GB2637334A - A nanoporous electrode for electrochemical detection of nucleic acids - Google Patents

Abstract

A nucleic acid sensing electrode comprising at least one electrically conductive layer 101 comprising a surface, a nano porous layer 102 made of a non-electrically conductive material which is chemically bonded to the surface of the electrically conductive layer, and at least one single stranded nucleic acid probe 103 attached to the nano porous layer. The electrically conductive layer can comprise an inorganic oxide e.g., fluorine doped tin oxide, indium tin oxide or indium zinc oxide. The nano porous layer can comprise an interconnected network of nanopores with diameters of 5 nm to 100 nm. The nano porous layer can be fabricated using a sol-gel material. A further aspect is a nucleic acid sensing device comprising the electrode and at least a second electrode configured to contact a test liquid. The second electrode can be a counter electrode. The device can have a third electrode which is a reference electrode. The device can comprise a measurement unit configured to record changes in the electrochemical signal. A further aspect is a method of detecting nucleic acids in a sample using the device comprising measuring a baseline electrochemical signal, introducing a liquid sample to be tested, and measuring changes in electrochemical signal.

Description

A Nanoporous Electrode for Electrochemical Detection of Nucleic Acids

Technical field:

The present invention relates to the field of electrochemical sensors and more particularly to a nucleic acid-sensing electrode and a device comprising the same for detecting specific nucleic acid sequences, and a method for using such a device in the detection of nucleic acids present within a medium, including but not limited to bodily fluids and liquid samples. One exemplary utility of this invention is the electrochemical identification of specific nucleic acid sequences, such as those indicative of pathogen presence.

Background art:

The detection of nucleic acids is of utmost importance in the identification of pathogens, including viruses and bacteria, and in the diagnosis of various diseases. For example, during the COVID-19 pandemic, the identification of nucleic acid-based biomarkers was found to be critical in disease management because facilitated the identification of infected individuals, thereby helping to slow the spread of the disease. Moreover, the significance of nucleic acid detection extends to the diagnosis of non-infectious diseases, for example, in the utilization of liquid biopsies to detect mutations in circulating cell-free DNA or in traditional biopsies to detect the overexpression of oncogenes. Beyond diagnostics, nucleic acid detection is also important in various fields where bacteria and viruses are routinely measured, including water treatment evaluation, environmental and agricultural monitoring. While PCR stands as the benchmark technique for nucleic acid detection, its inherent limitations, including extended result times, the necessity for skilled personnel to perform the procedure, elevated material costs, and lack of portability of the necessary equipment (preventing point-of-care use), impede its universal adoption. Thus, there is a need for a more straightforward method for nucleic acid detection that offers swift, portable, and on-site nucleic acid detection.

Electrochemical nucleic acid detection emerged as a viable PCR alternative, offering quantitative outputs with rapid, cost-effective, and minimal manual sample preparation. In this context, the physical assembly of a membrane with nanometric pores (nanopores) directly onto the working electrode allowed the development of a nucleic acid detection methodology based on the nanopore blockage, so called nanochannel blockage (nanoblockage). 1-3 This detection mechanism measures the electrochemical signal changes produced upon the hybridization of a target nucleic acid with its complementary strand within the pore walls. The electrochemical output is typically expressed as a decrease in oxidation current and generally correlates with the analyte concentration, allowing quantitative detection.

Current nanoblockage-based sensors predominantly employ a multilayer stack sensor fabrication approach, which requires the mechanical attachment of a nanoporous layer onto an electrically conductive substrate, making them challenging to produce. First, pores with nanometric dimensions are induced in a solid non-conductive precursor membrane by elimination of segments of the precursor, which is generally achieve by chemical or electrochemical etching. Second, the obtained nanoporous membrane is mechanically mounted onto an electrically conductive electrode forming a stack of layers and is fixed in place by using mechanical fastenings. Third, a biological capture probe is chemically attached to the nanoporous layer. Historically, monocrystalline silicon wafers and high-purity aluminium materials have been utilized to generate the nanoporous membrane via electrochemical anodization. This top-down fabrication technique yields membranes of micrometric thickness. However, challenges derived from the sensor assembly and material robustness have restricted the effective deployment of these nanoblockage-based biosensors beyond research laboratories.' For instance, the integration of the nanoporous membrane onto a working electrode by mechanical means is time-consuming, and the brittle membrane is prone to fracture upon manipulation. Moreover, the fabrication of a nanoporous membrane from silicon wafers requires the use of hydrofluoric acid (HF), a highly toxic chemical, and the fabrication of nanoporous membranes from aluminium risks the generation of an explosive intermediate by-product. Therefore, there is a pressing need for a new way to integrate a nanoporous layer directly onto a working electrode for nanoblockage-based electrochemical sensors that can be used for nucleic acid detection.

References: (1) Vlassiouk, I.; Takmakov, P.; Smirnov, S. Sensing DNA Hybridization via Ionic Conductance through a Nanoporous Electrode. Langmuir 2005, 21 (11), 4776-4778. https://doi.org/10.1021/Ia0471644.

