EP4133111A1

EP4133111A1 - Device and method for photothermal enhanced plasmonic biosensing

Info

Publication number: EP4133111A1
Application number: EP21709706.2A
Authority: EP
Inventors: Guangyu QIU; Jing Wang
Original assignee: Eidgenoessische Technische Hochschule Zurich ETHZ; Eidgenoessische Materialprufungs und Forschungsanstalt EMPA
Current assignee: Eidgenoessische Technische Hochschule Zurich ETHZ; Eidgenoessische Materialprufungs und Forschungsanstalt EMPA
Priority date: 2020-04-11
Filing date: 2021-03-09
Publication date: 2023-02-15
Also published as: WO2021204479A1

Abstract

The present invention is directed to a dual-functional plasmonic detector comprising (a) a photothermal unit configured to generate a monochromatic light beam; (b) a plasmonic sensing unit configured to introduce a localized surface plasmon resonance effect and to transduce a target sample by means of a polychromatic radiation beam; and (c) a detection unit configured to detect plasmonic characteristics by recording and processing the radiation beam altered by the plasmonic sensing unit. The invention also relates to the use of the dual-functional plasmonic detector of the invention for detecting, quantifying and/or characterizing a target sample, in particular for detecting, quantifying and/or characterizing a nucleic acid or a virus, e.g. a coronavirus such as SARS-CoV or SARS-CoV-2.

Description

DEVICE AND METHOD FOR PHOTOTHERMAL ENHANCED PLASMONIC BIOSENSING

For detecting viral infections, the molecular method reverse transcription polymerase chain reaction (RT-PCR) is routinely used to detect causative viruses using patient samples, e.g. from respiratory secretions (Corman et al., Eurosurveillance 2020, 25). However, RT-PCR can also fail for various reasons, such as its amplification of spurious nucleic acid contaminations. For example, the RT-PCR assays for SARS-CoV-2 detection have reported a number of false-negative results on confirmed infection cases (Xie et al., Radiology 2020, 200343). In clinical diagnosis, a single negative PCR result does not rule out COVID-19 infection as the reported positive rate was only 30-50% for laboratory-confirmed COVID-19 cases at the early stage of the outbreak (Zhang & Zhao, 2020, BioRxiv, 2020.02.20.958785). This is particularly true if the sample is from an upper respiratory tract (URT) specimen. A recent study of 167 COVID-19 infection patients showed that 5 (3%) patients had positive chest computed tomography (CT) findings but false-negative results from the RT-PCR testing. These 5 patients were eventually confirmed with COVID-19 infection by repeated swab tests (Xie et al., Chest CT for Typical 2019- nCoV Pneumonia: Relationship to Negative RT-PCR Testing. Radiology 2020, 200343). In addition, the current RT-PCR-based detection methods demand high manpower and long processing time.

Biosensors can provide an alternative and reliable solution to the clinical diagnosis, real-time detection and continuous monitoring. Among the different known biosensing techniques, localized surface plasmon resonance (LSPR) biosensing systems are applicable to different classes of analytes of clinical interests (Haeset al., J. Am. Chem. Soc. 2005, 127, 2264-2271). LSPR is a strong photon- driven coherent oscillation of the surface conduction electrons, which can be modulated when coupling occurs at the surface of the plasmonic materials (Willets & Van Duyne, Annu. Rev. Phys. Chem. 2007, 58, 267-297). Owing to the enhanced plasmonic field in the vicinity of nanostructures, LSPR sensing systems demonstrate high sensitivity to local variation, including the refractive index change and molecular binding.

However, the challenge remains that detection of biological particles such as, e.g. viral RNA, by LSPR detection of antibody binding does not always lead to satisfactory results (see Qiu et al., Adv. Funct. Mater. 2019, 29, 1806761) while the LSPR detection of hybridization of RNA with complementary bases suffers from the same essential drawbacks as RT-PCR (see above).

It is the objective of the present invention to provide a device and method for reliably detecting, characterizing and/or quantifying target samples, optionally biological samples, optionally viruses.

In a first aspect, the present invention is directed to a dual-functional plasmonic detector comprising:

(a) a photothermal unit (100) configured to generate a monochromatic light beam;

(b) a plasmonic sensing unit (200) configured to introduce a localized surface plasmon resonance effect and to transduce a target sample by means of a polychromatic radiation beam; and

(c) a detection unit (300) configured to detect plasmonic characteristics by recording and processing the radiation beam altered by the plasmonic sensing unit.

As used herein, the term "to detect plasmonic characteristics" means and is used interchangeably for all aspects and embodiments with "to detect localized surface plasmon resonance (LSPR)" or "to detect localized surface plasmon resonance (LSPR) response".

The term "dual-functional", as used herein, refers to the combined function of the photothermal unit 100 and the plasmonic sensing unit 200, wherein the photothermal unit 100 is configured to generate a plasmonic photothermal effect on the plasmonic sensing unit 200 and to control, e.g. elevate or heat, the local temperature thereon.

It was surprisingly found that the combination of heating, e.g. by means of the plasmonic photothermal effect (PPT), with LSPR detection leads to excellent detection thresholds and essentially excludes false positive or negative results, in particular for nucleic acid, optionally RNA detection. For example, the criteria for RNA-DNA hybridization are based on nucleic acid strand melting. Two complementary strands can specifically hybridize with each other when the temperature is slightly lower than their melting temperature, while a single mismatch can cause the melting temperature to decrease significantly. It was surprisingly found that heat, e.g. the plasmonic photothermal (PPT) heat energy, also known as the thermoplasmonic effect, can be used as a stable in situ heat source for controllable and uniform thermal processing of the hybridization process which is detected by LSPR. FIG. 1 depicts a representative embodiment of a dual-functional plasmonic detector according to the present invention comprising: a photothermal unit 100 that is capable of generating adjustable heating, optionally thermoplasmonic heating and controlling the local temperature, e.g. on an LSPR matrix; a plasmonic sensing unit 200 configured to modulate the LSPR effect and produce the optical sensing response; and a detection unit 300, determining plasmonic characteristics by detecting a spectrum of the light signal.

It is noted that "heating with the photothermal unit" or similar wordings can mean any way of heating, e.g. direct irradiation of the sample with the photothermal unit, wherein the sample is directly heated by the irradiated light, or alternatively, if an LSPR matrix is installed, by the LSPR matrix's conversion of light energy to heating energy (i.e. optical energy to thermal energy) based on the plasmonic photothermal effect.

As shown in FIG.2, the photothermal unit 100 for use in the present invention may comprise a radiation source 101, and an objective lens 102.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the photothermal unit (100) further comprises

(a-i) a source (101) configured to generate the monochromatic light beam; and (a-ii) an objective lens (102) configured to control the beam divergence.

The monochromatic electromagnetic radiation source 101 may comprise a light-emitting diode (LED), a laser diode (LD) or any other monochromatic light sources that have a designated wavelength and/or a high optical power. The monochromatic radiation source 101 may be sufficient to heat a sample of interest by the optical energy alone. Alternatively and if an LSPR matrix is installed, the designated wavelength of the monochromatic radiation source 101 preferably matches the peak absorption of the LSPR matrix 205. For example, a 532 nm LD can be employed when the LSPR matrix has a peak absorption at about 532 nm in the normal incident angle. The objective lens 102 is utilized to control the focal condition of the radiation source 101.

The plasmonic sensing unit 200 enables the plasmonic modulation and direct detection of target analytes as illustrated in FIG. 3.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the plasmonic sensing unit (200) further comprises

(b-i) a source (201) configured to generate the polychromatic radiation beam;

(b-ii) a polarizer (202) configured to linearly polarize the radiation beam;

(b-iii) a birefringent crystal (203) configured to generate retardation to the linearly-polarized radiation beam; (b-iv) a prism (204) arranged to couple the incident radiation beam onto an LSPR matrix (205) in an inclined incident angle;

(b-v) an LSPR matrix (205) configured to perform local plasmonic photothermal heating and localized surface plasmon resonance sensing transduction;

(b-vi) a probe cell (206) arranged to deliver the target sample onto the LSPR matrix;

(b-vii) a secondary polarizer (207) arranged to combine the radiation beam components from the LSPR matrix; and/or

(b-viii) a pass filter (208) configured to remove a scattered light component of the photothermal unit's (100) monochromatic light beam.

The polychromatic electromagnetic radiation source 201 may compromise a quartz tungsten halogen (QTH) lamp, a white light-emitting diode (WLED) or any other suitable polychromatic radiation source. The polychromatic radiation source 201 emits a collimated beam of light containing non-polarized components and the wavelength range should optionally be wide enough to include the response wavelength of the LSPR matrix 205. For example, a WLED with broad- spectrum from 500 nm to 750 nm can be employed when the LSPR matrix 205 has an adjustable plasmonic resonance wavelength from 560 nm to 640 nm in an inclined angle of incidence.

The polarizer 202 is employed to convert randomly polarized light to linearly polarized light containing p- and s- polarized components. The birefringent crystal 203 with designated thickness introduces a specific optical path difference to the p- and s- polarized components to produce the interference spectrum. The glass prism 204 is used to couple the light onto the LSPR matrix 205 with the attenuated total reflection (ATR) configuration.

The LSPR matrix 205 may, e.g. comprise a layer of two-dimensional distributed plasmonic nanoparticles, which have a strong plasmonic resonance at different wavelengths in the normal incident angle and an inclined ATR angle. The plasmonic nanoparticles for use in the present invention may be, e.g. nanoparticles comprising plasmonic-active materials, e.g. made by noble metals, e.g. gold, silver or their alloys, metal oxides, e.g. tin oxide, cadmium oxide, zinc oxide, transition metal nitrides or any combinations thereof.

For example, two-dimensional gold nanoislands (AuNIs) randomly distributed on a transparent substrate can be the plasmonic nanoparticles, which have a strong plasmonic resonance at 532 nm in the normal incident angle and 580 nm in an inclined ATR angle.

Optionally, the peak absorption of the LSPR matrix 205 matches the monochromatic light source in the photothermal unit 100. For example, the LSPR matrix can comprise a glass surface, or the LSPR matrix can be a transparent substrate as described below, with randomly distributed AuNIs. The AuNIs can have, e.g., a peak absorption at about 532 nm in the normal incident angle and a plasmonic resonance wavelength at about 580 nm in the inclined ATR angle. Therefore, the 532 nm light from the photothermal unit 100 can be absorbed by the AuNIs and converted to the PPT heating to elevate the local temperature, while the interference spectrum from the LSPR sensing unit 200 can be modulated by the AuNIs at 580 nm. The probe cell 206 for use in the present invention may, e.g. comprise a transparent, optionally an essentially fully transparent chamber or a chamber with a transparent window, e.g. a, optionally essentially fully, transparent poly (methyl methacrylate) chamber, a polydimethylsiloxane or a Teflon chamber with a transparent glass top. Alternatively, the probe cell 206 may be a single channel or a multichannel, e.g. microfluidic, flow cell arranged to transfer the target substance to the LSPR matrix 205, e.g. in nuclease-free water.

Another polarizer 207 can recombine the p- and s- polarized components after the glass prism 204. The recombined light beam contains the spectral interference signal carrying the plasmonic phase response introduced at the surface of LSPR matrix 205. The pass-filter 208 may indicate an optical filter that has a sharp cut-on wavelength to remove the interference light from the monochromatic light source 101. For example, a long-pass filter with a cut-on wavelength of 552 nm can be used to eliminate the scattered light from photothermal unit 100 with the 532 nm monochromatic LD source 101, while the plasmonic sensing beam containing the adjustable resonance wavelength, e.g. from 560 nm to 640 nm, is able to pass the filter.

The skilled person appreciates that the above-noted features (b-ii), (b-iii), (b-iv), (b-vii) and (b- viii) are optionally not essential for practicing the present invention. Some or all of these features may be omitted as long as the device comprises a photothermal unit 100, a plasmonic sensing unit 200 and a detection unit 300 as defined herein and configured such that they allow LSPR detection as described herein.

As shown in FIG. 4, the detection unit 300 may comprise an iris aperture 301 and an optical probe 302.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the detection unit (300) further comprises

(c-i) an iris aperture (301) arranged to select the radiation beam from the plasmonic sensing unit; and

(c-ii) an optical probe (302) arranged to detect the radiation beam signal selected by the iris aperture.

The iris aperture 301 is provided for selecting the spatial position of interest from the ATR light beam. The optical probe 302 may indicate a spectrometer that is able to resolve the spectral interference from the plasmonic sensing unit 200. In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the optical probe (302) further comprises an optical probe unit configured to obtain the intensity spectrum of the radiation beam; and a processing unit configured to determine the characteristics based on the intensity spectrum.

As shown in FIG. 5, the dual-functional plasmonic detection system is based on the PPT heating effect from the photothermal unit 100 and the in situ LSPR sensing transduction from the plasmonic sensing unit 200. The final transducing signal collectively effected by the photothermal unit 100 and the plasmonic sensing unit 200 are recorded by the detection unit 300.

