WO2009027708A2

WO2009027708A2 - Pulsed electrochemical sensor

Info

Publication number: WO2009027708A2
Application number: PCT/GB2008/002952
Authority: WO
Inventors: Timothy David Gibson; John Griffiths; Daniel James Lonsdale
Original assignee: Elisha Systems Limited
Priority date: 2007-08-30
Filing date: 2008-08-29
Publication date: 2009-03-05
Also published as: GB0716877D0; WO2009027708A3

Abstract

A method of detecting a biological affinity agent, comprising the steps of: providing a sensor comprising an antibody or other affinity reagent immobilised on a conductive substrate; the sensor further comprising one or more working electrodes, a counter- electrode and a reference electrode; contacting the sensor with an analyte containing the affinity agent; allowing the affinity agent to bind with the reagent; applying a voltage pulse to the sensor to polarise the sensor; disconnecting the voltage from the sensor and allowing the sensor to equilibrate in open circuit; measuring the rate of potential relaxation at the sensor; and using the measured rate of relaxation to calculate the amount of the affinity agent.

Description

PULSED ELECTROCHEMICAL SENSOR

This invention relates to an electrochemical sensor, particularly, but not exclusively to a sensor for use in affinity sensor detection of an analyte. This invention also relates to a method of detecting a biological affinity agent.

Various techniques have been used for voltametric assay, including pulsed potential step voltametry and pulsed amperometric detection. In both techniques, a continuous voltage waveform is applied to a sample and the characteristics of current flow through the sample are measured to determine the properties of the sample. A depolarising pulse may be used to restore the sample to the original state before the process is repeated.

Sensors employing biological affinity agents such as antibodies or nucleic acids represent a simple means of measurement of the concentration of an affinity ligand. Antibody sensors, hereinafter termed immunosensors, are particularly important because antibodies can be produced having specificity to most analytes. It is possible to produce an extremely diverse range of immunosensors.

Production of a measurable electrochemical signal from an immunosensor can be difficult since there is usually no appreciable electron transfer on affinity binding. Using labels such as enzymes to produce electroactivity intoduces significant chemistry into the manufacture of immunosensors. This can alter the affinity of the antibody for the ligand target. Nevertheless, there are several disclosures of techniques comprising interrogation and subsequent measurement.

US 6,300,123 discloses the use of alternating current impedance as an interrogation tool to probe both enzymatic and affinity reactions of anti-leutenising hormone, where the enzyme (lactate dehydrogenase) or antibody (anti-LH) is adsorbed or physically entrapped into a conducting polymer matrix. The measurement outputs were recorded as changes in the 'real'and 'imaginary' parts of a complex plane impedance spectra obtained and these are related to the concentration of the analytes lactate or leutenising hormone respectively. This method has also been reported in the scientific literature. A key disadvantage of the technique is the instability of the conductive polymer, e.g. polypyrrole when electrochemically polarised in the liquid used for the measurements. In addition the amounts of antibody used to make such systems operational are quite large, so that diagnostic devices are uneconomically expensive.

US 5,494,831 and application EP-A-640832 also disclose an electrochemical interrogation method for measurement of antibody binding. In this case antibodies are adsorbed onto gold or platinum electrodes and an AC output superimposed onto a DC voltage sweep is used to excite the electrode. The second harmonic and/or the phase angle changes are used as measurement outputs for antibody binding. No examples are provided.

CN 1588028 discloses the use of sol-gel immobilised antibodies immobilised to the electrodes by gold-alkylthiol matrices. Changes in capacitance are used as a measurement of antibody binding.

US 5,391,272 discloses use of pre-electrode bound antibody with analyte- enzyme complexes to produce a catalytic current at an electrode. Binding of free analyte removes a proportion of the analyte conjugate, giving a measure of the free analyte by a measured change in the analyte current. This system also uses an electroactive enzyme label to generate the measurement. In a similar disclosure, in WO2005/026689 a labelled antibody is used to generate an electrochemical signal proportionate to the quantity of affinity ligand (analyte) in solution. However in this case two surfaces were used as a means of correcting for interferences from background electroactive substances. WO2003/105660 discloses the use of graphite electrodes and antibodies adsorbed thereon to produce metal oxidation. This is used as a detection reaction for the affinity ligands bound. Anode voltametry is used as the electrochemical detection method.