(2) De La Escosura-Muniz, A.; Merkoci, A. Label-Free Voltammetric Immunosensor Using a Nanoporous Membrane Based Platform. Electrochem. commun. 2010, 12 (6), 859-863. https://doi.org/10.1016/j.elecom.2010.04.007.

(3) Huo, Q.; Wang, K.; Xu, J.-J.; Li, J.; Li, S.-J.; Wang, C.; Chen, H.-Y.; Xia, X.-H. A Nanochannel Array-Based Electrochemical Device for Quantitative Label-Free DNA Analysis. ACS Nano 2010, 4 (11), 6417-6424. https://doi.org/10.1021/nn101050r.

(4) de la Escosura-Muniz, A.; Merkoci, A. Nanochannels for Electrical Biosensing. TrAC -Trends Anal. Chem. 2016, 79,134-150. https://doi.org/10.1016/j.trac.2015.12.003.

Summary of Invention:

The present invention provides a nucleic acid-sensing electrode capable of detecting specific nucleic acid sequences in a liquid sample. The electrode includes an electrically conductive layer, a nanoporous layer, and single-stranded nucleic acid probes. The invention further encompasses a device comprising the nucleic acid-sensing electrode and additional components facilitating the detection process, and a method of utilizing this device for nucleic acid detection.

In one aspect, the invention features a nucleic acid-sensing electrode comprising an electrically conductive layer selected from a group consisting of fluorine-doped tin oxide, indium tin oxide or indium zinc oxide, a nanoporous layer made of a non-electrically conductive material bonded to the electrically conductive layer, and nucleic acid molecules attached to the nanoporous layer. The nanoporous layer is chemically bonded to the electrically conductive layer as a result of the fabrication process, such that the attachment of the nanoporous layer to the electrically conductive layer is not a result of mechanical fastening. The functional nanoporous layer can be created using different materials, including silicates and metal oxides, and specific fabrication processes, including direct fabrication onto the electrically conductive layer by means of block copolymer and sol-gel nanoparticles co-assembly, and the single-stranded nucleic acid probe can be attached to the nanoporous layer using particular chemical methods, including covalent bonding, electrostatic attachment, and physical adsorption.

In another aspect, the invention features a nucleic acid-sensing device comprising the nucleic acid-sensing electrode and at least one additional electrode configured to contact the test liquid. The device is designed for enhanced sensitivity and specificity in detecting nucleic acid sequences complementary to the nucleic acid probe in the non-electrically conductive nanoporous layer.

In yet another aspect, the invention provides a method for detecting nucleic acids using the aforementioned device. The method involves measuring a baseline electrochemical signal in a reference liquid, introducing a liquid sample, and observing changes in the electrochemical signal with respect to the baseline measurement, indicative of the presence of target nucleic acids.

The present invention provides a novel and effective solution to the limitations and challenges inherent in existing electrochemical nucleic acid detection methodologies based on the nanochannel-blockage, making it valuable for various applications in diagnostics, research, and environmental monitoring.

Brief description of the drawings:

The invention will now be described solely by way of example and with reference to the accompanying drawings in which: Figure 1: This figure illustrates a schematic cross-section of the nucleic acid-sensing electrode. It shows the electrically conductive layer, the nanoporous, and the single-stranded nucleic acid probes. These probes are covalently attached to the nanoporous layer. The nanoporous layer is depicted as being bonded to the electrically conductive layer. The schematic also includes a detailed representation of the interconnections between the nanopores, which are represented with a hexagonal close packed configuration. This figure highlights the unique structure and functionality of this invention, showcasing the integral structure of the electrode.

Figure 2: This figure presents a schematic detailing the fabrication process of the nanoporous layer directly onto the electrically conductive layer. The process involves the co-assembly of block copolymer micelles with sol-gel nanoparticles, followed by calcination. This method results in the nanoporous layer being chemically bonded to the surface of the electrically conductive layer. Therefore, the attachment of the nanoporous layer to the electrically conductive layer is not due to mechanical fastening, but rather a result of this specialized chemical process. The figure highlights the intricate steps and the significance of each component in achieving the desired chemical bonding.

Figure 3: This figure illustrates the optical characterization techniques employed: spectroscopic ellipsometry and ellipsometric porosimetry, which are used to determine the film thickness, refractive index, and pore size distributions of a nanoporous aluminosilicate thin film. Part (a) involves measuring and fitting the ellipsometric angles 4; and A within the visible spectral range, enabling the precise calculation of the film's thickness and refractive index. Part (b) displays the ellipsometric porosimetry adsorption-desorption isotherms, obtained using toluene as an adsorptive molecule. Part (c) presents the pore size distribution derived from the isotherms shown in part (b). This comprehensive set of techniques and results collectively provide a detailed characterization of the nanoporous thin film properties.

Figure 4: This figure showcases two critical aspects of nanoporous thin film analysis. Part (a) presents a grazing-incidence small-angle scattering (GISAXS) pattern of the nanoporous thin film, captured after the calcination process and prior to its functionalization with nucleic acids. This characterization technique is crucial for understanding the film structure and properties post-calcination. Part (b) displays the in-plane line-cuts integration (qy) of the GISAXS pattem, revealing various peaks. These peaks are indicative of the ordered porous structure within the film. The presence and nature of these peaks provide valuable insights into the film's porous arrangement and uniformity.

