Field of the Invention This invention relates to the field of electrochemical methodology, polynucleotide sequences and the use of such methodology in the detection of target substrates in a sample.

Background As part of medical diagnosis, quality assurance, and environment testing, there is an ongoing need for methods that rapidly detect and quantify the presence of target molecules in samples. For example, molecules, such as, endocrine disrupting compounds and hormones, are often found as contaminants in the environment. Such contaminants can be found in waterways, soils, biological samples, including both plant and animals, as environmental pollutants from residential, agricultural, commercial and/or industrial applications. It is known, in some cases, that these small molecular weight compounds, together with their metabolites and/or synthetically modified variants pose a threat to the health of human and wildlife populations by blocking, mimicking, stimulating or inhibiting the production and function of natural hormones. The organic residues that mimic these endogenous steroidal hormones, and metabolites are lipid soluble, and thus have the ability to bio-accumulate in living systems of mammals and marine species. Evidence of this has been identified in human blood plasma, breast milk, foetal tissues and biological fluids [Allmyr ef a/., Anal. Chem., 78, 6542-6546, 2006; Hileman, Chemical and Engineering News, 85: 31 -33, 2007; Van-Pelt et al. Endocrinology, 140, 478-483, 2001 ; Skakkebaek et al., Human Reproduction 16: 972-978, 2001 ; Vandenberg et al., Endocrine Reviews, 33(3), 2012] In humans and other animals, hormone levels play an important role in regulating physiology, behaviour and reproduction. Therefore, quantifying hormone levels in humans and animals is important for medical diagnostics, including detecting and monitoring diseases and managing fertility.

DNA sequences tethered to electrodes, including nanoporous conducting polymer electrodes, have been used to detect charged complementary sequences at low nM levels using Electrochemical Impedance Spectroscopy (EIS) [Kannan, B., Williams, D. E., Booth, M. A. & Travas-Sejdic, J., Anal. Chem. 83, 3415-3421 , 2011; Booth, M. A., Harbison, S. & Travas-Sejdic, J., Biosensors and Bioelectronics 28, 362-367, 2011 ; Booth, M. A., Harbison, S. & Travas-Sejdic, J., Electroanalysis 24, 131 1-1317, 2012.] Kannan et al have also shown that charged DNA sequences tethered to nanoporous electrodes are more sensitive in detection when compared with planar electrodes. However, the same model cannot be directly applied to aptamers because the target substrate may be neutral, carry a small charge, which may be substantially lower than a DNA sequence, or the target substrate may be a singly charged ion.

Aptamers are single-stranded nucleic acids (ssRNA, ssDNA), which unlike traditional nucleic acids, possess unique binding characteristics to specific targets with high affinity and specificity analogous to antibodies [Tuerk, C. Gold, L, Science, 1990, 249(4968), 505-510; Ellington, A.D., Szostak, J.W., Nature, 1990, 346(6287), 818-822.] Aptamers are isolated in vitro from combinatorial oligonucleotide libraries, typically containing 10 ¹² to 10 ¹⁵ oligonucleotides, and are synthetically evolved by a process known as SELEX. The oligonucleotides are subjected to a selection process for their ability to bind a specified target. Over a number of selection rounds (typically 8-16 rounds), the most specific nucleic acid sequences are isolated. Depending on the techniques used in SELEX, the process might take from days to months [Cho, E. J., Lee, J.W., Ellington, A.D., Ann. Rev. Anal. Chem., 2009, 2(1 ), 241-264; Ellington, A. D., Ann. Rev. Anal. Chem., 2009, 2(1 ), 241-264.]

Aptamers have been generated for a wide range of targets, ranging from ions to entire cells [for proteins see de la Escosura-Muhiz, A.; Maltez-da Costa, M.; Guix, M.; Ozsoz, M.; Merkoci, A. Biosens. Bioelectron., 2010, 26, 1715; for hormones see Lin, Z.; Chen, L.; Zhang, G.; Liu, Q.; Qiu, B.; Cai, Z.; Chen, G. Analyst, 2012, 137, 819; for small molecules refer to Zayats, .; Huang, Y.; Gill, R.; Ma, C; Willner, I. J. Am. Chem. Soc, 2006, 28, 13666; for cells see Pan, C; Guo, M.; Nie, Z.; Xiao, X.; Yao, S. Electroanalysis 2009, 21, 1321 ; Feng, L; Chen, Y.; Ren, J.; Qu, X. Biomaterials 2011 , 32, 2930]. The use of an in vitro process enables the generation and selection of aptamers that can bind toxic targets, which is not possible by immunologically initiated recognition elements, such as antibodies. The small size of aptamers (generally <3 nm in a coiled conformation) also makes them more readily applicable to surface-based aqueous sensing purposes in comparison to antibodies (approximately >10 nm in size) [Song, S., et ai, Trends in Analytical Chemistry, 2008, 27(2), 108-1 17] and also permits the aptamers to be incorporated into the pores of porous electrodes.

Furthermore, aptamers can be readily modified or coupled to other materials from which binding signals can be engineered in different environments, and this makes aptamers ideal for use in sensor devices. In low molecular weight (MW) targets (including narcotics and steroidal hormones), binding is generally associated with conformational switching of the aptamer. For example, a polynucleotide sequence (aptamer) may change from a loose linear conformation to a more tightly folder structure around the target. The binding signals induced from a conformational change can be transduced via colorimetric, fluorescence, size, and electrochemical responses to indicate the presence or absence of the low molecular weight target molecules to be detected.

Electrochemical signal transduction methods offer the ability to detect the presence of the target, and quantification of the concentration over a wide range. EIS is particularly useful because it is sensitive to surface interactions and quantifies the interfacial charge transfer resistance (RCT) that is associated with charged redox probes. R _CT is strongly affected by changes in charge distributions near the electrode-solution interface in a sample solution, which results in surface sensitivity. This surface sensitivity can be exploited by attaching biological recognition elements, such as, DNA aptamers to electrode surfaces. Lin ef a/, covalently tethered (terminally tethered) an E2 binding aptamer sequence to a planar gold electrode through a thiol linker and achieved detection levels of 2 pM utilising EIS where the signal on binding of the E2 to the aptamer is caused by the re-distribution of the charges on the aptamer surface. [Lin, Z. et al. Analyst (Cambridge, U. K.) 137, 819-822, 2012]. Chen ef al. constructed a similar sensor by tethering a K ⁺ binding sequence to a planar gold electrode and employing EIS to achieve detection levels of 100 pM [Chen, Z., Chen, L., Ma, H., Zhou, T. & Li, X. Biosensors and Bioelectronics 48, 108-112, 2013].

DNA aptamers specific for small molecules or ions have been tethered to planar electrodes, and the interaction between the two has been detected through electrochemical signal transduction methods, such as, electrochemical impedance spectroscopy (EIS). However, such systems as this require additional steps, such as, passivation of the system to prevent spurious signals and/or prevent interference from non-target agents/molecules that may be present.

While many polynucleotide sequence based sensor devices are able to detect the presence of targets at a particular concentration level, they are unable to easily quantify the concentrations of the targets present in the sample. For example, the colour change of a colorimetric polynucleotide sequence sensor can be saturated within a narrow target concentration range and does not necessarily give an accurate indication of the concentration of the target substrate present.

Electrochemical polynucleotide sequence sensors offer high sensitivity, as well as rapid, label-free quantification of target molecules over a wide range of target concentrations. Therefore, there is a requirement for sensors that rapidly and accurately quantify small molecules at extremely low concentration levels, in particular, levels at which these small molecules elicit biological responses.

However, the nucleic acids in polynucleotide sequence-based electrochemical sensors are modified to facilitate tethering of the polynucleotide sequence, for example with a thiol group that facilitates a strong covalent tether to the electrode. Sensor systems currently used in the prior art, for example, those systems where a polynucleotide sequence is tethered to an electrode surface via a linker (such as a thio linker, or a nucleotide sequence adapted as a linker), do not substantially cover or substantially coat the electrode surface. Additionally, the polynucleotide sequences in tethered electrochemical systems are tethered in a manner such that the polynucleotide sequence is orientated so that the polynucleotide sequence extends away from the electrode surface (in a linear-type manner). This leaves parts of the electrode surface exposed to interfering agents. These interfering agents affect the signal generated when a polynucleotide sequence binds to a target substrate. Consequently, to remove any spurious signals, the electrode requires backfilling with an additional agent that has some affinity for the noble metal surface (such as an alkyl thiol) to prevent any interfering agents from influencing the signal on the binding of the polynucleotide sequence to the target substrate. This extra step of backfilling and blocking spurious signals adds additional steps and expense to the process. Further, the polynucleotide sequence does not dissociate from the electrode surface on association with the target substrate. The lack of surface dissociation upon target recognition means that tethered DNA probes generate a binding signal if binding induces a substantially different surface charge distribution.

Polynucleotide sequences that are tethered to the noble metal surface such that the polynucleotide sequence extends away from the electrode surface are perceived to be undesirable in electrochemical sensors using nucleic acid sequences (e.g. DNA or RNA) probes because the tethered polynucleotide sequences are not well defined, and result in non-specific and poorly defined changes in measureable parameters such as charge- transfer resistance. Furthermore, sensors known in the art that are derived from noble metal nanoparticle electrodes require additional crosslinking between the noble metal nanoparticles; require surface passivation steps in order to prevent de-stabilisation and loss of the nanoparticles from the electrode surface; and also to prevent interference from interfering agents.

At present, detection and diagnosis is often accomplished by chromatographic and mass spectrometric analytical techniques which involve expensive equipment and is time consuming. Therefore, there is a need for simple low cost and effective methods to detect target substrates in a sample.

It is an object of the present invention, to provide a method of detecting target substrates in a sample, such as nucleic acids, peptides, proteins, small target molecules, or ions, using polynucleotide sequences (such as dsDNA, ssDNA, and RNA sequences) together with electrochemical methodology, that will provide for target selectivity and wide dynamic ranges, or to at least provide the public with a useful alternative. Summary of the Invention

The present invention exploits EIS signals and intrinsic passivation of the disclosed detection system by substantially adsorbing polynucleotide sequences (such as dsDNA, ssDNA, and RNA sequences) or aptamers substantially along their length substantially along electrodes (such as glassy carbon; metal oxide; conducting polymer; and noble metal electrodes including gold, ruthenium, rhodium, palladium, platinum and silver). This provides for the surface of the electrode to be substantially covered or substantially coated with the polynucleotide sequence. Unlike EIS detection systems that are known in the art that probe conformational changes of a tethered sequence, the present invention resolves changes in R _CT when target-bound polynucleotide sequences (such as dsDNA, ssDNA, and RNA sequences) dissociate from the electrode surface to provide good sensitivity and low detection levels for a target substrate. The detection system of the present invention therefore provides for a quick and convenient system to detect target substrates in a sample. Sensitivity and detection limits of the present invention provide detection limits of target substrates at concentrations down to and around the fM level, while also providing for wide dynamic ranges, and showing selectivity toward target substrates, without the detection system being influenced by interfering molecules and requiring further passivation steps.

The substantial adsorption of the polynucleotide sequence substantially along the surface of the electrode surface in the present invention fulfils the role of surface passivation by substantially coating the electrode and preventing other interfering agents such as nucleotide sequences, molecules or ions from interfering with the electrode creating spurious signals.

Further, when the electrode comprises a noble metal nanoparticle surface, the adsorption of the polynucleotides substantially along the surface of the noble metal electrode in the present invention, provides the further benefit that no additional cross-linking of the noble metal nanoparticles is required because the polynucleotides are adsorbed and arranged along the electrode surface instead of a tethered fashion, away from the surface. This provides benefits including stabilisation of the nanoparticles onto the electrode surface; strong perturbation of the charge distribution within the confined double layer region (and concomitantly reduced R _CT values); and passivation of the noble metal surface when it is saturated with the polynucleotide sequence.