US 5,403,451 discloses a method where an antibody electrode formed by entrapping an antibody into a polypyrrole matrix during polymerisation is interrogated using a periodic alternating voltage. This may be referred to as pulsed amperometric detection or pulsed electrochemical detection. The changes in the current output as the affinity ligand (analyte) binds to the electrode gives a measure of the concentration of the analyte. It is also disclosed that the system is reversible and that the antibody electrode can be used repeatedly because the analyte is not irreversibly bound to the antibody. A disadvantage of this system is the instability of the polypyrrole matrix to the continuous pulsing the periodic current applied to the electrode and the fact that reproducible deposition of antibody into polypyrrole is extremely difficult, with large amounts of antibody being required.

It is an object of the present invention to provide a simple immunosensor with a suitable method of interrogation method to provide unequivocal measurements.

According to a first aspect of the present invention, a method of detecting a biological affinity agent comprises the steps of: providing a sensor comprising an antibody or other affinity reagent immobilised on a conductive substrate; the sensor further comprising one or more working electrodes, a counter- electrode and a reference electrode; contacting the sensor with an analyte containing the affinity agent; allowing the affinity agent to bind with the reagent; applying a voltage pulse to the sensor to polarise the sensor; disconnecting the voltage from the sensor and allowing the sensor to equilibrate in open circuit; measuring the rate of potential relaxation at the sensor; and using the measured rate of relaxation to calculate the amount of the affinity agent.

According to a second aspect of the present invention, an electrochemical sensor for detecting a biological affinity agent comprises: a sensor comprising an antibody or other affinity agent immobilised on a conductive substrate; the sensor further comprising one or more working electrodes, a counter- electrode and a reference electrode; the sensor being adapted to be contacted with an analyte containing an affinity agent; means for applying a voltage pulse to the sensor to polarise the sensor; means for disconnecting th e voltage from the sensor to allow the sensor to equilabrate in open circuit; means for measuring the rate of potential relaxation of the sensor; and means for calculating the amount of the affinity agent on the measured rate of relaxation.

The biological affinity agent may comprise an antibody, binding protein or nucleic acid.

In a preferred embodiment, the potential is measured after a predetermined interval following the end of the pulse. The decrease in potential of the interval serves as a measure of the rate of potential relaxation.

This invention may provide a novel method of electronic interrogation of a sensor, for example, an electroactive thin film which contains or supports by means of immobilisation, antibodies or other affinity reagents. An electroactive thin film may be supported on various types of conducting surfaces including metal wires, thin metal films, carbon rods, carbon layers, screen printed matrices, photolithographically deposited metal films, conductive ceramics, conductive plastics and the like.

The conductive substrate is preferably supported by an insulating support layer. Preferred support layers may be composed of polyethylene terephthalate or other insulating polymer, aluminium oxide or silicon with a silicon dioxide coating layer.

The electrode may be formed by a variety of techniques. Preferred techniques include screen-printing, photolithography or by formation of a microarray. Preferred electrodes comprise screen-printed carbon, screen-printed gold, screen-printed platinum, photolithographic gold or a microarray, for example as disclosed in EP-A- 1749205. A preferred method for silicon based electrodes comprises deposition of gold on silicon, using a titanium adhesion layer and insulated with a silicon dioxide layer under the metal.

The conductive substrate may be formed from conductive polymers including polyaniline, polypyrrole/polyacetic acid or poly(3-aminophenylboronic acid).