Figure 5: This figure presents detailed images of the nanoporous layer. Part (a) shows a scanning electron micrograph (SEM) of the top view of the nanoporous layer, with a 2D spatial distribution function depicted in the inset to illustrate pore ordering. The scale bar representing 1 pm is included in part (a). Part (b) features a high magnification image obtained with a focused ion beam (FIB), showcasing nanopores and pore interconnections, alongside an inset schematic of a perfect hexagonal close-packed pore configuration. The scale bar for this FIB image is 100 nm. In part (c), the SEM image illustrates a nanoporous layer bonded to an inorganic substrate by block copolymer co-assembly, displaying ellipsoidal-shaped nanopores with dark contrast and surrounding nanopore walls in light contrast. The scale bar in part (c) is 500 nm. These images collectively highlight the nanoporous architecture of the thin film used for this invention.

Figure 6: This figure details a chemical method for covalently attaching single-stranded nucleic acids to a nanoporous layer. Initially, the method may involve the modification of the nanoporous layer using oxygen plasma. Following this, the surface of the layer can be sequentially modified: first with (3-aminopropyl)triethoxysilane (APTES) and subsequently with glutaraldehyde. This process prepares the surface for attachment by introducing aldehyde groups via glutaraldehyde. A single-stranded nucleic acid, modified with an amino group (ssDNA-NH2), can then be covalently attached to these aldehyde groups. Finally, to ensure stability and specificity of the attachment, unreacted aldehyde groups may be blocked using the chemical ethanolamine. This step-by-step illustration outlines the precise chemical interactions and modifications necessary for the successful attachment of nucleic acids to the nanoporous layer.

Figure 7: This figure demonstrates the real-time monitoring of surface modification and functionalization with single-stranded DNA (ssDNA) using quartz crystal microbalance with dissipation (QCM-D) sensors. The figure exemplifies this process by showcasing the frequency and dissipation responses (at the 5th harmonic) of two different QCM-D sensor setups: a flat QCM-D sensor and a sensor coated with a nanoporous film. These responses are compared during surface modification in two different environments -in toluene (displayed in part (a) and part (c)) and in PBS buffer (presented in part (b) and part (d)). This comparison illustrates the effective chemical modification of the nanoporous surface as demonstrated by changes in frequency and dissipation of the sensor.

Figure 8: This figure illustrates the process of DNA hybridization between ssDNA molecules. Part (a) shows the mass changes observed in both flat and nanoporous-coated QCM-D sensors functionalized with ssDNA capture probes. These changes occur upon exposure to complementary ssDNA. In part (b), QCM-D sensors functionalized with the ssDNA capture probe are exposed to the target ssDNA modified with the fluorescent molecule 6FAM. The sensors are imaged with a 30-second exposure time using a 535 nm filter. Part (c) presents a comparative analysis of the fluorescence intensity between the flat and nanoporous-coated QCM-D sensors, as shown in part (b), which have been hybridized with ssDNA-6FAM. This comparison highlights the higher number of ssDNA probes that can be attached to a nanoporous sensor in comparison with a flat surface.

Figure 9: This figure presents a schematic of the electrochemical detection of nucleic acids using the current invention. In part (a), the working electrode of a three-electrode electrochemical sensor is coated with a nanoporous film and an ssDNA (single-stranded DNA) capture probe. In part (b), the initial electrical signal, specifically the impedance, is measured before starting the detection assay. Part (c) illustrates the nucleic acid detection process: hybridization with the target sequence (which is complementary) leads to an increase in the impedance of the system. This increase is indicative of successful nucleic acid detection, demonstrating the functionality of the sensor in identifying specific nucleic acid sequences.

Figure 10: This figure presents the electrochemical measurements that indicate ssDNA detection using the nucleic acid-sensing electrode described in the current invention. Part (a) displays a Nyquist plot for a DNA sensor functionalized with a capture probe that is complementary to the target DNA. In part (b), a Nyquist plot is shown for a DNA sensor functionalized with a negative control probe, which is non-complementary to the target DNA. Part (c) illustrates the concentration-response curves for nanoporous sensors functionalized with both complementary and non-complementary capture probes. The error bars represent the standard deviation based on measurements from at least three sensors. These plots collectively demonstrate the sensors ability to differentiate between complementary and non-complementary DNA sequences to the capture probe attached to the sensor, showcasing its specificity and sensitivity in DNA detection.

Detailed Description of Various Embodiments:

The present invention can be embodied in various forms without departing from the spirit or essential attributes thereof. Accordingly, the following descriptions of certain embodiments, including their specific components and functionalities, are provided for illustrative purposes within the scope of the invention.

Nucleic acid-sensing electrode In one embodiment, the nucleic acid sensing electrode comprises an electrically conductive layer 101, containing a chemically bonded nanoporous layer 102, and ssDNA captures probes 103 attached to the nanoporous layer 102. The nanoporous layer 102 contains nanopores 104 that are interconnected through pore interconnections 105, as schematized in Figure 1. These interconnections 105 create open pathways between the external surface of the nanoporous thin film 102 and the electrically conductive layer 101 beneath this nanoporous layer.

