and methylenehydrazino backbones; sulfonate and sulfonamide backbones; amide backbones; and others having mixed N, 0, S and CH₂ component parts.

In other suitable nucleic acid analogues, both the sugar and the internucleoside linkage, i.e., the backbone, of the nucleotide units may be replaced with novel groups. The base units are maintained for hybridization with the analyte nucleic acid. For example, in preferred embodiments, the nucleic acid analogue is peptide nucleic acid (PNA) or a derivative thereof.

Peptide nucleic acid (PNA) is a non-naturally occurring nucleic acid analogue in which the sugar-phosphate backbone of the nucleic acid is replaced with a pseudopeptide backbone composed of repeating N- (2- aminoethyl) -glycine units (Egholm et al., 1993). The nucleobases are retained and are bound directly or indirectly to aza-nitrogen atoms of the amide portion of the backbone, for example by methylene carbonyl bonds. Because it lacks phosphate groups, PNA is uncharged. The synthesis and use of PNA is well known in the art (Nielsen, 2004). PNA molecules suitable for use as probe molecules are commercially available (e.g. Panagene, S.Korea).

A PNA probe molecules may be at least 5, at least 6, at least 7, at least 8, at least 9, or at least 10 bases in length. A PNA probe molecules may be up to 20, up to 23, up to 25 or up to 30 bases in length. For example, a PNA probe molecule may be from 7 to 16 bases.

The base sequence of the PNA probe molecule allows specific hybridization to the nucleic acid analyte. For example, the sequence of the PNA probe molecule may be complementary to the nucleotide sequence of the nucleic acid analyte or may include 5 or less, 4 or less, 3 or less, 2 or less or 1 base mismatches.

In other embodiments, the probe molecule may be negatively charged and the analyte may be a positively charged molecule. Binding of the positively charged analyte to the negatively charged probe molecule leads to a reduction in the negative charge on the capture electrode surface. This decreased negative charge on the electrode surface may for example, reduce the R_ct of the capture electrode for a negatively charged redox couple. Probe molecules suitable for detecting positively charged analyte may have negatively charged backbones and may include nucleic acids such as DNA. Examples of suitable probe molecules include DNA aptamers and other DNA molecules that bind to non-nucleic acid analytes.

The capture electrode may be contacted with the sample solution under conditions which allow the probe molecule to specifically bind to analyte, if present in the sample solution. Optionally, the conditions may be sufficiently stringent to prevent non-specific binding of molecules in the sample solution to the electrode surface.

After contact with the sample solution, the capture electrode may be washed under stringent conditions to remove molecules which are bound non-specifically to the electrode surface. Suitable conditions for removal of non-specific binding are well-known in the art and may readily be determined by the skilled person for any particular biosensor arrangement. For example, stringent conditions for specific binding of a PNA probe to a nucleic acid analyte may include room temperature in 20OmM phosphate buffer, 40OmM K₂SO₄ pH 7.0; 42°C in

0.25M Na₂HPO₄, pH 7.2, 6.5% SDS, 10% dextran sulfate and a final wash at 55°C in 0. IX SSC, 0.1% SDS; or 65°C in 0.25M Na₂HPO₄, pH 7.2, 6.5% SDS, 10% dextran sulfate and a final wash at 60⁰C in 0.1X SSC, 0.1% SDS.

The probe molecule may be immobilised on the surface of the capture electrode by any convenient technique. The probe molecule may be immobilized directly onto the electrode, for example in a single chemical step. Alternatively, a functional group may be immobilized on the electrode, for example within a monolayer and the probe may then be attached to the functional group. Preferably, the probe molecule comprises a terminal attachment group which facilitates attachment of the probe molecule to the surface of the capture electrode. Suitable terminal attachment groups include thiol groups which allow direct immobilisation of the probe molecule on the capture electrode surface by chemisorption. Other suitable terminal attachment groups may allow attachment of the probe molecule to functional groups immobilised on the capture electrode surface.

In preferred embodiments, the probe molecules are uniformly distributed on the capture electrode surface i.e. the probe molecules are immobilised at a substantially constant density all over the surface of the electrode.

The probe molecules are separated on the surface of the capture electrode by spacer molecules. The spacer molecules control the density of the probe molecule on the electrode surface and support charge transfer from redox molecules in the solution to the electrode. The separation of the probe molecules creates nano-channels between the probe molecules which allow redox molecules to approach sufficiently close to the surface of the capture electrode to enable the transfer of charge from the solution to the capture electrode, for example, within tunnelling distance. This charge transfer may be modulated, for example by changes in charge or charge distribution at the electrode surface or by the blocking or opening of conductance paths, which are caused by the presence of probe:analyte complex.

Charge may be transferred from redox molecules in the solution to the electrode by tunnelling, conductive spacer molecules, or other conductive molecules in the mixed monolayer on the electrode surface. Charge may also be directly transferred at pinholes/defect sites on the electrode surface.

In some embodiments, the spacer molecule may support the conductance of electrons between the solution and the electrode by being itself conductive or may be comprised within a mixed monolayer which contains conductive molecules which support the conductance of electrons between the solution and the electrode.

In other embodiments, the spacer molecule may be non-conductive and support charge transfer between the solution and the electrode by tunnelling. A suitable spacer molecule is preferably sufficiently small to allow tunnelling and may, for example, comprise a substituted or unsubstituted alkyl chain of less than 12 carbon atoms.

The spacer molecule may also be useful in preventing or reducing interactions between the probe molecule and the electrode and nonspecific binding of analyte or other molecules (e.g. proteins) in the sample solution and the electrode surface.

Suitable spacer molecules may be uncharged or charged and conductive or non-conductive, depending on the context.