In the representative thermoplasmonic experiments carried out in the context of the present invention, the direct absorption of laser radiation 101 at 532 nm decayed nonradiatively by generating more hot electrons in AuNIs of the LSPR matrix 205. The photoexcited highly energetic electrons quickly dissipated and released thermal energy to heat the ambient environment. Conversely, the PPT induced temperature increase was also responsible for a refractive index variation of the surrounding environment, which was in situ detected by the plasmonic sensing unit 200 as shown in FIG 7d.

In the Examples, the local PPT temperature on the LSPR matrix 205 was calibrated based on a series of tests at different ambient temperatures and the thermo-induced LSPR response was recorded accordingly as shown in FIG 7c. Based on this calibrated LSPR-temperature regression, the localized photothermal temperatures under different laser powers were retrieved as shown in FIG

7d.

To further evaluate the radiation-induced PPT effect and the local temperature profile on the LSPR matrix 205, the detection unit 300 was used to scan the heating area for mapping the LSPR phase responses and actual temperature distribution on the AuNIs.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the source (101) configured to generate the monochromatic light beam is configured to generate adjustable output power and further comprises a controller unit configured to supply adjustable power to the source (101).

For example, a monochromatic radiation source 101 with, e.g., 32 mW power can be applied to the optimal AuNIs absorbers with a peak absorption at, e.g., 532 nm. At each point of interest, the detection unit 300 can be used to record two interferometric spectra: one reference spectrum without PPT heating and one spectrum with PPT heating. By scanning the laser spot and surrounding area, e.g. with a 0.5 mm step interval, the spatial distribution of LSPR phase changes can be retrieved. At each scanning pixel, the retrieved phase response can subsequently be converted to the local temperature based on the calibration curve shown in FIG 7c. Therefore, the corresponding temperature distribution around the PPT heating can be obtained and illustrated, as e.g. shown in FIG 7f. Under the radiation of 32 mW, the local temperature of LSPR matrix 205 was significantly elevated from 21 °C (room temperature) to 41 °C at the center of the laser spot. In an embodiment, the local temperature of the LSPR matrix is elevated to about 35 to 90, 35 to 60, or 35 to 45, optionally about 38 to 42, optionally about 41 °C.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the LSPR matrix (205) further comprises a layer of plasmonic-active nanostructures arranged to be functional in plasmonic photothermal effect and localized surface plasmon resonance sensing transduction; a transparent substrate configured to support a layer of plasmonic-active nanostructures and coupled to the prism (204); and an optical liquid configured to match the refractive index of the transparent substrate and the prism (204).

Plasmonic-active nanostructures for use in the present invention can be, e.g., nanospheres, nanorods, nanopillars, nanowires, nanoholes, nanoribbons, nanorings, nanoflowers, nanocages, nanofoams, nanoplatelets, nanoshells, nanotips, nanobipyramids or any combinations thereof.

The transparent substrate configured to support a layer of plasmonic-active nanostructures for use in the present invention can be, e.g., a transparent silicate glass plate, e.g. soda-lime glass, borosilicate glass, lead glass, aluminosilicate glass; or a transparent polymer plate, e.g. cyclic olefin copolymer, poly(methyl methacrylate), polydimethylsiloxane; or a transparent metal oxide plate, e.g. tin oxide, indium oxide, zinc oxide, titanium oxide; or a transparent ceramic plate, e.g. yttrium aluminum garnet, yttria, vanadium oxide; or other composite plates combined thereof.

Specifically, the transparent substrate may be an LSPR matrix 205 which may comprise a layer of plasmonic active nanoparticles, which have a strong plasmonic resonance at a specific wavelength. The LSPR matrix 205 may further comprise plasmonic active materials as described herein.

The optical liquid configured to match the refractive index of the transparent substrate and the prism for use in the present invention may be, e.g., microscope immersion oil, or optical coupling liquids as known in the art.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the plasmonic-active nanostructure further comprises a probe arranged to selectively interact with a target substance of interest. The LSPR matrix 205 and/or the transparent substrate may comprise a selective probe arranged to interact with a target of interest in a target sample. The target and/or probe may be or comprise any single stranded nucleic acid sequences which are capable of hybridizing to each other in suitable aqueous solutions and/or under suitable physiological conditions, optionally in nuclease- free aqueous solutions, optionally nuclease-free water and/or under physiological conditions, e.g. an RNA or a single stranded DNA (e.g. cDNA), optionally a viral nucleic acid sequence, e.g. a viral RNA or cDNA thereof. The target can be applied as a solution in optionally nuclease-free aqueous solutions.

The term hybridizing and hybridizing in aqueous solutions and/or under physiological conditions means for all aspects and embodiments herein that two single strands of complementary or partially complementary nucleic acids, optionally with at least 5 complementary base pairs, optionally with 10 to 30 or more complementary base pairs, stably bind to each other to form a double stranded structure to the extent that this binding can be detected with the detector of the present invention. The description and Examples provided herein also demonstrate and inform the skilled person how to determine and verify hybridization, e.g. also as a function of temperature, e.g. at different temperatures which can be varied by the photothermal unit 100.

For example, the ability of hybridization of two single stranded nucleic acids (e.g. probe and target) can be determined experimentally as commonly known in the art, optionally by hybridizing under the conditions provided in the Examples below, and detecting by using the detector of the present invention or by other known, e.g. optical, methods.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein

- the plasmonic sensing unit (200) comprises an LSPR matrix (205) configured to perform local plasmonic photothermal heating and localized surface plasmon resonance sensing transduction;

- the LSPR matrix (205) comprises a layer of plasmonic-active nanostructures arranged to be functional in plasmonic photothermal effect and localized surface plasmon resonance sensing transduction; and

- the plasmonic-active nanostructures comprise a probe that is a single stranded nucleic acid configured to hybridize with an at least partially complementary single stranded nucleic acid target substance, optionally RNA, optionally a single stranded nucleic acid probe hybridizing to a viral RNA-dependent RNA polymerase (RdRp) sequence, optionally selected from SEQ ID NO. 1 (RdRp-COVID) and 2 (RdRp-SARS), a viral open reading frame lab (ORFlab) sequence, optionally SEQ ID NO. 3 (ORFlab-COVID), or non-structural protein 13 (nspl3) nucleic acid sequence having SEQ ID NO. 4. In the context of the present invention, the term "at least partially complementary" means that the strands are complementary to the extent that hybridization, as defined herein, is possible and can be detected by the detector of the invention. For example, at least partially complementary may be that at least 5, or optionally 10 to 30 complementary nucleic acids, or at least 50% of the nucleic acids in the sequence are complementary, and are optionally all consecutive in the sequence.

For example, a specific sequence from severe acute respiratory syndrome coronavirus (SARS- CoV or SARS) or severe acute respiratory syndrome coronavirus 2 (COVID or SARS-CoV-2), such as the RNA-dependent RNA polymerase (RdRp), optionally having SEQ ID NOs.l and 2, the open reading frame lab (ORFlab) sequence, optionally having SEQ ID NO.3, or the non-structural protein 13 (nsp-13) nucleic acid sequence, optionally having SEQ ID NO: 4, as a target is capable to and can be used to hybridize to the specific probe. The target substance and probe can be complementary strands as shown in Table 1 below, or at least comprise these complementary parts in their strands. This means that a probe for use in the present invention which hybridizes to a target of, e.g., SEQ ID NO. 1, comprises or has the complementary sequence of, e.g., SEQ ID NO. 5. The same applies for probes that hybridize with targets of SEQ ID NOs. 2, 3, 4 and 9 which probes then can, e.g., comprise or have SEQ ID NOs. 6, 7, 8 and 10, respectively.

Alternatively, other at least partially complementary probes, or sequence selective units may be included for detecting target strands other than the listed sequence or for detecting other temperature-responsive substances. Also, the probes may be synthetic products or isolated natural probes.

Sequence Function Strands SEQ ID NO.

RdRp- Target 5'-CAGGT GGAAC CTCAT CAGGA GATGC-3' 1

COVID Probe (-C) Thiol-5'-GCATC TCCTG ATGAG GTTCC ACCTG-3' 5

RdRp- Target 5'-C CAGGT GGAAC ATCAT CCGGT GATGC-3' 2

SARS Probe (-C) Thiol-5'-GCATC ACCGG ATGAT GTTCC ACCTG G-3' 6

ORFlab- Target 5'-CCGTC TGCGG TATGT GGAAA GGTTA TGG-3' 3

COVID Probe (-C) Thiol-5'-CCA TAACC TTTCC ACATA CCGCA GACGG-3' 7

Ta rget 5'-ACAC TAGCC ATCCT TACTG CGCTT CG-3' 9

E

Probe (-C) Thiol-5'-CGAAG CGCAG TAAGG ATGGC TAGTG T-3' 10

5'- CTACTGTACG TGAAGTGCTG TCTGACAGAG 4

„ „ „ Target AATTACATCT TTCATGGGAA GTTGGTAAAC CTAGACC-

Nsp-13 ⁵ ₃,

Probe (-C) Thiol- 5’-GGT CTA GGTTTA CCA ACT TCC C-3’ 8 Table 1 - Exemplary Probe and Target Sequences

The probes can be functionalized to adhere to the LSPR matrix in any suitable way, e.g. by a thiolate-metal bonding, an aldehyde-hydroxyl crosslinking or a physical adsorption. Optionally, the probe comprises a linker, optionally a hydrocarbon linker, optionally a C6-linker, between the LSPR matrix or, e.g. the thiol group, and the probe, optionally between the thiol group and the 5'-end of the, e.g. nucleic acid, probe. Based on the synthetic oligonucleotide probes with a thiol group, the LSPR matrix 205 or transparent substrate can be directly functionalized by forming the Au-S bond between the thiol-cDNA probe and AuNIs.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the probe cell (206) further comprises a single channel or a multichannel, optionally microfluidic, flow cell arranged to transfer a target substance to the plasmonic-active nanostructure.

Optionally, a total of 1 nmol thiol-cDNA probes can be utilized to functionalize the single channel microfluidic cell 206 and/or AuNIs LSPR matrix 205, e.g. for SARS-CoV-2 sequence detection. The proper surface functionalization that is enough to functionalize essentially the entire AuNIs sensing surface of LSPR matrix 205 can increase the sensitivity and suppress non-specific binding events. Optionally, the AuNIs LSPR matrix 205 can be functionalized with about 0.1 to 10 nmol probes, e.g. cDNA probes.

Referring to Fig. 5, the LSPR matrix 205 can further comprise a functionalized layer of cDNA probes to hybridize the specific target, e.g. a viral sequence. The target sequences can be brought onto the LSPR matrix 205 through the microfluidic flow cell 206. Before the injection of the target sequence, nuclease-free aqueous solutions can be flown across the microfluidic flow chamber 206 and the photothermal radiation 101 can be turned on to establish a steady phase reference and baseline. Optionally, the interaction of probe and target sequences can be from about 1 to about 600 s, optionally about 600 s. Then the dual-functional AuNIs LSPR matrix 205 can be further flushed with nuclease-free water to remove non-specific binding items and to check the final LSPR phase response from the detection unit 300. The localized photothermal effect generated by the photothermal unit 100 is capable to significantly improve the hybridization kinetics of the target sequence and its probe, e.g. cDNA probe. Due to the faster hybridization kinetics, the differential phase response retrieved from the plasmonic sensing unit 200 is elevated for detecting the target sequences at different concentrations.

More importantly, the PPT heating generated by the photothermal unit 100 is capable to inhibit the spurious binding of non-matching sequences by elevating the local temperature in the vicinity of LSPR sensing matrix 205 above room temperature, e.g. to about 25-90 °C, 25-60°C or to about 41 °C.

For example, in the selected gene sequences shown above, only three fixed nucleotide bases were different between RdRp-COVID and RdRp-SARS. The plasmonic sensing unit 200 without the aid of photothermal unit 100 reported a false positive response signal when detecting the RdRp- SARS sequence (see below), which indicated that a similar but not fully complementary sequence was also able to interact and partially hybridize with the cDNA probes at room temperature in the absence of photothermal unit 100. At the elevated temperature by including the photothermal unit 100, the standard free energy of hybridization was weaker due to the mismatched base-pairs. Thus, the similar but not fully matched sequences of SARS-CoV could be distinguished by the detector of the present invention.

In an example operation of the dual-functional plasmonic detector to detect RdRp-COVID target sequence, the LSPR matrix 205 was functionalized with the complementary probe RdRp- COVID-C; the 532 nm laser radiation in photothermal unit 100 was applied on the LSPR matrix 205; the interferential phase response of the plasmonic sensing unit 200 was recorded by the detection unit 300. The dual-functional plasmonic detections of RdRp-COVID were investigated over the concentration range from 0.01 pM to 50 mM as shown in FIG 10a. The dual-functional plasmonic detector exhibited a limit in the range from 0.1 pM to 1 mM for detecting oligonucleotides, covering seven orders of magnitude. The detectable RdRp-COVID sequence concentration corresponding to the systematic limit of detection was about 0.22 pM.