Fabrication of a nanoporous layer onto FTO-coated glass In another embodiment, a nanoporous layer 102 can be fabricated directly onto an electrically conductive layer 101 by the co-assembly of block copolymer micelles with sol-gel nanoparticles 107 and subsequent calcination, as schematically shown in Figure 2. A working electrode 110, made of fluorine tin oxide (FTO) coated glass 101, was covered with a nanoporous thin film 102. First, the amphiphilic block copolymer poly(isoprene)-block-poly(ethylene oxide) (Pl-b-PEO) was mixed with aluminosilicate nanoparticles in an azeotropic mixture of toluene and 1-butanol. The repulsive interaction of the hydrophobic PI block with the solvent and PEO block induces the self-assembly of the block copolymer in solution to form spherical micelles with nanometric dimensions 106. On the other hand, the nanoparticles coordinate with the PEO block by hydrogen bonding without disturbing the geometric configuration of the micelles. The hybrid mixture was subsequently spin-coated onto the working electrode 110 (FTO-coated glass 101) to generate a continuous and homogeneous hybrid thin film 107. This mixture contained 50w% of block copolymer (BCP) aiming to fabricate pores 104 in the upper range of the mesoscale with large interconnections 105 to facilitate the molecular transport through the nanoporous film 102 i.e. to allow the diffusion of nucleic acids 103 used for the functionalisation and sensing within the porous structure. The modified electrode was first calcined in argon to carbonise the BCP concurrently with the condensation of the aluminosilicate. Finally, the electrode 110 was calcined in air to remove the carbonised BCP, revealing the nanopores 104 of approximately 50 nm in diameter.

The pore dimensions and porosity of the thin films can be obtained by spectroscopic ellipsometry and ellipsometric porosimetry. For instance, a nanoporous aluminosilicate thin film 102 was fabricated onto a reflective substrate made of a polished silicon wafer and spectroscopic ellipsometry and ellipsometric porosimetry (EP) were measured. A Cauchy dispersion law was fitted to the measured ellipsometric angles tit and A (Figure 3a), indicating a film thickness of approximately 150 nm and a refractive index of 1.12. Toluene was used as an adsorptive molecule for EP as its high vapour pressure allows for the measurement of pores in the upper range of the mesoscale (near -50 nm). A porosity of 65.3% was obtained from the maximum volume adsorbed ratio (dashed reference line in Figure 3b). The type IVb shape of the isotherm suggests that the nanopores 104 are interconnected with pore necks 105 of smaller diameter. The nanopore 104 size and nanopore interconnection 105 size distribution were calculated by applying the modified Kelvin equation to the EP isotherm (Figure 3c). A Cauchy fitting was used to adjust the size distribution to a normal distribution (mean size ± standard deviation), resulting in a nanopore 104 size of 44±12 nm and nanopore interconnection 105 size of 23±10.5 nm, from the adsorption and desorption isotherms, respectively. A surface area of 140 m2cm3 was calculated using the BET method and a toluene molecule cross-section of 0.343 nm2.

