Nanomaterials as efficient platforms for sensing DNA

Biomaterials 214 (2019) 119215

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Biomaterials journal homepage: www.elsevier.com/locate/biomaterials

Review

Nanomaterials as efficient platforms for sensing DNA Kumar Vikrant

a,1

, Neha Bhardwaj

b,1

, Sanjeev K. Bhardwaj

c,d,1

T a,∗

, Ki-Hyun Kim , Akash Deep

c,d,∗∗

a

Department of Civil and Environmental Engineering, Hanyang University, 222 Wangsimni-Ro, Seoul, 04763, Republic of Korea Department of Biotechnology, University Institute of Engineering and Technology, Panjab University, Sector 25, Chandigarh, 160036, India Central Scientific Instruments Organization (CSIR-CSIO), Sector 30 C, Chandigarh, 160030, India d Academy of Scientific and Innovative Research, CSIR-CSIO, Sector 30 C, Chandigarh, 160030, India b c

A R T I C LE I N FO

A B S T R A C T

Keywords: Nanotechnology Genosensors DNA Biosensing Biomedicine Cancer

The advent of modern biomedical science has led to various accomplishments such as the early detection of genetic disorders. To pursue further advancement in this field, the development of highly specific, sensitive, and economical probes for DNA may be an emerging imperative. Due to the diverse merits of nanomaterials (e.g., cost-effective, rapid, and precise detection capabilities with improved detection limits), nanomaterial technology has made itself a viable option for designing new and advanced sensors. At present, the techniques for sensing DNA are primarily based upon biosensing approaches. This review article highlights the recent developments in nanotechnology as a potential platform for the detection of DNA. We further identify the present bottlenecks and future scope for the advancement of nanostructure-based DNA sensors and describe the research needs in associated areas.

1. Introduction Deoxyribonucleic acid (DNA) is widely regarded as the cornerstone of biological heredity and a conveyor of genetic information [1,2]. The highly sensitive detection of distinct genetic sequences plays a pivotal role in gene therapy, medical diagnostics, food safety, biodefense applications, and environmental monitoring [3,4]. Highly sensitive DNA probes may thus be used as an effective tool for the early stage detection of genetic disorders (e.g., cancer) [5]. Notably, cancer is a global health menace and is reported to be the second leading cause of death in the United States of America with 1,735,350 new estimated cases in the year 2018 [6]. The sheer volume of the number of reported cases of genetic disorders and the huge death toll indicates the importance of the development of efficient strategies for the early detection of these diseases to safeguard human health and further the advancement of allied research areas [6,7]. The available techniques for DNA detection are presented in Fig. 1. At present, the conventional detection of DNA is a very tedious operation and suffers from low accuracy [8]. Essentially, the application of conjugated double-bonded systems of the derivatives of pyrimidine and purine rings (nucleic acids, nucleosides, and nucleotides) results in strong absorption of light in the near-ultraviolet region [9]. This strong light absorption tendency is exploited for DNA quantification via

spectrophotometry [8]. A more accurate DNA amplification/detection strategy is based on the polymerase chain reaction (PCR) [10]. The PCR has been widely used for the early stage detection of pathogenic DNA [11]. Although these techniques are being used extensively in clinical and analytical laboratories, their field and point-of-care use is still in its infancy [12]. The major bottlenecks faced by these conventional methods are the requirement of complex and expensive instrumentation coupled with the absence of a simple transduction mechanism for the generated detection signal [10,12]. An ideal DNA sensor should display swift response, heightened sensitivity, good economics, and easy miniaturization [13]. In this regard, biosensors have attracted widespread attention due to their simple, efficient, and easy operation [14]. Essentially, biosensors consist of a biological recognition element along with a transducer component, which transforms the interactions with the target analyte into sensible signals [15,16]. To obtain a low limit of detection, the transducer element needs to be sensitive enough to detect even the smallest modulations in the signal [16]. Also, the generation of reproducible and stable signals generally requires a transducer surface with properties like robustness, chemical inertness, uniformity, and easily accessibility to the target analyte [17,18]. In this respect, a variety of nanomaterials (NMs) have been suggested as potential candidates for DNA biosensing due to their biocompatibility, great specific surface, innocuous nature,