The experimental Examples disclosed herein demonstrate that the dual-functional plasmonic detector of the present invention can provide up to picomolar-level sensitivity of the target sequence, which can be useful in applications such as clinical diagnosis of viral diseases, detecting the virus in the environment and long-term monitoring.

Optionally, the dual-functional plasmonic detector described herein is based on a photothermal enhancement and a highly specific DNA hybridization in the label-free LSPR sensing unit. Therefore, it is an accurate and reliable device and/or method for analytical detection. In addition, the system is applicable to various nucleic acid sequences by simply employing different complementary probe sequences. For example, the dual-functional plasmonic detector is validated by performing the selective hybridization detection on several different genome sequences, i.e. the ORFlab-COVID sequence and the envelope (E) sequence from SARS-CoV-2, and the RdRp-SARS sequence from SARS-CoV. The corresponding LSPR phase sensing responses with the enhancement of photothermal unit 100 are illustrated in FIG 10c. As the concentration increased from 1 pM to 1 nM, the mean LSPR response levels of each sequence climbed in a proportional manner, which proved the feasibility of this dual-functional plasmonic detector for the quantitative analysis of viral nucleic acids.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the detector further comprises a source that comprises and/or is configured to add to the probe before, during or after hybridization with the single stranded nucleic acid target substance:

(i) a further fluorescent single stranded nucleic acid that

- is at least partially complementary to the nucleic acid target substance when bound to the probe, and

- comprises a quencher suitable for quenching the fluorescence;

- comprises a recognition site for a cleaving enzyme; and

(ii) a cleaving enzyme suitable for recognizing the recognition site and cleaving the fluorescent nucleic acid when at least partially hybridized with the nucleic acid target substance, and unquenching the fluorescence; wherein the detection unit (300) is configured to also detect the fluorescence-induced LSPR response

The source may be internal or external to the detector of the present invention and it may be detachable or integrated into the detector. Delivery from the source may be performed via the probe cell 206 described herein which may comprise or be fluidly connected to the source.

For easier reading, the further fluorescent single stranded nucleic acid described above is also termed the fluorescent probe or fluorescent nucleic acid. For clarity's sake it is noted that this fluorescent single stranded nucleic acid is termed "fluorescent" although it is not actually fluorescent before its cleavage by the cleaving enzyme due to the quenching action of the quencher which is comprised in the fluorescent single stranded nucleic acid.

The further fluorescent single stranded nucleic acid is at least partially complementary to the nucleic acid target substance when the target substance is bound to the probe, which means that the fluorescent single stranded nucleic acid binds to the nucleic acid target substance, i.e. hybridizes with it. The terms "at least partially complementary" and "hybridizing" are used as defined above. The cleavage enzyme is selected such that it only cleaves the fluorescent single stranded nucleic acid when bound to the target substance.

As shown in FIG. 21a, the single-stranded region of the probe-captured target nucleic acid sequence can combine, i.e. hybridize, with a an at least partially complementary fluorescent single stranded nucleic acid to form a double-stranded duplex structure. Under the action of a cleavage enzyme that is site- or structure- specific for the recognition site of the further fluorescent single stranded nucleic acid - when target-sequence-hybridized - the fluorescent single stranded nucleic acid can be cleaved, e.g. at the recognition site, to release a fluorescent signal. This process can be cyclically performed (termed cyclic fluorescence probe cleavage (CFPC)) due to the assistance of the photothermal unit 100, thereby forming an amplification-based detection signal. In other words, the optionally cyclic, fluorescent signal emitted after cleavage of the fluorescent nucleic acid is used to stimulate an additional LSPR readout.

As used herein, CFPC detection means the detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe.

As shown in FIG. 21b for a representative non-limiting example, the combination of (I) direct LSPR detection based on the target substance binding to the probe and (II) secondary CFPC detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe, constitutes a self-validating biosensing method. These two detections can be completed on the same plasmonic sensing unit (200), e.g. also on the same LSPR matrix (205).

The dual-functional plasmonic detector for use in the present invention can be configured to provide the desired heating temperature for nucleic acid hybridization, enzyme cleavage and cyclic fluorescent probe release.

The plasmonic effects, which provide the desired temperature for nucleic acid hybridization, enzyme cleavage and cyclic fluorescent probe release can be experimentally characterized by measuring the thermo-induced variation of the refractive index. For example, the two-dimensional LSPR matrix 205 can be initially optimized to have the peak absorption matching the irradiation wavelength of the monochromatic light beam of the photothermal unit 100, e.g. homogenized laser 101. Typically, within a time period in the order of nanoseconds, the photoexcited plasmon is quickly transferred to the metallic lattice and eventually releases the thermal energy to the surrounding medium, e.g. water, through electron-photon collisions. This thermal dissipation process is believed to be responsible for a local temperature elevation and a thermally induced variation of the refractive index in the vicinity of the plasmonic-active nanostructures, e.g. AuNIs. By gradually increasing the irradiation power of photothermal unit 100, e.g. with a control unit configured to control this power, a fast and stable in-situ temperature adjustment can be achieved (see, e.g. FIG .25). Simultaneously, the LSPR response, e.g. the ATR light in prism 204 which transduces the spectral LSPR phase shift, can be used to quantify the variation of the refractive index and local temperature.

In all embodiments and aspects disclosed herein, the photothermal unit 100 may comprise a power control unit which may comprise one or more objective lenses 102, attenuator or chord for controlling the total output power of the photothermal unit 100, and therefore controlling the temperature on the plasmonic sensing unit 200, e.g. on the LSPR matrix 205.

Typically, at a temperature below the melting threshold ( T_m ), two complementary nucleic acid strands can hybridize and form a thermodynamically favoured double-helix (duplex) structure by establishing the non-covalent and sequence-specific base-pairs interactions (hybridization as used herein). For example, in the first-step hybridization test, the plasmonic sensing unit 200 probes the on-chip hybridization of two complementary sequences and provides a quantitative result for the binding kinetics as shown in FIG. 26.

As example, thiolate DNA receptors can be covalently immobilized onto the LSPR matrix 205 as probes by forming metal-thiol bonds. When the fully matched sequences reach the DNA receptors (probes), they are able to capture and hybridize with these target sequences. This binding reaction changes the local refractive index surrounding the LSPR matrix 205 and is directly sensed by the detection unit 300, thereby producing a fast and real-time phase response.

In order to further understand the thermodynamic and binding kinetics of the hybridization, the photothermal unit 100 can be used to assist the study of nucleic acids hybridization and dehybridization. A far-field laser excitation was found suitable to control the near-field thermoplasmonic heating on the LSPR sensor surface to reach the desired temperature, therefore optimizing the hybridization condition for a specific sequence. Generally, higher temperature can increase the free diffusion of single-stranded nucleic acid sequences and make the hybridization thermodynamically favourable, thereby improving hybridization kinetics. However, temperatures approaching the melting threshold cause the dehybridization of double-stranded nucleic acid sequences.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the detection unit (300) is configured to detect the LSPR response of the target-probe binding and the fluorescence-induced LSPR response and is configured to detect, quantify and/or characterize the target substance based on a comparison of the LSPR responses.

For accomplishing the comparison of the LSPR responses, the detection unit 300 may be used or the detector may comprise an additional self-validating unit for comparing the responses and validating these in view of their combination. In the detection unit 300 or the self-validating unit, the detection results of LSPR-probe-target compound-detection and CFPC fluorescence detection can be quantitatively and qualitatively analyzed through real-time detection values, e.g. as shown in FIG. 22 and FIG. 23, and through regression curves as shown in FIG. 24a and FIG. 24b, respectively. The detection unit 300, the self-validating unit or an additional processing unit can analyze and process the results of the two detections to give the final detection report. Two different detection processes, namely amplification-free nucleic acid hybridization detection and amplification-based CFPC LSPR detection, can analyze the same sample on a single detection platform or chip. These two detection approaches both use LSPR for sensing modulation, but they have complementary characteristics. For instance, the direct hybridization detection is fast, while the CFPC method has higher sensitivity. The direct communication of these two tests surprisingly makes the self-validating test results have the outstanding characteristics of fast, accurate, sensitive and reliable.

- the further fluorescent single stranded nucleic acid comprises a fluorophore at one end of the strand and a quencher at the other end; and/or

- the recognition site for a cleaving enzyme is selected from the group consisting of a dehydrothymine site spacer, an inosine site spacer, a hydroxymethyluracil site spacer, an 8- oxogunanie site spacer, and an apurinic/apyrimidinic site spacer.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the further fluorescent single stranded nucleic acid or the fluorophore has a maximum absorption at a wavelength that overlaps with the excitation wavelength of the photothermal unit source (101), optionally a maximum absorption between and including 512 and 545 nm, and/or the fluorophore is selected from the group consisting of ATT0532, Rhodamine 6G, Alexa Fluor 532, Yakima Yellow and Dyomics-530.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the quencher is able to decrease the fluorescence intensity of a given substance, optionally the quencher is selected from a metal nanoparticle, molecular oxygen, iodide ions, acrylamide, a black hole quencher (BH ) and a Dabcyl quencher. Optionally, the quencher is selected from a black hole quencher (BHQ) and a Dabcyl quencher.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the enzyme suitable for recognizing the recognition site and cleaving the fluorescent nucleic acid is selected from the group consisting of a nucleic acid enzyme, optionally a restriction endonuclease, optionally endonuclease III, endonuclease IV, Tth endonuclease IV, endonuclease V, and endonuclease VIII.

In an embodiment, the dual-functional plasmonic detector described herein is one, wherein the probe cell (206) further comprises a single channel or a multichannel, optionally microfluidic, flow cell arranged to transfer a target substance, and/or the second fluorescent single strand nucleic acid and the enzyme suitable for cleaving the fluorescent nucleic acid to the plasmonic- active nanostructure. In an example operation of the dual-functional plasmonic detector of the present invention to detect an nsp-13 target sequence (SEQ ID NO: 4), the single-stranded region of the nsp-13 target sequence bound to the exemplary probe can combine, i.e. hybridize, with the designed and exemplary further fluorescent single stranded nucleic acid as shown in Table 2 below and form a double-stranded duplex structure. Under the action of site- and/or structure- specific enzymes (e.g. a site-specific enzyme, e.g. endonuclease IV), the further fluorescent single stranded nucleic acid can be cleaved to release a fluorescent signal. This process can be cyclically performed due to the assistance of the photothermal unit 100, thereby forming an amplification-based detection signal.

Sequences Sequence Length Melting temperature

5'- CTACTGTACG TGAAGTGCTG

Nsp-13 target

TCTGACAGAG AATTACATCT sequence 67 Tm: 68.7°C TTCATGGGAA GTTGGTAAAC (SEQ ID NO: 4) CTAGACC-3'

Probe (-C) THIOL-5'-GGT CTA GGT TTA CCA

22 Tm: 63.9°C (SEQ ID NO: 8) ACT TCC C-3 Fluorescent

ATT0532-5' - GAC AGC ACT TC(AP- Probe 22 Tm: 54.0°C site)CGT ACA GTA G -3' - BHQ1

(SEQ ID NO: 11) Fluorophore

ATTQ532-5' - GAC AGC ACT TC-3' 11 Tm: 34.0°C (SEQ ID NO: 12)

Quencher

5'- CGT ACA GTA G -3' - BHQ1 10 Tm: 30.0°C

(SEQ ID NO: 13)

Table 2 - nsp-13 target sequence, respective probe and further fluorescent single stranded nucleic acid (termed fluorescent probe)

FIG. 27 shows a series of exemplary in-situ dehybridization tests on the DNA duplexes noted in Table 2 by gradually increased thermoplasmonic heating with photothermal unit 100. For example, the thermoplasmonic induced local temperature gradually increased from 33.2 °C to 45.6 °C. After each heating cycle, the photothermal unit 100 irradiation laser was switched off, and the local temperature quickly reduced to a typical room-temperature at about 25 °C. Under this laser- off condition, the real-time phase response was recorded and used to determine whether the plasmonic photothermal (PPT)-induced dehybridization occurred. When the thermoplasmonic temperature was below 41.6 °C, the dehybridization responses were weak, which indicated that the nucleic acid duplexes maintained the double-stranded state. However, in a PPT near-field with the temperature higher than 42.5 °C, a negative LSPR phase signal appeared, which indicated that part of the nucleic acid duplexes had dehybridized by the AuNI thermoplasmonic field. This dehybridization process was substantial at temperatures higher than 42.5 °C for this target sequence, showing that higher temperatures were not favourable for effective in-situ detection of the selected nucleic acid target.

The photothermal unit 100 can be used for the detection of the binding and cleavage of the further fluorescent single stranded nucleic acid as disclosed herein to assist in regulating enzyme activity, cleavage rate, nucleic acid sequence hybridization and dehybridization, and fluorescence signal excitation.