Use of polyaniline is preferred for many applications. The polyaniline may be deposited by cyclic voltammetery between -0.2V and +0.8 V vs Ag/ AgCl using 0.1 M aniline dissolved in 1.0M HCl at 20 cycles and 50 mv.sec^"1. Use of polyaniline on screen-printed carbon or microarray electrodes is especially preferred.

Also preferred are self assembly monolayer substrates. Particularly preferred self-assembled monolayers may be selected from: mercaptohexadecanoic acid (MHDA), 4-mercaptoaniline, biotin caproyl, phospho-ethylanolamine and 1,2- dioleoyl-sn-glycero-3- { [N(5 -amino- 1 -carboxypentyl)iminodiacetic acid]succinyl } nickel salt (DOGS), 4-mercaptophenyl boronic acid, thiophene-3-boronic acid and nitriloacetic acid-DOGS. The electrodes may be arranged on the substrate in any suitable configuration. In a preferred embodiment, the working electrodes are arranged on a laminar substrate on either side of the reference electrode and a counter-electrode is arranged to surround the working electrodes and reference electrode.

The electrodes and conductive substrate may be coated with a soluble coating or glaze, for example composed of sugar, to protect the electrodes before use, the coating being arranged to dissolve when the electrode is inserted into an analyte solution.

The sensor assembly may be provided with an upper, absorbent or bibulous layer to absorb and retain an analyte which is dropped onto the sensor in contact with the electrode assembly. However, it is preferred that the sensor is immersed in a body of analyte solution in use so that an adsorbent layer is not required.

Analytes which may be determined using the sensor include: haemoglobin, myoglobin, prostatic specific antigen (PSA), CA125, SlOO protein, myelin basic protein, neuron specific enolase, bovine serum albumen, fluoroquinoline antibiotics, chitin and herbicides, for example atrizine.

The methods disclosed in EP-A- 1749205 may be employed. In a preferred method, a layer of poly- 1,3 phenylene diamine may be perforated with sono chemically ablated apertures into which polyaniline "mushrooms" may be grown.

The electroactive films may be prepared by electropolymerisation of a suitable monomer species, for example, selected from aniline, substituted anilines, substituted thiophenes and substituted pyrroles. A biotin moiety may be linked to a pyrrole monomer. Electropolymerisation of this monomer may give a biotin loaded electroactive film that can be used to bind an avidin linked protein. A matrix employing biotin was disclosed in patent application WO2006/018643 where a biotin substituted thiophene monomer (ethylenedioxy-thiophene) was used. Other matrices include mixed lipid self-assembled monoayers (lipid-SAMS). Sol-gels, related silicon based matrices and non-intrinsically conductive polymer films (e.g. polycarbonate, PVC) may be used provided that they are rendered electronically conductive.

Loading an antibody into or onto an electroactive film may be carried out by a number of methods. Preferably the antibody or other affinity reagent is immobilised onto or into a conducting matrix of the electrode after the conducting matrix has been deposited. This is referred to as post-electropolymerisation immobilisation. Also, the antibody may be added before polymerisation. However, the deposition process was found to be difficult to reproduce and a large amount of antibody was needed, leading to high costs in manufacture.

Post-electropolymerisation immobilisation has the advantages that it gives reproducible results, uses minimal amounts of antibody, is extremely gentle in chemical terms and produces highly active immobilised antibody films.

Immobilisation routes most preferably lead to either a covalent attachment of antibody to the electroactive matrix or to a high affinity attachment via molecular complexing surface chemistry. Examples are: biotin-avidin affinity, cis-diol-boronic acid complexation and metal chelation using hexa-histidine tags or nickel - nitriloacetic acid (NTA) groups. Such chemistries are standard for protein chemistry and molecular biology procedures where expressed proteins can be engineered to contain avidin, biotin and hexa-histidine tags.

Exposing such antibody loaded films to a solution containing the respective analyte allows an affinity reaction to take place between the antibody and its antigen and a concentration dependant response may be observed.