∗

Corresponding author. Corresponding author. Academy of Scientific and Innovative Research, CSIR-CSIO, Sector 30 C, Chandigarh, 160030, India. E-mail addresses: [email protected] (K.-H. Kim), [email protected] (A. Deep). 1 These authors are considered as co-first authors as they contributed equally to this work. ∗∗

https://doi.org/10.1016/j.biomaterials.2019.05.026 Received 26 January 2019; Received in revised form 13 May 2019; Accepted 16 May 2019 Available online 21 May 2019 0142-9612/ © 2019 Elsevier Ltd. All rights reserved.

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Fig. 1. Strategies for DNA sensing.

the basis of the nucleic acid hybridization event in which complementary single stranded DNA hybridize with each other to form a stable duplex molecule [24]. Such a sequence specific event of hybridization is converted into an analytical signal, which can then be displayed by the biosensors. The principles of such sensing strategies, i.e., electrical and optical sensing, are described in detail in the following sections.

chemical stability, good conductivity, and appreciable catalytic activity [4,19]. NMs are touted to provide stable and highly efficient detection probes for DNA with elevated sensing capabilities. This paper was designed to investigate the realm of DNA biosensing which holds a prominent position in biomedical engineering because of its ability to efficiently diagnose serious diseases (e.g., cancer, tuberculosis, and HIV) in their early stages. This review article was structured to accentuate the recent advances in research and development of the NM-based detection of DNA. To this end, the biosensing principles and mechanisms of NMs were explored to cover both their pristine and engineered forms with a particular focus on their framework and prevalent functionalities for sensing applications. This review is further dedicated to a critical analysis on the performance of various NMs. As such, it will offer better knowledge on their practical applications in this highly challenging area of research. Moreover, the discussion made in this review should strive to uncover the underlying hurdles and barriers for the advancement of nanotechnology towards DNA sensing applications.

2.1. Electrochemical DNA biosensors The electrochemical sensing of nucleic acids is an important strategy for converting the DNA hybridization event into an analytically quantifiable signal. This technology is of particular interest in the fields of clinical, environmental, and forensic investigations [25,26]. The blending of nanotechnology, molecular biology, and electrochemistry has been projected as a major advancement when developing devices to monitor sequence-specific hybridization events [27,28]. Different electrochemical techniques have been employed for DNA sensing based on the collection of voltammetric, amperometric, and impedimetric signals using different transducers (such as carbon paste electrodes, screen printed electrodes, gold electrodes, graphite electrodes, mercury film electrodes, and platinum electrodes) [29]. The electrochemical sensing mechanism for DNA as the target analyte varies with the type of transducer surface (nanomaterial) used for sensor development [29]. The analytical application of nanomaterials in biosensors has become an emerging field of research. The potential utility of nanomaterials have been validated in improving the sensitivity of electrochemical sensors as signaling labels in tracing biomolecules or through electrode surface modifiers for various targets [30–36]. The NM-based modification of the electrode surfaces is found to be an efficient option in increasing the effective surface area of electrode and its electron transfer rate for the generation of maximum output signals. In such assays, the NMs are also conjugated with probe DNA for hybridization with complementary analyte DNA, thus enabling direct wiring of hybrids to electrode surface for the final realization of targets by voltammetric and/or amperometric approaches. Also, because of their high surface-to-volume ratio, nanomaterials are advantageous in signal amplification of bio-recognizable targets (e.g., bioreceptor molecules) by accommodating them at large quantities. Becasue the stability of nanoparticles (NPs) is higher than that of enzymes (generally used as labels), they are being utilized as electrochemical labels. Also, to obtain the maximum output signal as signaling molecules, the nanomaterials are used as vehicles of electrochemical labels by conjugating themselves with enzymes or other redox-active compounds for direct detection of analytes at the electrode surface