For example, Endo-IV and the fluorescent probe (22nt) possessing a synthetic AP site were used for the cyclic gain-assisted amplification and in-situ secondary CFPC detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe. Endo-IV specifically recognizes and cleaves the AP site within a DNA duplex by generating a hydroxyl group at the 3'-terminus. In the CFPC detection, 50 pL freshly prepared reaction solution containing 1 unit of Endo-IV and 1 mM fluorescent DNA probes was applied onto the biosensor surface loaded with the captured target sequences. The local heating by the photothermal unit 100 provided a suitable nanoscale temperature on LSPR matrix 205 (T=41.3 ± 0.2 °C) for both DNA hybridization and enzyme cleavage. First, the synthesized fluorescent probes hybridized with the captured sequences in the section which was still single-stranded (FIG. 21a). The melting temperature Tm of the fluorescent probe was determined to be 54 °C with the nearest-neighbour method. Then, Endo-IV recognized the AP site in the middle of the probe and cleaved the probe into two short pieces, which contained an ATT0532 fluorophore and a BHQ1 quencher, respectively. After the cleavage, the two short oligonucleotides possessed much lower melting temperatures at 34 °C and 30 °C, respectively. Since the PPT-induced local temperature was evidently higher than their Tm thresholds, the two short oligonucleotides dissociated from the target sequence and departed from the single-stranded viral sequence, leaving it available for the next cyclic binding, cleavage and dissociation (BCD) reaction.

In the CFPC test, the released fluorophores were simultaneously excited by the 532 nm laser and emitted photons of a low energy as shown in FIG 28. Before diffusing into the bulk solution, the released fluorophores were in the vicinity of the AuNI biosensing interface (<15 nm) and directly stimulated a transient LSPR phase response through the near-field interaction with AuNIs as shown in the real-time sensorgram shown in FIG. 28. In comparison with the background phase response (without laser excitation) in the first 500s of the reaction, the near-field interaction (500s to 1200s) between the released fluorophores (gain molecules) and LSPR AuNIs produced transient phase jumps with high amplitudes. Different from the molecular binding, the instantaneous phase responses caused by this energy transfer and gain medium stimulation demonstrated a wider phase distribution. As a signal distinct from the background LSPR phase response, this transient phase jump provided qualitative proof for verifying the existence of the target sequence.

This cyclic BCD process proceeded continuously on the AuNI biosensor surface under the cooperation of Endo-IV and local thermoplasmonic heating with photothermal unit 100. Subsequently, the fluorescent DNA probes diffused into the bulk solution and the overall fluorescence intensity gradually increased over time. As shown in FIG. 29, the accumulated fluorescent irradiation was able to continuously stimulate the LSPR response as an active gain medium and continuously shifted the LSPR phase response.

As a control experiment, the gain-enhanced detection was further verified by using laser assisted detection of Rhodamine 6G (R6G) fluorophores in the liquid. As shown in FIG. 30, laser excitation of the photoactive medium (e.g. R6G) significantly amplified the LSPR phase detection response compared to that without laser irradiation. Since the ATT0532 molecules have similar emission properties as R6G, this set of control experiments proved that the cumulated fluorescent probes in the liquid effectively provided sensitive and reliable biosensing results.

At the same temperature (corresponding to a constant cleavage rate by the Endo-IV) and time duration, the accumulation rate of the released fluorescent DNA probe was proportional to the number of target nucleic acid sequences on the AuNI plasmonic sensing unit. Therefore, the differential LSPR phase response can be used for the quantitative measurement of the viral nucleic acid concentration.

In another aspect, the present invention is directed to a use of the dual-functional plasmonic detector as described herein for detecting, quantifying and/or characterizing a target sample.

In an embodiment, the use of the present invention is one, wherein the target sample is a nucleic acid, e.g. DNA or RNA, optionally an RNA, optionally a viral RNA, optionally a positive sense, single stranded viral RNA, or a virus, e.g. corona virus, optionally SARS-CoV or SARS-CoV-2, optionally the target sample comprises a viral RNA-dependent RNA polymerase (RdRp) sequence, optionally selected from SEQ ID NO. 1 (RdRp-COVID) and 2 (RdRp-SARS), a viral open reading frame lab (ORFlab) sequence, optionally SEQ ID NO. 3 (ORFlab-COVID), or non-structural protein 13 (nspl3) nucleic acid sequence having SEQ ID NO. 4.

In another aspect, the present invention is directed to a method for detecting, quantifying and/or characterizing a target sample comprising the steps:

(i) providing a dual-functional plasmonic detector as described herein; (ii) providing a target sample of interest and positioning the sample in the plasmonic sensing unit (200), optionally on an LSPR matrix;

(iii) generating, from the photothermal unit (100), a monochromatic light beam;

(iv) heating the target sample, optionally the LSPR matrix (205), with the light beam, optionally based on a plasmonic photothermal effect;

(v) generating, from a plasmonic sensing unit (200), a polychromatic radiation beam containing randomly polarized components;

(vi) modulating the radiation beam in the plasmonic sensing unit (200) to reach the target sample, optionally the LSPR matrix (205), optionally in an inclined incident angle;

(vii) introducing an LSPR effect associated with the target sample to the radiation beam in the plasmonic sensing unit;

(viii) recording and processing the radiation beam from the plasmonic sensing unit in the detection unit (300) to determine an LSPR response; and

(ix) detecting, quantifying and/or characterizing the target substance based on the LSPR response, optionally based on a calibrated regression with LSPR responses at different target concentrations.

Heating with the monochromatic light beam can be done, e.g., for about 1 to 50 s. All definitions provided above in the context of the device of the present invention also apply to the use and the method of the present invention.

In an embodiment, the method of the present invention is one, wherein after step (viii), the method further comprises steps

(viii-a) contacting the sample in the plasmonic sensing unit (200) with

- a further fluorescent single stranded nucleic acid that is at least partially complementary to the nucleic acid target substance under conditions that allow for binding to the probe, wherein the further fluorescent single stranded nucleic acid comprises a quencher that quenches the fluorescence; and

- an enzyme which cleaves the fluorescent nucleic acid when at least partially hybridized with the nucleic acid target substance, and unquenches the fluorescence;

(viii-b) generating, from the photothermal unit (100), a monochromatic light beam;

(viii-c) heating the target sample, optionally the LSPR matrix (205), with the light beam, optionally based on a plasmonic photothermal effect;

(viii-d) generating, from a plasmonic sensing unit (200), a polychromatic radiation beam containing randomly polarized components; (viii-e) modulating the radiation beam in the plasmonic sensing unit (200) to reach the target sample, optionally the LSPR matrix (205), optionally in an inclined incident angle;

(viii-f) introducing an LSPR effect associated with the target sample to the radiation beam in the plasmonic sensing unit; and

(viii-g) recording and processing the radiation beam from the plasmonic sensing unit in the detection unit (300) to determine a fluorescence-induced LSPR response; and wherein in step (ix), the detecting, quantifying and/or characterizing of the target substance is based on the LSPR response of the target-probe binding and the fluorescence-induced LSPR response, optionally based on a calibrated regression with LSPR responses at different target concentrations.

For example, steps (viii-b) to (viii-g) and (ix) can be seamlessly repeated or cycled, e.g., because when the photothermal unit 100 is switched on, the photothermal heating can cause the cleavage enzyme to continue cleaving the further fluorescent single stranded nucleic acid bound to the nucleic acid target substance. The detection unit 300 can continue to record the ongoing reaction (signal amplification) as a function of time.

In an embodiment, the method of the present invention is one, wherein in step (ix), the LSPR response of the target-probe binding and the fluorescence-induced LSPR response are compared and the detecting, quantifying and/or characterizing of the target substance is based on the comparison of the LSPR responses.

As noted in the context of the detector of the present invention, for accomplishing the comparison of the LSPR responses, the detection unit 300 may be used or the detector may comprise an additional self-validating unit. In the detection unit 300 or the self-validating unit, the detection results of LSPR-probe-target compound-detection and CFPC fluorescence detection can be quantitatively and qualitatively analyzed through real-time detection values, e.g. as shown in FIG. 22 and FIG. 23, and through regression curves as shown in FIG. 24a and FIG. 24b, respectively. The detection unit 300, the self-validating unit or an additional processing unit can analyze and process the results of the two detections to give the final detection report. In the following representative scenarios are given for the comparison of the LSPR response of the target-probe binding and the fluorescence-induced LSPR response, e.g. in the context of self-validation.

In a first exemplary scenario, when the direct LSPR detection and the CFPC detection (i.e. the detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe) both report positive results, self-verification of these dual mode readouts confirm the presence of the target substance, e.g. SARS-CoV-2 viruses. In a second exemplary scenario, when the direct LSPR detection and the CFPC detection both report negative results, self-verification of the two readouts double confirmed the sample is most likely negative.

In a third exemplary scenario, when a weak positive direct LSPR detection signal and a negative CFPC detection signal are reported, it can be concluded that the analysed sample was negative. The positive direct detection readout may be caused by the spurious hybridization or the non-specific bindings.

In a fourth exemplary scenario, when a negative direct detection response and a positive CFPC detection signal are reported, it can be inferred that there are only trace amounts of the target substance, e.g. SARS-CoV-2 viral RNA in the sample.

In an embodiment, the method of the present invention is one, wherein after step (vii), the method further comprises the step of obtaining an intensity spectrum of the radiation beam component by capturing an interference spectrum, wherein the capturing comprises:

- measuring the spectral interferential intensity for each wavelength; and

- calculating the LSPR response based on an indicator of phase change, intensity change, wavelength shift or a combination thereof.

In an embodiment, the method of the present invention is one, wherein in step (iv) and/or (viii-c):

- the monochromatic light beam is converted into a plasmonic photothermal heating at the peak absorption wavelength of the LSPR matrix; and

- the polychromatic radiation beam is modulated with a localized surface plasmon resonance effect at a wavelength different from the wavelength of the monochromatic light beam.

In an embodiment, the method of the present invention is one, wherein the output power levels of the monochromatic light beam are adjusted to alter the temperature generated by the plasmonic photothermal effect.

Power levels for use in the present method can be, e.g. in the range of about 0 to 10 W.

In an embodiment, the method of the present invention is one, wherein the method further comprises: detecting a specific target substance at a specific, optionally elevated, temperature; and discriminating different target substances at an adjustable temperature range.

The term "elevated temperature" as used herein refers to a temperature above room temperature, optionally to a temperature between and including room temperature and 95 °C or 60 °C, or between and including room temperature and 33 °C, or to a temperature between and including 33 °C and 45.6 °C, optionally between and including room temperature and 41 °C or between and including 33.2 °C and 41 °C or 41.5 °C.

For example, the specific target RdRp-COVID sequence can be sensitively detected at an elevated PPT-induced temperature below the melting temperature of the sequence. The similar sequence RdRp-SARS, or any other similar sequences with one or more base-variations can be discriminated by adjusting the PPT induced temperature to approach the melting temperature of the similar sequence.

In an embodiment, the method of the present invention is one, wherein the target sample is a nucleic acid, e.g. DNA or RNA, optionally an RNA, optionally a viral RNA, optionally a positive sense, single stranded viral RNA, or a virus, e.g. a corona virus, optionally SARS-CoV or SARS-CoV-2, optionally wherein the target sample comprises a viral RNA-dependent RNA polymerase (RdRp) sequence, optionally selected from SEQ ID NO. 1 (RdRp-COVID) and 2 (RdRp-SARS), a viral open reading frame lab (ORFlab) sequence, optionally SEQ ID NO. 3 (ORFlab-COVID), or non-structural protein 13 (nspl3) nucleic acid sequence having SEQ ID NO. 4.

The following Figures and Examples serve to illustrate the invention and are not intended to limit the scope of the invention as described in the appended claims.

Fig. 1 is a block diagram of a dual-functional plasmonic detector according to the present invention.

Fig. 2 is a schematic view of the photothermal unit illustrated in FIG.l, which consists of a monochromatic light source and an objective lens.

Fig. 3 is a schematic view of the plasmonic sensing unit according to the present invention.

Fig. 4 is a schematic view of a detection unit to analyze the dual-functional plasmonic sensing results.

Fig. 5 is a schematic view of the detector for photothermal enhanced plasmonic biosensing.

Fig. 6 shows an experimental setup and system optimization (a) Schematic and (b) experimental setup of the dual-functional PPT enhanced LSPR biosensing system. In the LSPR sensing path, the collimated wide spectrum beam passes through the aperture-iris (11/12), the linear polarizers (P1/P2), the birefringent crystal (BC), and totally reflected at the interface of AuNIs- dielectric for LSPR detection. In the excitation unit, a laser diode (LD) generates the PPT effect on the AuNIs in the normal incident angle (c) and (d) demonstrate the normalized absorbances of the AuNIs sensor chips showing a fine-tune peak absorption from 523.4 nm to 539.7 nm (± 0.2 nm). (e) Plasmonic resonance wavelength at about 580 nm under the ATR (attenuated total reflection) configuration for LSPR sensing transduction. Fig. 7 shows the in situ characterization of local PPT heating on AuNIs. (a) Periodic laser excitation and the PPT induced plasmonic phase response (b) Temperature variations and real-time LSPR responses (c) Calibration curve illustrating the relationship between the temperature and LSPR phase response (d) Real-time LSPR responses caused by the laser induced PPT effect under different laser powers (e) Scanned local LSPR responses around the PPT heat source on AuNIs. (d) Mapping the temperature distribution around the PPT heat source.