In some preferred embodiments, spacer molecules may be uncharged non- conductive molecules, for example molecules having an alkyl chain and a polar end group, such as alkyl alcohols. A spacer molecule may have the formula CH₃ (CH₂) _nR, where n is 2 to 12, more preferably 2 to 9, for example CH₃(CH₂)_SR, and R is a polar group. For example, an alkyl alcohol spacer molecule may have the formula CH₃(CH₂J_nOH, where n is 2 to 12, more preferably 2 to 9, for example CH₃ (CH₂) ₅0H.

The spacer molecule preferably comprises a terminal attachment group which facilitates direct attachment to the surface of the capture electrode. A spacer molecule may have the formula R₂CH₂ (CH₂) _nRχ, where n is 2 to 12, and Ri is a -H or a polar group and R₂ is an attachment group. Suitable terminal attachment groups include thiol groups which allow immobilisation by chemisorption A suitable spacer molecule may have the formula SH (CH₂) _nOH, where n is 2 to 12, for example SH (CH₂) ₆0H (MCH) . Preferably, the spacer molecules do not bind or associate with the analyte or other molecules in the sample solution. In some embodiments, the spacer molecule may comprise a terminal group which reduces or prevents non-specific adsorption. Suitable terminal groups include oligo (ethylene glycol) (OEG) groups, methyl (CH₃) groups, carbonyl (COOH) groups or amine (NH2) groups.

The data herein shows that the relative densities of the probe and spacer molecules on the capture electrode surface exert a significant effect on the change in electrical properties, such as R_ctr which occurs in the presence of analyte. Preferably, the ratio of probe molecules to spacer molecules on the electrode surface is less than 10%, less then 5%, less than 2% or less than 1%. The ratio of probe molecules to spacer molecules on the electrode surface may be greater than 0.01%, 0.02%, 0.05%, or 0.1%. In some preferred embodiments, the optimum surface ratio of probe molecules to spacer molecules on the electrode is about 0.6%.

The probe molecules and spacer molecules may form a mixed monolayer on the capture electrode surface i.e. a single layer which comprises both the probe and the spacer molecules. Typically, the probe molecules are longer than the spacer molecules and extend beyond the spacer molecules in the monolayer.

In some preferred embodiments, the changes in the electrical properties of the electrode caused by the formation of the probe : analyte complex are detected without additional amplification steps. For example, no additional signal amplification molecules may be bound to the probe: analyte complex before or during detection to amplify the changes in the electrical signal. In other words, the presence on the electrode surface of the probe: analyte complex itself (i.e. a complex consisting of the probe molecule and the analyte) may be detected. Signal amplification molecules are well known in the art and include bulky or charged particles which form secondary structures on the probe : analyte complex and enzymes which catalyse the formation of signal amplifying products (e.g. insoluble products which precipitate on the electrode to increase R_ct) (see for example Bardea et al 1999, Patolsky et al 1999, Patolsky et al 2000) .

A suitable mixed monolayer may be produced on the surface of the capture electrode by co-immobilising probe molecules and spacer molecules onto the surface. Preferably, the probe molecules and spacer molecules are co-immobilised simultaneously i.e. in a single step. For example, a monolayer preparation solution comprising probe and spacer molecules may be contacted with the surface of the electrode under appropriate conditions such that the probe molecules and spacer molecules react with the electrode surface and become immobilised on the surface. For example, probe molecules and spacer molecules containing thiol groups may be immobilised by chemisorption onto a metal capture electrode surface by incubating the surface in the monolayer preparation solution at room temperature.

The monolayer preparation solution may be produced by mixing probe molecules and spacer molecules in an appropriate solvent at a mole fraction suitable to provide the desired surface ratio of probe to spacer on the electrode. For example, for a PNA probe molecule, the mole fraction of probe in the monolayer solution may be 1 to 20% or 5% to 15%, preferably about 10%.

Following the simultaneous co-immobilisation of the probe molecules and spacer molecules onto its surface, the electrode may be washed and/or further treated. For example, the surface of the capture electrode may be backfilled with spacer molecules after immobilisation of the monolayer to ensure complete coverage of the electrode surface and/or ultrasonicated to remove non-specifically adsorbed probe or spacer. A biosensor as described herein may comprise multiple capture electrodes for detecting analyte. A biosensor may comprise two or more capture electrodes which detect the same analyte. For example, a biosensor may comprise 2, 3, 4, 5, 6, 7, 8, 9, or 10 or more capture electrodes for an analyte. In some embodiments, all the capture electrodes in the biosensor may detect the same analyte.

A biosensor may comprise two or more capture electrodes which detect different analytes. For example, the biosensor may comprise one or more electrodes which detect a first analyte and one or more electrodes detect a second analyte. The number of different analytes which may be detected by different capture electrodes in the biosensor will depend on the particular application of the biosensor. For example, 5 or more, 10 or more, 20 or more, 50 or more, 100 or more or 1000 or more different analytes may be detected. The presence of a particular analyte in a sample solution may be detected by a change in the electrical properties at the surface of the capture electrode (s) which comprises probe molecules specific for that analyte.

The capture electrode (s) may be positioned in a circuit which comprises additional electrodes and is adapted to form an electrochemical cell when the electrodes are contacted with suitable electrolyte solutions. Any suitable type and arrangement of electrodes may be employed and the skilled person is familiar with suitable electrochemical systems. For example, the biosensor may comprise a counter electrode and/or a reference electrode. The use of counter electrodes and reference electrodes in two and three electrode cells is well-known in the art. Additional electrodes may also be added for other purposes.

The biosensor may further comprise one or more microfluidic devices for dispensing sample and/or measurement solutions, mixing, and washing and/or rinsing the capture electrode. Suitable devices are well known in the art (Nguyen and Wereley, 2006) . The biosensor may further comprise a detector which is connectable or connected to the circuit comprising the capture electrode (s) . The detector may detect or measure changes in the electrical properties of the capture electrode which are caused by the formation of the probe : analyte complex on its surface.