2. Nanomaterial-based DNA sensing principles Advances in the usage of NMs for DNA biosensing has prompted the development of efficient sensors with superior signal processing and micro fabrication capabilities. The attractive properties of nanomaterials such as quantum confinement phenomena, high surface to volume ratio, elevated optical/magnetic/electrical characteristics, and high surface reactivity have made them popular in biosensor development. Further, the large surface area to volume ratio of nanomaterials results in a higher loading of bioreceptors over their surface, which assists in achieving higher sensitivity values [20,21]. Nanomaterial-based biosensors have found various applications in food, environment, and health industries for the detection of several analytes. With respect to DNA sensing, various biosensors have been developed using unique electrochemical or optical properties of many different forms of nanomaterials like metal nanoparticles, carbon nanotubes (CNTs), quantum dots (QDs), graphene and its derivatives, and other carbon-based materials. Such DNA nanosensors are important tools in the field of genetics, pathology, criminology, pharmacogenetics, and the medical/ food industry [22,23]. In recent years, there has been increasing interest among researchers with respect to the analysis of mutant genes and nucleic acids for precise and early diagnosis of several contagious agents and the prevention of many serious ailments among humans. Thus, many nanomaterial-based electrochemical and optical sensors have been reported in the literature for DNA detection. These sensors function on 2

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assay times may be needed for diagnosis. Also, the high cost acts as a limiting factor for its use in clinical applications. 2.1.2. Label-free electrochemical detection of DNA The label-free method of DNA detection involves the collection of electrochemical signals after DNA hybridization in which purine/pyrimidine bases are oxidized after sequence specific hybridization processes [29]. Subsequently, the process of DNA hybridization can be realized within the reaction medium using probe functionalized nanoparticles followed by measurement of the oxidation signals of purines/ pyrimidine bases (e.g., adenine, guanine, and cytosine) using an electrochemical technique. The magnetic particles have been employed in such assays for increasing the sensitivity of the sensor. The magnetic particles have been used for signal amplification in electrochemical assays in several ways such as the direct contact with the electrode surface, using them as a carrier for transport of redox-active species to the surface of electrode, and by forming thin films on electrode surface [43]. In most of the reported sensor works, the magnetic particles have been used either by integrating them with the transducer material or in dispersed form in the sample followed by their separation using an external magnetic field onto the surface of detection platform [44,45]. Magnetic particles are generally labelled with single stranded probe DNA molecules that can be hybridized with target DNA. As such, such particles are useful for the efficient concentration and separation of target analytes. The event of hybridization is then detected using different carbon or graphite transducer electrodes. In such magnetic beadbased assays, the hybridized target DNA/capture probe is first magnetically separated from the analyte mixture. Next, the target DNA is detached from the MBs and hydrolysed using perchloric acid to release purine bases. The synthesized purines are at last determined using stripping voltammetric techniques. Also, the use of electrocatalytic reporter molecules, [Ru(NH3)6]3+ and [Fe(CN)6]3, provides increased sensitivity to the label-free electrochemical sensors [46]. Further, a dual strategy based on the detection of oxidation signals of both metal ions (e.g. silver) and purine base (e.g. guanine) has also been reported in literature using disposable pencil graphite electrodes (PGE) [47]. In the assay, the silver nanoparticles were first conjugated with amine functionalized probe DNA and were then passively adsorbed over the electrode surface. Next, the hybridization of probe DNA occurred with its complementary sequence, which was detected through electrochemical oxidation of silver and guanine base. The assay offered advantages of facile preparation, good sensitivity (10 μg/ml), and selectivity. Carbon nanotubes (CNTs) offer great potential in many sensing applications due to the presence of terminal functional groups [29,48–50] (. CNTs (labelled with probe DNA) have been used to modify suitable electrode surfaces (such as carbon electrodes or graphite electrodes) and were then used for the detection of target DNA adsorption/hybridization through electrochemical oxidation of nucleic acids. For example, a multiwalled carbon nanotube (MWCNT)-modified glassy carbon electrode (GCE) was reported for voltammetric detection of breast cancer BRCA1 gene (LOD = 100 f mol) using the electrochemistry of guanine and adenine bases [48]. The incorporation of CNTs into a carbon paste matrix provided better sensitivity for electrochemical sensing of oxidation signals of bases through signal enhancement [49]. The developed sensor obtained higher current values (29 and 61-fold) compared to the sensors formed using bare carbon electrodes for sensing of ssDNA and short oligonucleotide molecules, respectively. In addition, a nanoelectrode array has been developed by aligning (vertically embedded) multi-walled CNTs (MWCNTs) in a silica matrix for sensing of DNA hybridization using ruthenium bipyridine (RuBy) mediated guanine oxidation [51]. The assay offered ultra-sensitive detection of a few attomoles of DNA molecules. Further, the attractive properties of nanowires have also been reported for the development of highly sensitive and specific biosensors [46,52]. Probe DNA labelled