Fig. 8 shows selected viral sequences for SARS-CoV-2 detection (a) Selected sequences and their relative positions used for SARS-CoV-2 and SARS-CoV detection. M: membrane protein gene;

N: nucleocapsid protein gene; S: spike protein gene. The numbers below the sequences are genome positions according to GenBank, SARS-CoV-2 NC_045512. (b) Schematic illustration of AuNIs functionalization based on the reaction with thiol-cDNA ligands (c). Real-time monitoring of AuNIs functionalization dynamics. 10 mί, solution containing 0.1 nmol cDNA was injected in each step (d) Calibrated surface functionalization efficiency to retrieve the optimal cDNA amount.

Fig. 9 shows PPT enhancement in LSPR biosensing (a) Schematic illustration of the hybridization of two complementary strands with SEQ ID NOs. 1 (RdRp-COVID) and 2 (RdRp-COVID- C). (b) Real-time hybridization of RdRp-COVID and its cDNA sequence (RdRp-COVID-C) with or without the thermoplasmonic enhancement (c) PPT enhancement on RdRp-COVID sequence detection at different concentrations. The error bars refer to the standard deviations of LSPR responses after reaching steady conditions following the buffer flushing (d) Schematic illustration of inhibited hybridization of two partially matched sequences (SEQ ID NOs. 2 (RdRp-SARS) and 6 (RdRp-SARS-C)). The red-arrows indicate the mismatch bases of RdRp-SARS and functionalized cDNA of RdRp-COVID. (e) Discrimination of two similar sequences with PPT heat. The laser was applied at 200 s and switched off at 700 s. (f) The RdRp-SARS sequence dissociation from the immobilized RdRp-COVID-C sequence. The original phase responses (red dots) and the corresponding smoothed means (black curve) are shown.

Fig. 10 shows the evaluation of the dual-functional LSPR biosensor performance on detecting viral nucleic acids (a) The plot of LSPR phase responses versus RdRp-COVID oligos concentrations using the PPT enhanced LSPR biosensor (b) Zoom-in view of the low concentration range for LoD identification (c) Concentrations of various viral oligos measured using the dual-functional LSPR biosensors (d) Detection comparison of single analyte RdRp-COVID and mixture of multiple sequences. The error bars refer to the standard deviations of LSPR responses after reaching steady conditions following the buffer flushing.

Fig. 11 shows the absorbance spectra of AuNIs. The peak wavelengths of the AuNIs spectral absorbances, or optical density (OD) were optimized in a range from 523.4 nm to 539.6 nm. The typical absorbance ranges from 0.335 (absorbance unit) AU to 0.361 AU, with a median of 0.354 AU.

Fig. 12 shows the temperature profiles of PPT heating (a) Real-time LSPR responses of laser off-toon switching. In the thermoplasmonic heating processes, the rapid "heat-up" in step-i can be completed within one second (the acquisition time was 1 second). The subsequent "dynamic equilibrium" in step-ii took about 11 seconds before finally entering the temperature "steady" condition in step-iii. (b) Zoom-in responses from the temperature "dynamic equilibrium" (step-ii) to "steady" state (step-iii). (d) Real-time LSPR responses of laser on-to-off switching. In the cooling process, the rapid "cooling-down" process also finishes within one second. The following "dynamic equilibrium" of cooling to room temperature requires less than 5 seconds.

Fig. 13 shows the PPT heating system for characterizing the temperature distribution (a) PPT unit with a 532 nm laser was applied to the LSPR sensor chamber. Under the normal incident angle, AuNIs showed a peak absorption at 532.2 ± 0.2 nm. (b) The side-view of the AuNIs sensing chamber. The total reflected light from the AuNIs-water interface showed a red-orange color due to the plasmonic absorption, i.e. the peak absorbance at 532.2 ± 0.2 nm. (c) Scanning demonstration for mapping the PPT temperature distribution around the laser spot. The ATR signals from the optimal AuNIs chip were analysed by the LSPR sensing system pixel by pixel (with 0.5 mm interval). At each detection point, the excitation laser was switched off-to-on repeatedly and two interferometric spectra were recorded accordingly. Then, the phase differences between "on" and "off" states were calculated to retrieve the PPT induced LSPR response and map the PPT heating distribution.

Fig. 14 shows the microfluidic dual-functional LSPR detection system (a) Dual-functional plasmonic sensing system with a microfluidic sensing chamber (b) The microscopic image of AuNIs microfluidic sensing cell. The AuNIs sensing chip was bonded to the PDMS embedded in the microfluidic channels. The viral sequences detections were mainly conducted based on the dual functional plasmonic system with a microfluidic sensing chamber. Before injecting into the detection chamber, the solution passed through a serpentine microchannel to further disperse the target sequences. The excitation laser (532 nm with 32 mW optical power) was applied to the center of the microsensing channel (2 mm in width). In the LSPR transduction system, the aperture- iris (0.5 mm in diameter) screened the sensing beam to make the spectrometer only receiving the effective sensing light from the microfluidic channel.

Fig. 15 shows a comparison of AuNIs surface functionalization. The AuNIs chip surface was oversaturated when functionalized with 200pL 50 mM cDNA (dotted line). The AuNIs chip surface was insufficiently functionalized by using 200pL 0.5 mM cDNA (solid line). Fig. 16 shows the PPT effects on the real-time LSPR detection (a) With the in situ PPT enhancement, the RdRp-COVID sequences at concentration levels from 1 pM to 1 nM were detected by the AuNIs biosensors with the immobilized RdRp-COVID-C probes (b) Without the PPT enhancement, the same concentration levels were tested. Based on these comparison results, the hybridization rates and the LSPR sensing response levels were obviously suppressed when the PPT unit was shut down (c) PPT heating enhancement on the association rate constants k_a of the RdRp- COVID sequences.

Fig. 17 shows the discrimination of two similar sequences without the PPT heat. Real-time LSPR sensorgram of detecting (a) specific RdRp-COVID sequences and (b) non-complementary RdRp-SARS sequences by using the RdRp-COVID-C.

Fig. 18 shows the dissociation rate constant of RdRp-SARS. (a) Dissociation of RdRp-SARS from the RdRp-COVID-C probes (b) Dissociation of RdRp-COVID from the RdRp-COVID-C probes.

Fig. 19 shows the blank measurement for LoD. The distribution fitting of blank measurement signals based on the water buffer solution. The mean response level was 2.92 x 1CT³ radian, with a standard deviation of 3.118 x 1CT³ radian.

Fig. 20 shows the selected nsp-13 probes, viral target sequence and their relative positions in SARS-CoV-2 (NC_045512.2, Genbank) corresponding to SEQ ID NOs. 4, 8 and 11.

Fig. 21a is a schematic view of the cyclic fluorescent probe cleavage based biosensing flowchart according to the present invention. Fig 21b shows a combined detection based on LSPR detection and CFPC detection (i.e. the detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe): (A) Target substance (e.g. virus), (B) target substance (e.g. RNA), (C) plasmonic sensing unit 200 (e.g. gold AuNI with nucleic acid probe attached), (D) LSPR detection, (E) cleavage enzyme (e.g. Endo-IV), (F) further fluorescent single stranded nucleic acid, (G) CFPC detection.

Fig. 22 shows the real-time sensorgrams of the target sequence direct LSPR detection, with a series of concentrations from 0.1 pM to 10 nM.

Fig. 23 shows the real-time sensorgrams of the CFPC detection (i.e. the detection based on the fluorescent nucleic acid which is cleaved after binding to the target substance already bound to the probe) by using the Endo-IV and fluorescent probe, with a series of concentrations from 0.01 pM to 10 nM.

Fig. 24a shows the calibrated regression curve of the direct LSPR detection responses versus the target sequence concentrations using the detector of the invention. Fig. 24b shows the calibrated regression curve of the CFPC detection responses versus the target sequence concentrations using the detector of the present invention. Fig. 25 shows the real-time differential phase responses stimulated by increased irradiation power. The laser power increased step-by-step from 13.44 W/cm² to 94.56 W/cm² with 15 complete laser on/off irradiation cycles.

Fig. 26 shows a real-time LSPR phase change triggered by the hybridization of the nucleic acid probe and complementary target.

Fig. 27 shows the phase responses caused by nucleic acid dehybridization at different temperatures. When the AuNI thermoplasmonic field generated local temperatures higher than 42.5 °C, part of the hybridized complementary strands underwent thermal dissociation.

Fig. 28 shows the transient phase response signal obtained by subtracting the accumulated change. The round scattered points (ls-500s) represent the background signal without laser irradiation, and the triangular scattered points (500s-1200s) represent the transient phase responses stimulated by the near-field cyclic released fluorophores.

Fig. 29 shows he real-time sensorgram of the cyclic amplification, with a typical laser-off background and an amplified phase response under thermoplasmonic assisted CFPC detection.

Fig. 30 shows the direct detection of gain molecules (Rhodamine 6G) to investigate the amplification efficiency of gain-assisted plasmonic detection.

Example 1: Materials

All chemicals were purchased from commercial suppliers and used without further purification. Nuclease-free water was purchased from ThermoFisher and used as the buffer for oligonucleotides dilution and LSPR detection. All selected oligonucleotides, including the RdRp- COVID, RdRp-SARS, ORFlab-COVID, E, non-structural protein (nsp) gene sequence, their thiol-cDNA probes, including the RdRp-COVID-C, RdRp-SARS-C, ORFlab-COVID-C, E-C, Nspl3-C, and fluorescent probe were synthesized and provided by Microsynth (Balgach, Switzerland). Endo-IV (2 U/pL, EN0591, Thermo-Fisher) was prepared in the reaction buffer (with 50 mM Tris-acetate, 50 mM NaCI, 50 mM KCI, ImM EDTA, 0.05 v/v% TritonX-100, pH 7.5). All AuNIs sensor chips and fluidic sensing chambers were cleaned using absolute ethanol followed by rinsing with Milli-Q water before testing.

Example 2: Synthesis of dual-functional AuNIs chip.

The AuNIs sensor chips were synthesized based on the self-assembly process of thermal dewetted Au nanofilm. The original magnetron sputtered Au nanofilms were optimized in a thickness range from 5.0 nm to 5.2 nm. Then the Au nanofilm was thermal annealed at 550 °C for 3 hours. The AuNIs were self-assembled on the BK7 glass surface. The visible light absorption of each AuNIs sensor chips was measured to retrieve the optimal plasmonic resonance condition.

Example 3: Dual-functional LSPR system. In the interferometric LSPR phase sensing system, a white light sensing beam was generated by a LED source and subsequently linearly polarized by a polarizer (PI). The thin birefringent crystal (BC) added sufficient retardation into the two orthogonal components of the linearly polarized light, i.e. the s- and p-components to create the spectral interferogram. The BK7 prism was able to couple the incident light into the AuNIs dielectric interface at an inclined nominal incident angle of 72° and excited the local electromagnetic fields in the vicinity of the AuNIs by the Kretschmann configuration. The plasmonic resonance wavelength for LSPR sensing transduction was found to be 580 nm. The interferometric spectra were screened by an aperture-iris (11/12, Thorlabs) with a hole diameter of 0.5 mm and finally recorded by the spectrometer (AvaSpec, Avantes). In addition to this plasmonic transducing unit, a high-power 532 nm laser diode (LD, 532 nm peak wavelength, DJ532- 40 Thorlabs) was used for PPT heating by illuminating the AuNIs chips in the normal incident angle. A long-wavelength pass filter (LPF, 552 nm cut-on wavelength) was used to block the excitation signal before the spectrometer. The ambient temperature was measured and recorded with digital temperature sensors (SHTC1, Sensirion) for LSPR-temperature calibration.

Example 4: Surface functionalization with thiol-cDNA.

The AuNIs surface functionalization was investigated based on the step-by-step injection of O.lnmol thiol-cDNA. In the sensing chamber, 90 pL nuclease-free water was initially injected to build the phase reference baseline for 400 s. Then, each time a 10 pL solution which contained 0.1 nmol thiol-cDNA, e.g. the RdRp-COVID-C sequence was injected into the sensor chamber in every 200 s, until no further phase changes were recorded. Based on the optimal result, the solution containing 1 nmol cDNA was utilized to functionalize the AuNIs chips for the following SARS-CoV-2 sequences detection.

Example 5: Detection of SARS-CoV-2 viral sequences.

After the probe immobilization, the desired concentration of target DNA in nuclease-free water (200 pL) was introduced into the AuNIs microfluidic chamber for 800 s, and the hybridization reaction was allowed under the PPT heat (32 mW optical power at 532 nm). In the LSPR sensing path, an aperture-iris with a hole diameter of 0.5 mm was used to screen the sensing beam entering the spectrometer, which corresponded to the ATR light from the center of the PPT heat. Experiment on the mismatched nucleic acids and multi-sequence mixtures were also conducted based on the dual-functional LSPR biosensors as described above. A stringent buffer flushing with nuclease-free water was conducted after the hybridization. The whole testing process was real-time recorded by the spectrometer for plasmonic phase detection.