Furthermore, the adsorption of the polynucleotide sequences substantially along their length to substantially along the surface of the electrode leads to displacement of a dense layer of ions that are found directly at the electrode interface at high electrode potentials, providing enhanced charge transfer, and greater sensitivity to subsequent changes in the interfacial region (such as target binding), when compared to polynucleotide sequences that are tethered (terminally tethered) to an electrode surface, such that the polynucleotides extend away from the electrode surface.

The enhanced charge transfer signals of polynucleotide sequences adsorbed to an electrode surface, substantially along the surface of the electrode, and substantially along the length of the polynucleotide sequence are exploited in the present invention using substantially adsorbed polynucleotide sequences on electrodes (such as glassy carbon; metal oxide; conducting polymer; and noble metal including gold, ruthenium, rhodium, palladium, platinum and silver).

Strong electrochemical responses to a target substrate, and intrinsic passivation by the polynucleotide sequence make this detection system an appealing alternative to, or complementary system for, other detection systems of the prior art. The present invention may also find use in the dual development of methods that offer both visual detection, and sensitive electronic quantitation using polynucleotide sequences. In a first aspect of the invention, the invention provides for an electrochemical sensor to detect a target substrate in a sample, wherein the electrochemical sensor comprises: i. substantially adsorbing a target specific polynucleotide sequence to an electrode to form a coated electrode that is substantially coated with the target specific polynucleotide sequence,

wherein the adsorbing of the target specific polynucleotide sequence to the electrode surface is sufficient to adsorb the sequence to the electrode, but to also provide complete dissociation of the sequence from the electrode on association of the target specific polynucleotide sequence with the target substrate; and

ii. a means of detecting dissociation of the polynucleotide from the substantially coated electrode surface and/or detecting association of the polynucleotide sequence with a target substrate in a sample.

In a second aspect of the invention, the invention provides for an electrochemical method of detecting a target substrate in a sample, wherein the electrochemical method comprises: i. substantially adsorbing a target specific polynucleotide sequence to a noble metal surface to form a coated electrode surface substantially coated with the target specific polynucleotide sequence,

wherein the adsorbing of the target specific polynucleotide sequence to the electrode surface is sufficient to adsorb the sequence to the electrode, but to also provide complete dissociation of the sequence from the electrode on association of the target specific polynucleotide sequence with the target substrate; and

ii. exposing the target specific polynucleotide sequence coated noble metal surface to a sample;

wherein, when the target substrate is present in the sample, the polynucleotide sequence dissociates from the surface of the electrode and the polynucleotide sequence forms an associative interaction with the target substrate. In an embodiment of the second aspect, the method further comprises a means of detecting dissociation of the target specific polynucleotide from the coated electrode surface and/or detecting association of the target specific polynucleotide sequence with a target substrate in a sample. In an embodiment of any of the aspects of the invention, the dissociation of the target specific polynucleotide sequence from the coated electrode surface and/or the association of the target specific polynucleotide sequence with the target substrate in a sample is measurable by means known in the art.

The means to detect the dissociation of the target specific polynucleotide sequence from the coated electrode and/or the association of the target specific polynucleotide sequence with the target substrate in a sample includes but are not limited to techniques, such as, cyclic voltammetry, differential pulsed voltammetry and electrical impedance spectroscopy.

When the target specific polynucleotide sequence dissociates from the electrode surface and associates with the target substrate, there is a change in the R _CT value measured by electrochemical impedance spectroscopy when compared to the bare electrode (i.e. electrode without a polynucleotide sequence). Preferably, this change in R _CT that results from the dissociation of the target specific polynucleotide sequence from the electrode and/or the association of the target specific polynucleotide sequence with the target substrate is a decrease, or a reduction in R _CT when compared to the bare electrode (i.e. electrode without a polynucleotide sequence).

Preferably, the dissociation of the target specific polynucleotide sequence from the electrode and/or the association of the polynucleotide sequence with the target substrate in a sample is detected using a technique selected from cyclic voltammetry and electrical impedance spectroscopy.

In an embodiment of any of the aspects of the invention, adsorption of the target specific polynucleotide sequence to the electrode surface comprises an attractive force between the target specific polynucleotide sequence and the electrode surface. The attractive force is sufficient to hold the target specific polynucleotide sequence to the electrode surface but also sufficient to provide (or allow) complete dissociation of the target specific polynucleotide sequence when the target specific polynucleotide sequence associates with the target substrate. Adsorption of the target specific polynucleotide sequence to the electrode surface includes, but is not limited to ion-ion interactions, for example between attractive charges; covalent interactions; electrostatic interactions; dipole interactions; ion-dipole interactions; hydrogen bonding interactions; van der Waals interactions; pi-stacking interactions; sharing of electron density or combinations thereof between the target specific polynucleotide sequence. The adsorption of the target specific polynucleotide sequence to the electrode excludes an interaction of the target specific polynucleotide sequence with the electrode, wherein the interaction is sufficient to bind the sequence to the electrode, and wherein the binding prevents complete dissociation of the target specific polynucleotide sequence from the electrode when the target specific polynucleotide associates with the target substrate. Such an interaction may be present in polynucleotides that are tethered to the electrode surface. More preferably, the adsorption of the target specific polynucleotide sequence to the electrode surface is due to a soft-soft Lewis acid-base interaction. The adsorption of the target specific polynucleotide to the electrode is sufficient to hold the target specific polynucleotide sequence to the electrode, but sufficient to also provide complete dissociation of the target specific polynucleotide sequence when the target specific polynucleotide sequence associates with the target substrate. Such adsorption interactions would be readily understood by those of normal skill in the art.

In an embodiment of any of the aspects of the invention, the target specific polynucleotide sequence is specific for a target substrate. The target specific polynucleotide sequence is selected from DNA polynucleotide sequences and RNA polynucleotide sequences. Where the target specific polynucleotide is a DNA sequence the DNA sequence is selected from a dsDNA polynucleotide sequence and an ssDNA polynucleotide sequence. The target specific polynucleotide sequence may also be a target specific polynucleotide sequence probe.

Alternatively, the target specific polynucleotide sequence is an aptamer. Preferably, the aptamer is selected from a ssDNA polynucleotide sequence and an RNA polynucleotide sequence. Those of skill in the art will realise that a target specific polynucleotide sequence may be modified without substantially altering the selectivity of the target specific polynucleotide sequence for the target substrate.

In a further embodiment of any of the aspects of the invention, the target specific polynucleotide sequence comprises a ligand binding domain. Preferably, the target specific polynucleotide sequence has been optimised for use in the electrochemical method to provide a target specific polynucleotide sequence comprising nucleotide bases in addition to the ligand binding domain. The additional nucleotide bases are selected to provide optimal signal transduction. Preferably, the additional nucleotide bases are situated at an end selected from the 5' end, the 3' end, or the 5' and 3' ends of the polynucleotide sequence. In an embodiment of any of the aspects of the invention, the optimisation comprises comparing two or more target specific polynucleotide sequences, wherein at least one target specific polynucleotide sequence has one or more additional nucleotide bases in addition to the ligand binding domain, and selecting a preferred target specific polynucleotide sequence based on optimal sensitivity toward the small target molecule and/or other target substrates in the electrochemical sensor. The additional nucleotide bases comprise from 1 to 10 nucleotide bases at the selected end or ends. Preferably, the additional nucleotide bases comprise from 2 to 8 additional nucleotide bases at the selected end or ends. Most preferably, the additional nucleotide bases comprise from 4 to 7 nucleotide bases at the selected end or ends. The additional nucleotide bases may be selected from 1 , 2, 3, 4, 5, 6, 7, 8, 9 and 10 nucleotide bases at the 3' end and/or 1 , 2, 3, 4, 5, 6, 7, 8, 9 and 10 nucleotide bases at the 5' end.

In an embodiment of any of the aspects of the invention, the electrode may be selected from a glassy carbon electrode; a metal oxide electrode; a conducting polymer electrode; and a noble metal electrode (including flat electrodes and nanoparticle electrodes). In another embodiment of any of the aspects of the invention, the noble metal of the noble metal electrode is stable to air, oxygen and solvents (such as organic solvents and water). Preferably, the noble metal is a conductive noble metal.

Preferably, the noble metal is selected from gold, ruthenium, rhodium, palladium, silver, and platinum. Preferably, the noble metal is gold. The noble metal electrode may comprise a flat noble metal electrode or a glassy carbon electrode coated with a noble metal. Where the noble metal electrode is a glassy carbon electrode coated with a noble metal, the noble metal may take the form of noble metal particles. The noble metal particles may be selected from noble metal microparticles, nanoparticles or quantum dots. Preferably the noble metal microparticles, nanoparticles or quantum dots are selected from gold, ruthenium, rhodium, palladium, silver, and platinum.

In a further embodiment of any of the aspects of the invention, the electrode is substantially covered (or substantially coated) by the target specific polynucleotide sequence. The target specific polynucleotide sequence substantially coats the electrode by means of adsorption. The adsorption of the target specific polynucleotide sequence to the electrode to substantially coat the electrode is sufficient to adsorb the sequence to the electrode, but also provides complete dissociation of the sequence from the electrode on association of the target specific polynucleotide sequence with the target substrate. Preferably, the target specific polynucleotide sequence when adsorbed to the electrode surface is substantially in contact with the electrode substantially along the length of the target specific polynucleotide sequence. The adsorption of the target specific polynucleotide sequence to the electrode may substantially prevent interfering particles from accessing the electrode by passivating the electrode (to interfering polynucleotide sequences, molecules or ions). Preferably, the target specific polynucleotide sequence substantially passivates the electrode. Preferably, the target specific polynucleotide sequences substantially adsorb and substantially contact the electrode and do not substantially form tethers that extend linearly away from the electrode. More preferably, the target specific polynucleotide sequences do not substantially form tethers such that the target specific polynucleotide sequence extends linearly away from the electrode and such that the target specific polynucleotide sequences substantially contact and/or (coat), substantially along the surface of the electrode.

In another embodiment of any of the aspects of the invention, the binding or association between the target specific polynucleotide sequence and the target substrate involves the attractive binding (i.e. non-repelling) of two or more species held together by attractive forces. Such binding comprises an interaction between the target specific polynucleotide sequence and the target substrate that pulls (or draws) the target specific polynucleotide sequence and target substrate in a sample together. Binding or association includes, but is not limited to covalent interactions; electrostatic interactions; ion-ion interactions, for example between attractive or opposite charges; dipole interactions; ion-dipole interactions; hydrogen bonding interactions; van der Waals interactions; pi-stacking interactions; the sharing of electron density or combinations thereof.

In another embodiment of any of the aspects of the invention, the sample containing the target substrate to be detected by the method of the present invention may be an environmental sample, for example a water sample, soil sample, or even a plant sample. Alternatively, the sample may be a biological sample. The biological sample may be from an animal, for example a tissue sample, a hair sample, wool sample, a urine sample, a blood sample, a serum sample, a saliva sample, a semen sample, or a faecal sample. In yet another embodiment of any of the aspects of the invention, the target substrate may be selected from a small target molecule or a target polynucleotide sequence or a target protein or peptide.