Fig. 2. Strategies for electrochemical DNA sensing.

through stimulation of the electrochemical reaction (e.g., amperometric or voltammetric sensors). The electrochemical DNA sensing mechanism based on nanomaterials can broadly be classified as label-based methods and label-free methods as explained below (Fig. 2).

2.1.1. Label-based electrochemical detection of DNA Label-based methods involve the use of an electroactive indicator system for signal detection using either DNA intercalators (metal coordination complexes, dyes, antibiotics, etc.) that bind dsDNA in an intercalative mode or some metal tag labelled nanoparticles [25,37,38]. The intercalating molecules (such as metal ions or enzymes) can aid in differentiation of perfect and imperfect DNA hybridization processes. The approach can thus facilitate the detection of genetic mutations for medical applications [39]. For example, a bioassay was recently reported using carbon dot-modified gold electrodes for the detection of F508del mutation in the cystic fibrosis transmembrane regulator (CFTR) gene [40]. For the development of assay, 100-mer DNA capture probes were conjugated to the surface of CD-gold electrodes for hybridization with specific sequence present in a 373-bases PCR amplicon of exon 11 in CFTR gene. The hybridization event was monitored using safranine dye (specifically that can bind to ds DNA) as redox indicator in the electrochemical assay with enhanced reproducibility and selectivity. Also, a direct RNA-DNA hybridization-based amperometric assay was introduced for the detection of miRNA-21 (a relevant oncomiRNA in breast cancer patients) using gold nanoparticles modified carbon electrodes [41]. In the assay, the thiolated probe DNA molecules were bound to gold NPs that specifically hybridized with complementary target miRNA to form a heterohybrid (DNA-RNA) structure. The developed heterohybrid was further detected using heterohybrid specific antibody molecules labelled with horse radish peroxidase enzymes (Ab-HRP). The overall binding of target miRNA-probe DNA/AbHRP was later electrochemically monitored in the presence of the H2O2/hydroquinone (HQ) system with a detection limit of 29 fM. The principles of surface chemistry have also been explored for development of electrochemical DNA sensors using metal tags. Nanoparticles labelled with metal tags have been utilized for the detection of hybridization signal by monitoring the acid dissolution of metal ions at the electrode surface using different voltammetric techniques [29]. For example, gold nanoparticle labelled probes were employed for the detection of a 406-base pair human cytomegalovirus DNA sequence (HCMV DNA) by anodic stripping voltammetry (ASV) [42]. In their work, the probe DNA labelled with gold nanoparticles aided in capturing of target DNA analyte through hybridization with probe DNA. The gold metal atoms were then released from the hybrid through oxidative dissolution induced by an acidic bromine–bromide solution. The solubilized gold ions were later subject to indirect determination of analyte DNA in the assay based on ASV approach (detection limit 5 pM). Thus, the label-based approach in electrochemical sensors provided highly specific and sensitive detection of selected targets. However, since an additional step of labelling (with enzymes or metal tags) is required for carrying out the detection procedures, longer 3