Example 6: Exemplary device and method for SARS-CoV-2 detection. The dual-functional plasmonic performances were systematically studied in the aspects of LSPR sensing transduction and PPT heating. The common-path differential phase-sensitive LSPR system, as shown in Figure 6a, was adopted to measure the local refractive index changes or the binding events. In the LSPR sensing transduction unit, the sensing beam was generated by a wide spectrum LED source and operated in the ATR (attenuated total reflection) mode at the interface between the glass substrate and liquid environment. When reaching the two-dimensional AuNIs sensing layer, the measured optical power of the beam was found to be 32.58 pW. The local plasmonic responses were retrieved from the ATR spectral interferograms by using the windowed Fourier transform phase extraction method, as described elsewhere (Qiu et al., Differential Phase- Detecting Localized Surface Plasmon Resonance Sensor with Self-Assembly Gold Nano-Islands. Opt. Lett. 2015, 40, 1924-1927.) This phase response, reported in the unit of radian, is more prominent than the conventional spectral and angular responses. Therefore, it has been utilized for improving the sensitivity of plasmonic sensors (Smolyaninov et al., Programmable Plasmonic Phase Modulation of Free-Space Wavefronts at Gigahertz Rates. Nat. Photonics 2019, 13, 431-435). In order to generate a stable and intense thermoplasmonic field, an excitation laser with 532 nm peak wavelength and 40 mW maximum optical power was applied onto the AuNIs chip in the normal incident angle (Figure 6b). In addition, optimizing the AuNIs chip so that its peak absorbance wavelength was exactly at 532 nm can significantly improve the conversion efficiency of thermoplasmonic. By adjusting the Au nanofilm thickness before dewetting, the absorption peak (under normal incident angle) can be accurately controlled within a wavelength range from 523.4 nm to 539.7 nm as shown in Figure 6c, 6d, and Figure 11. In this work, the AuNIs that matched the laser excitation wavelength at 532.2 nm (± 0.2 nm) were utilized for the PPT heating (Chen et al., Imaging Local Heating and Thermal Diffusion of Nanomaterials with Plasmonic Thermal Microscopy. ACS Nano 2015, 9, 11574-11581). It is worth noting that under the ATR condition with a 72° inclined incident angle, the plasmonic resonance wavelength for LSPR sensing transduction red-shifted to 580 nm due to the prism coupling and the inclined angle of incidence (Figure 6e).30 The phase changes caused by a local variation of LSPR conditions were confined in a narrow wavelength region from 578 nm to 582 nm. Moreover, after adding a long-pass filter (LPF) with a cut-on wavelength at 552 nm, the 532 nm photothermal excitation laser from the PPT unit did not influence the stability of the realtime LSPR sensing transduction.

In the thermoplasmonic testing, the direct absorption of laser irradiation at 532 nm decayed nonradiatively by generating more hot electrons in AuNIs (Baffou et al., Nanoscale Control of Optical Heating in Complex Plasmonic Systems. ACS Nano 2010, 4, 709-716). The photoexcited highly energetic electrons quickly dissipated and released thermal energies to heat the ambient environments. Conversely, the PPT induced temperature increase was also responsible for a refractive index variation of the surrounding environment, which can be in situ detected by the LSPR detection system as shown in Figure 7a. Specifically, the AuNIs chip was exposed to laser excitation for 50 s, as indicated by the shaded region. Then the laser was switched off to re-attain the baseline. The generation and equilibrium of local photothermal heating were relatively fast. According to the laser switching tests as shown in Figure 12, the rapid heating process was completed within one second after turning on the laser excitation. Subsequently, the dynamic equilibrium process took another 11 s before finally entering the steady state. In the examples, the LSPR phase response was calibrated under different ambient temperatures. The in situ temperature arising from the PPT effect was characterized based on the measurement of the thermal-induced refractive index variation in the vicinity of AuNIs (Chen et al., Imaging Local Heating and Thermal Diffusion of Nanomaterials with Plasmonic Thermal Microscopy. ACS Nano 2015, 9, 11574-11581 and Baffou et al., Thermal Imaging of Nanostructures by Quantitative Optical Phase Analysis. ACS Nano 2012, 6, 2452-2458). During the ambient temperature variation, the real-time LSPR phase responses and temperature values were recorded in parallel (Figure 7b), and the correlation was established as shown in Figure 7c. Based on this calibrated LSPR-temperature regression, the localized photothermal temperatures under different laser powers were retrieved as shown in Figure 7d.

To further evaluate the laser induced PPT effect and the local temperature profile, the spectrometer was used to scan the heating area for mapping the LSPR phase responses and actual temperature distribution on the AuNIs sensor chips. In the experimental setup as shown in Figure 13, the excitation laser with 32 mW power was applied to the optimal AuNIs absorbers with a peak absorption at 532.2 nm (± 0.2 nm). At each point of interest, the LSPR transducing unit was used to record two interferometric spectra: one reference without PPT heating and one spectrum with PPT heating. By scanning the laser spot and surrounding area with a 0.5 mm step interval, the spatial distribution of LSPR phase changes was retrieved as shown in Figure 7e. At each scanning pixel, the retrieved phase response was subsequently converted to the local temperature based on the calibration curve in Figure 7c. Therefore, the corresponding temperature distribution around the PPT heating was obtained and illustrated in Figure 7f. The local temperature was significantly elevated from 21.47 °C (room temperature) to 41.08 °C at the center of the laser spot.

The full genome sequence data of the viruses, i.e. SARSCoV-2 and SARS-CoV, have been retrieved from the GISAID platform. The selected oligonucleotides for specific SARS-CoV-2 detection and their relative positions were given in Figure 8a and Table SI. These viral oligonucleotides refer to sequences used in different countries for COVID-19 diagnosis, and some of them have been published in the latest researches (Corman et al., Detection of 2019 Novel Coronavirus (2019-nCoV) by Real-Time RT-PCR. Eurosurveillance 2020, 25; World Health Organization, Laboratory Testing for Coronavirus Disease 2019 (Covid-19) in Suspected Human Cases: Interim Guidance, 2 March 2020; WHO/COVID-19/laboratory/2020.4; World Health Organization: Geneva, 2020. https://www.who.int/emergencies/diseases/novel-coronavirus-

2019/technicalguidance/laboratory-guidance (accessed March 2, 2020) and Zhu et al., China Novel, C, A Novel Coronavirus from Patients with Pneumonia in China, 2019. N. Engl. J. Med. 2020, 382, 727-733). The basic local alignment search tool (BLAST) was used to compare these viral sequences with the library of SARS-CoV-2 to confirm their representativeness and specificity. In the present case of COVID-19, SARS-CoV-2 isolates or samples from infected patients are challenging to obtain and handle. Thus, the corresponding DNA sequences were artificially synthesized for representative LSPR sensing demonstration of SARS-CoV-2 and SARS-CoV. According to the WHO guideline and local alignment searching results, two specific sequences from SARS-CoV-2 were selected, i.e. the RdRp and the ORFlab as shown in Figure 8a. Validation and proof of selectivity were demonstrated by choosing the closely related nucleic acid sequence from RdRp of SARS-CoV. In addition, an oligonucleotide sequence from the coronaviral envelope protein gene (E) was also synthesized and tested to aid the virus identification.

Based on the synthetic oligonucleotide probes with a thiol group (Table S2), the LSPR sensing chips were directly functionalized by forming the Au-S bond between the thiol-cDNA probe and AuNIs as illustrated in Figure 8b. The surface functionalization process was initially optimized on its amount and concentration in order to achieve proper surface coverage and high sensitivity. During the real-time surface functionalization as shown in Figure 8c, step-by-step injections of 0.1 nmol thiol-cDNA of RdRp-COVID, (RdRp-COVIDC) caused continuous phase jumps due to the covalent binding between AuNIs and thiol-cDNA. After a total immobilization of 1 nmol (10 x 0.1 nmol) RdRp- COVID-C as shown in Figure 8c and 8d, the LSPR response stopped growing and indicated the appropriate amount of cDNA probes for AuNIs functionalization. Hereafter, the solution containing 1 nmol thiol-cDNA was utilized to functionalize the AuNIs microfluidic sensor chips for SARS-CoV-2 sequences detection (Figure 14). The proper surface functionalization that is sufficient to functionalize the entire AuNIs sensing surface can increase the sensitivity and suppress the non specific binding events. In contrast, the AuNIs sensor chip was over-saturated when functionalized with 10 nmol cDNA, and insufficiently covered by using 0.1 nmol cDNA (Figure 15).

The surface functionalized AuNIs chips were subsequently installed in the LSPR systems for specific viral sequences detection (Figure 9a). The impacts of the localized thermoplasmonic heating on nucleic acids hybridization and LSPR detection were systematically studied. According to the temperature profile shown in Figure 7f, the excited PPT heat with approximately 41 °C nominal temperature was generated on the AuNIs sensor. Before the injection of RdRp sequence, nuclease- free water was flown across the microfluidic sensing chamber and the thermoplasmonic laser (32 mW) was turned on to establish a steady phase reference and baseline. According to the phase sensing diagram in Figure 9b and 16a, the LSPR response of the dual-functional AuNIs biosensor started to increase when the RdRp-COVID genes were injected into the microfluidic chamber at about 200 s and attained the maximum phase value after about 800 s hybridization. The dual functional AuNIs sensing chip was further flushed with the nuclease-free water to remove the non specific binding items and to check the final LSPR phase response. In the comparison with and without the PPT effect, the hybridization rate and the LSPR sensing response level were obviously suppressed when the PPT unit was shut down as shown in Figure 9b. It proved that the localized photothermal effect can significantly improve the hybridization kinetics of the RdRp-COVID and its cDNA. Thus, the response-slope of the photothermal enhanced LSPR was much steeper than that without the photothermal assistance. Due to the faster hybridization kinetics, the differential phase response levels were also elevated for the RdRp-COVID sequence at different concentrations as shown in Figure 9c and Figure 16. The PPT effect and its derived local heat can effectively promote the fast and sensitive detection of nucleic acids by improving the hybridization kinetics of fully matching strands.

More importantly, the PPT heating was capable to inhibit the spurious binding of nonmatching sequences by elevating the local temperature at the vicinity of AuNIs. SARS-CoV and SARS-CoV-2 viruses are similar beta-coronavirus and their genetic similarities are high. The specific SARS-CoV-2 genetic target recommended by WHO, i.e. the RdRp-COVID sequence as shown in Table SI, is very closely related to that of SARS-CoV. Specifically, in the selected gene sequences, only three fixed nucleotide bases were different between RdRp-COVID and RdRp-SARS. A real-time LSPR detection was conducted on the two closely related sequences. The LSPR sensor without the aid of photothermal unit reported a false positive response signal when detecting the RdRp-SARS sequence (Figure 17), which indicated that a similar but not fully complement sequence was also able to interact and partially hybridize with the cDNA probes at room temperature. Although the hybridization kinetics of RdRp-SARS sequence from SARS-CoV was clearly slower than that of SARSCoV-2, the non-matching spurious binding of any closely related sequence can affect the accurate virus detection and discrimination. Therefore, the local heat based on the proposed PPT effect was employed to improve the specificity of hybridization. At the elevated temperature of 41 °C as illustrated in Figure 9d, the standard free energy of hybridization was weaker due to the mismatched base-pairs. Thus, the similar but not fully matched sequences of SARS-CoV can be distinguished. In detail, the calculated association rate constant k_a Of RdRp-COVID with PPT heating enhancement was found to be 1.11 x 10^s M¹ s¹. The detailed discussion and calculation were given in Figure 16c. For a typical biological sensing system, k_a ranges between 10³ to 10⁷ M ¹ s ¹ and a higher associate rate indicates a stronger binding affinity (Xu et al., Real-Time Reliable Determination of Binding Kinetics of DNA Hybridization Using a Multi-Channel Graphene Biosensor. Nat. Commun. 2017, 8 and Schreiber et al., Fundamental Aspects of Protein-Protein Association Kinetics. Chem. Rev. 2009, 109, 839-860). In a comparison experiment including the PPT heat with 32 mW optical power, the 532 nm laser was applied onto the surface of the AuNIs sensor from 200 s to 700 s as shown in Figure 4e. The local PPT heat was generated immediately to make the LSPR phase jump to about 1.76 radian. After turning off the laser at 700 s, the LSPR phase response of the mismatching RdRp-SARS gene was fully suppressed to the ground state of blank measurement ( i.e . the responses from 0 to 200 s) as shown by the black curve in Figure 9e. Since the RdRp-SARS sequences reported a weak response of 0.002 radian, it was determined that its association rate constant was lower than 10³ M ¹ s ¹ under the PPT heating. At the same time the fully matching RdRp-COVID sequence from SARS-CoV-2, showed an apparent phase difference before and after the laser excitation (orange curve in Figure 4e). Thus, it is believed that a similar but not fully matched sequence could be distinguished based on their different binding affinity and the PPT heating.