Small target molecules to be detected as target substrates may be selected from pollutants, chemicals that mimic hormones, hormones, naturally occurring phytoestrogens, narcotics, and metabolites thereof. Preferably, the target substrate is an endocrine disrupting compound, a steroidal sex hormone, metabolites, or synthetic variants thereof. More preferably, the target substrate is selected from endocrine-disrupting compounds, and metabolites thereof. More preferably, the target substrate belongs to the estrogenic family of compounds. Even more preferably, the target substrates are selected from 17 ?-oestradiol (E2); oestrone; oestriol; androstenedione; testosterone; dihydrotestosterone; pregnenolone; progesterone; 17a-hydroxyprogesterone, 17a-ethynylestradiol; isoflavones; lignans; coumestans; organohalides including organochlorines, polychlorinated organic compounds, polychlorobiphenyl (PCB); alkylphenols; alkylphenol ethoxylates; phthalates; bisphenol-A (BPA); Bis (4-hydroxyphenyl)methane; cholesterol; adenosine; triclosan; or synthetic steroids such as d iethy Isti Iboestrol (DES); cocaine, heroin and any metabolites of the mentioned compounds thereof. Even more preferably, the target substrate to be detected is selected from 17 ?-oestradiol, testosterone, progesterone, adenosine and BPA.

Alternatively, the target substrate to be detected may also be hormone or a marker of a condition of disease in a body. For example, the target specific polynucleotide sequence (suitable for adsorption to the electrode) could be selective for the detection of hormones and/or metabolites to establish fertility, or status in an animal. Alternatively, the target specific polynucleotide sequence (that is suitable for adsorption) can be selected for the detection of known markers of disease, for example overexpression of a cancer gene to detect cancer, detection of molecules associated with infection, or to establish levels of specific metabolites associated with a particular condition.

Target substrate polynucleotide sequences may be selected from target substrate DNA sequences or target substrate RNA sequences. Target substrate DNA sequences include but are not limited to single stranded DNA sequences (ssDNA), double stranded DNA sequences (dsDNA), extracellular DNA (eDNA), complementary DNA (cDNA); gene sequences. Alternatively, target substrate RNA sequences include, but are not limited to messenger RNA (mRNA), transfer RNA (tRNA), ribosomal RNA (rRNA).

The examples and figures described herein are for the purposes of illustrating embodiments of the invention. Other embodiments, methods, and types of analyses are within the capabilities of persons of ordinary skill in the art and need not be described in detail herein. Other embodiments within the scope of the art are considered to be part of this invention. Description of the Figures

Figure 1 a) EIS spectra, presented as Nyquist plots for a flat Au electrode (O) modified with a 35-mer aptamer via Au-S (Δ) (SEQ ID No: 3) and non-Au-S couplings (□) (SEQ ID No: 4).

b) Same as a) but with AuNP electrode.

c) Cyclic voltammetry conducted at a scan rate of 20 mV/s for AuNP electrode (1 ) and Au-S (3) (SEQ ID No: 3) and non-Au-S (2) (SEQ ID No: 4) coupling of a 35-mer aptamer.

Measurements conducted in phosphate-buffered saline solution with 5.0 mM

Fe(CN) ₆ ^3_/4~ . The Randies equivalent circuit model shown as an inset in a) was used to fit the experimental data (symbols) of EIS.

Figure 2 a) SEM image for glassy carbon electrode (top) and AuNP modified electrode

(bottom).

b) EIS spectra, presented as Nyquist plots for bare GCE electrode (O) and after electrochemical deposition of AuNPs (□).

c) Current vs. square root of scan rate for bare GCE (□), AuNP/ GCE modified electrode (O) for surface area calculation.

d) Representative cyclic voltammetry for GCE (1 ) and AuNP/ GCE (2) at scan rate of 10 mV/s.

Figure 3 EIS spectra presented as Nyquist plots for:

a) EIS spectra presented as Nyquist plots for tethered (Δ) and substantially adsorbed 35-mer aptamer (SEQ ID Nos: 3 and 4 respectively) (O) on a AuNP electrode deposited on flat Au (□).

b) EIS spectra presented as Nyquist plots for 75-mer tethered (O) and substantially adsorbed (SEQ ID Nos: 2 and 1 respectively) (Δ) on a AuNP electrode (□) grown on GCE.

c) EIS spectra presented as Nyquist plots for random 35-mer substantially adsorbed (SEQ ID No: 7) (Δ) on a AuNP electrode (□) grown on GCE. d) EIS spectra presented as Nyquist plots for random 75-mer substantially adsorbed (SEQ ID No: 6) (Δ) on a AuNP electrode (□) grown on GCE e) EIS spectra presented as Nyquist plots for 30-mer substantially adsorbed (SEQ ID Nos: 8, 9, 10, and 11 ) on an AuNP electrode grown on GCE. Symbols represent AuNP electrode (□), polyA (*), polyG (Δ), polyT (+), and polyC (O). Figure 4 EIS spectra presented as Nyquist plots characterising control experiments of exposing the 35-mer aptamer (SEQ ID No: 4) on GCE for 20 mins.

Figure 5 a) Substantially adsorbed 35-mer aptamer (SEQ ID No: 4) density before and after exposure to 20 μΜ E2 measured via chronocoulometry. The extrapolated y- intercepts shown by solid lines indicate the surface excess charge for measurements with and without the RuHex for the substantially adsorbed 35-mer aptamer (SEQ ID No: 4) with no E2.

b) Summary of surface densities of 75-mer aptamer (SEQ ID No: 1 ) and 35-mer aptamer (SEQ ID No: 4) before and after incubation with 20 μΜ E2. The respective random aptamer (SEQ ID Nos: 6 and 7) results are also shown.

Figure 6 a) Charge transfer resistance (RCT) for AuNP electrode, tethered 35-mer aptamer

(SEQ ID No: 3), substantially adsorbed 35-mer aptamer (SEQ ID No: 4) and treatment with MCH (3), MHA (2), BSA (4), and Lysozyme (5).

b) Examination of loss of AUNP from the electrodeposited surface after washing the electrode with buffer. AuNP electrodes were either not coated, coated with the tethered 35-mer aptamer (SEQ ID No: 3), or substantially adsorbed (SEQ ID No: 4).

c) Buffer washing cycles of a flat Au electrode modified with the tethered 35-mer aptamer (SEQ ID No: 3) and the substantially adsorbed 35-mer aptamer (SEQ ID No: 4).

Figure 7 EIS spectra presented as Nyquist plots characterising the stability of different electrodes

a) EIS spectra presented as Nyquist plots characterising the stability of AuNP/ GCE electrode against washing cycles of BWB. Symbols represent GCE (□), NPs (Δ), 1 ^st cycle (H), 2 ^nd cycle ( ), 3 ^rd cycle (O), 4 ^th (* ), 5 ^th cycle (+), and 6 ^th cycle (#).

b) EIS spectra presented as Nyquist plots characterising the stability of tethered

35-mer on AuNP (SEQ ID No: 3) / GCE electrode against washing cycles of BWB. Symbols represent GCE (□), NPs (Δ), 1 ^st cycle (HI), 2 ^nd cycle (O), 3 ^rd cycle (O), 4 ^th (*), 5 ^th cycle (+), 6 ^th cycle (*), and 7 ^th cycle (·); and c) EIS spectra presented as Nyquist plots characterising the stability of adsorbed 35-mer on AuNP (SEQ ID No: 4) / GCE electrode against washing cycles of

BWB. Symbols represent GCE (□), NPs (Δ), 1 ^st cycle (S), 2 ^nd cycle (O), 3 ^rd cycle (O), 4 ^th (*), 5 ^th cycle (+), and 6 ^th cycle (#). Figure 8 a) Differential change in R _CT vs. concentration of E2 with the tethered and substantially adsorbed 35-mer aptamer (SEQ ID Nos: 3 and 4 respectively) on a flat Au electrode, along with the response of the bare electrode (error bars indicate the average of two experiments).

b) EIS spectra presented as Nyquist plots characterising the response towards E2 of a substantially adsorbed 35-mer aptamer (SEQ ID No: 4) on an AuNP electrode. Experimental data (symbols) are fitted to the Randies cell equivalent circuit (lines).

c) Differential change in R _CT vs. concentration of E2, 1 : 1 mixture of E2/bisphenol-

A (BPA), progesterone (P4), and bisphenol-A (BPA) for a substantially adsorbed 35-mer aptamer (SEQ ID No: 4) on an AuNP electrode. The response of a randomised 35-mer (SEQ ID No: 7) towards E2 is also shown as a control experiment. Error bars represent the range of two experiments. Error bars represent the standard deviation of five experiments for E2 sensing and the range of two for specificity and control experiments.

Figure 9 EIS spectra presented as Nyquist plots characterising the response towards E2 with the 35-mer aptamer (SEQ ID No: 4) tethered to a flat Au electrode. Experimental data (symbols) are fitted to the Randies cell equivalent circuit

(lines). Symbols represent 0 E2 (□), 20 fM E2 (Δ), 20 pM E2 (O), 20 nM (O), and 20 μΜ (★).

Figure 10 a) EIS spectra presented as Nyquist plots characterising the response towards

E2 with the 35-mer aptamer (SEQ ID No: 4) substantially adsorbed on a flat Au electrode. Experimental data (symbols) are fitted to the Randies cell equivalent circuit (lines). Symbols represent 0 E2 (□), 20 fM E2 (Δ), 20 pM E2 (O), 20 nM ( ), and 20 μΜ (* ).

Figure 1 1 a) Nyquist plots of the 35-mer aptamer (SEQ ID No: 4) adsorbed on AuNPs and sequentially exposed to spiked rat urine with a range of E2 concentrations, b) Differential RCT as a function of E2 addition to rat urine. The response towards E2 is also shown for a control electrode with a random 35-mer DNA (SEQ ID No: 7).

Figure 12 Differential change in R _C T VS. concentration of E2, progesterone (P4), bisphenol- A (BPA) characterising the response of the 35-mer aptamer (SEQ ID No: 4) (with random 35-mer control (SEQ ID No: 7)) and 75-mer aptamer (SEQ ID No: 1 ) (with random 75-mer control (SEQ ID No: 6)).

Figure 13 CV experiments characterising the stability of different electrodes

a) CV experiments characterising the stability of AuNP electrode after 6 washing cycles of BWB. Numbers indicate GCE (1 ), NPs (2), NPs after washing (3), b) CV experiments characterising the stability of AuNP with tethered 35-mer aptamer (SEQ ID No: 3) electrode after 6 washing cycles of BWB. Numbers indicate GCE (1 ), NP-tethered 35-mer aptamer (SEQ ID No: 4) (2), NP-tethered 35-mer aptamer after washing (SEQ ID No: 4) (3), and

c) CV experiments characterising the stability of AuNP with adsorbed 35-mer aptamer (SEQ ID No: 4) electrode after 6 washing cycles of BWB. Numbers indicate GCE (1 ); NP-adsorbed 35-mer aptamer (2); NP-adsorbed 35-mer aptamer after washing (3).

Figure 14 a) EIS spectra presented as Nyquist plots characterising the response towards

E2 with the substantially adsorbed 35-mer aptamer (SEQ ID No: 4) on a flat Au electrode. Experimental data (symbols) are fitted to the Randies cell equivalent circuit (lines).

Figure 15 a) Nyquist plots of the 75-mer BPA aptamer (SEQ ID No: 5) adsorbed on AuNPs and sequentially exposed to a range of BPA concentrations in buffer. Symbols represent 0 BPA (□), 1 fM (Δ), 20 fM (Ξ), 1 pM (O), 20 pM (O), 1 nM (★), 20 M (+), 1 μΜ (*), and 20 μΜ (·).

b) Differential R _CT as a function of exposure to different BPA concentrations using the 75-mer BPA aptamer (SEQ ID No: 5) (□). The response towards BPA is also shown for a control electrode with a random 75-mer DNA (0)(SEQ ID No: 6). Figure 16 Electrode cleaning procedures; 1 ^st to 4 ^th steps were completed for blood samples and 1 ^Et to 2 ^nd steps for PBS measurements.