The change in electrical properties may be measured by comparing the electrical properties of the capture electrode before and after contact with the sample solution. The electrical properties of the capture electrode may be determined from the electrical signal which is generated by the reaction between the electrode and redox molecules in an electrochemical cell.

In some preferred embodiments, the detector measures changes in charge and/or charge distribution at the capture electrode surface. Changes in charge and/or charge distribution may be measured, for example, by measuring changes in impedance, current or potential in an electrochemical cell comprising the capture electrode.

The detector may be adapted to detect changes in R_ct, impedance, capacitance, current, potential or field effect in the cell which are caused by the formation of the probe : analyte complex at the capture electrode surface. For example, the detector may comprise a potentiostat, current measurement device, field effect device, electrometer, instrument amplifier or a capacitance measurement device .

The detector may be connected or connectable to a processor. The processor may be programmed to record, process and/or analyse the changes in potential, field effect, impedance or current detected by the detector. Changes in any of these parameters are attributable to the interaction of probes molecules on the capture electrode surface with analyte. For example, the processor may determine from the detected changes the R_ct of the capture electrode and/or the presence, identity and/or amount of analyte in the sample.

To detect analyte in the sample solution, the capture electrode is first contacted with the sample solution. Analyte in the sample solution specifically binds to the probe molecule on the capture electrode surface to form a probe : analyte complex. As described above, the formation of the probe : analyte complex causes a change in the electrical properties of the capture electrode which may then be detected in any convenient manner. For example, changes in R_ct, impedance, capacitance, current, potential or field effect in an electrochemical cell comprising the capture electrode may be detected.

Changes in the electrical properties of the capture electrode (e.g. one or more of R_ct, impedance, capacitance, current, potential or field effect) may be caused by changes in the surface charge and/or charge distribution of the capture electrode.

Alternatively, changes in the electrical properties of the capture electrode (e.g. one or more of R_ctr impedance, capacitance, current, potential or field effect) may be caused by conductive paths being obstructed or opened as a consequence of the probe: analyte complex formation, without change in charge on the electrode surface.

In some embodiments, after contacting the capture electrode with the sample solution under conditions which allow formation of the probe : analyte complex, changes in the electrical properties of the capture electrode indicative of the presence of probe : analyte complex (for example caused by changes in surface charge and/or charge distribution) are detected whilst the capture electrode is still in contact with the sample solution.

In other embodiments, after contacting the capture electrode with the sample solution under conditions which will allow formation of the probe : analyte complex, the capture electrode is removed from the sample solution, optionally washed, for example at high stringency, and contacted with a measurement solution. Changes in the electrical properties of the capture electrode indicative of the presence of probe : analyte complex (for example caused by changes in surface charge and/or charge distribution) may be detected when the solution is in contact with a measurement solution.

The use of separate sample and measurement solutions allows the individual optimisation of each solution. The sample solution may be optimised for specific binding between the probe molecule and analyte. For example, for nucleic acid analytes, a high ionic strength may be preferred to overcome electrostatic repulsion between analyte molecules. For example, the ionic strength of the sample solution may be greater than 5OmM, greater than 10OmM, greater than 20OmM, greater than 30OmM or greater than 40OmM. Typically for a PNA probe molecule, the ionic strength of the sample solution may be greater than 40OmM.

The measurement solution may be optimised for detection of probe: analyte complex at the capture electrode surface. For example, decreasing the ionic strength of the measurement solution results in an increased Debye charge screening length. This increases the change in electrostatic barrier at the capture electrode surface when an uncharged probe molecule (e.g. PNA) binds to a negatively charged analyte (e.g. DNA), thereby amplifying the change in R_ct caused by analyte binding. The ionic strength of the measurement solution may be less than 2OmM, more preferably 0. ImM to 1OmM.

As described above, the sample solution and the measurement solution may be the same (i.e. changes in the surface charge and/or charge distribution are detected when the capture electrode is in the solution which is being tested for the presence of analyte) or the sample solution and the measurement solution may be different (i.e. changes in the surface charge and/or charge distribution are detected when the capture electrode is in a different solution to the sample solution being tested for the presence of analyte) .

The solution which is in contact with the capture electrode when the changes in electrical properties are detected (i.e. the sample solution or more preferably the measurement solution) may comprise one or more redox active molecules which transfer charge to and from the electrode. Suitable redox active molecules include single redox species or redox couples .

For example, in embodiments in which R_ct is measured, the solution may comprise a redox couple which transfers charge to and from the electrode .

A redox couple comprises an oxidized and a reduced form of a charged redox molecule. The reduced form may lose an electron and convert to the oxidised form and the oxidised form may gain an electron and convert to the reduced form.

The charge of the redox couple may be the same as the charge of the probe : analyte complex (e.g. negative for PNA/DNA). However, when a charged spacer molecule of opposite sign to the redox couple is employed, the redox couple may be the same sign as the probe: analyte complex.

The amount of charge of the redox couple affects the sensitivity of detection, since the effect of the electrostatic barrier increases as the charge of the redox couple increases.

Many suitable redox couples are known in the art and include ferri/ferrocyanide [Fe (CN) ₆] ^3~/4~, [Mo(CN)₆]^3"74" and ruthenium hexamine .

In embodiments in which cyclic voltammetry is measured, the solution which is in contact with the capture electrode when the changes in electrical properties are detected (i.e. the sample solution or more preferably the measurement solution) may comprise a single redox species which transfers charge to and from the electrode.

Another aspect of the invention provides an array comprising multiple biosensors as described herein. An array is an arrangement of multiple biosensors, each biosensor suitable for detection of a different analyte. Each of the biosensors in the array may comprise one electrode or multiple electrodes for detecting their respective analytes, as described herein.

Each of the biosensors in the array may be connected to a single processor for recording, analysis and characterisation of the detected signals. The analyte (s) in the sample may be identified from the biosensor (s) in the array which produces the signals.