The long-range nanopore arrangement can be determined by using GISAXS scattering patterns and the corresponding in-plane line-cuts integration of a nanoporous film 102 after calcination (Error! Reference source not found. 4). The numerous Bragg peaks observed in the in-plane integration of the GISAXS pattern line cuts in Figure 4b suggest a long-range porous order after calcination. The first Bragg peak (q1=0.102 nm-1) and the higher order Bragg peaks q,, qz, q3, and qa, in the positions q/q*= ^I4, -si7, and -0, respectively, are coincident with the formation of a hexagonal close-packed (HCP) arrangement of the nanopores 104. A pore centre-to-centre distance, D0. c, of -71 nm was calculated from the first Bragg peak. ;Figure 5a shows a representative SEM micrograph of a nanoporous thin film 102 surface used in a nucleic acid-sensing electrode. The concentric hexagonal rings of the 2D spatial distribution function (inset in Figure 5a) confirmed the high degree of order of the nanomaterial. The high magnification FIB micrograph of the nanoporous layer surface 102 allowed to visualise the nanopores 104 and the nanopore interconnections 105 (Figure 5b). In a perfect HCP configuration, each nanopore in the surface should be connected to three pores in the layer beneath, as schematised in the inset of Figure 5b. The FIB image confirmed that the top layer of the nanoporous film 102 link to three outof-plane nanopores 104 via nanopore necks or nanopore interconnections 105. The cross-section of the film shown in figure 5c confirmed the thickness of the nanoporous layer 102 was about 150 nm. ;Functionalization of the nanoporous layer with single stranded nucleic acid capture probes In yet another embodiment of the invention, an aluminosilicate nanoporous layer 102 fabricated onto FTO-coated glass 101 was chemically modified with single-stranded DNA capture probes 103. Using nanopores for electrochemical biosensing requires attaching a biorecognition element to the nanopore walls with the ability to bind the target analyte. This nucleic acid sensor uses the intrinsic lock-and-key property of DNA to hybridise complementary strands. This characteristic provides high sensitivity, specificity, and reproducibility to electrochemical DNA sensors. The nanoporous layer 102 was functionalised with an ssDNA 103 that served as a capture probe. ;The immobilisation protocol followed to attach a single-stranded DNA 103 modified with an amino group (ssDNA-NH2) via covalent bonds to the nanopore layer 102 is schematised in Figure 6. The nanoporous material 102 was first oxygen plasma treated to introduce hydroxyl groups on the aluminosilicate oxide surface (Figure 6a), enabling surface modification with the amino silane (3-aminopropyl)triethoxysilane (APTES) dissolved in toluene as a result of an SN2 reaction with the hydroxyls of the surface (Figure 6b). The exposed NH2 groups of the APTES molecules were subsequently modified with glutaraldehyde (GA) (Figure 6c) in PBS buffer, a bis-aldehyde homobifunctional crosslinker. Next, the exposed aldehyde groups of GA reacted with the NH2 modification of the ssDNA capture probes 103 (Figure 6d), thereby fixing the capture probe 103 to the nanopore 104 walls in the nanoporous layer 102. Finally, unreacted aldehyde groups were blocked with ethanolamine (Figure 6e). ;A quartz crystal microbalance with dissipation monitoring was used to show in real-time the kinetics and efficiency of one surface modification protocol, as shown in Figure 7. To this end, in-situ changes in frequency and dissipation of flat silica-coated and nanoporous-coated 102 QCM-D sensors were measured when exposed to the binding molecules. The flat sensor served as a reference to understand the effect of the nanoporous structure in the frequency and dissipation response, i.e. to distinguish whether changes observed are related to surface-bound or pore-confined molecules. ;Figures 7a and 7b show the frequency changes of the fifth harmonic of QCM-D sensors during aminosilane modification performed in toluene and glutaraldehyde (GA) followed by ssDNA-NH2 103 modification performed in PBS buffer (pH 7.3), respectively. In QCM-D measurements, negative changes in frequency indicate mass adsorption and vice versa. Therefore, the negative changes in frequency observed in both sensors upon exposure to APTES, GA, and ssDNA 103, suggest that all types of molecules were effectively adsorbed to the sensor surface. Rinsing the sensors with the respective liquid matrix facilitates the removal of non-covalently bound molecules. Thus, the permanent net negative changes in frequency observed after rinsing indicate the effective immobilisation of APTES, GA, and nucleic acid capture probes 103. ;Figures 7c and 7d show the respective dissipation changes measured in ppm (10-6). Energy dissipation changes (> 1 ppm) were measured in both sensors upon exposure to APTES and glutaraldehyde, suggesting the formation of a soft layer on the sensor surface. The APTES and GA molecules that were non-covalently adsorbed to the sensors were removed during rinsing, resulting in negligible residual energy dissipation (< 1 ppm). The absent dissipation and the negative frequency changes measured in the same period suggest that a rigid monolayer of APTES and GA remained attached to the flat sensor. Finally, flat sensor modification with the ssDNA 103 showed a constant energy dissipation (-1 ppm) before and after rinsing, displaying the soft nature of the DNA layer. DNA 103 functionalised in the nanopores 104 dissipated lower energy (< 1 ppm) than in the flat surface. This may be related to the strict confinement of DNA probes in the nanopores. ;Timescale necessary for hybridization between complementary ssDNA in the nanoporous electrode In another embodiment of the invention, the timescale necessary for nucleic acid detection assays was determined by using QCM-D sensors functionalised with complementary and non-complementary probes. Figure 8a shows the mass changes related to the ssDNA hybridisation calculated from the frequency and dissipation changes of the QCM-D sensor when exposed to 1 pM ssDNA 103 target sequence in 0.1 M PBS buffer. ;The flat QCM-D sensor functionalised with the capture probe 103 (complementary ssDNA) showed that hybridisation with the target ssDNA occurred upon contact with the sensor. On the contrary, the lack of mass changes measured by the control sensor, i.e. the flat sensor functionalised with the non-complementary ssDNA 