Figure 17 A) Schematic representation of non-specific adsorption of aptamer sequences onto Au electrode. B) Nyquist diagrams of bare and aptamer functionalized Au electrodes proving successive adsorption. Figure 18- A) Nyquist diagrams of gold electrodes with physically adsorbed DNA on the surface upon E2 incubation with a concentration range of 1fM to 100 nM. Experimental data are presented as symbols and the fitting curves to the equivalent circuit as solid lines. B) Normalized sensor response (ARQT/RCJ ⁰ ) versus E2 concentration of electrodes functionalised with physically adsorbed

DNA. Error bars represent n=3 measurements.

Figure 19 Nyquist diagrams obtained from measurements with A) %1 C) %5 human blood in the target solutions containing externally added E2 from 1fM to 1 pM. Experimental data are presented as symbols and the fitting curves to the equivalent circuit as solid lines. Normalized sensor responses, AR _C T/RCJ°, of the electrodes that were incubated with B) %1 and D) 5% human blood. Black symbols represent the measurements after incubating electrodes with externally added E2, white coloured symbols represent the measurements performed after incubating without any added E2. n=2

Detailed Description Definitions

The chemical terminologies as used herein have their standard meanings known in the art, in accordance with the lUPAC Goldbook, unless explicitly stated. Unless the context clearly requires otherwise, throughout the description and the claims, the words "comprise", "comprising" and the like, are to be construed in an inclusive sense as opposed to an exclusive or exhaustive sense, that is to say, in the sense of "including, but not limited to". The term "animal" is intended to mean human and non-human subjects. For example, humans; domesticated stock including cows, sheep, goats, horses, pigs, deer; domesticated pets including cats, dogs; wild animals including monkeys, birds, amphibians, reptiles; aquatic life forms such as fish. The term "aptamer" as described herein is intended to mean a single strand of RNA or DNA that specifically binds to particular target molecules. The term "aptamer" relates to polynucleotide or oligonucleotide sequences. The aptamers used in the present invention are specific for a particular target substrate. Those skilled in the art will also readily understand that minor variation of the sequence code of the aptamer may be made by standard methodology without substantially affecting the binding of the substrate to the aptamer.

The term "association", "attraction" or "attractive" means assembling of separate molecular or ionic entities into an aggregate. The entities include but are not limited to oppositely charged free ions, ion pairs, charged molecules, compounds, proteins, and clusters of ions held together by attractive (non-repelling) forces, such as, electrostatic interaction or covalent interactions.

The term "conformational changes" means a change in the conformational form of the aptamer, for example, a change from a tightly folded structure to a loose linear-type structure that results in opening up of the binding site, or from a loose linear-type structure to a tightly folded structure. This type of alteration would be readily understood by those skilled in the art.

The term "dissociation" means the separation of an aggregate comprising two or more entities into separate entities.

The term "estrogenic family" of compounds means compounds that are chemically related to estrogens. Estrogenic compounds may be natural or synthetic, steroidal or non-steroidal, and includes metabolites of such compounds. The term "flanking sequences" means non-binding portion of the nucleotide sequences at one or both ends of the molecule and may include primers.

The term "interfering agents" means particles, molecules, compounds, biomarkers, antibodies, antigens, ions, polynucleotide sequences, or combinations thereof that the polynucleotide sequence that is substantially adsorbed onto the electrode surface is not selective for and does not form a substantial associative (or binding) interaction with. For example, such interfering agents may adversely affect the selection of the substrate by interfering with the signalling process. The terms "adsorption" or "binding" are as defined according to the lUPAC Goldbook. That is that an attractive interaction draws or pulls the two or more species together and does not include the repelling of the two or more species. Such interactions include, but are not limited to covalent interactions; electrostatic interactions; ionic interactions, for example between attractive or repelling charges; dipole interactions; ion-dipole interactions; hydrogen bonding interactions; and van der Waals interactions; pi-stacking interactions or combinations thereof.

In respect of substantial adsorption of the target specific polynucleotide sequence to the electrode, the target specific polynucleotide adsorbs to the surface of the electrode such that the length of the target specific polynucleotide is substantially adsorbed substantially along the surface of the electrode to substantially cover or substantially coat the surface of the electrode. The adsorption is such that it is sufficient to hold or bind the target specific polynucleotide sequence to the electrode, but provides complete dissociation of the sequence from the electrode when the polynucleotide sequence binds the target substrate.

The term "polynucleotide sequence" means a nucleic acid sequence selected from double stranded DNA (dsDNA), single stranded DNA (ssDNA) or RNA. Preferably, the polynucleotide sequence has selectivity and is specific for a target substrate.

The term "particle" is intended to encompass nanoparticles, microparticles and quantum dots. That is, nanoparticles, microparticles or quantum dots may be used, and use of one term throughout the specification is not intended to exclude the others, unless expressly stated.

The term "sample" is intended to mean a sample isolated from or collected from an environmental or biological source and that sample is located ex vivo. The sample may be of biological origin, isolated from an animal or may be collected from the environment. Sources of samples may include without limitation, for example soils, waterways, tissue, blood, serum, urine, saliva, faeces, hair and wool.

The term "salt" is intended to apply to salts derived from inorganic or organic acids, including, but not limited by the following salts: halides (chloride, bromide, iodide fluoride), acetate, adipate, alginate, aspartate, benzoate, benzenesulfonate, bisulfate, butyrate, citrate, camphorate, camphorsulfonate, cyclopentanepropionate, digluconate, dodecylsulfate, ethanesulfonate, formate, fumarate, glucoheptanoate, hydrochloride, hydrobromide, hydroiodide, 2-hydroxyethanesulfonate, lactate, maleate, malonate, methanesulfonate, nitrate, oxalate, persulfate, phosphate, picrate, pivalate, propionate, p- toluenesulfonate, salicylate, succinate, sulfate, tartrate, thiocyanate, and undecanoate, and may comprise the cations selected from ammonium (NH ₄ ⁺ ), calcium (Ca ²⁺ ), lithium (Li ⁺ ), magnesium (Mg ²⁺ ), potassium (K ⁺ ), Sodium (Na ⁺ ).

The term "small molecules" is intended to mean compounds of simple molecular structure with a Mw of from about 60 to about 2000 g mol ^"1 , alternatively in the range of from about Mw 100 to 500 g mol ^"1 , more alternatively of from about 150 to 350 g mol ^"1 . The molecular weight of such compounds and the calculation of the molecular weights are well known to those of skill in the art. Such compounds include, without limitation, pollutants, hormone mimics, hormones, naturally occurring phytoestrogens, narcotics and metabolites thereof, organohalides and compounds, such as, 17 ?-oestradiol (E2); oestrone; oestriol; androstenedione; testosterone; dihydrotestosterone; pregnenolone; progesterone; 17a- hydroxyprogesterone, 17a-ethynylestradiol; isoflavones; lignans; coumestans; organohalides including organochlorines, polychlorinated organic compounds, polychlorobiphenyl (PCB); alkylphenols; alkylphenol ethoxylates; phthalates; bisphenol-A (BPA); Bis (4-hydroxyphenyl) methane; cholesterol; adenosine; triclosan; or synthetic steroids including but not limited to diethylstilboestrol (DES); cocaine, heroin and any metabolites of the mentioned compounds thereof.

In respect of the term "tethered" in relation to a target specific polynucleotide sequence, the target specific polynucleotide sequence is tethered to the electrode through a linking moiety to the electrode. The tethering moiety prevents complete dissociation of the target specific polynucleotide sequence from the electrode. Such tethering (or linking) moieties include but are not limited to amino, carboxylate, thio, carbamate, carbonyl moieties, and polynucleotide sequences that act as a tether. The tether provides the linking or contacting of the sequence to the electrode such that the bulk of the sequence extends away from the electrode Such tethering (or linking) moieties would be readily understood by those of normal skill in the art.

As used in the specification and the appended claims, the singular forms "a," "an" and "the" include plural referents unless the context clearly dictates otherwise.

Abbreviations

AuNP Gold nanoparticles

BPA bisphenol A

BPF Bis(4-hydroxyphenyl methane)

DLS Dynamic Light Scattering D _t particle diffusion coefficient

E2 17/?-oestradiol

ECP Electrochemical conducting polymer

EIS Electrical impedance spectroscopy

FWHM Full Width at Half Maximum

GCE Glassy carbon Electrode

HAuC Chloroauric acid

MP microparticle

NP nanoparticle

P4 Progesterone

PBS Phosphate buffered saline

PSS Poly(4-styrenesulfonic acid) solution

Py Pyrrole

RCT Charge transfer resistance ssDNA single strand DNA

SCE saturated calomel electrode

SEM Scanning electron microscopy

T Testosterone

Zim Imaginary component of impedance

Real component of impedance

Discussion

The present inventors have advantageously established that target specific polynucleotide sequences (such as dsDNA, ssDNA, and RNA sequences) that are substantially adsorbed onto an electrode surface (for example glassy carbon electrodes; metal oxide electrodes; conducting polymer electrodes; and noble metal electrodes including gold, ruthenium, rhodium, palladium, platinum and silver) show different interfacial properties (such as interfacial charge distribution) when compared to target specific polynucleotide sequences that are tethered (Figure 1 D). Consequently, this adsorption leads to pronounced differences (such as surface charge distribution and surface potential) between the tethered and adsorbed systems that can be resolved by techniques including, but not limited to EIS, cyclic voltammetry and differential pulsed voltammetry or combinations thereof. Such techniques would be readily known to those of skill in the art. This finding has led to the use of these properties in the development of methods for detecting target substrates in a sample using polynucleotide sequences.

The electrochemical detection method of the present invention is found to provide effective and selective detection methods for target substrates in a sample; particularly when compared with electrode systems comprising target specific polynucleotide sequences tethered to an electrode through linkers or tethers. This link/tether can be achieved through conventional tethering means, such as, a covalent linker including amino, carboxylate, thio, carbamate, carbonyl moieties to electrodes, or through the use of other polynucleotide sequences, although those of skill in the art will realise that tethered orientation of the target specific polynucleotide sequence may be formed by other interactions.

The present invention works by providing a method utilising polynucleotide sequences specific for a target substrate and electrochemical methodology (i.e. a target specific polynucleotide sequence). Electrodes, such as glassy carbon electrodes; metal oxide electrodes; conducting polymer electrodes; and noble metal electrodes (including gold, ruthenium, rhodium, palladium, platinum and silver) or noble metal particles deposited on a glassy carbon electrode to provide a noble metal coated surface) may be used. The noble metal (such as gold, ruthenium, rhodium, palladium, platinum and silver) may be deposited onto an electrode surface in particle form, for example, in the form of microparticles, nanoparticles or quantum dots. Those of skill in the art will realise that use of one type of particle is not intended to exclude the use of any of the other types of particles. The electrode surface is exposed to a polynucleotide sequence to which the polynucleotide sequence substantially adsorbs (substantial association or interaction) to substantially contact and/or substantially coat the surface of electrode to provide a target specific polynucleotide sequence coated electrode. Coating the electrode surface with the target specific polynucleotide sequence substantially passivates the electrode surface and prevents other species from interacting with the electrode. The target specific polynucleotide sequence coated electrode can then be exposed to the sample having a target substrate. In doing so, the target specific polynucleotide sequence dissociates from the electrode surface and interacts with the target substrate. This dissociation can be measured by electrochemical techniques known in the art but not limited solely to, techniques, such as EIS or cyclic voltammetry, and can be applied to the detection of target substrates in a sample. Those of skill in the art will realise that other electrochemical detection methods may be used without deviating from the spirit and scope of the invention. For example, in the case of noble metal nanoparticle (NP) electrodes (such as Au), substantial adsorption of the target specific polynucleotide sequence (aptamers) to the noble metal surface has the benefit of stabilising any loosely interconnected AuNPs. The stabilisation occurs because the target specific polynucleotide sequence adsorbs substantially along the surface of the noble metal electrode. This adsorption along the noble metal electrode surface is not in a tethered-type fashion that extends away from the noble metal surface, but the polynucleotide sequence is substantially contacted and adsorbed, substantially along the noble metal electrode surface (See Figure 1 D). Conversely, target specific polynucleotide sequences tethered to the electrode through a linking moiety (e.g. via a thiol linker, or another polynucleotide sequence) move the charge distribution of the polynucleotide sequence further away from the electrode interface, and leave some exposed electrode surface between the strands of the polynucleotide sequence (Figure 1 ). Consequently, the exposed surface on the tethered polynucleotide sequence electrode needs to be further passivated to avoid spurious signals from being detected. Polynucleotide sequences substantially adsorbed onto the electrode surfaces such as glassy carbon; metal oxide; conducting polymer; and noble metal electrodes including, flat, nanoparticle, gold, ruthenium, rhodium, palladium, platinum and silver, are found to show decreased R _CT values, which provides a strong signal for measurement when the target specific polynucleotide sequence dissociates from the electrode surface; lower total aptamer content under saturation; and effective passivation that substantially prevents interference from interfering molecules. It is also found that polynucleotide sequences substantially adsorbed onto noble metal nanoparticle electrodes provide substantial stabilisation of the noble metal nanoparticle electrodes so that the noble metal nanoparticles are retained on the electrode surface when compared with tethered polynucleotide sequences (aptamers).