The array may be contacted with a sample solution under conditions suitable for the binding of probe molecule and analyte. After a predetermined length of time, changes in the charge transfer resistance of the capture electrodes of the biosensors of the array may be detected. A change detected at a specific biosensor in the array identifies the analyte in the sample. The sample may be tested for the presence of multiple analytes simultaneously using the array.

Another aspect of the invention provides a method of producing a biosensor for detecting an analyte comprising; co-immobilising probe molecules and spacer molecules onto the surface of an electrode to form a mixed monolayer, wherein the mole fraction of probe molecules in the monolayer is less than 10% and the probe molecules specifically bind to analyte in an electrolyte solution to form a probe : analyte complex.

Preferably, probe molecules and spacer molecules are co-immobilised simultaneously. Probe molecules and spacer molecules may be simultaneously co-immobilised onto the surface of the electrode by contacting the surface with a monolayer preparation solution comprising probe and spacer molecules such that the probe molecules and spacer molecules react with the electrode surface and become immobilised. The mole fraction of probe in the monolayer preparation solution may be 5% to 15%, preferably about 10%.

Suitable electrodes are described above. Before immobilisation of the monolayer, the surface of the electrode may be prepared by conventional techniques, including polishing, sonication and electrochemical cleaning.

Alternatively, a metal electrode may be freshly evaporated or chemically cleaned or etched (e.g. piranha cleaned: 10 mins in 1:1 96% H₂SO₄: 30% H₂O₂)

Probe and spacer molecules are described elsewhere herein. The probe and spacer molecule may be immobilised through a terminal attachment group which forms a bond with the surface of the electrode. Suitable terminal groups include thiol groups which may bind by chemisorption to the surface of a metal electrode, preferably a gold electrode.

In some embodiments, the surface of the electrode may be backfilled with spacer molecules after immobilisation of the monolayer to ensure complete coverage of the electrode surface. A method may comprise the step of contacting the surface having the mixed monolayer immobilised thereon with a backfill solution comprising spacer molecules such that the spacer molecules react with any exposed electrode surface (i.e. surface not covered by the monolayer) and become immobilised.

The capture electrode thus produced may be connected or connectable to appropriate circuitry, additional electrodes and other components as described herein to produce the biosensor. In some embodiments, the probe molecule which is immobilised on the electrode may be bound to analyte in a probe :analyte complex. The analyte may then be removed to leave the immobilised probe molecule. A method of producing a biosensor for detecting an analyte may comprise; co-immobilising spacer molecules and complexes comprising probe molecules and analyte onto the surface of an electrode to form a mixed monolayer, wherein the mole fraction of probe : analyte complexes in the monolayer is less than 10% and; removing the analyte from the immobilised complexes, such that probe molecules remain immobilised in the mixed monolayer on the surface .

Preferably, the complexes and spacer molecules are co-immobilised simultaneously.

The analyte may be removed from the probe : analyte complex by any suitable technique. For example, the electrode surface may be washed under regeneration conditions which may include high temperature, high or low pH and/or surfactants.

Following removal of the analyte, the probe molecules on the surface are available to specifically bind to analyte in a sample solution.

Various further aspects and embodiments of the present invention will be apparent to those skilled in the art in view of the present disclosure. All documents mentioned in this specification are incorporated herein by reference in their entirety.

"and/or" where used herein is to be taken as specific disclosure of each of the two specified features or components with or without the other. For example "A and/or B" is to be taken as specific disclosure of each of (i) A, (ii) B and (iii) A and B, just as if each is set out individually herein. Unless context dictates otherwise, the descriptions and definitions of the features set out above are not limited to any particular aspect or embodiment of the invention and apply equally to all aspects and embodiments which are described.

Certain aspects and embodiments of the invention will now be illustrated by way of example and with reference to the figures and tables described below.

Figure 1 shows EIS characteristics for PNA/MCH functionalized electrode and after incubation with 1 μM fully complementary (FC) target DNA. Electrode prepared by co-immobilization with 5% PNA mole fraction, measured with 0.1 ioM K₄[Fe(CN)₆] + 0.I mM K₃[Fe(CN)₆] in 10 mM PB. Z_reai and Z_imaq are the real and imaginary parts of the impedance, respectively.

Figure 2 shows R_ct for PNA/MCH functionalized electrodes and fractional change in R_ct upon incubation with 1 μM fully complementary target DNA, as a function of fraction of PNA to total thiol concentration in immobilization solution. Measurements were taken in 10 mM PB supporting electrolyte. Error bars show the mean and spread for at least two samples at each PNA mole fraction.

Figure 3 shows charge transfer resistance for PNA/MCH mixed SAM and after incubation with fully complementary (FC) target DNA, as a function of measurement buffer ionic strength. Electrodes prepared with 5% PNA mole fraction. Error bars show the mean and spread for 3 samples at each ionic strength.

Figure 4 shows charge transfer resistance for ssDNA/MCH mixed SAM and after incubation with fully complementary (FC) target DNA, as a function of measurement buffer ionic strength. Electrodes prepared with 10% ssDNA mole fraction. Error bars show the mean and spread for 3 samples at each ionic strength.

Figure 5 shows fractional changes in R_ct upon incubation with fully complementary (FC) or non-complementary (NC) target, as a function of target concentration. Electrodes prepared with 5% PNA mole fraction and measured in 10 mM PB supporting electrolyte. Error bars show the mean and spread for 2 samples at each target concentration.

Figure 6 shows Nyquist plots for the detection of PCR products generated from a 1OnM template. Black squares indicate signal from a PNA/MCH functionalized electrode before exposure to the target; white circles indicate signal after exposure to the symmetric PCR products and white squares indicate signal after exposure to the asymmetric PCR products.