103, suggests that no hybridisation occurred with the target ssDNA. The target ssDNA modified with the fluorescent molecule 6-carboxyfluorescein (6FAM) confirmed the hybridisation between complementary nucleic acids (Figure 8b and Figure 8c). ;Hybridization in the nanoporous film 102 with the complementary ssDNA showed rapid hybridisation kinetics in the first minutes that slowed to a plateau, resembling an exponential association model. Thus, a time constant r=5.36±0.63 min was calculated by fitting an exponential association model to the hybridisation kinetics measured in three nanoporous sensors. ;Accordingly, the hybridised mass (Hb) at a given time (t) can be calculated using the time constant (r) by using the following equation 1: Hb(t) = Hb,11",(1-e9, where Hbmax denotes the maximum hybridised mass at equilibrium for a given concentration of target ssDNA. Figure 8a (red dashed line) shows the exponential fit of the hybridisation mass in a nanoporous QCM-D sensor using equation presented. ;Setting the hybridisation time t = T sets equation 1 to Hb(t) = Hbmax(1-e1) = 0.63Hbmax, which means that after 5.36 min, 63% of the capture probes hybridised to the target ssDNA 103. ;Similarly, setting t = [2r; 3r; 4r; 5r] for hybridisation time sets Hb(t) = [0.86; 0.95; 0.98; 0.99]Hbmax, which means after a time t = [11; 16; 21; 27] min, 86%, 95%, 98%, and 99% of the capture probes were hybridised, respectively. ;The hybridisation time for nucleic acid detection assays was set to a minimum of 3r (-16 min), as this is considered a trade-off between time and hybridisation near equilibrium (95% Hbmax). ;Electrochemical detection of single-stranded DNA In another embodiment of the invention, the nucleic acid-sensing electrode was employed to detect single-stranded DNA. In particular, to detect ssDNA sequences derived from the 16S ribosomal RNA of the bacteria E. coli, i.e. the target sequence 112. The target sequence 112 used was CCA AGT CGA CAT CGT TTA CGG CGT GGA C. In addition, different nanoporous films 102 were functionalised with two nucleic acid sequences: a capture probe 103 entirely complementary to the target sequence 112 (NHZC6 -GTC CAC GCC GTA AAC GAT GTC GAC TTG G) and a non-complementary negative control (NH2 C6 -CAC AAA TTC GGT TCT ACA GGG TA). The electrochemical detection of nucleic acids was done using the nanoporous material platform fabricated by block copolymer co-assembly. Figure 9 illustrates the nucleic acid-sensing concept based on the nanopore blockage. In this electrochemical nucleic acid sensor, the nanoporous transducer 102 modified with the ssDNA capture probe 103 was used as the working electrode 110 of a three-electrode setup 111 (Figure 9a), along with a reference electrode 109 and a counter electrode 108. The detection assay included measuring the electrochemical impedance of the three-electrode setup 111 before testing (Figure 9b). This measurement allowed to obtain a baseline, i.e. the electrical resistance to the electronic flow by the nanoporous film 102 before detection. The liquid sample with the analyte was then dispensed in the working electrode and left to react for 20 min. The electrochemical impedance was measured again. In a positive result, the capture probe 103 bound to the nanopore 104 walls hybridised with the target sequence, thereby blocking the nanopores 104. As a result, the ionic transport through the nanoporous structure 102 was reduced, thus increasing the system's electrical resistance, which is visually characterised by an increase in the diameter of the typical semicircle of faradic Nyquist plots, as illustrated in Figure 9c. In a negative result, no nanopores were blocked, preserving the initial configuration of the sensor with no changes in the electrical signal. ;In another embodiment, electrochemical measurements were performed using ferricyanide/ferrocyanide ([Fe(CN)6]314-) as the redox mediator 113, the nanoporous transducer as the working electrode 110, a platinum wire as the counter electrode 108 and a silver/silver chloride reference electrode 109. The three-electrode setup 111 was connected to a potentiostat to record the electrochemical signals. Electrochemical impedance spectroscopy was measured, aiming to track changes in the electrical resistance of the system due to DNA hybridisation. However, experimental impedance values are affected by coupling effects between ohmic resistance, capacitance, constant-phase elements and Warburg impedance. Therefore, an equivalent circuit of the nanoporous biosensor can be used to approximate the experimental data with the different elements that can contribute to the measured impedance, thus allowing to determine changes related to electrical resistance specifically. Figures 10a and 10b show typical Nyquist plots of sensors functionalised with the capture probe 103 (complementary ssDNA) and the negative control (non-complementary ssDNA) upon consecutive incubation from 1 picomolar to 1 nanomolar of the target ssDNA 112, respectively. After incubation with the target ssDNA 112, the sensor functionalised with the capture probe 103 showed a clear increase in the impedance from the baseline (denoted as 0 pM). Incubation with higher target ssDNA 112 concentrations produced more significant changes in impedance. On the contrary, the sensor functionalised with the negative control (i.e. a non-complementary ssDNA attached to the sensor) displayed negligible changes in impedance upon incubation with the same target ssDNA, proving that the specific hybridisation between the capture probes 103 functionalised in the nanopore walls with the target DNA 112 produced the changes in impedance. ;To compare the response between different sensors upon incubation with the target DNA 112, the charge transfer resistance after incubation (Rd) was normalised by the initial charge transfer resistance (Re?) using the following equation 2: Rctl-R," 4R et 1100112= 0 * Ra Figure 10c shows the ARdnorm for increasing concentrations of the target ssDNA 112 from 1 picomolar to 1 nanomolar on a logarithmic scale. ARanorm changes measured in complementary sensors (n = 4) increased linearly with the concentration of the target ssDNA 112. Equation 3 shows the linear fit (R2 = 0.985) in logarithmic scale that allows reliable quantification of the target DNA 112 present in a liquid sample with a sensitivity of 0.08 Aar (slope of the calibration curve): AR ct, noun 0.18655 + 0.08021 Jog (x), where x is the concentration of the target ssDNA 112.