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of fluorescence of the donor molecules present in its vicinity at the Förster distance (typically < 10 nm). The observed variation in fluorescence intensity (during the FRET process) is then directly correlated to the analyte DNA concentration. The FRET process can lead to either an increase or decrease in fluorescence intensity of the fluorophore. The fluorescence quenching properties of various nanomaterials (such as carbon nanotubes, gold NPs, metal organic frameworks, and graphene oxide) have been exploited for the development of turn OFF/ ON type DNA biosensors [56]. In such sensors, fluorophore labelled ssDNA molecules are conjugated to the nanostructures through hydrophobic and pi-stacking interactions, resulting in fluorescence quenching due to FRET. In the presence of target DNA, the binding between the target and probe DNA leads to the formation of a hybrid and the alteration of probe DNA conformation. This structural alteration in probe DNA leads to its release from the nanomaterial surface to restore the fluorescence. Similarly, aptasensors have been developed based on target DNA/aptamer binding-induced conformational changes, resulting in a change in the fluorescence intensity of nanomaterials (e.g., quantum dots).

silicon nanowires have also been reported for label-free and sequencespecific conductometric sensing of DNA molecules [52]. Thus, the labelfree electrochemical DNA sensing using nanomaterials provides good sensitivity and specificity with reduced assay times [53,54]. In literature, the approach has been employed for several applications such as mutation analysis, determination of gender, and detection of tumors [55]. 2.2. Optical DNA biosensors The optical biosensors function by measuring the changes in the optical properties on the sensor surface upon binding of a target analyte. They possess unique advantages of high sensitivity, relatively rapid detection, ease of operation, and adaptability to a range of assay conditions, which facilitate their usage at commercial scale in the form of dipstick/microfluidic assays. The advancements made in the field of nanotechnology have further assisted researchers in increasing the sensitivity and specificity of optical biosensors. The optical properties (such as absorbance, reflectance, resonance, and luminescence) of several nanomaterials have been exploited for development of optical DNA diagnostic systems. Most of these systems are based on the detection of the DNA hybridization process and measure the change in resonance, plasmonic alterations, or fluorescence of nanomaterials. Different optical techniques used for development of nanomaterialbased DNA biosensors are summarized in Fig. 3 and are described in the following sub-sections.

2.2.2. Surface plasmon resonance-based DNA biosensors In a Surface Plasmon Resonance (SPR) system, plane-polarized light is passed through a glass prism, the bottom of which remains in contact with the bioconjugated transducer (nanomaterial) surface. The refractive index of the transducer surface changes upon analyte binding to change the angle of light exiting the prism (known as the SPR angle), which is measured and correlated with analyte concentration. The plasmonic properties of nanomaterials (such as gold NPs and graphene) have been exploited for development of SPR-based optical DNA biosensors [57–59]. Recently, a graphene-based portable DNA sensor was reported for detection of Mycobacterium tuberculosis DNA using a SPR technique [58]. In their work, graphene layers were first drop-casted over SPR chips and then single stranded probe DNA molecules were immobilized over it through pi-pi interactions. In the presence of target complementary ssDNA (cssDNA), the probe DNA molecules were released from the graphene surface to change the refractive index, which is measured to determine analyte DNA concentration. The results demonstrated that cssDNA (around 28 fM) can be measured sensitively in salt buffer. The SPR-based detection offered label-free detection with high sensitivity. However, the problem of a false positive might be encountered at times due to unwarranted variations in refractive index (e.g., due to changes in sample temperature or composition).