The mode of detection of concentration in accordance with this invention is beneficial for the long term stability of the antibody films since for the majority of the time during analysis the film remains unpolarised. Stable transient measurements are used to ascertain the concentration of the analyte.

The sensor is preferably connected to a circuit which delivers a narrow pulse of voltage to the sensor surface followed by a switch to an open circuit, allowing the potential to decay. The binding between antibody and antigen is detected as a perturbation in the decay of the potential. This decay is dependant on the antigen concentration. The results may be plotted as a power spectrum. Transforming of these into the first derivative can give a clear analysis of the binding event.

The working electrodes are initially at a low potential with respect to the open circuit potential or in open circuit. The voltage pulse is usually a square wave rising sharply to a predetermined voltage which is maintained at a constant value for a predetermined period before the electrode is disconnected to form an open circuit. However, any convenient wave form may be used to polarise the electrode.

In a preferred embodiment, the voltage is ± 0.4v, although the voltage may be selected to suit the reagents employed.

The duration of the pulse may be about 10 milliseconds, although the preferred value will depend on the specifics of any sensor, including, but not limited to: snesor construction, electrode area, the characteristics of the sensor surface and the solution in which it is immersed. In one embodiment the decay of voltage after disconnection of the pulse is measured and the data Fourier transformed to give a power spectrum.

Measurement of the voltage at a frequency where the FT response is above the noise floor is required. This will depend on the sensor, the pulse and the acquisition parameters, but a preferred value lies in the 0.1 Hz to 100 Hz region.

In this way the sensor detects low frequency mediation or passivation of the electrode surface by the immobilised analyte. In preferred embodiments, the analyte is contained in a solution comprising a suitable buffer, for example phosphate buffered saline (PBS). A preferred buffer contains 25 mmol phosphate, and/or 8g NaCl and 0.2g KCl per litre.

In preferred embodiments, an electrochemical mediator is added to the solution. A preferred mediator comprises equimolar amounts of potassium feπϊcyanide and potassium ferrocyanide, for example in amounts of 5mmol of potassium ferricyanide and 5mmol of potassium ferrocyanide. Alternative mediators include ruthenium salts, ferrocene derivatives, for example ferrocene carboxylic acid and other mediators known to those skilled in the art. Mediators have been found to be useful to stabilise the open circuit potential of the sensor.

Non-specific binding, which is a problem with affinity reactions in general, may be minimized during the sensor construction using various blocking agents such as bovine serum albumin (BSA), casein and various other additives such as polyethylene glycol. Problems due to non-specific binding may be overcome using dual electrode systems, where two working electrodes are included in the immunosensors. A nonspecific antibody is deposited onto one working electrode and the specific antibody onto the other. Incubation of the two electrodes with antigen gives two signals due to non-specific binding and antibody-antigen binding. Subtraction of one from the other gives the specific response due to specific binding.

The invention will now be described by means of examples, but not in any limitative sense, with reference to the accompanying drawings, of which:

Figure 1 is a diagrammatic view of a sensor in accordance with this invention;

Figure 2 is a diagram of the voltage of the pulse applied in accordance with this invention;

Figure 3 is a graph of voltage against time for a sensor in accordance with this invention, showing the wave form of the voltage of the pulse applied and subsequent relaxation using a HB concentration of 10^" gl^" Figure 4 is a fast-Fourier transformation of the decay portion of the curve shown in Figure 3;

Figure 5 shows corrected sensor response versus haemoglobin concentration for a dual electrode format using poly-3-aminophenyl boronic acid immobilisation in accordance with this invention;

Figure 6 is a log response in dual response mode for the sensor shown in Figure 5;

Figure 7 is a graph showing the voltage of the pulse applied and subsequent relaxation due to HB concentration of 5 x 10^"7M;

Figure 8 is a fast-Fourier transform of the curve shown in Figure 7;

Figure 9 shows the sensor response 33 Hz for the sensor;

Figure 10 is a power spectrum analysis of a SAMS gold electrode; and

Figure 11 is a differential power spectrum for the spectrum shown in Figure 10.