It has been established by the present inventors that having greater electrode surface contact area for a target specific polynucleotide sequence (such as dsDNA, ssDNA, and RNA sequences) to adsorb to, leads to a more lateral binding mode along the electrode surface with a large fraction of the bases directly at the electrode interface (within the Debye screening length) when compared to tethered polynucleotide sequences that extend away from the electrode surface (in a linear type fashion), rather than spreading along surface. The substantial adsorption of the target specific polynucleotide sequence to the electrode surface leads to displacement of a dense layer of ions that are found directly at the electrode interface at high electrode potentials. Therefore, enhanced charge transfer and greater sensitivity to subsequent changes are observed when compared to target specific polynucleotide sequences that are tethered to an electrode surface. The enhanced charge transfer signals of substantially adsorbed polynucleotide sequences compared to tethered target specific polynucleotide sequences, and the intrinsic passivation by the target specific polynucleotide sequences are exploited in the present invention using substantially adsorbed target specific polynucleotide sequences on electrodes, including noble metal electrodes (such as gold, ruthenium, rhodium, palladium, platinum and silver).

Because the current method does not require passivation, the sensor can be utilised in an initial colorimetric type assay, but can then be directly utilised in the claimed electrochemical assay if the requirement dictates a more sensitive and quantitative result. The present invention resolves changes in R _CT when target-bound polynucleotide sequences dissociate from the noble metal electrode. This differs from current EIS sensors in the art because current EIS sensors are tethered to the electrode and probe conformation changes near the electrode. The present invention has been particularly exemplified using a noble metal such as gold or gold nanoparticles for the detection of E2 using aptamers, in particular using DNA aptamers (Table 1 ). Those of skill in the art will realise and understand that any other noble metal can be used as an alternative to gold without deviating from the spirit and scope of the invention. Polynucleotide sequences defined in Table 1 that comprise a SH moiety, are polynucleotide sequences that may be tethered to the noble metal electrode surface. Table 1

On contact of an electrode with an electrolyte solution, an electric double layer is created by strongly polarised electrolyte ions found at the charged noble metal electrode surface. The distance that the double layer region extends from the electrode can be estimated by the Debye screening length A _D :

where e ₀ is the permittivity of free space;

e _r the dielectric constant of the electrolyte,

k is Boltzmann's constant,

T is the temperature in degrees Kelvin,

N _A is Avogadro's number,

e is the charge of an electron, and I is the ionic strength.

A 1.7 M ionic strength electrolyte solution is used in the present examples to exemplify the present invention, the Debye length is 0.7 nm. A substantially significant fraction of the electric potential will be compensated by a dense layer of chloride anions at the noble metal electrode surface and is attributed to the ionic strength of the electrolyte solution, and also because the EIS measurement is conducted at a potential of +0.23 V. This potential is considered a high enough electrode potential such that there will be a very dense layer of ions at the electrode to compensate the electrode potential. The distribution of anions close to the electrode surface strongly affect the R _CT experienced by the anionic [Fe(CN) ₆ ] ^3" ' ^4" probe approaching the electrode solution interface because the impedance signal is dependent on the accessibility of the [Fe(CN) ₆ ] ^3-/4~ probe to the electrode: solution interface. A high concentration of ions at this interface will change the potential energy of approaching the electrode, and therefore change the R _CT measured.

The EIS data presented in Figure 1a shows comparative Nyquist plots for a flat Au electrode functionalised with a tethered thiolated 35-mer aptamer (SEQ ID No: 3) with the same aptamer lacking the thiol group (SEQ ID No: 4), together with the bare electrode. The Nyquist plots are fit to the Randies cell equivalent circuit (shown in the inset of figure 1 a). The interfacial charge-transfer resistance (RCT) is extracted from the data as the diameter of the characteristic semicircle. It can be seen in Figure 1 a that covalently coupling the 35-mer aptamer (SEQ ID No: 3) increased the R _CT from -332 Ω to 1700 Ω, and shows consistency with EIS investigations employing tethered aptamer [M. Gebala and W. Schuhmann, Chemphyschem, 2010, 11 , 2887-95; L. Fan, G. Zhao, H. Shi, M. Liu, and Z. Li, Biosens. Bioelectron., 2013, 43, 12-8; X. Liu, Y. Qin, C. Deng, J. Xiang, and Y. Li, Talanta, 2015, 132, 150-4; J. Tymoczko, W. Schuhmann, and M. Gebala, ACS Appl. Mater. Interfaces, 2014, 6, 21851-8]. The observed R _CT increase (when compared to the bare electrode) can be explained by the negative charge of the aptamer introducing an electrostatic barrier and impedes the access of negatively charged ferricyanide ([Fe(CN) ₆ ] ^{3_ 4~} ) redox probe molecules near the surface. However, the substantially adsorbed 35-mer aptamer (SEQ ID No: 4) results in significantly reduced R _CT , from ~ 332 Ω to 32 Ω when compared to the bare electrode surface.

The intrinsic negative charge density of an aptamer, such as an ssDNA aptamer can be approximated as 2.7 x 10 ²¹ charges/cm ³ . This intrinsic negative charge density is based on the structure of non-flexible double-stranded DNA, and assuming a cylinder with 2 nm diameter, 0.23 nm spacing, and two charges per base pair. The intrinsic charge density of a given aptamer is significantly higher than the charge density of the bulk electrolyte beyond the double layer (9 x 10 ¹⁹ charges/cm ³ , based on an electrolyte concentration of 0.15 M). The tethered thiol 35-mer aptamer (SEQ ID No: 3) can extend up to -12 nm away from the surface of the electrode. This is essentially bulk solution for such a short Debye length. If a polynucleotide sequence is in a more relaxed conformation (e.g fully extended or linear), most of the polynucleotide's sequence negative charges are distanced far enough from the noble metal surface (beyond the Debye length) such that they increase the negative charge density, thereby impeding access of negatively charged redox species to contacting the electrode surface and increasing the measured R _CT value.

Conversely, substantial adsorption of the target specific polynucleotide sequence to the noble metal surface via weak base interactions requires significant surface contact of the target specific polynucleotide sequence with the noble metal surface. This therefore requires that the volume of the target specific polynucleotide sequence must lie in direct proximity to the electrode. In order to achieve sufficient contact with the noble metal surface, such as gold, ruthenium, rhodium, palladium, platinum and silver, the target specific polynucleotide sequence (aptamer) may displace more surface charge density from the electrode:solution interfacial region than the intrinsic charge density that it brings to the surface, as depicted in Figure 1d. Therefore, concentration of counter-ions in the diffuse region of an electrode interface can be estimated by the Boltzmann equation: c¾ = c ₀ . e RT

(2)

where c,- and c ₀ represent the ion concentration at the electrode: solution interface and bulk respectively, z is the charge of the counter-ion, Ψ is the electrostatic potential, T is the temperature, F is Faraday's constant, and R is the gas constant. An interfacial charge density equivalent to that of the intrinsic aptamer charge density (-2.7 x 10 ²¹ charges/cm ³ ) corresponds to a potential difference of only 7.3 mV in the diffuse interface region relative to the bulk for a 1.7 M ionic strength solution as calculated according to equation (2).

Taking into account the applied electrode potential of 230 mV, and the higher ion density in the Stern layer (a dense layer of ions immediately at the electrode:solution surface) compared with ion density in the diffuse region, the inventors postulate that the anion density within the first nanometre of a bare electrode surface under EIS conditions must be substantially higher than the intrinsic charge density of aptamer. Conversely, when target specific polynucleotide sequences (such as ssDNA) displace surface charge to adsorb to the electrode, the net surface charge density is diluted by the polynucleotide sequence. Further, the [Fe(CN) ₆ ] ^3""/4- redox probe will experience a lower barrier to approaching the surface occupied by the substantially adsorbed polynucleotide sequences than for a bare electrode surface, which can account for the lower R _CT value measured.

Coverage and surface area of the electrode The EIS response in Figure 1 b of the same pair of DNA aptamers (SEQ ID Nos: 3 and 4) functionalised on an electrode comprising of electrodeposited noble metal nanoparticles (AuNPs). Noble metal nanoparticle electrodes, in particular AuNP electrodes, are versatile in that they offer the benefit over non-particle electrodes of providing a higher surface area than flat electrode surfaces [X. Li, H. Qi, L Shen, Q. Gao, and C. Zhang, Electroanalysis, 2008, 20, 1475-1482], as well as the ability to be grown onto any conductive substrate, including glassy carbon electrodes (GCE) and conducting polymers.

AuNPs were electrodeposited on a bare GCE for 30 seconds from a solution containing 1 mM HAuCI ₄ . The AuNP based electrode has a roughened morphology as shown by SEM images in Figure 2a. A surface roughness factor of approximately 2 was obtained by applying the Randles-Sevcik equation to scan-rate dependent cyclic voltammograms for the flat gold electrode versus AuNP electrodes (Figure 2c) [X. Zhang, M. R. Servos, and J. Liu, J. Am. Chem. Soc, 2012, 134, 7266-9. S. Hrapovic, Y. Liu, K. B. Male, and J. H. T. Luong, 2004, 76, 1083-1088; G. Jarzabek and Z. Borkowska, Electrochim. Acta, 1997, 42, 2915- 2918]. Figure 2b also shows that modifying a GCE with AuNPs leads to decreased R _CT (from -4000 Ω for GCE to -1000 Ω for AuNP coated). This decrease in R _CT can be attributed to the increased surface area of the AuNP electrode compared with the GCE electrode, and also reflected in the increased peak anodic and cathodic currents measured via cyclic voltammetry (CV) (Figure 2d) for the AuNP electrode compared with the GCE electrode. Similar behaviour is observed on the AuNP electrode (Figure 1 b) as described above for the flat Au electrode (Figure 1a) when comparing the EIS response of thiol tethered aptamers versus substantial adsorption of DNA aptamers. The measured R _CT of the bare AuNP surface (~ 1000 Ω) increased to -2500 Ω when the thiol tethered aptamer was added. However, R _CT decreased to 190 Ω for the substantially adsorbed aptamer on the AuNP electrode compared to the bare surface (~ 1000 Ω). The same trends were observed when AuNPs were electrodeposited on a flat Au electrode (Figure 3) and such behaviour is intrinsic to electrode surface interactions, rather than whether the electrode is flat or nanoparticulate.