The negligible changes in andodm measured in the non-complementary sensor upon incubation with the target ssDNA 112 indicate the high selectivity of the nucleic acid detection platform, provided by the inherent characteristic of complementary nucleic acids to hybridise.

The limit of detection of a sensing platform is defined as the ability to differentiate a positive result from the noise of a blank measurement. The LOD was calculated from the values of the negative control by using equation 4: LOD=yeammi+kSb, where v control is the average value from 1 pM to 1 nM of the normalised La( of the negative control, Sr, is the standard deviation of the negative control measurements, and k is a numerical factor that is chosen according to the confidence level desired, with a recommended value of 3 for the LOD. A LOD of 0.0614 Aar was estimated for the DNA sensor, meaning that changes in the normalised Aar smaller than 0.0614 cannot be differentiated from noise with a statistical confidence of 99.7%. The concentration at the LOD (CLOD) was calculated by equalizing the LOD obtained with Equation 3 to Equation 2, resulting in a CLOD of 30 fM. The limit of quantification (LOQ) of a detection platform is defined as the minimum amount of the target analyte that can be quantified with acceptable precision. LOQ can be calculated similarly to the LOD by using k=10, which resulted in an LOQ of 0.1635 Aar. A corresponding concentration at the LOQ (CLOD) of 500 fM was calculated by equalizing the LOQ obtained with equation 3. This provides evidence that this nucleic acid-sensing platform can confidently quantify target DNA 112 concentrations as small as 500 fM (0.5 picomolar). The ability to detect electrochemically a target ssDNA that is complementary to the ssDNA attached to the nanoporous layer, indicate that this invention can be used to detect any target ssDNA by attaching the complementary ssDNA to the sensor and measuring changes in the impedance, as described in this embodiment.

In order to further elucidate the invention, reference is made to the accompanying Examples. Example 1: Fabrication of a nanoporous layer directly onto an electrically conductive layer by block copolymer co-assembly with sol-gel aluminosilicate nanoparticles Two stock solutions with a fixed concentration of aluminosilicate and block copolymer (BCP) were formulated. First, the aluminosilicate stock solution was prepared by mixing a silica precursor (2.8 g of GLYMO), an aluminium precursor (0.32 g of aluminium tri-sec-butoxide), and 20 mg of KCI in an iced bath. After 15 min under constant stirring, 135 pl of 10 mM HCI was added dropwise to the blend. The mixture was stirred for another 15 min at room temperature. Next, dropwise, 850 pl of 10 mM HCI was added and stirred for 20 min to complete the hydrolysis. Finally, 2.135 ml of the azeotrope toluene/1-butanol (72.84/27.16 wt%) was added to the solution to get a concentration of 1 g m1-1 of aluminosilicate. The mixture was then filtered with a 0.2 pm syringe filter and kept refrigerated at 5 °C for use. Second, a BCP stock solution was prepared by dissolving 40 mg m1-1 of the block copolymer polyisoprene block polyethylene oxide (Pl-b-PEO) in the azeotrope toluene/1-butanol.

A so-called hybrid mixture of BCP was made by mixing 60 pl of the aluminosilicate stock solution with 750 pl of the BCP stock solution in a glass vial. The mixture was left in a shaker for 30 min prior to use.

Hybrid thin films were prepared by spin-coating at 2,000 rpm during 20 s 40 pl of the hybrid solution onto electrically conductive substrates made of fluorine-doped tin oxide (FTO) coated glass. Hybrid thin films were subsequently reactive ion etched in CHF3 during 2 min using a 15/50 mixture of CHF3/Ar 15/50 at 215 W and 40 mbar in a plasma etcher to remove the upper layer of aluminosilicate aiming to obtain a fully open superficial porous structure. Samples were then calcined in argon at 450 °C during 30 minutes with a heating ramp of 5 °C min-1 in a tubular furnace, and later air calcined at 450 °C during 30 minutes with a heating ramp of 5 °C min-1. All films were left to cool inside the furnace. The high temperature of calcination eliminates the block copolymer micelles, generating the nanopores in an inorganic matrix of condensed aluminosilicate.

Example 2:

Immobilisation of ssDNA capture probes on the nanoporous layer: Immobilisation of amino-modified ssDNA capture probes (ssDNA-N H2) in a nanoporous layer, as the one described in example 1, was achieved following a five step functionalisation procedure. First, the nanoporous layer was oxygen plasma-treated during 15 s at 100 W and 0.33 mbar to introduce OH groups in the surface. Second, the nanoporous surface was aminosilanised by dip coating for 20 min at room temperature in a solution of 5 v/v% of APTES in anhydrous toluene under an argon atmosphere. The functionalised nanoporous layer was then sonicated two times for 5 min in toluene and then one time in ethanol to remove unreacted material from the surface. Third, the nanoporous layer was dip-coated in a solution of 10 v/v% of glutaraldehyde in 0.1 M PBS buffer for 30 min at room temperature and ambient conditions, to attach the homobifunctional crosslinker to the amine groups. The nanoporous layer was then rinsed and sonicated two times for 5 min in PBS to remove unreacted glutaraldehyde molecules. Fourth, the modified nanoporous layer was incubated overnight at 4 °C in a 1 pM ssDNA-NH2 in 0.1 M PBS buffer. The nanoporous surface was then rinsed three times with a 0.1 M PBS solution. Finally, the remaining aldehyde groups were blocked with 0.1 M ethanolamine in 0.1 M PBS buffer for 30 min. Please note that the PBS buffer used for immobilisation and sensing was prepared using nuclease-free water.