Figure 1 is a diagrammatic view of an electrode for use in a sensor in accordance with this invention, where the circle represents the electrochemical cell containing the sensor and the solution, the four arms are the electrodes and the electronics are outside the cell. The sample or samples to be analysed are located on the surface of the working electrodes or in the analytes surrounding the electrodes. Part A of Figure 1 is a circuit adapted to provide the potential and current to reach a desired set point. The set point can be either a defined potential between the reference electrode RE and the working electrodes WE or a defined current flow through the working electrodes. Component A can be switched out of the circuit. When switched out of the circuit, A provides no influence on the electrochemical cell or the remaining electronics.

Where a combined counter electrode CE and reference electrode RE are used, the circuit components A and B are connected together outside the electrochemical cell (circle in Fig. 1) and the switch that connects A to the electrochemical cell or otherwise will only act on A and have no effect on B. Component B shown in Figure 1 measures the potential set across B and the working electrode or electrodes. If a single working electrode is used, then there is only a single potential reported by the electronics circuit in B, whereas if there are numerous working electrodes, then there will be a potential reported for each electrode by the electronics in component B. The potential reported may be defined as the potential between the reference electrode and the associated working electrode.

Component C of Figure 1 measures the current flowing through the working electrode. If a single working electrode is used, a single current is reported by the electronic circuit in C, whereas if numerous working electrodes are used then there will be a separate current reported by the electronics circuit in C. Component C can be switched in or out of the circuit. When switched into the circuit, C measures the current flowing through its associated working electrode. When switched out of the circuit, C provides no route for current, allowing the working electrode to float to its natural potential. Switching C in and out of the circuit does not prevent the operation of the electronics in B, which reports the potential between the reference electrode and the working electrode.

Example 1 : Atrazine

Anti-atrazine was immobilised onto a gold electrode by sequential addition of mercaptohexadecanoic acid (MHDA) and l,2-dioleoyl-sn-glycero-3-{[N(5-amino-l- carboxypentyl)iminodiacetic acidjsuccinyl} nickel salt (DOGS), to provide chelation of histidine tagged antibodies.

Addition of these components was followed by the pulsed interrogation of the electrode surfaces. Adding histidine tagged anti-atrazine gave a further response. The immunosensor was fully constructed. Adding antigen i.e. atrazine showed a large response after 2 minutes incubation, which was only slightly increased after a further 13 minutes. Removal of the anti-atrazine layer may be achieved by titrating the surface with free histidine in solution in order to compete for the NTA binding sites, forcing off the anti-atrazine-atrazine complex and thus providing a regenerable sensor. This may be used in flow system analysis.

Example 2 - Anti-Haemoglobin on a Poly 3-Amino-Phenylboronic Acid surface:

Pulsed wave spectroscopy can be used to measure for the antibody binding of haemoglobin on a poly 3-amino-phenylboronic acid surface. This was achieved by monitoring the responses of specific and non-specific carbon screen-printed electrodes in a sequential manner. The resulting specific response was then corrected by subtracting the non-specific (or background) response from the specific response.

The following procedure was used on a screen-printed sensor with four electrodes, where the counter electrodes and working electrodes (1 & 2) were carbon based and the reference electrode was a screen-printed Ag/ AgCl reference. The approximate areas of the current carrying electrodes exposed to solution were as follows:

Counter electrode 90mm²

Working electrode 27mm²

Working electrode 2 27mm²

Hydrochloric acid was added to pH1.5 or less until the solution was clear. The solution was then made up to 10cm³ with water.

A dual screen printed carbon electrode with a planar carbon surface was immersed in the solution making sure all the working electrode surfaces were covered. Both working electrodes were connected to the potentiostat and the counter and reference electrodes were connected to make a standard electrochemical circuit. The potential was cycled between -0.3V and +1 V vs the Ag/ AgCl reference electrode. 20 scans were carried out at a scan rate of SOmV.second^'1 to deposit a thin even film of poly 3-amino phenylboronic acid onto both working electrodes.