Figure 1c shows cyclic voltammograms for an AuNP electrode compared with electrodes functionalised with thiol-tethered (SEQ ID No: 3) and substantially adsorbed 35-mer aptamers (SEQ ID No: 4). A pair of well-defined peaks at E ₀ = 225 mV when compared with SCE corresponds to the formal potential of the [Fe(CN) ₆ ] ^{3_ 4~} couple (Figure 1c). The cyclic voltammograms of the three electrodes differ in their peak currents, and are related to the heterogeneous electron transfer rate in that the larger the electron transfer rate, the larger the peak current. The thiol tethered aptamers (SEQ ID No: 3) leads to a suppressed interfacial electron transfer rate (lower peak current in CV), compared with the bare electrode whereas the adsorbed aptamers (SEQ ID No: 4) enhance the rate (higher peak current) compared with the bare electrode. The decrease in R _CT when DNA aptamers substantially adsorbed to Au electrodes was not limited to the 35-mer sequence (SEQ ID No: 4), but a general phenomenon and can be reasonably considered to apply to other substantially adsorbed polynucleotide sequences (aptamers). As shown in Figure 8, decreased R _CT values were also observed to different degrees for substantial adsorption of a 75-mer aptamer for E2 (SEQ ID No: 1 ); randomised versions of the E2 35-mer (SEQ ID No: 7) and E2 75-mer (SEQ ID No: 6); as well as 30-mer homonucleotides of poly-T (SEQ ID No: 8); poly-C (SEQ ID No: 10); and poly-G (SEQ ID No: 11 ). Only poly-A (SEQ ID No: 9) caused the R _CT value to increase slightly.

It is believed that the poly-A sequence (SEQ ID No: 9) behaves more like a thiol tethered sequence, where a small segment of the aptamer is in strong contact with the surface of the electrode and the remaining strand extends (or dangles) away from the surface of the electrode. This has been shown to be the case for 5-mer homo-oligonucleotides affinity to Au surfaces, poly-A sequence was found to strongly dominate over the other bases and could compete effectively against thiol binding [H. Kimura-Suda, D. Y. Petrovykh, M. J. Tarlov, and L J. Whitman, J. Am. Chem. Soc, 2003, 125, 9014-5].

Without wishing to be bound by theory, it is thought that in the method of the present invention substantial target specific polynucleotide sequence adsorption saturates the surface of the electrode and eliminates the need for additional passivation steps. The coupling densities of the polynucleotide sequence were investigated for the tethered sequences and substantially adsorbed sequences using chronocoulometry. The target specific polynucleotide sequence is exposed to a positively charged redox probe, Ru(NH ₃ ) ₆ , which electrostatically binds to the phosphate groups of the target specific polynucleotide sequence. The target specific polynucleotide sequence surface coverage density on the electrode is related to the density of bound probes, where one probe binds per phosphate, i.e., one per nucleic acid base. For a given sequence the amount of bound probes are proportional to the amount of target specific polynucleotide sequence on surface which is measured via chronocoulometry and separated from the unbound probe concentration via time-dependent currents as shown in Figure 5. Tethered thiol 35-mer aptamer (SEQ ID No: 3) has a surface density of 6.5 x 10 ¹³ molecule/cm ² or 1.5 nm ² per molecule, which corresponds to the maximum steric close-packing of DNA aptamer [D. Y. Petrovykh, H. Kimura-suda, L. J. Whitman, and M. J. Tarlov, 2003, 111 , 3011-3016]. Conversely, the substantially adsorbed aptamer (SEQ ID No: 4) had a surface density of 4 x 10 ¹³ molecule/cm ² , or 2.5 nm ² per molecule. The higher total density measured for thiol tethered aptamers (SEQ ID No: 3) is consistent with previous measurements using X-ray photoelectron spectroscopy, fluorescence labelling, and chronocoulometry [T. M. Heme and M. J. Tarlov, J. Am. Chem. Soc, 1997, 119, 8916-8920.; R. Lao, S. Song, H. Wu, L. Wang, Z. Zhang, L. He, and C. Fan, Anal. Chem., 2005, 77, 6475-80.; P. Sandstrom, M. Boncheva, and B. Akerman, Langmuir, 2003, 19, 7537-7543.]. This difference in coverage density is a consequence of polynucleotide sequence aligned along the surface versus away from the surface, however, this measurement does not resolve whether any Au surface remains exposed in either case.

The surface density of tethered polynucleotide sequence and substantially adsorbed polynucleotide sequence (before and after E2 detection) on a AuNP electrode was determined using the chronocoulometry method developed by Steel et al [Steel, A. B.; Heme, T. M.; Tarlov, M. J. Anal. Chem. 1998, 70, 4670-4677.]. The aptamer functionalised AuNP electrode was immersed in a low ionic strength electrolyte, 10 mM tris-HCI buffer at a pH 7.4, the potential stepped from 200 to -500 mV versus (Ag/AgCI) for 500 ms, and the resulting charge flow was measured. The electrode was then immersed for 20 min in a solution of 150 μΜ RuHex (hexaamineruthenium(lll) chloride (Ru(NH ₃ ) ₆ ³⁺ )) in tris buffer, and the measurement was repeated. In the low ionic strength tris-CI, the trivalent redox marker RuHex electrostatically associated to the negatively charged DNA phosphate groups in the ratio 1 :3. The charge Q as a function of time t from the potential step is the sum of the reduction of RuHex diffusing from solution, the double layer charge and the charge due to reduction of surface confined RuHex and is given by the integrated Cottrell equation:

2nFAD ^l2 c _n ^*

- ¹ 1/2 ^• t ^1/2 + Q _d i + nFAr ₀

Q is charge; n is the number of electrons per molecule for reduction;

F is the Faraday constant (C/eq);

A is the electrode area (cm ² ),

D ₀ is the diffusion coefficient (cm ² /s);

C ^* o is the bulk concentration of RuHex (mol/cm ² ),

Qdi is the capacitive charge,

nFA Γ ₀ is the charge from the reduction of adsorbed redox marker [C ₀ (mol/cm ² )],

r ₀ is the excess surface density of RuHex.

By plotting Q versus f ^{1 2} , the linear part of the charge release response was used to determine the intercept at time zero, which corresponds to (¾ + ηΡ>4Γ ₀ . By assuming the double layer capacitance to be equal in measurements with and without RuHex Q _d i for the fixed voltage step is constant and nFAf ₀ is calculated as the difference in intercepts, as shown in Figure 5.

The DNA surface density is determined from the surface excess of RuHex (Γ ₀ ), as shown in Figure 7a.

ΪΟΝΑ = ΓοΗ m ¾

I~ _DNA is the probe surface density (molecules/cm ² ),

m is the number of phosphate groups on the aptamer (75 and 35 groups for different aptamers),

z is the charge on the redox molecule (3 charges) and

N _A is Avogadro's number. Negatively charged 6-mercaptohexanoic acid (MHA) and neutral 6-mercaptohexanol (MCH) are known to have a strong affinity for noble metal surfaces (e.g. gold, ruthenium, rhodium, palladium, silver, and platinum); and form a dense self-assembled monolayer on the noble metal surface. In order to probe if the tethered or substantially adsorbed polynucleotide sequence leave accessible Au surfaces on the electrode, the polynucleotide sequence are exposed to MHA, as shown in Figure 6. The R _CT of a bare AuNP electrode substantially increases when it is exposed to the MHA. On exposure of the thiol-tethered polynucleotide sequence electrodes to MHA and MCH, an increase in R _CT is resolved compared to the bare electrode because the densely packed polynucleotide sequence leaves small exposed pores on noble metal electrode surface, allowing MHA and MCH to access the surface of the electrode. Conversely, MHA shows no effect on the R _CT of the substantially adsorbed polynucleotide sequence electrode. The inventors consider that MHA is prevented from accessing the surface of the noble metal electrode and may be aided by electrostatic repulsion between the negatively charged aptamer and MHA. MCH produced an R _CT response of the substantially adsorbed polynucleotide sequence electrode. The surfaces of the noble metals were also exposed to negatively charged bovine serum albumin (BSA) and positively charged lysozyme (LYS) to represent the interactions found in physiological medium. The RCT of the bare AuNP and terminally tethered thiol- polynucleotide sequence electrodes are substantially increased upon exposure to BSA and LYS. The invariance of R _CT for the substantially adsorbed polynucleotide sequence electrode confirms that the surface is passivated against BSA and LYS adsorption. It is for this reason that it is believed that that the substantially adsorbed polynucleotide sequences substantially cover/coat the noble metal surface of the electrode and passivate the noble metal surface against interaction with potentially interfering molecules.

Polynucleotide Sequence Stabilisation

Substantially adsorbed polynucleotide sequences play an additional role in stabilising the AuNP electrodes as can be seen in Figure 6. It is known that electrodeposited noble metal NP electrodes, such as gold, are instable due to loose electrostatic connectivity between the AuNPs. This problem can be overcome by crosslinking the nanoparticles together with dithiol molecules to Au electrodes or by derivation of GCE with sulfhydryl-terminated monolayer [L. Zhang, X. Jiang, E. Wang, S. Dong, Biosens. Bioelectron., 2005, 21, 337-45]. However, this adds further synthetic steps, expense and time into the process.

The loss of AuNPs from the surface of the AuNP coated electrode is seen via increases in RCT (towards the flat GCE electrode value) (Figure 7) when the AuNP electrode is sequentially washed with buffer (Figure 6), and when compared to freshly prepared AuNP electrode. The R _CT increase upon sequential washing steps was observed for both the bare AuNP electrode and the tethered thiol polynucleotide sequence (SEQ ID No: 3) AuNP electrode when compared to the freshly prepared AuNP electrode, preventing its use as a sensor. However, the substantially adsorbed polynucleotide sequence (SEQ ID No: 4) suppresses this undesired drift in R _CT after multiple cycles of washing with buffer. Repeating the measurement on a flat Au electrode (Figure 6) confirms that the effect is due to the instability of the AuNP electrode. While sensors known in the art that are derived from AuNP electrodes require additional AuNP crosslinking steps as well as surface passivation steps, the substantially adsorbed polynucleotide sequence of the present invention fulfils the roles of surface passivation and does not require cross-linking of the noble metal nanoparticles.

EIS data in Figure 6a shows loss of AuNPs during washing steps and is also supported by CV studies (Figure 13). The current associated with the GCE is almost recovered for the bare AuNP and thiol tethered polynucleotide sequence electrodes after washing (6x15 min). However, the current of the substantially adsorbed polynucleotide sequence electrode remained unchanged. It can therefore be concluded that that substantially adsorbed polynucleotide sequences produce electrode interfaces with fundamentally different properties (such as surface charge distribution, particle cross-linking, and surface passivation) than tethered thiol polynucleotide sequences do. Substantially adsorbed polynucleotide sequence spread along the Au surface with multiple base contact points because of a lack of covalent bonds with the electrode surface. This provides important benefits of the substantially adsorbed over tethered binding mode such as stronger perturbation of the charge distribution within the confined double layer region (and concomitantly reduced R _CT values), passivation of the Au surface when it is saturated with polynucleotide sequence, and ensuring the connectivity of AuNP electrodes when compared to those electrodes having tethered polynucleotide sequences.