Experiments

Material and Methods

Sample Preparation DNA oligonucleotides were purchased from the Protein and Nucleic Acid

Chemistry facility of Cambridge University Biochemistry Department.

Thiol modified PNA oligonucleotides were purchased from Panagene,

Korea. All other chemicals used in the sample preparation were purchased from Sigma-Aldrich, UK. All aqueous solutions were prepared using 18.2 MΩ.cm ultra-pure water (Millipore, Billerica MA, USA) with a Pyrogard filter (Millipore) to remove nucleases.

Gold disk working electrodes with a radius of 1.0 mm (CH Instruments, Austin TX, USA) were polished with 0.3 urn aluminium oxide particles (Buehler, Lake Bluff IL, USA) on a polishing pad (Buehler) . Particles were removed by sonication in ultra-pure water, polishing on a blank polishing pad, and sonication in ultra-pure water. Electrodes were subsequently electrochemically cleaned in 0.5 M H₂SO₄ by scanning the potential between the oxidation and reduction of gold, -0.05 V and +1.1 V versus an Hg/Hg₂SO₄ reference electrode, for 60 cycles until there was no further change in the voltammogram. Electrodes were rinsed with deionised water, dried in a stream of nitrogen, and exposed to 25 μl of mixed PNA/MCH or DNA/MCH immobilization solution for 16 hours in a humidity chamber.

Probe single-stranded PNA (ssPNA) had the 15-base sequence AAT CCT CTT TGA CGA and was modified on the N-terminal (equivalent to 5 '-terminal of DNA) to give HS-(CH₂)_II-CO-NH-(CH₂CH₂O)₃-CH₂-CO-NH-SSPNA. A range of molar fractions of ssPNA and MCH were prepared in 50% dimethyl sulfoxide (DMSO) , 50% ultra-pure water (v/v) , with the concentrations chosen to ensure an excess of thiol molecules over surface sites. Electrodes were also prepared with only 1 μM thiol modified ssPNA, and with only MCH. A minimum of two samples were prepared at each PNA mole fraction.

After immobilization of PNA, electrodes were rinsed in 50% DMSO, 50% ultra-pure water. To ensure complete thiol coverage of the gold surface, the electrodes were backfilled with MCH by immersion in 1 mM MCH in 50% DMSO, 50% ultra-pure water for 1 hour. Electrodes were rinsed with 50% DMSO, 50% ultra-pure water and then ultra-pure water.

To enable comparison between PNA and DNA probes, electrodes functionalized with single-stranded DNA (ssDNA) probes were prepared using co-immobilization of thiol modified DNA and MCH (Keighley et al., 2008). Probe ssDNA had the 18-base sequence AAT CCT CTT TGA CGA CTC and was modified on the 5 '-terminal to give HS-(CH₂J₆-PO₄-(CH₂CH₂O)₆- ssDNA. The DNA immobilization buffer consisted of 0.8 M PB + 1.0 M NaCl + 5 mM MgCl₂ + 1 mM ethylene diamine tetraacetic acid (EDTA) pH 7.0. The high ionic strength and Mg²⁺ ions screen the DNA charge and allow high probe surface densities (Estrela et al., 2005; Boon et al., 2002) . After immobilization of DNA, electrodes were sequentially rinsed in: immobilization buffer, 200 mM PB, 10 itiM PB and finally 10 itiM PB + 10 mM EDTA to remove any remaining Mg²⁺ ions. Electrodes were backfilled with MCH using immersion in 1 mM MCH in ultra-pure water for 1 hour. Electrodes were then rinsed with ultra-pure water.

Prior to hybridization, electrodes were rinsed with 200 mM PB + 400 mM K₂SO₄ pH 7.0, and incubated with 25 μl of target DNA in 200 mM PB + 400 mM K₂SO₄ pH 7.0 for 2 hours. Target DNA was of 18-base fully complementary or unrelated sequence, GAG TCG TCA AAG AGG ATT and AGC ACA GGC TGA AAT GGT, respectively. Although the PNA-DNA duplex stability is almost independent of ionic strength, a high ionic strength hybridization buffer is used to overcome electrostatic repulsions between target oligomers as they hybridize to probes on the surface. Electrodes were then rinsed with 200 mM PB + 400 mM K₂SO₄ pH 7.0, and in PB of the concentration to be used for the subsequent characterization. Electrodes were characterized using EIS in the presence of ferri/ferrocyanide, and the charge transfer resistance determined before and after interaction with target DNA.

Sample Characterization

All measurements used a three-electrode cell, with an Hg/Hg₂SO₄ (K₂SO₄ sat.) reference electrode (Radiometer Analytical, Lyon, France) against which all potentials are quoted, and a Pt counter electrode (BAS, West Lafayette IN, USA) .

Electrodes were characterized using electrochemical impedance spectroscopy. Measurements were taken using either an Autolab PGSTAT302 potentiostat (Eco Chemie, Utrecht, Netherlands) or a Gamry Instruments Femtostat (Gamry Instruments, Warminster PA, USA) . The electrochemical impedance spectrum was measured in a solution of 0.1 mM K₄[Fe(CN)₆] + 0.I mM K₃[Fe(CN)₆] in phosphate buffer pH 7.0 in a range of concentrations. The reference electrode was connected via a salt bridge filled with PB of concentration equal to that in the cell. The impedance spectrum was measured over the frequency range 10 kHz to 10 mHz, with a 10 mV a.c. voltage superimposed on a d.c. bias set equal to the formal potential of the redox couple.

The charge transfer resistance R_ct was determined by fitting data to a Randies equivalent circuit, with the double layer capacitance C_dl replaced with a constant phase element (non-ideal capacitance) , in parallel with #_ct and a Warburg element that models diffusion (Bard and Faulkner, 2001) .