Anti-haemoglobin was then immobilised on the surface of one of the working electrodes by pipetting a buffered solution of anti-Hb onto the surface making sure that no solution contacted the second working electrode. On the other working electrode a non-specific anti human IgG antibody was immobilised in the same manner. The electrode was then incubated for 1 hour in a damp atmosphere to prevent drying out. This step was completed by rinsing the electrode in buffer and storage in buffer.

The prepared electrode was immersed in phosphate buffered saline solution in an electrochemical cell and connected to the respective poles of the pulse circuit, where the working electrode connection mates with the specific antibody electrode, in the first instance.

The electronic control circuit was then programmed to apply the following voltage waveform to the sample as shown in Figure 2. The waveform was applied twice in rapid succession and the voltage-time results were averaged.

In Figure 2 the electrochemical cell has an equilibrium potential as denoted by the signal level T. At the start of the experiment (at 't=0') the electronic circuit pulsed the sample (or working electrode) to the value defined at level '2', and maintained this potential applied to the circuit for the time at level '3'. Finally, at the end of level '3' the electronics 'A' and 'C (defined in Figure 1) were disconnected from the circuit and the potential relaxation was measured for a period '4' (and suggested by the response shown as '5' in Figure 2.)

For this example, the following are the configuration parameters of the pulsed signal used in this embodiment of the invention as shown in Figure 2: Pulse Voltage: (level 2) +0.4volts re Ag/AgCl

Pulse Time: (level 3) 10 milli-seconds

Acquisition Time: (period 4) 30 milli-seconds

Sample Rate in period '4': 4OkHz

Number of repeat runs to average: 2.

After the acquisition had taken place (at the end of the 2^nd application of the pulsed waveform spectroscopic pulse-train) and the 2-runs had been averaged, the data in section '4' of the acquisition was analysed with fast-Fourier transformation. Figure 3 below shows the voltage waveform acquired in an experiment with anti- HB antibodies immobilised on a poly 3-amino-phenylboronic acid surface when exposed to BSA (background) and a haemoglobin concentration of lxlθ^"8gl^"1.

The response given in Figures 3 and 4 constituted the specific sensor response for the anti-HB sensor under known conditions. The system was then configured to excite the remaining sensor and obtain non-specific, or background readings for the

BSA and HB concentration (IxIO^" gl^" in this instance).

The decay curve shown above was typical of the sensor response for these conditions and varied with HB concentration. The final step made in the analysis procedure used here was to differentiate the frequency spectrum shown in Figure 4 and take the value at 33Hz as a sensitive and representative indicator for the combined specific and non-specific response to HB for the specific sensor. The same analysis was applied to the non-specific sensor and was used to remove the background from the first result by subtraction.

Using the technique outlined above and providing increasing levels of haemoglobin, a calibration plot of sensor response vs HB concentration was produced and corrected for background (or non-specific) responses. Figures 5 and 6 show the trend for a range of HB concentrations after the non-specific response had been removed. Example 3 - Anti-Haemoglobin on an Activated Poly-Acrylic Acid surface

This example shows use of the pulsed waveform spectroscopy technique to interrogate for the antibody binding of haemoglobin on an activated polyacrylic acid surface. This was achieved by monitoring the response of a specific antibody immobilised in a polymer on a bare gold electrode. Any non-specific response was tested for by measuring the response of a background analyte, in this case Bovine Serum Albumen (BSA).

Immobilisation protocols were devised based on selective chemical derivitisation of conducting polymer surfaces. Such methods were found to be more controllable than simple antibody entrapment.

Films of polypyrrole were produced using electropolymerisation with a polymeric counterion: polyacrylic acid. This gave numerous carboxylic acid groups available for derivatisation, using Woodwards reagent.

Planar gold electrodes were used (type P3, with lmm diameter circular gold electrode open to solution). These were prepared by sputtering gold onto masked silicon wafers.