Example 3:

Electrochemical detection of single-stranded DNA Electrochemical measurements for the detection of ssDNA were performed in a three-electrode configuration inside a cell made of polytetrafluoroethylene (PTFE) using a silver/silver chloride (Ag/AgCI) reference electrode of 4 mm in diameter, a platinum (Pt) wire counter electrode of 0.4 mm in diameter, and a working electrode of 0.5 cm in diameter. The working electrode was an FTOcoated glass containing the nanoporous layer that was previously functionalised with the single-stranded DNA capture probes, as described in the example 2. Electrochemical impedance spectroscopy (EIS) measurements starting at a frequency of 100 kHz to 0.1 Hz with an amplitude of mV and applying 0 V versus the open circuit potential were carried out using a potentiostat. Electrochemical measurements were analysed and fitted with using a software.

First, the working electrodes functionalised with ssDNA probes, complementary and non-complementary to a target DNA, were mounted in a PTFE cell. The PTFE cell was designed to allow exposing the working electrode to a liquid sample of 500 pl without loss of containment. The counter and the reference electrodes were introduced in the cell to be in direct contact with the liquid sample. An initial electrochemical impedance spectroscopy (EIS) measurement was performed in 500 pl of 2 mM [Fe(CN)6]3/4-in PBS buffer, pH 7.4. This measurement served as the initial baseline. The PTFE cell was then rinsed with PBS, and a new EIS measurement was performed. This process was repeated until two consecutive EIS measurements were identical to ensure stability in the measurements. Then, target ssDNA in the following concentrations1 pM, 10 pM, 100 pM, and 1000 pM were prepared in 0.1 M PBS buffer and they were sequentially incubated by applying 200 pl on the working electrode surface for 20 min. EIS measurements were performed before and after the target ssDNA incubation. Then, the impedance measured before the incubation with the ssDNA and after the incubation with the ssDNA were compared.

Claims (15)

1. Claims 1) A nucleic acid-sensing electrode, comprising: a. At least one electrically conductive layer comprising a surface, and b. A nanoporous layer bonded to the surface of the electrically conductive layer, wherein the nanoporous layer is made of a non-electrically conductive material, wherein the nanoporous layer is chemically bonded to the surface of electrically conductive layer as a result of the fabrication process, such that the attachment of the nanoporous layer to the electrically conductive layer is not a result of mechanical fastening, and c. At least one single-stranded nucleic acid probe attached to the nanoporous layer.
2. 2) The nucleic acid-sensing electrode of claim 1, wherein the electrically conductive layer comprises an inorganic oxide.
3. 3) The nucleic acid-sensing electrode of claim 2, wherein the inorganic oxide of the electrically conductive layer is selected from the group consisting of fluorine-doped tin oxide, indium tin oxide and indium zinc oxide.
4. 4) The nucleic acid-sensing electrode of any one of claims 1 to 3, wherein the nanoporous layer comprises an interconnected network of spaces defined as nanopores having a generally ellipsoidal or spherical shape with a diameter ranging between about 5 nm to about 100 nm.
5. 5) The nucleic acid-sensing electrode of any one of claims 1 to 4, wherein the nanoporous layer is fabricated directly onto the electrically conductive layer using a block copolymer and solgel nanoparticle co-assembly fabrication process.
6. 6) The nucleic acid-sensing electrode of any one of claims 1 to 5, wherein the nanoporous layer is fabricated using a sol-gel material selected from the group consisting of silicates and metal oxides.
7. 7) The nucleic acid-sensing electrode of any one of claims 1 to 6, wherein the single-stranded nucleic acid is attached to the nanoporous layer using a chemical method.
8. 8) The nucleic acid-sensing electrode of the claim 7, wherein the chemical method comprises a process selected from the group consisting of covalent bonding, electrostatic attachment, and physical adsorption.
9. 9) The nucleic acid-sensing electrode of any one of claims 1 to 8, wherein the single-stranded nucleic acid is a nucleic acid sequence complementary to a target sequence.
10. 10)A nucleic acid-sensing device comprising: a. the nucleic acid-sensing electrode of any one of claims 1 to 9; and b. at least a second electrode, wherein the second electrode is configured to contact a test liquid.
11. 11) The nucleic acid-sensing device of claim 10, wherein the second electrode is a counter electrode.
12. 12) The nucleic acid-sensing device of claims 10 or 11, wherein a third electrode is configured to contact the test liquid, wherein the third electrode is a reference electrode.
13. 13)The nucleic acid-sensing device of any one of claims 10 to 12, further comprising a measurement unit configured to record changes in the electrochemical signal.
14. 14)A method for detecting nucleic acid in a liquid sample using the nucleic acid-sensing device of any one of claims 10 to 13, the method comprising: a. measuring a baseline electrochemical signal of the device in a reference liquid; b. introducing a liquid sample to test to the device; and c. measuring changes in the electrochemical signal with respect to the baseline measurement in the reference liquid, wherein changes in the electrochemical signal are indicative of the presence of a nucleic acid complementary to the single-stranded nucleic acid in the nucleic acid-sensing electrode of any one of claims 1 to 9.
15. 15) The method of claim 14, further comprising comparing the changes in the electrochemical signal with a predefined threshold to determine the presence of the nucleic acid.