Results

Optimization of PNA probe surface density

Thiol modified PNA probes and mercaptohexanol were co-immobilized onto gold electrodes. The mole fraction of PNA to total thiol (PNA + MCH) in the immobilization solution was used to control the probe surface density. The charge transfer resistance for the PNA/MCH functionalized electrode was determined using electrochemical impedance spectroscopy. Typical EIS characteristics before and after incubation with target DNA are shown in Figure 1. The PNA/MCH SAM has a charge transfer resistance of 8.8 kΩ. This reflects formation of a good MCH SAM, and sufficient spacing of PNA probes such that the ferri/ferrocyanide can freely penetrate to the MCH SAM. Hybridization results in a 24-fold increase in R_ct to 219 kΩ, due to formation of an electrostatic barrier for the negative ferri/ferrocyanide.

The effect of PNA mole fraction on R_ct for the PNA/MCH SAM and on the fractional change in R_ct with target DNA hybridization is shown in Figure 2. Increasing the PNA mole fraction will increase the PNA surface density. As the PNA fraction is increased up to 50%, R_ct for the PNA/MCH SAM increases only slightly. The inclusion of the MCH spacer thiol allows the redox molecules to penetrate freely between the PNA probes, with the densely packed MCH SAM resulting in an approximately constant R_ct. However, when PNA is immobilized by itself, the resulting R_ct is around 350-fold larger than for a mixed SAM formed with 50% PNA. This is due to the PNA forming a relatively dense SAM in the absence of a spacer thiol (Mateo-Marti et al., 2005) that prevents charge transfer to the redox couple.

The fractional change in R_ct upon incubation with fully complementary target DNA increases significantly as the PNA fraction is increased, up to a 10% fraction of PNA. The increasing probe surface density results in an increasing surface density of hybridized target DNA and a larger electrostatic barrier for the ferri/ferrocyanide . Further increase of the PNA fraction results in a decrease of the fractional change in R_ct upon hybridization. This is presumably due to steric hindrances which reduce target hybridization. Although PNA forms a relatively compact SAM when immobilized by itself, some target DNA hybridization still occurs, as reported by Mateo-Marti et al. (2007). However, the fractional change in R_ct that results is small due to the large initial R_ct.

The maximum fractional change in R_ct with hybridization occurs at a 10% PNA fraction. A similar trend is observed when the concentration of the supporting electrolyte is reduced to 1 inM. In this case, the fractional change in R_ct is increased, with a maximum 385-fold increase in R_ct for 10% PNA. However, variability in R_ct is lower at PNA fractions of 5% or below. A 5% PNA fraction results in a 157-fold increase in R_ct upon hybridization, so is selected for further study.

Ionic strength can also be seen to be a key variable in the sensor response, and is investigated in more detail in the following section.

Optimization of measurement ionic strength As the ionic strength I is decreased, the Debye charge screening length increases, varying as I^'112. This increases the electrostatic barrier for ferri/ferrocyanide, decreasing the concentration of redox couple that can exchange charge at the electrode. This decreases the exchange current density J₀ and increases the charge transfer resistance. The charge transfer resistance is defined from the exchange current density (Schmickler, 1996) as:

where R is the molar gas constant, T is the absolute temperature and F is the Faraday constant. The exchange current density varies with the surface concentrations of redox species as:

where k₀ is the rate constant, c^s _red and c^s _ox the surface concentrations of the reduced and oxidized redox species respectively, and a is the anodic transfer coefficient.

The surface concentration c^s of redox species with charge number z that can undergo charge transfer varies with the interstitial potential φ_L in the regions between oligomers according to a Boltzmann factor:

c =c_oexp (3) kT

where C₀ is the bulk redox species concentration, e₀ is the electronic charge, k is the Boltzmann constant and potentials are taken with respect to the bulk electrolyte.

Since R_ct will vary approximately inversely with the surface concentration of redox species, it will vary exponentially with the electrostatic barrier. Although a theoretical analysis of the electrostatic barrier is difficult due to its three-dimensional nature, R_ct is expected to increase as I is reduced for the duplex electrode, but to be independent of ionic strength for the neutral PNA/MCH SAM.

The effect of measurement buffer ionic strength on charge transfer resistance for the PNA/MCH SAM and after DNA hybridization is shown in Figure 3. Decreasing ionic strength results in minimal R_ct change with the neutral PNA/MCH SAM, but a significant increase in i?_ct after hybridization, due to the increase in the electrostatic barrier for the ferri/ferrocyanide to reach the electrode surface. At high ionic strength, the DNA charge is mostly screened and there is little electrostatic barrier in the regions between probes. At low ionic strength, the screening length is greatly increased, resulting in a greater electrostatic barrier in the regions between probes (Piunno et al., 1999). For example, at 700 mM ionic strength a 24% change in R_ct is observed upon hybridization. By reducing the ionic strength to 2 mM at an optimized probe density, the fractional change is increased by a factor of 600. The change in R_ct is significantly amplified compared to previously reported results with DNA or PNA probes, and is achieved without the use of additional biochemical amplification steps.

For the case of ssDNA probes, the effect of measurement buffer ionic strength on charge transfer resistance for an ssDNA/MCH SAM and after DNA hybridization is shown in Figure 4. Electrodes were prepared by co-immobilization of ssDNA and MCH with a 10% DNA mole fraction. As the supporting electrolyte concentration is decreased, R_ct for both the ssDNA/MCH SAM and after hybridization increase similarly. Target DNA hybridization increases the charge at probe sites, but does not significantly change the spacing between charged sites on the electrode. Therefore, the change in screening length affects the electrostatic barrier for ferri/ferrocyanide approximately equally for both ssDNA and dsDNA. This results in the similar trends in R_ct. Decreasing the measurement ionic strength results in an initial increase in the fractional change in charge transfer resistance upon hybridization, but for ionic strengths of 60 mM or less the change remains approximately constant at 55±10%. The large enhancement of the fractional change in R_ct upon hybridization obtained by reducing the measurement ionic strength when using PNA probes is not achieved with DNA probes.