Polypyrrole / polyacrylic acid deposition was carried out at constant potential of 1.1 V vs Ag/ AgCl from 0.1 M aqueous solution of pyrrole, 0.4M aqueous solution of polyacrylic acid, pH 6.5. The total charge passed was 20mC.cm² in each case.

The antibody was immobilised by first reacting the polypyrrole/polyacrylic acid films with Woodwards reagent to give activated surfaces and the antibody was immobilised by simple incubation of a buffered solution of antibody with the activated electrodes. Antibody electrodes were interrogated using an electrochemical cell having an external reference electrode, a platinum counter electrode and the antibody prepared electrode (type P3) as the working electrode.

The electronics were programmed to apply a voltage waveform to the sample as shown in Figure 2. In this case, the waveform was applied repeatedly in rapid succession and the voltage-time results were averaged.

In Figure 2 the electrochemical cell has an equilibrium potential as denoted by the level ' 1 '. At the start of the experiment (at 't=0') the electronics pulses the sample (or working electrode) to the value defined at '2', and keeps this potential applied to the circuit for the time '3'. Finally, at the end of '3' the electronics 'A' and 'C (defined in Figure 1) were disconnected from the circuit and the potential relaxation was measured for a period '4' (and suggested by the response shown as '5' in figure 2.)

For this example, the following are the pulsed waveform spectroscopy configuration parameters as shown in Figure 2:

Pulse Voltage (Level 2): -0.4volts re Ag/AgCl.

Pulse Time (Level 3): lO milli-seconds

Acquisition Time (Period 4): 30 milli-seconds

Sample Rate in period '4': 4OkHz

Number of repeat runs to average: 9.

After the 9-runs had been averaged, the data in section '4' of the acquisition was analysed with Fourier techniques (i.e. fast-Fourier transformed). Figure 7 shows the voltage waveform acquired in an experiment with anti-HB antibodies immobilised on an activated poly-acrylic acid surface when exposed to BSA (background) and a haemoglobin concentration of 5x10^"7M. The response given in Figure 8 constituted the specific sensor response for the anti-HB sensor under known conditions.

The decay curve shown was typical of the sensor response for these conditions and varied with HB concentration. The final step made in the analysis procedure used here was to differentiate the frequency spectrum shown in Figure 11 and take the value at 33Hz as a sensitive and representative indicator for the combined specific and non-specific response to HB for the specific sensor.

Using the technique outlined above and providing increasing levels of haemoglobin, a calibration plot of sensor response vs HB concentration was produced. The background responses were seen for the antibody and antibody-BSA (bovine serum albumen) as separate responses giving approximately the same response of —17% of max signal (defined below as that taken at 50OnM).

Figure 9 has the sensor response measured at 33Hz in the derivative of the frequency spectrum, expressed as a percentage of the largest response, given here as 50OnM. Here, the response seen for the antibody (AB) and the AB-BSA were around 17%. Increasing the HB concentration produced a change in response of around 80% signal strength.

Using the pulse waveform electronics to give a power spectrum analysis, with anti-haemoglobin as the selective antibody, the surface was interrogated in the absence of haemoglobin, in the presence of a large concentration of BSA and then with progressively larger amounts of haemoglobin from 5nM to 50OnM. The concentration curve is shown in Figure 9.

Claims

1. A method of detecting a biological affinity agent, comprising the steps of: providing a sensor comprising an antibody or other affinity reagent immobilised on a conductive substrate; the sensor further comprising one or more working electrodes, a counter- electrode and a reference electrode; contacting the sensor with an analyte containing the affinity agent; allowing the affinity agent to bind with the reagent; applying a voltage pulse to the sensor to polarise the sensor; disconnecting the voltage from the sensor and allowing the sensor to equilibrate in open circuit; measuring the rate of potential relaxation at the sensor; and using the measured rate of relaxation to calculate the amount of the affinity agent.