In another set of verification experiments, the RdRp-SARS genes were initially bound to the RdRp-COVID-C probes at room temperature. Then, the 532 nm laser (32 mW) was applied on the AuNIs surface to stimulate the local thermoplasmonic effect. In the real-time LSPR sensorgram showed in Figure 9f, the dissociation of the RdRp-SARS genes from the RdRp-COVID-C probes was observed after the temperature rise. The calculated dissociation rate constant was 8.287 ^c 10³ s ¹ as shown in Figure 18. The dissociation half-life ti/2 which indicated the time to dissociate half of the hybridized sequences was 83.3 s. In contrast, the complementary sequence of RdRp-COVID showed a much lower dissociation rate constant at 3.5 ^c 10⁶ s ¹ and long dissociation half-life time of 1.97 x 10⁵ s. These results further verified that the thermoplasmonic effect can eliminate the non-matching hybridization quickly and promote the selective detection of the target sequence, so as to achieve highly accurate nucleic acid detection and virus differentiation. Compared with the conventional plasmonic biosensing system, it was demonstrated how this proposed dual-functional plasmonic sensing system can be the basis of a reliable and easy-to-implement thermoplasmonic biosensing technique: it can significantly reduce the false-positive-rate and enhance the reliability in genetic diagnosis. To quantify the sensing performance, the dual-functional plasmonic detections of RdRp-

COVID were further investigated over the concentration range from 0.01 pM to 50 mM as shown in Figure 10a. The AuNIs sensing system started to attain the saturation condition when the concentration of the RdRp-COVID sequence reached 1 mM. In contrast, the low RdRp-COVID concentration, i.e. 0.1 pM, only resulted in a weak phase response by 2.90 x 1CT³ radian (Figure 10b), which was close to the system blank measurement of 2.92 x 1CT³ radian. Thus, as illustrated in the sensing calibration curve in Figure 5b, the dual-functional LSPR sensing system exhibited a limit of the range from 0.1 pM to 1 mM for detecting oligonucleotides, covering seven orders of magnitude. The calibrated regression curve was further used to estimate the limit of detection (LoD), which is defined by lUPAC (International Union of Pure and Applied Chemistry) as the sum of the blank measures, i.e., 2.92 x 1CT³ radian with the nuclease-free water buffer and triple of its standard deviation (Figure 19). Thus, the LoD of the photothermal enhanced LSPR sensing system was found to be (2.92 x 1CT³) + 3 x (3.12 x 1CT³) = 0.0123 radian as shown by the dashed line in Figure 10b. Therefore, the detectable RdRp-COVID sequence concentration corresponding to the systematic LoD was about 0.22 ± 0.08 pM (Figure S9). 200 mί analyte solution at this LoD concentration contained about 2.26 x 10⁷ copies of the RdRp-COVID sequence. The actual size of SARS-CoV-2 is about 29.9 kilobases in length, which is 1000 times longer than the RdRp-COVID sequence used in this study. Thus, based on the LSPR signal - target size relationship, the estimated LoD for detecting the entire RNA strands from SARS-CoV-2 could be approximately 2.26 x 10⁴ copies (Qiu et al., Total Bioaerosols Detection by a Succinimidyl-Ester-Functionalized Plasmonic Biosensor to Reveal Different Characteristics at Three Locations in Switzerland. Environ. Sci. Technol. 2020, 54, 1353-1362). A recent study reported the viral loads of SARS-CoV-2 from different respiratory trace samples including the throat/nasal swabs and the sputum. Based on these clinical specimens collected from 82 infected individuals, the overall viral load soon after onset was higher than 1 x 10^s copies/mL.39 This indicated that the dual-functional detector of the present invention has the potential for direct analysis of SARS-CoV-2 sequences in respiratory samples.

In addition to the RdRp-COVID sequence, the dual-functional detector was also validated by performing the selective hybridization detection on several different genome sequences from both SARS-CoV-2 and SARS-CoV, i.e. the ORFlab-COVID sequence and the E sequence from SARS-CoV-2, the RdRp-SARS sequence from SARS-CoV. The corresponding LSPR phase sensing responses with the in situ PPT enhancement were illustrated in Figure 10c. The complementary cDNA sequences, i.e. ORFlab-COVID-C, E-C, and RdRp-SARS-C were functionalized onto the AuNIs chips respectively, for the detection of specific viral sequence. Since the physical length and molecular weight were roughly same, the hybridization of these target sequences reported a similar LSPR phase response (Figure 10c). As the concentration increased from 1 pM to 1 nM, the mean LSPR response levels of each sequence also climbed in a proportional manner, which further proved the feasibility of this dual-functional LSPR sensing system for quantitative analysis of viral nucleic acids. Among them, the ORFlab-COVID sequence produced the strongest responses due to its high molecular weight (8715.6 g/mol) and long length (28 bases). While the responses for E sequence were slightly lower.

In clinical diagnosis, the respiratory trace samples after viral lysis and RNA extraction may contain multiple nucleic acid sequences from the same viral source of SARS-CoV-2. Thus, detecting the accurate concentration of a specific sequence under the interference of multiple non-specific sequences was beneficial to demonstrate its potential for real clinical applications. In the experiments as shown in Figure lOd, the multi-sequences mixture containing RdRp-COVID sequences (100 pM), E sequences (100 pM), and ORFlab-COVID sequences (100 pM), was prepared to simulate an actual sample after virus lysis. The ORFlab-COVID and E sequences in the mixture showed extreme low spurious binding with the immobilized RdRp-COVID-C probes. Compared with the standard detection of 100 pM RdRp-COVID as shown in Figure lOd, the calculated recovery rate based on the dual-functional LSPR biosensors was found to be 96% in the mixture sample. This experimental result further demonstrated that the dual-functional detector of the invention with the in situ PPT enhancement can perform accurate detection of the target sequence and facilitate the highly accurate SARS-CoV-2 detection.

Example 7: Selected target sequences from SARS-CoV-2 and SARS-CoV

Table SI. Selected target-sequences from SARS-Cov-2 and SARS-CoV.

Name of the Target Sequence Molecular Length Melting Oligonucleoweight (bases) temperature, tide (g/mol) (NN method °C)

RdRp-COVID 5'-CAGGT GGAAC CTCAT 7716 25 59.5

SEQ ID NO: 1 CAGGA GATGC-3' RdRp-SARS 5'-C CAGGT GGAAC ATCAT 7996.2 26 61 SEQ ID NO: 2 CCGGT GATGC-3' ORFlab-COVID 5'-CCGTC TGCGG TATGT 8715.6 28 61.8 SEQ ID NO: 3 GGAAA GGTTA TGG-3' E 5'-ACAC TAGCC ATCCT 7842.1 26 63

SEQ ID NO: 9 TACTG CGCTT CG-3' Nsp-13 5'- CTACTGTACG 20774.5 67 68.7 SEQ ID NO: 4 TGAAGTGCTG TCTGACAGAG AATTACATCT TTCATGGGAA GTTGGTAAAC CTAGACC-3'

*ORF: open reading frame; RdRp: RNA-dependent RNA polymerase gene; E: envelope protein gene; M: membrane protein gene; N: nucleocapsid protein gene; S: spike protein gene.

RdRp sequences selected from SARS-CoV and SARS-CoV-2 represented two similar related sequences from the beta-coronavirus family. These two sequences, i.e. the RdRp-COVID and RdRp- SARS were utilized to exam the discrimination performance of dual-functional LSPR biosensors. According to the latest version of "WHO interim guidance for laboratory testing for 2019 novel coronavirus (COVID-19) in humans" (accessed 2 March 2020), these selected sequences can be used as the probe-target in RT-PCR for the clinical diagnosis of the coronavirus induced diseases. Among them, RdRp-COVID and ORFlab-COVID may detect only the novel SARS-CoV-2 virus, RdRp-SARS only targets on the SARS-CoV, while E sequence is applicable to both viral strains (i.e. SARS-CoV and SARS-CoV-2). The sequences' information, including the molecular weight, oligonucleotide length, and the calculated melting temperature were also provided in the table. The melting temperatures were calculated using the nearest-neighbour (NN) method.

Table S2. Selected target sequences and their complementary thiol-cDNA for LSPR functionalization

The selected complementary sequences, including the RdRp-COVID-C, ORFlab-COVID-C, RdRp-SARS-C, E-C and Nsp-13-C were synthesized with a thiol group (HS-) at the head (5'-) of each strand. In the surface functionalization step, the prepared cDNA solution (diluted in nuclease-free water) was directly injected into the sensor chamber and reacted with the dual-functional AuNIs chip. The thiol group can covalently bind to the AuNIs surface by forming Au-S bonds, and the whole single-stranded cDNA probes were also anchored on the two-dimensional distributed AuNIs. Example 8: Microfluidic dual-function LSPR detection system.

The viral sequences detections were mainly conducted based on the dual-functional plasmonic system with a microfluidic sensing chamber. Before injecting into the detection chamber, the solution passed through a serpentine microchannel to further disperse the target sequences. The excitation laser (532 nm with 32 mW optical power) was applied to the center of the micro sensing channel (2mm in width). In the LSPR transduction system, the aperture-iris (0.5 mm in diameter) screened the sensing beam to make the spectrometer only receiving the effective sensing light from the microfluidic channel.

Example 9: Comparison of AuNIs surface functionalization.

In the comparison experiments of different surface functionalization, two concentrations, i.e. 200 pL 50 mM and 200 pL 0.5 mM RdRp-COVID-C sequence solutions were used as the negative references. When functionalized with the 50 mM RdRp-COVID-C oligonucleotide solution, the AuNIs surface was fully covered by the thiol-cDNA ligand within 200 sec. The saturation condition reported a maximum response at 2.84 radian. In this case, most of the cDNA sequences were wasted as it could not continue to react with AuNIs. Furthermore, the sensitivity of AuNIs can be diminished due to excessive cDNA coverage.

In contrast, under a lower concentration condition with 200 pL 0.5 mM RdRp-COVID-C, the surface of AuNIs cannot be completely covered within the entire reaction time. This insufficient surface functionalization not only limits the sensing dynamic range of AuNIs, but also raises the possibility of non-specific binding due to these naked AuNIs.

The dynamic observation and optimization of surface modification were discussed above (see Figure 8c and 8d), and the final condition was determined to be 200 pL of 5 mM, which containing 1 nmol RdRp-COVID-C sequences

Example 10: PPT effects on the real-time LSPR detection. In order to estimate the different association rate constant k_a, the real-time hybridization responses of RdRp-COVID at 1 nM were fitted as shown below:

Where R_eq = 1.57 radian was the equilibrium response from the saturated LSPR sensors as shown in Figure 5a, k_a is the association rate constant. Based on the curve fittings, the calculated k_a at InM are found to be 1.11 x 10^s M ^s ¹ with the PPT heating enhancement and 1.41 x 10⁵ M ^s ¹ without PPT heat.

Example 11: Discrimination of two similar sequences without the PPT heat.

The LSPR detection of two similar sequences, i.e. RdRp-SARS and RdRp-COVID. According to Table SI, there are only three bases different between these two similar oligonucleotides. Under the room temperature (without thermoplasmonic enhancement), both RdRp-COVID and RdRp-SARS reported positive responses, which indicated that the non-fully matched sequence, i.e. RdRp-SARS can hybridize with the RdRp-COVID-C probes in a non-specific manner. In actual detection, the appearance of similar sequences is unavoidable. Therefore, without a proper temperature condition, the non-specific hybridization can significantly influence the reliability in clinical diagnosis. The proposed thermoplasmonic effect and local PPT heat can elevate the local temperature to facilitate the in situ LSPR detection of the target sequences.

Example 12: Dissociation rate constant of RdRp-SARS.

The AuNIs sensor was functionalized with RdRp-COVID-C probes, which can specifically hybridize to the RdRp-COVID sequence. Before the dissociation test, the solution contained RdRp- SARS sequences (1 nM) was allowed to react with RdRp-COVID-C at room temperature. According to Figure 17b, the non-specific RdRp-SARS sequence can be captured by the RdRp-COVID-C probe at room temperature. Then the laser excitation was applied to the chip and the real-time dissociation responses were recorded as shown in Figure 9f. The dissociation part, i.e. 400 sec to 1400 sec were plotted in Figure 18 for the dissociation analysis. The experimental data were fitted to the integrate-rate-equation as shown below:

Where Ro = 0.02436 radian was the LSPR response at t = to = 0 as shown in Figure 18, k_d is the dissociation rate constant. Based on the curve fitting, the calculated k_d is found to be 8.3 x 10³ s ¹.

In contrast, the fully matched sequences RdRp-COVID did not show observable dissociation due to the strong binding affinity. In the dissociation fitting, the Ro = 0.9280 radian was the initial response at t = to = 0. The fitted dissociation rate constant based on the integrated-rate-equation was 3.5 x 10⁶s ¹. Then the dissociation half-life time was calculated, which indicated the time need to remove half of the sequences. The retrieved half-life time of RdRp-SARS was 83.3 sec, while that of RdRp- COVID was 1.97 x 10⁵ sec. In the present experiment, the testing time for SARS-CoV-2 sequences was generally 800 sec. Thus, the fast dissociation of the non-specific sequences allows the LSPR biosensor to accurately detect the target sequence.