A tethered thiol polynucleotide sequence AuNP system could not be reliably used for E2 sensing due to this AuNP instability that is identified above. However, the response of a thiolated polynucleotide sequence (SEQ ID No: 3) to E2 on a flat Au electrode was investigated. It is known in the art that the 35-mer E2 polynucleotide sequence (SEQ ID No: 3) has a dissociation constant of 1 1 nM and shows excellent selectivity towards E2 and does not interact with other potentially interfering molecules. It is found that sequential addition of E2 to buffered water results in increased R _CT values, as shown in Figure 8a (Nyquist plots are provided in Figure 9). This increase compared to the absence of a target is consistent with a previous ElS-based E2 sensor employing thiol-tethered polynucleotide sequence (Analyst, (2014) 137, 819-822; doi:10.1039/c1 an15856b), and can be understood as resulting from an increased negative charge density when the target-bound aptamer adopts a compact conformation compared to an extended or random conformation. Figure 8a also reveals the high background signal from E2 directly interacting with a bare Au surface and highlights the need to backfill the electrode with a passivating agent. This is because covalent tethering of polynucleotide sequences to the noble metal electrode leaves exposed surface, which results in spurious signals. Nonetheless, accounting for the background signal leads to a level of E2 detection of approximately 20 nM for the thiolated aptamer (SEQ ID No: 3). The substantially adsorbed 35-mer polynucleotide sequence (SEQ ID No: 4) was also investigated in the detection of E2. Substantially adsorbed polynucleotide sequences may still recognise the E2 target, but the signal transduction mechanism is expected to be different when compared to the tethered polynucleotide sequence. In a colorimetric AuNP sensor, aggregation of the polynucleotide sequences to the target substrate is triggered when target-bound polynucleotide sequences dissociate from noble metal nanoparticle (AuNP) surfaces. It is expected that target binding should be associated with a signal concomitant with reduced polynucleotide sequence surface coverage (compared to the bare electrode) because target-bound polynucleotide sequences dissociate from AuNP surfaces. Since adsorption of the polynucleotide sequences to the target substrate substantially reduce the R _CT relative to no target binding, the target binding should therefore result in an increase in R _CT values. Figure 8a shows that E2 addition to the polynucleotide sequence coated electrode results in substantially increased R _CT values when the 35-mer polynucleotide sequence (SEQ ID No: 4) is adsorbed on a flat Au electrode when compared to the polynucleotide sequence coated electrode without E2. The differential increase in R _CT upon E2 addition is substantially higher than the baseline response of a bare Au electrode at the lowest E2 concentration measured, 20 fM. Nyquist plots are provided in Figure 10.

The response of the AuNP electrode functionalised with substantially adsorbed 35-mer polynucleotide sequence (SEQ ID No: 4) is shown in Figure 8b. The diameter of semicircles in the Nyquist plot increased gradually as the concentration of E2 increased, reflecting an increased R _CT with increased E2 concentration. Analysing the normalised change in AR _C j (Figure 8c) shows that there is a reproducible logarithmic correlation (based on five individual experiments) between the change in R _CT and the concentration of E2 down to 1 fM with a wide dynamic range up to 20 μΜ. The increase in R _CT signal with increasing E2 concentration is a result of polynucleotide sequence dissociation from the surface and can be established using chronocoulometry to determine the amount of polynucleotide sequence adsorbed on the surface of the noble metal electrode before and after E2 exposure (see Figure 7). Exposure to 20 μΜ E2 results in 73 % reduction in surface coverage of the electrode by the 35-mer polynucleotide sequence (SEQ ID No: 4) when compared to the absence of target substrate (E2) present. Polynucleotide sequence dissociation from the surface of the Au electrode upon E2 recognition could result in subsequent loss of deposited AuNPs that are not strongly adsorbed to the electrode surface because of the stabilisation role played by the polynucleotide sequence as described herein. Provided that the polynucleotide sequence responds to its target by dissociating from the noble metal surface (Au), this additional effect may amplify the R _CT increase with E2 concentration for the AuNP electrode.

In addition, the specificity of the R _CT response by changing both the polynucleotide sequence probe and the analyte was also investigated. The E2 sensitivity is not retained when a randomised polynucleotide sequence (SEQ ID No: 7) is replaced with a random 35- mer polynucleotide sequence (SEQ ID No: 7) that comprises the same nucleotide bases as (SEQ ID No: 4) but lacks the E2 ligand binding domain due to the random organisation of the nucleotide base order. Chronocoulometry confirms that the randomised sequence (SEQ ID No: 7) remains adsorbed to the Au surface of the electrode after exposure to E2 (Figure 7). Therefore, confirming that the randomised polynucleotide sequence (SEQ ID No: 7) has no affinity for E2. The sensing steps were repeated using the E2 binding polynucleotide sequence (SEQ ID No: 7) with structurally similar target substrates that can either co-exist with E2 in biological samples such as progesterone or P4; or could trigger a similar endocrine response to E2, such as bisphenol-A or BPA. As shown in Figure 8c, the 35-mer polynucleotide sequence (SEQ ID No: 4) is specific for the detection of E2, with only a baseline response to the potentially interfering molecules examined even at concentrations as high as 1 μΜ, supporting that the 35-mer E2 (SEQ ID No: 4) sensor is specific for E2. Moreover, the response towards E2 is unaffected by mixing with BPA (circled data points in Figure 8c). These findings show that the R _CT signal arises from the 35-mer polynucleotide sequence (SEQ ID No: 4) binding with its target E2. The observed specificity is achieved as a result of the surface passivation established above, as well as the intrinsic specificity of this polynucleotide sequence

It can also be seen in Figure 8 that a simple and reliable electrochemical sensor can be constructed with substantial adherence of the desired polynucleotide sequence (along its length) onto a noble metal surface (such as gold, ruthenium, rhodium, palladium, platinum and silver; or noble metal nanoparticle surface such as gold, ruthenium, rhodium, palladium, platinum and silver) together with EIS and monitoring the EIS signals upon exposure of the noble metal or noble metal coated electrode to the test sample containing the target substrate. The level of detection of E2 using the electrochemical method of the present invention is over five orders of magnitude lower than known colorimetric E2 sensors which are based on a dispersion of AuNPs coated with polynucleotide sequence (SEQ ID No: 4). Furthermore, levels of detection of E2 using the method of the present invention is at least seven orders of magnitude lower than all previous colorimetric aptamer (with 43 nM level of detection) and small molecule aptasensors (200 pM level of detection) of the prior art [H. Li, L. Rothberg, Proc. Natl. Acad. Sci. U. S. A., 2004, 101 , 14036-9]. Additionally, the 1 fM levels of detection provided by the method of the present invention are approximately five orders of magnitude lower than recorded concentrations of E2 in human urine [Z. Lin, L. Chen, G. Zhang, Q. Liu, B. Qiu, Z. Cai, and G. Chen, Analyst, 2012, 137, 819-22. ; E. Taioli, A. Im, X. Xu, T. D. Veenstra, G. Ahrendt, and S. Garte, Reprod. Biol. Endocrinol., 2010, 8, 93.] The sensory method described herein is effective in the detection of target substrates in biological samples. In Figure 11 , the method of detecting target substrates described herein has been applied to rat urine. Figure 1 1 shows that a very similar response is observed for samples where rat urine has been spiked with E2 when compared to buffered water spiked with E2 in Figure 8. Similarly, an increase in R _CT with target concentration arose from specific interaction by the lack of response obtained from a sensor functionalised with the randomised 35-mer polynucleotide sequence (SEQ ID No: 4).

Shortened Sequences It is generally acknowledged that the closer identity that a polynucleotide sequence has to a ligand binding domain (LBD) of a target substrate, the higher the affinity and the better the interaction between the polynucleotide sequence and the target substrate. That is, by having non-binding nucleotide sequences present on the target specific polynucleotide sequence, the sequence is expected to have a high binding affinity because non-binding nucleotide sequences may interfere with the association of the target specific polynucleotide sequcene with the target substrate. This is because shortening of a target specific polynucleotide sequence facilitates the polynucleotide sequence's substantial dissociation from the electrode upon binding to the target. However, truncated target specific polynucleotide sequences consisting essentially of the ligand binding domain do not have sufficient enough nucleotide bases to interact with the electrode and with the target substrates to be detected. That is, when there are a reduced number of bases on a polynucleotide sequence, a target specific polynucleotide sequence consisting essentially of the LBD adsorbs to an electrode and this reduces the ability of the LBD to access the target substrate. Therefore, the ligand binding domain requires some additional nucleotides at the 5' end, the 3' end or both the 3' and 5' end of the sequence (flanking or primer sequences). These extra nucleotide bases (flanking or primer sequences) facilitate adsorption and/or association to the electrode and the target substrate without compromising or suppressing the signal transduction. Specifically, the presence of additional residues allows for the optimisation of the colorimetric assay. Essentially, the method provides for the production of a multiple number of polynucleotide sequences that contain one or more additional nucleotides bases at one or both ends, in addition to the LBD. Ideally, the polynucleotide sequences contain between one and ten additional residues, in addition to the LBD, at one end (3' or 5'), or both ends (3' and 5') of the LBD sequence. A shortened 35-mer polynucleotide sequence (SEQ ID No: 4), that lacks the primer nucleotides of the parent 75-mer E2 polynucleotide sequence (with a K _D of 25 nM) (SEQ ID No: 1 ), was found to produce a better colorimetric sensor as a result of the shortening.

To establish if the shortening of an polynucleotide sequences length is applicable to the present EIS sensory system, the AuNP EIS methodology described herein is applied to a 75- mer equivalent E2 polynucleotide sequence (SEQ ID No: 1 ) that comprises flanking sequences at the 3' and 5' ends. As shown in Figure 12, the 75-mer polynucleotide sequence (SEQ ID No: 1 ) exhibits an EIS response in detecting E2. This response is further confirmed via a control experiment with randomised sequence (SEQ ID No: 6) and interfering molecules. However, the EIS signal of the 75-mer system (SEQ ID No: 1 ) towards E2 addition is significantly suppressed by a factor of 10 (based on the slope Figure 12) in comparison with the 35-mer polynucleotide sequence (SEQ ID No: 4). The difference in signal sensitivity likely relates to the lesser degree of surface dissociation when SEQ ID No: 1 binds to E2 when compared to the dissociation of SEQ ID No: 4. Chronocoulometry measurements (Figure 5) show that only 57% of the 75-mer aptamer (SEQ ID No: 1 ) dissociates under saturation with E2, compared with 75% for the 35-mer (SEQ ID No: 4). Comparing the colorimetric and EIS responses for the polynucleotide sequences (SEQ ID Nos: 1 and 4) to the E2 target confirms that both methods depend upon the dissociation of target-bound polynucleotide sequence from Au surface.

The present methodology using EIS affords over five orders of magnitude of enhancement in the sensitivity of detecting the dissociation of polynucleotide sequence adsorbed on Au surfaces compared with the colorimetric method. The dependence of the detection method on polynucleotide sequence length for sensors based on dissociation from Au surfaces contrasts with knowledge for E2 polynucleotide sequences terminally tethered to conducting polymer electrodes where a stronger EIS response was observed for the 75-mer (SEQ ID No: 1 ) as a result of its higher charge content being redistributed near the conducting polymer surface.

Other Substrates

A BPA polynucleotide sequence (SEQ ID No: 5) was adsorbed to a AuNP electrode, following the same protocol used for the E2 detection as described above and in the Experimental detail. A decreased R _C T signal from - 1000 Ω to 224 Ω was produced when compared to a bare AuNP electrode. The sensor system was exposed to BPA and produced EIS signals similar to those observed when detecting E2, and detecting the target at concentrations as low as 1 fM (Figure 15a). The sensor did not exhibit a response when the BPA polynucleotide sequence (SEQ ID No: 5) was replaced by a random 75-mer polynucleotide sequence (SEQ ID No: 6), as shown in Figure 15a. Where the description has been made to integers having known equivalents thereof, those equivalents are herein incorporated as if individually set forth. Although the invention has been described in connection with specific preferred embodiments, it should be understood that the invention as claimed should not be unduly limited to such specific embodiments. It is appreciated that further modifications may be made to the invention as described herein without departing from the spirit and scope of the invention.