Detection limit of optimized sensor

A calibration curve for the optimized PNA probe sensor is shown in Figure 5, for 10 mM PB supporting electrolyte. A similar curve is obtained for 1 mM PB. To obtain a baseline for the minimum detectable specific interaction, samples were saturated with a high concentration (1 μM) non-complementary strand. The detection limit is below 1 nM for measurement in both 10 mM PB and 1 mM PB. With 10 mM PB, R_ct changes by 27+7% with 1 nM FC DNA, while only 2±1% change is observed with 1 μM NC DNA. The detection limit corresponds to 25 fmol target DNA in the 25 μl sample volume.

Reduction in electrode size and the consequent reduction of sample volume may further improve the detection limit.

Application of PNA assay to the detection of asymmetric PCR products

The detection of single stranded targets may be advantageous in signal maximization in the assays described herein.

We have demonstrated such an advantage by performing asymmetric PCR of the sequence; 5'- CAA AGA CAT CTT CAA GTC TCT GCG CGA TCT CGG CTT TGA GGG GGC CTG A- 3', which is part of the Avian Influenza M-gene sequence. The forward and reverse primer sequences were 5'- TCA GGC CCC CTC AAA GCC GAG ATC - 3' and 5'- CAA AGA CAT CTT CAA GTC TCT GCG - 3', respectively. The forward primer was chosen to amplify the sequence complementary to the PNA probe. The concentrations of template molecules and primers used are indicated in Table 1. All reagents were supplied by Sigma Aldrich. Electrochemical Impedance Spectroscopy measurements were carried out on bulk electrodes prepared as described above. The Nyquist plots for the 1OnM template are shown in Figure 1 for three cases: 1) Before exposure to the target (black squares); 2) After exposure to the symmetric PCR products (white circles); and 3) After exposure to the asymmetric PCR products (white squares) . The values of the charge transfer resistance in each case were extracted by fitting the experimental data to a Randies circuit, by using a commercial computer program. The ratio of the charge transfer resistances R_ct before and after exposure of the electrode to the target are shown in Table 1. These results demonstrate the advantage of asymmetric PCR using the biosensor described herein. Furthermore, in addition to the target DNA, PCR products include species such as dNTP, primers, enzymes that could potentially interfere with the detection process. These results demonstrate the high selectivity and robustness of the assay described herein.

Co-immobilization of thiol-modified PNA and mercaptohexanol may be used to form mixed self-assembled monolayers on gold. Specifying the solution mole ratio of the thiol components provides an effective and easily implemented method for probe surface density optimization. The maximum change in charge transfer resistance upon hybridization occurs at around 10% PNA mole fraction.

The electrostatic barrier for charge transfer to the ferri/ferrocyanide redox couple is approximately independent of measurement buffer ionic strength for neutral PNA/mercaptohexanol SAMs, but greatly increases with reduced ionic strength after hybridization with fully complementary target DNA. By reducing the measurement buffer ionic strength from 700 mM to 2 mM, the fractional change in R_ct upon hybridization is increased by a factor of 600.

The use of PNA probes, the optimization of PNA surface density and the optimization of measurement ionic strength are key to the success of electrochemical impedance spectroscopy for label-free detection of DNA hybridization. Optimization results in massive enhancement of the fractional change in R_ct upon hybridization, without the use of additional biochemical amplification steps. A fractional change 100- fold larger than previously reported results is achieved.

Optimization of PNA surface density in a PNA/mercaptohexanol SAM results in a small initial charge transfer resistance, defined by the mercaptohexanol SAM. A smaller charge transfer resistance gives a greater current density for a given electrode area and overpotential . For a given sensitivity of the detection electronics, both reducing the initial charge transfer resistance per unit area and increasing the fractional change in charge transfer resistance with hybridization enable a reduction of the minimum sensing electrode area. Reduction of sensor area and the accompanying reduction in the required sample volume, enable significantly enhanced detection limits.

A detection limit of 25 fmol target is demonstrated. This may be further improved by reduction of the electrode area and sample volume. For example, reducing the electrode diameter from 2 mm to 100 μm, a 400-fold decrease in area, would increase the initial R_ct to around 4 MΩ. 10 mV a.c. overpotential would result in an approximately 2.5 nA a.c. current, feasibly measured in a portable detection system. This would allow the sample volume to be scaled to 3 nl, reducing the detection limit to 3 amol. This shows electrochemical impedance spectroscopy to be a promising technique for portable DNA detection applications .

The novel immobilization strategy is relevant to a wide range of biosensors. In particular PNA probes may be used in biosensors to detect DNA. The results set out herein are also relevant for other detection techniques that rely on the variation of the intrinsic charge of DNA with hybridization, for example potentiometric detection with field-effect devices.

Table 1

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Claims

Claims :

1. A biosensor for detecting an analyte comprising; a capture electrode having probe molecules and spacer molecules immobilised at the surface thereof, wherein the probe molecules are separated on the surface by the spacer molecules and the proportion of probe molecules to spacer molecules is less than 10%, and; wherein the probe molecules specifically bind to the analyte to form a probe : analyte complex, the capture electrode being connectable to a detector which measures changes in the electrical properties of the capture electrode surface indicative of the formation of the probe: analyte complex.

2. A biosensor according to claim 1 wherein one of the probe molecule and the probe: analyte complex is charged and the other is uncharged or the probe molecule and the probe: analyte complex have opposite charges.

3. A biosensor according to claim 1 or claim 2 wherein the detector measures changes in charge and/or charge distribution at the capture electrode surface.

4. A biosensor according to claim 3 wherein the detector measures changes in impedance, current or potential which are indicative of said changes in charge and/or charge distribution.