Example 12: Blank measurement for LoD.

The uncertainty was further discussed in two basic aspects: the standard deviation of LSPR sensing response represented the 'random uncertainty', while the response shift of a blank measurement indicated the 'systematic uncertainty'. The systematic uncertainty is related to the faults in the LSPR sensing instrument, e.g. the LED source and the spectrometer. According to the definition of lUPAC (International Union of Pure and Applied Chemistry), the LoD is the sum of the blank measures (2.92 x 1CT³ radian) and triple of its standard deviation (3.12 x 1CT³ radian). Thus, the propagated error can be calculated to be 0.00426 radian

The corresponded sequence concentration of this propagated uncertainty of ± 2.92 x 1CT³ radian was found to be ± 0.08 pM. Thus, the final LoD of the system was found to be 0.22 ± 0.08 pM.

Example 13: site specific probe cleavage biosensors

In cyclic fluorescent probe cleavage (CFPC) detection, 50 pL reaction solution containing a unit of Endo-IV and ImM fluorescent DNA probes was injected into the reaction chamber for the cyclic reaction. The homogenized laser beam provided thermoplasmonic assistance in this rection process: providing a high temperature suitable for Endo-IV to cleave the AP sites in the DNA duplex and making the short oligo sequences dissociate from the viral target sequences. The excitation of surface plasmons manifested itself as the phase shift near the resonance wavelength at approximate 584 nm. By employing the windowed Fourier transform calculation, the phase change was extracted from the spectral interferometric patterns and utilized for quantifying the local refractive index change and molecular bindings.

Example 14: site-specific fluorescent probe cleavage for SARS-CoV-2 detection.

The PPT assisted site specific probe cleavage sensing was validated with the nsp-13 viral sequence of SARS-CoV-2. According to the reference SARS-CoV-2 full genomic sequence (NC_045512.2) in the Genbank, the viral target sequence in the helicase/non-structural protein 13 (nsp-13) gene region was selected as shown in FIG 20. This viral target sequence (VTS-nspl3) with the length of 65nt has good representativeness and specificity compared with the corresponding location of other six human coronaviruses. Based on this target nsp-13 sequence, the complementary DNA receptor (nsp-13-C) was further designed for direct detection of the virus sequence, as well as the AP-site modified fluorescent probe containing a 5'-ATT0532 head and a 3'- BHQ1 quencher terminus for CFPC reaction as shown in FIG. 20. With the thiol-modification, the AuNI biosensing chips were directly functionalized with the viral receptor through Au-S covalent bonding. The functionalized microfluidic AuNI biosensing chip was initially used for the direct viral target sequence quantification. The real-time direct detections of VTS-nspl3 with a concentration range from 0.1 pM to 10 nM were shown in FIG. 22. Prior to the detection, a baseline was established. In each detection, the standard VTS-nspl3 sample was injected into the microfluidic PTAPS detection chamber. The thermoplasmonic assistant unit constructed a uniform two- dimensional photothermal field on the surface of the AuNI sensor, thereby promoting the rapid hybridization between the nsp-13 and the complementary DNA receptor. After the direct hybridization stage, the microfluidic AuNI sensor was flushed with nuclease free water to remove non-specific binding substances on the surface and acquired the final differential phase responses.

After the direct viral sequence detection, the mixture of Endo-IV and fluorescent probes was used to perform the secondary CFPC detection as shown in Fig. 21. The fluorescent probes initially hybridized to the captured nsp-13 sequences by forming the DNA duplex. Then the Endo-IV recognized the AP site in the DNA probe and cleaved the fluorescent probes into two short strands. Simultaneously, the local thermal field constructed by the homogenized laser beam dissociated the two short strands, and 'switched-on' the fluorescent gain medium to amplify the plasmonic response. As this process can be circularly performed under the thermoplasmonic effect, the LSPR phase response was observed to continuously increase as shown in FIG. 23. Different from the hybridization-based direct detection, this gain-assisted CFPC detection signal continued to grow in a linear manner over a long period of time. In the concentration range from O.OlpM to lOnM, the LSPR phase increment was positively correlated with the concentration of the nsp-13 sequence. The CFPC reaction time was initially set to be 500s, and the regression fitting of the TP-DMT detection was given in FIG. 24a and b. It was found that the CFPC approach reported enhanced phase responses when detecting the viral sequence with concentrations lower than 30 pM. Using the same LOD calculation method as described above, the LOD of CFPC detection was estimated as 0.275 ± 0.051 fM, which was more than two orders of magnitude lower than that of the direct detection approach.

Claims

1. A dual-functional plasmonic detector comprising:

2. The dual-functional plasmonic detector according to claim 1, wherein the photothermal unit (100) further comprises

3. The dual-functional plasmonic detector according to claim 1 or 2, wherein the plasmonic sensing unit (200) further comprises

(b-i) a source (201) configured to generate the polychromatic radiation beam;

(b-ii) a polarizer (202) configured to linearly polarize the radiation beam;

(b-iii) a birefringent crystal (203) configured to generate retardation to the linearly- polarized radiation beam;

(b-iv) a prism (204) arranged to couple the incident radiation beam onto an LSPR matrix (205) in an inclined incident angle;

(b-vii) a secondary polarizer (207) arranged to combine the radiation beam components from the LSPR matrix; and

4. The dual-functional plasmonic detector according to any of claims 1 to 3, wherein the detection unit (300) further comprises

1 (c-i) an iris aperture (301) arranged to select the radiation beam from the plasmonic sensing unit; and

5. The dual-functional plasmonic detector according to any of claims 2 to 4, wherein the source (101) configured to generate the monochromatic light beam is configured to generate adjustable output power and further comprises a controller unit configured to supply adjustable power to the source (101).

6. The dual-functional plasmonic detector according to any of claims 3 to 5, wherein the LSPR matrix (205) further comprises a layer of plasmonic-active nanostructures arranged to be functional in plasmonic photothermal effect and localized surface plasmon resonance sensing transduction; a transparent substrate configured to support a layer of plasmonic-active nanostructures and coupled to the prism (204); and an optical liquid configured to match the refractive index of the transparent substrate and the prism (204).

7. The dual-functional plasmonic detector according to claim 6, wherein the plasmonic-active nanostructure further comprises a probe arranged to selectively interact with a target substance of interest.

8. The dual-functional plasmonic detector according to any of claims 1 to 7, wherein

- the plasmonic-active nanostructures comprise a probe that is a single stranded nucleic acid configured to hybridize with an at least partially complementary single stranded nucleic acid target substance, optionally RNA, optionally a single stranded nucleic acid probe hybridizing to a viral RNA-dependent RNA polymerase (RdRp) sequence, optionally

2 selected from SEQ ID NO. 1 (RdRp-COVID) and 2 (RdRp-SARS), a viral open reading frame lab (ORFlab) sequence, optionally SEQ ID NO. 3 (ORFlab-COVID), or non-structural protein 13 (nspl3) nucleic acid sequence having SEQ ID NO. 4.

9. The dual-functional plasmonic detector according to claim 8, wherein the detector further comprises a source that comprises and/or is configured to add to the probe before, during or after hybridization with the single stranded nucleic acid target substance:

(i) a further fluorescent single stranded nucleic acid that

- comprises a quencher suitable for quenching the fluorescence;

- comprises a recognition site for a cleaving enzyme; and

(ii) a cleaving enzyme suitable for recognizing the recognition site and cleaving the fluorescent nucleic acid when at least partially hybridized with the nucleic acid target substance, and unquenching the fluorescence; wherein the detection unit (300) is configured to also detect the fluorescence-induced LSPR response.

10. The dual-functional plasmonic detector according to any of claims 6 to 9, wherein the detection unit (300) is configured to detect the LSPR response of the target-probe binding and the fluorescence-induced LSPR response and is configured to detect, quantify and/or characterize the target substance based on a comparison of the LSPR responses.

11. The dual-functional plasmonic detector according to any of claims 9 to 10, wherein

- the recognition site for a cleaving enzyme is selected from the group consisting of a dehydrothymine site spacer, an inosine site spacer, a hydroxymethyluracil site spacer, an 8-oxogunanie site spacer, and an apurinic/apyrimidinic site spacer.

12. The dual-functional plasmonic detector according to any of claims 9 to 11, wherein the further fluorescent single stranded nucleic acid or the fluorophore has a maximum absorption at a wavelength that overlaps with the excitation wavelength of the photothermal unit source (101), optionally a maximum absorption between and including 512 and 545 nm, and/or the

3 fluorophore is selected from the group consisting of ATT0532, Rhodamine 6G, Alexa Fluor 532, Yakima Yellow and Dyomics-530.

13. The dual-functional plasmonic detector according to any of claims 9 to 12, wherein the quencher is selected from a metal nanoparticle, molecular oxygen, iodide ions, acrylamide, a black hole quencher (BHQ) and a Dabcyl quencher.

14. The dual-functional plasmonic detector according to any of claims 9 to 13, wherein the enzyme suitable for recognizing the recognition site and cleaving the fluorescent nucleic acid is selected from the group consisting of a nucleic acid enzyme, optionally a restriction endonuclease, optionally endonuclease III, endonuclease IV, Tth endonuclease IV, endonuclease V, and endonuclease VIII.

15. The dual-functional plasmonic detector according to any of claims 6 to 14, wherein the probe cell (206) further comprises a single channel or a multichannel, optionally microfluidic, flow cell arranged to transfer a target substance, and/or the second fluorescent single stranded nucleic acid and the enzyme suitable for cleaving the fluorescent nucleic acid to the plasmonic-active nanostructure.

16. The dual-functional plasmonic detector according to any of claims 4 to 15, wherein the optical probe (302) further comprises an optical probe unit configured to obtain the intensity spectrum of the radiation beam; and a processing unit configured to determine the characteristics based on the intensity spectrum.

17. A use of the dual-functional plasmonic detector according to any of claims 1 to 16 for detecting, quantifying and/or characterizing a target sample.

18. The use according to claim 17, wherein the target sample is a nucleic acid, optionally an RNA, optionally a viral RNA, optionally a positive sense, single stranded viral RNA, or a virus, optionally SARS-CoV or SARS-CoV-2.

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19. A method for detecting, quantifying and/or characterizing a target sample comprising the steps:

(i) providing a dual-functional plasmonic detector according to any of claims 1 to 16;

(ii) providing a target sample of interest and positioning the sample in the plasmonic sensing unit (200), optionally on an LSPR matrix;

(iii) generating, from the photothermal unit (100), a monochromatic light beam;

(iv) heating the target sample, optionally the LSPR matrix (205), with the light beam, optionally based a on plasmonic photothermal effect;

20. The method according to claim 19, wherein after step (viii), the method further comprises steps

(viii-a) contacting the sample in the plasmonic sensing unit (200) with

5 (viii-d) generating, from a plasmonic sensing unit (200), a polychromatic radiation beam containing randomly polarized components;

(viii-e) modulating the radiation beam in the plasmonic sensing unit (200) to reach the target sample, optionally the LSPR matrix (205), optionally in an inclined incident angle;

(viii-g) recording and processing the radiation beam from the plasmonic sensing unit in the detection unit (300) to determine a fluorescence-induced LSPR response; and wherein in step (ix), the detecting, quantifying and/or characterizing of the target substance is based on the LSPR response of the target-probe binding and the fluorescence- induced LSPR response, optionally based on a calibrated regression with LSPR responses at different target concentrations.

21. The method according to claim 20, wherein in step (ix), the LSPR response of the target-probe binding and the fluorescence-induced LSPR response are compared and the detecting, quantifying and/or characterizing of the target substance is based on the comparison of the LSPR responses.

22. The method according to any of claims 19 to 21, wherein after step (vii), the method further comprises the step of obtaining an intensity spectrum of the radiation beam component by capturing an interference spectrum, wherein the capturing comprises:

- measuring the spectral interferential intensity for each wavelength; and

23. The method according to any of claims 19 to 22, wherein in step (iv) and/or (viii-c): the monochromatic light beam is converted into a plasmonic photothermal heating at the peak absorption wavelength of the LSPR matrix; and the polychromatic radiation beam is modulated with a localized surface plasmon resonance effect at a wavelength different from the wavelength of the monochromatic light beam.

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24. The method according to any of claims 19 to 23, wherein the output power levels of the monochromatic light beam are adjusted to alter the temperature generated by the plasmonic photothermal effect.

25. The method according to any of claims 19 to 24, wherein the method further comprises: detecting a specific target substance at a specific, optionally elevated, temperature; and discriminating different target substances at an adjustable temperature range.

26. The method according to any of claims 19 to 25, wherein the target sample is a nucleic acid, optionally an RNA, optionally a viral RNA, optionally a positive sense, single stranded viral RNA, or a virus, optionally SARS-CoV or SARS-CoV-2.

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