2. An electrochemical sensor comprising: a sensor comprising an antibody or other affinity agent immobilised on a conductive substrate; the sensor further comprising one or more working electrodes, a counter- electrode and a reference electrode; the sensor being adapted to be contacted with an analyte containing an affinity agent; means for applying a voltage pulse to the sensor to polarise the sensor; means for disconnecting th e voltage from the sensor to allow the sensor to equilabrate in open circuit; means for measuring the rate of potential relaxation of the sensor; and means for calculating the amount of the affinity agent on the measured rate of relaxation.

3. A method or sensor as claimed in claim 1, wherein the biological affinity agent is an antibody, binding protein or nucleic acid.

4. A method or sensor as claimed in any preceding claim, wherein the potential is measured at a predetermined interval following the end of the pulse.

5. A method or sensor as claimed in any preceding claim, wherein the conductive substrate is supported by an insulating support layer.

6. A method or sensor as claimed in claim 5, wherein the support layer is composed of a polyethylene terephthalate insulating polymer, aluminium oxide or silicon with a silicon dioxide coating layer.

7. A method or sensor as claimed in any preceding claim, wherein the electrode is formed by screen-printing, photolithography, or by formation of a microarray.

8. A method or sensor as claimed in claim 7, wherein the electrode comprises screen-printed carbon, screen-printed gold, screen-printed platinum, photolithographic gold or a microarray.

9. A method or sensor as claimed in any preceding claim, wherein the conductive substrate is formed from a conductive polymer selected from the group consisting of: polyaniline, polypyrrole/polyacetic acid or poly(3-aminophenylboronic acid) or mixtures thereof. •

10. A method or sensor as claimed in any preceding claim, wherein the conductive substrate is a self-assembly monolayer.

11. A method or sensor as claimed in claim 10, wherein the monolayer is selected from the group consisting of: mercaptohexadecanoic acid (MHDA), 4- mercaptoaniline, biotin caproyl, phospho-ethylanolamine and DOGS, 4- mercaptophenyl boronic acid, thiophene-3-boronic acid and nitriloacetic acid- DOGS.

12. A method or sensor as claimed in any preceding claim, wherein an electroactive film is prepared by electropolymerrisation.

13. A method or sensor as claimed in claim 12, wherein an electroactive film is prepared by electropolymerisation of a monomer selected from the group consisting of: substituted anilines, substituted thiophenes and substituted pyrroles.

14. A method or sensor as claimed in claim 13, wherein a biotin moiety is linked to a pyrrole monomer.

15. A method or sensor as claimed in any preceding claim, wherein the antibody or other affinity reagent is immobilised onto or into a conductive matrix, the electrode.

16. A method or sensor as claimed in any preceding claim, wherein the voltage pulse is a square wave.

17. A method or sensor as claimed in any preceding claim, wherein the voltage pulse is ± 0.4v.

18. A method or sensor as claimed in any preceding claim, wherein the duration of the pulse is in the range of 1 to 10 milliseconds.

19. A method or sensor as claimed in claim 18, wherein the duration of the pulse is 10 milliseconds.

20. A method or sensor as claimed in any preceding claim, wherein the voltage is measured at a frequency of 0.1 to 100 Hz.

21. A method or sensor as claimed in any preceding claim, wherein the voltage is measured at a frequency of 0.1 to 10 Hz.

22. A method or sensor as claimed in claim 21, wherein the voltage is measured at a frequency of 1 to 10 Hz.

23. A method or sensor as claimed in any preceding claim, wherein analyte includes a mediator.

24. A method or sensor as claimed in claim 23, wherein the mediator comprises equimolar amounts of potassium ferricyanide and potassium ferrocyanide.

25. A method or sensor as claimed in any preceding, wherein the analyte is selected from the group consisting of: haemoglobin, myoglobin, prostatic specific antigen (PSA), CAl 25, SlOO protein, myelin basic protein, neuron specific enolase, bovine serum albumen, fluoroquinoline antibiotics, chitin and atrizine.