Examples Example 1

Phosphate buffered saline (PBS, pH 7.4, 137 mM NaCI, 2.7 mM KCI, 10 mM Na ₂ HP0 ₄ , 1.8 mM KH ₂ P0 ₄ ), chloroauricacid (HAuCI ₄ ), 6-mercaptohexanoic acid (MHA), 17 ?-estradiol (E2), progesterone (P4), and bisphenol-A (BPA) were purchased from Sigma-Aldrich. The thiolated and non-thiolated DNA sequences used in this study were synthesised and PAGE purified by Alpha DNA. The thiolated DNAs were treated with dithiothreitol (DTT) at the last step of the synthesis, by Alpha DNA, to reduce the S-S bond and DTT was removed by ethanol precipitation after which the DNAs were lyophilised and sealed under Argon. All DNAs were rehydrated with deionised water (Milli-Q, 18.2 MQcm) for subsequent dilutions. Sequences used in this study are given in Table 1. All aqueous solutions were prepared using deionised water (Milli-Q, 18.2 MQcm). Other chemicals in this study were of analytical grade and used as supplied unless otherwise stated. General Procedure for formation of AuNP based electrodes

A dense layer of AuNPs was electrochemically deposited on the surface of a glassy carbon electrode (GCE), Figure 1 b, or a gold electrode (where noted) [X. Lin, Y. Ni, and S. Kokot, Anal. Chim. Acta, 2013, 765, 54-62]. A 5mL three-electrode cell comprising a polished and cleaned glassy carbon working electrode (GCE) (eDAQ, 1.0 mm in diameter), or gold electrode (10 mm in diameter), Ag/AgCI (3 M NaCI, +0.197 V vs. SHE) reference electrode and Pt wire counter electrode (using Bio-Logic SP-300 instrument). The Au deposition was performed by immersing the electrodes in a 5 mL solution containing 1 mM HAuCI ₄ , 0.01 M Na ₂ S0 ₄ and 0.01 M H ₂ S0 ₄ at a constant potential of - 0.2 V (Ag/AgCI) for 30 s. The GCE or gold electrode was polished before deposition with 0.5 μνη alumina slurry, followed by electrochemical cleaning by subjecting to potential cycling between -1.0 and 1.0 V in 0.25 M H ₂ S0 ₄ , washing with acetone, ethanol and deionised water. All solutions were de-gassed for 15 min by purging with N ₂ gas.

General Procedure for polynucleotide sequence attachment to the electrode

For thiolated and non-thiolated DNAs, the surface of the electrode was rinsed thoroughly with 10 mM trisodium citrate buffer, pH 3. 0.1 nmole of the aptamer (16.7 μΜ, in 5 mM trisodium citrate buffer, pH 3) was incubated with the Au based surface of the electrode for 20 mins. This protocol was found to produce quantifiable monolayers of Au-S DNAs in 3 min on AuNPs, and was adopted here (a control experiment via exposing the aptamers on GCE resulted in no adsorption, Figure 4). Treatment with 6-mercaptohexanoic acid, SH- (CH ₂ ) ₅ COOH, (MHA) and 6-mercaptohexanol, SH-(CH ₂ ) ₅ CH ₂ OH, (MCH) was carried out under the same conditions as coupling the DNA but with a concentration of 10 pM and 1 mM, respectively. The chosen concentration of MHA was limited by its solubility in the coupling buffer. Bovine serum albumin (BSA) and lysozyme (LYS) were dissolved in deionised water (Milli-Q, 18.2 MQcm) at a concentration of 1.5 mM and exposed to the surface of AuNP electrodes for 20 min. Target sensing

Detection of E2 and interfering molecules (P4) progesterone, (BPA) bisphenol-A, was achieved by incubating the DNA-modified electrode in 30 μΐ binding washing buffer solution (BWB: 2 mM Tris-HCI at pH 7.5 containing 10 mM NaCI, 0.5 mM KCI, 0.2 mM MgCI ₂ , 0.1 mM CaCI ₂ and 5% ethanol) containing the desired concentration of the target for 15 min at room temperature. The electrode was washed three times before EIS or before subsequent measurements by using BWB buffer to remove any non-bound molecules of the targets. E2 detection in rat urine was the same as above except that E2 was spiked in rat urine after adjusting the content of ethanol to 5%, for solubility reasons. Rat urine was collected from sexually mature ship rats (Rattus rattus), then filtered with 0.22 μπ\ syringe-filters (control rat urine sample comprised blank rat urine containing 5% ethanol). EIS measurements

were recorded in a three-electrode cell containing 5 mL PBS solution with 5.0 mM of [Fe(CN) ₆ ] ^3~/4~ (1 : 1 , mol : mol) at an applied bias of 0.230 V (Ag/AgCI). EIS measurements were recorded on a Bio-Logic instrument (Bio-Logic SP-300). EIS measurements were carried on with a 10 mV sinusoidal amplitude and collected for harmonic frequencies between 100 mHz and 100 kHz at 12 steps per decade and analysed using Zfit (using the equivalent circuit shown in Figure 1c).

DNA surface densities

Aptamer surface densities on AuNPs/GCE were determined using the chronocoulometry method developed by Steel et al [A. B. Steel, T. M. Heme, M.J. Tarlov, Anal. Chem., 1998, 70, 4671 -4677]. Tethered thio polynucleotide sequence (SEQ ID No: 3) or substantially adsorbed polynucleotide sequence (SEQ ID No: 4 ) on AuNPs/GCE were immersed in a low ionic strength electrolyte, 10 mM tris-HCI buffer at a pH 7.4. The potential was stepped from 200 to - 500 mV versus (Ag/AgCI) for 500 ms (using BAS 100A electrochemical analyser), and the resulting charge flow was measured. The electrode was then immersed for 20 min in a solution of 150 μΜ Ru(NH ₃ ) ₆ ³⁺ (RuHex) in tris buffer, and the measurement repeated. By plotting the charge Q versus the square root of time (t ^{1 2} ), the excess of RuHex was determined which can be related to the surface density of the polynucleotide sequence. See Figure 5 and Figure 2.

Example 2: 17 ? -ESTRADIOL DETECTION FROM BLOOD SAMPLES

Reagents and materials

Phosphate buffer saline (PBS) pallets (137 mM NaCI, 2.7 mM KCI, 10 mM Na ₂ HP04, 1.8 mM KH ₂ P0 ₄ ), 17 ?-estradiol (E2), potassium ferricyanide and potassium ferrocyanide ([Fe(CN) ₆ ] ^3_/4" )were purchased from Sigma-Aldrich. Aptamer sequences (35 mer 5 ^* AAGG G ATG CCGTTTG G G CCCAAGTTCG G CATAGTG 3 ^' ; SEQ ID No: 4) used in this study were synthesised and PAGE purified by Alpha DNA™. 1.6 mm diameter gold disk (MF- 2014), standard Ag/AgCI (MF-2052) and platinum (Pt) spiral electrodes were purchased from BASI. Methods

Electrode Cleaning :

In the case of measurements in PBS media, gold (Au) electrodes were cleaned with a procedure of 1 ) electrochemical acid cleaning, 2) polishing with 0.5 μπι alumina slurry and ultra-sonicating in ethanol and deionized water (Milli-Q) for 5 minutes each. For electrochemical acid cleaning, cyclic voltammetry (CV) was carried out between -0.4 and 1.6 V of potential with a 50 mV/s scan rate in 0.5 M H ₂ S0 ₄ until stable oxidation and reduction signals were obtained. In these measurements, Au disk electrode was employed as working electrode (WE), Pt spiral electrode as counter electrode (CE) and standard Ag/AgCI as reference electrode (RE) in a three terminal electrochemical cell configuration. Upon the ultra-sonification electrodes were tested with electro chemical impedance spectroscopy (EIS) and 2 ^nd step was repeated if electrodes were not clean enough (Figure 16). In the case of measurements in blood media further cleaning was required due to the non-specific absorption of blood proteins. Thus following the 1 ^st and 2 ^nd step, -1.2 V potential was applied for 30 seconds to repel the adsorbed blood proteins (3 ^rd step). Then electrodes were polished and ultrasonicated in ethanol and de-ionised water subsequently (4 ^th step) and tested via EIS measurements. If the electrodes were not clean enough, 4 ^th step was repeated (Figure 16).

Aptamer attachment

Non-specific adsorption of DNA sequences was chosen as the surface functionalisation procedure. Prior to the attachment, gold disk electrode was rinsed thoroughly with 10 mM trisodium citrate buffer (pH: 3). Then 0.1 nmole of the aptamers (20 μΜ, in 5 mM trisodium citrate buffer, pH 3) was incubated to the Au surface for 20 mins. Thus, DNA sequences were physically adsorbed onto the electrode surface (Figure 17A). Efficiency of the surface functionalisation was investigated via electrochemical impedance spectroscopy (EIS) (Figure 17B).

Electrochemical measurements

All electrochemical measurements were performed in three terminals electrochemical cell where Au disk electrode was WE, Pt spiral wire CE and Ag/AgCI standard was RE. EIS measurements were performed in PBS (pH: 7.4) in the presence of 5mM ([Fe(CN) ₆ ^~/4~ ) each.

Results and Discussions

Target sensing measurements in PBS media

Upon surface functionalisation with aptamer sequences, electrodes were incubated with target solutions containing E2 with a concentration range of 1 fM- 100 nM in 50 μL· binding washing buffer solution (BWB: 2 mM Tris-HCI at pH 7.5 containing 10 mM NaCI, 0.5 mM KCI, 0.2 mM MgCI ₂ , 0.1 mM CaCI ₂ and 5% ethanol) containing the desired concentration of the target for 15 min at room temperature. Before EIS or subsequent measurements, electrodes were rinsed with BWB buffer to remove any non-bounded target molecules. Detection of the E2 targets, present in the buffer solutions, was investigated by Electrochemical Impedance Spectroscopy (EIS) in PBS (pH: 7.4) containing [Fe(CN) ₆ ] ^3"/4~ (5 mM each) (Figure 18A).

Normalised changes in the values were calculated by fitting the data in a simple Randless ^" electrical equivalent circuit of R ₁ +Q ₂ /(R2 ⁺ W ₂ ) where R1 is the solution resistance, Q2 is the constant phase element, R2 is the charge transfer resistance (R _CT ) and W2 is the Warburg diffusion constant. Then Δ/½-//½°. taken as the sensor response, was plotted versus target molecule concentration (Figure 18B). The prominent increase in R _CJ IR _C can be attributed to the fact that an increased negative charge density when the target-bound aptamer adopts a more compact conformation. Target sensing in the presence of human blood

Sensing performance of the DNA modified electrodes was also investigated in the presence of %1 and 5% human blood accompanying externally added E2 with a concentration range of 1fM to 1 pM (in BWB solution with % 5 ethanol) (Figure 19A&C and Table 2). Selectivity of the electrodes was also tested by incubating electrodes 1 % and 5% blood samples which were not containing any externally added E2 molecules (Figure 19B&D and Table 3).

Sensor response of the electrodes that were incubated %1 and 5% percent blood reveals a performance decrease when the blood amount increases. For instance, in the case of incubation with 1 pM E2;sensor response, AR _C J/RCT°, decreased from 3.45 (%1 blood) to 1.69 (%5 blood). This can be attributed to the amount of the adsorbed blood proteins blocking the aptamer sequences and decreasing the accessibility of the target molecules.

Table 2: Comparison of the sensor responses of electrodes that were incubated with externally added E2 (in BWB solution) in the presence of %1 and %5 blood.

Table 3: Comparison of the sensor responses of the electrodes that were incubated

Industrial Application

The methodology of the present invention may find application in the use of the analytical testing for the presence of target substrates in a sample.