5. A biosensor according to claim 1 or claim 2 wherein the detector measures changes in impedance, current or potential which are indicative of the probe-analyte complex obstructing or opening conductive paths at the capture electrode surface.

6. A biosensor according to any one of claims 1 to 5 wherein the probe molecules and spacer molecules form a mixed monolayer on the capture electrode surface.

7. A biosensor according to claim 6 wherein the mixed monolayer is produced by simultaneously co-immobilising probe molecules and spacer molecules onto the surface of the capture electrode.

8. A biosensor according to any one of claims 1 to 7 wherein the proportion of probe molecules to spacer molecules on the surface of the capture electrode is 0.01% to 10%.

9. A biosensor according to any one of the preceding claims wherein the probe molecules are uniformly distributed on the capture electrode surface.

10. A biosensor according to any one of the preceding claims wherein the probe molecules and spacer molecules are thiolated and immobilised on the capture electrode surface by chemisorption.

11. A biosensor according to any one of the preceding claims wherein the spacer molecules are uncharged.

12. A biosensor according to claim 11 wherein the spacer molecules are alkyl alcohols.

13. A biosensor according to claim 12 wherein the spacer molecules have the formula SH (CH₂) _nOH, where n is 2 to 12.

14. A biosensor according to any one of the preceding claims which is adapted such that the capture electrode surface is contacted with a measurement solution when changes in electrical properties are detected.

15. A biosensor according to claim 14 wherein the ionic strength of the measurement solution is less than 2OmM.

16. A biosensor according to claim 15 wherein the ionic strength of the measurement solution is 0. ImM to 10 mM.

17. A biosensor according to any one of claims 14 to 16 wherein the measurement solution comprises one or more redox active molecules.

18. A biosensor according to claim 17 wherein the one or more redox active molecules comprise ferri/ferrocyanide .

19. A biosensor according to any one of the preceding claims wherein the probe molecules are uncharged and probe : analyte complexes are negatively charged.

20. A biosensor according to claim 19 wherein the probe is an uncharged nucleic acid analogue and the analyte is a DNA or RNA molecule .

21. A biosensor according to claim 20 wherein the uncharged nucleic acid analogue is PNA.

22. A biosensor according to any one of the preceding claims comprising multiple capture electrodes.

23. A biosensor according to any one of the preceding claims comprising a reference and/or a counter electrode.

24. An array comprising a plurality of biosensors according to any one of claims 1 to 23.

25. A biosensor system comprising; a biosensor according any one of claims 1 to 23; and, a measurement solution having an ionic strength less than 2OmM.

26. A biosensor system according to claim 24 wherein the measurement solution comprises one or more redox active molecules.

27. A method of detecting an analyte comprising; providing a capture electrode having probe molecules and spacer molecules at the surface thereof, wherein the probe molecules are separated on the surface by spacer molecules and the proportion of probe molecules to spacer molecules is less than 10%; contacting the capture electrode surface with a sample solution such that analyte in the solution specifically binds to the probe molecules to form probe : analyte complexes at the capture electrode surface, and; measuring the electrical properties of the capture electrode, wherein changes in the electrical properties following contact with the solution are indicative of the formation of probe : analyte complex at the electrode surface.

28. A method according to claim 27 wherein one of the probe molecule and the probe : analyte complex is charged and the other is uncharged or the probe molecule and the probe: analyte complex have opposite charges .

29. A method according to claim 27 or claim 28 wherein the electrical properties are indicative of the charge and/or charge distribution at the capture electrode surface.

30. A method according to claim 29 comprising measuring changes in impedance, current or potential which are indicative of changes in charge and/or charge distribution.

31. A method according to claim 27 or claim 28 comprising measuring changes in impedance, current or potential which are indicative of the probe-analyte complex obstructing or opening conductive paths at the capture electrode surface.

32. A method according to any one of claims 27 to 31 wherein the probe molecules and spacer molecules form a mixed monolayer on the capture electrode surface.

33. A method according to claim 32 wherein the mixed monolayer is produced by simultaneously co-immobilising probe molecules and spacer molecules onto the surface of the capture electrode.

34. A method according to any one of claims 27 to 33 wherein the proportion of probe molecules to spacer molecules on the surface is

0.01% to 10%.

35. A method according to any one of claims 27 to 34 wherein the electrical properties are measured when the capture electrode is in contact with a measurement solution.

36. A method according to claim 35 wherein the ionic strength of the measurement solution is less than 2OmM.

37. A method according to claim 36 wherein the ionic strength of the measurement solution is 0. ImM to 10 mM.

38. A method according to any one of claims 35 to 37 wherein the measurement solution comprises one or more redox active molecules.

39. A method according to claim 38 wherein the one or more redox active molecules comprise ferri/ferrocyanide.

40. A method according to any one of claims 27 to 39 wherein the probe molecules are uncharged and probe: analyte complexes are negatively charged.

41. A method according to claim 40 wherein the probe is an uncharged nucleic acid analogue and the analyte is a DNA or RNA molecule.

42. A method according to claim 41 wherein the uncharged nucleic acid analogue is PNA.

43. A method of producing a biosensor for detecting an analyte comprising; simultaneously co-immobilising probe molecules and spacer molecules onto the surface of an electrode to form a mixed monolayer,

wherein the mole fraction of probe molecules in the monolayer is less than 10% and the probe molecules specifically bind to analyte in an electrolyte solution to form a probe : analyte complex.

44. A method according to claim 43 wherein the probe molecules immobilised onto the surface are comprised within probe : analyte complexes and the method further comprises; removing the analyte from the immobilised probe : analyte complexes such that unbound probe molecules remain immobilised on the surface

45. A method according to claim 43 or 44 wherein the probe molecule is a PNA molecule and the analyte is a DNA molecule.

46. A method according to claim 45, wherein the analyte is DNA molecules obtained by asymmetric PCR.