2.2.1. Fluorescence-based DNA biosensors The development of fluorescent sensors has gained significant attention due to their merits of high sensitivity, operational convenience, and low signal-to-noise ratio (background effects). The fluorescencebased optical nanobiosensors can record changes in fluorescence or luminescence of the transducer material upon analyte recognition. In such sensors, the probe DNA molecule (acting as recognition element) is conjugated with the transducer (a fluorescent nanomaterial) and therefore made to interact specifically with the analyte. A subsequent change in fluorescence of the nanomaterial after hybridization with target DNA is then measured using a fluorescence spectrophotometer. Several fluorescent DNA nanobiosensors exist based on a Fluorescence Resonance Energy Transfer (FRET) mechanism. FRET is an energy transfer mechanism between fluorophore (energy donor) and quencher (energy acceptor) molecules. The quencher molecules cause quenching

Fig. 3. Strategies for optical sensing of DNA. 4

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Fig. 4. Structure of a lateral flow gold NP-based DNA sensor; reproduced with permission from Ref. [60].

the detection of viral DNA within 30–50 s with a detection limit of 100 PFU/ml) [69].

2.2.3. Colorimetry-based DNA biosensors Colorimetric assays based on nucleic acid functionalized nanomaterials (particularly gold NPs) are another methodological option advocated for the detection of DNA. The aggregation effect in gold NPs causes changes in the color of NPs from red to blue/violet (with a decrease in average distance between particles). Such color change upon aggregation is widely exploited among researchers for the development of assays with a visual read out. One such lateral flow nucleic acid biosensor was reported for detection of nucleic acid sequences on a strip [60]. In the assay, the biotin-tagged probe ssDNA was conjugated to gold NPs. After hybridization with complementary target DNA, the NPs accumulated on a streptavidin coated test line. The generation of a red color at the test line indicated the presence of analyte DNA, which was further quantified using a strip reader (Fig. 4). The colorimetric assays are thus simple and easy to operate for visual detection of nucleic acids in real time. Also, the colorimetric assays have been widely accepted for commercial purposes due to their facile and convenient detection procedures. For example, Hain Lifescience has developed DNA.STRIP Technology that detects the DNA hybridization process via an alkaline phosphatase reaction [61]. Similarly, a lateral flow assay has been designed for on-site monitoring of genetically modified organisms (GMO) such as in wheat, cotton, and rice varieties [62]. Likewise, ABINGDON HEALTH has commercialized a lateral flow-based test strip (PCRD) for visual detection of amplified dsDNA within 10 min due to aggregation of carbon particles at the capture line [63].

3. Applications of nanomaterial-based biosensors The detection of DNA in various biological/environmental systems has commonly been conducted via conventional analytical methodologies (such as spectrophotometry and PCR) which involve high cost, complicated instrumentation, and tedious analytical procedures. These methods cannot be applied in point-of-care and field applications due to a number of limitations. DNA sensors have the potential to address such constraints by: (1) improving the characteristics associated with efficient detection (e.g., specificity, sensitivity, DNA type, and the nature of sample), (2) providing a swift response, (3) expanding real-world uses, (4) supplying quantitative analysis in addition to qualitative response, (5) providing cost-effective responses, and (6) allowing for miniaturization. In the subsequent subsections, we discuss and analyze the applications and performances of various NM-based DNA biosensors. 3.1. Carbon-based nanomaterials Carbon-based NMs consists of a major fraction of the reported literature on NM-based DNA biosensors (Tables 1–3). In recent times, CNTs have attracted interest across various disciplines due to their unique electrical, chemical, and physical properties [70,71]. Recently, hybrid biosensors based on CNT-composites (e.g., conducting polymers, redox mediators, or metal NPs) have also drawn great attention to maximize synergistic effects in such sensing systems [16,72]. For instance, MWCNTs were used as an immobilization platform for an amine-functionalized DNA sensing probe to make a lateral flow biosensor (LFB) [73] (Fig. 5). The immobilization was achieved via diimide-activated amidation between the amino groups present on the DNA probe and the carboxylic functionalities on the MWCNT surface. This MWCNT conjugated LFB was used for sandwich-type DNA hybridization reactions to visually detect ssDNA samples with satisfactory performance (detection limit of 40 pM with a linear range of 0.1–20 nM) [73] (Table 1). In a similar approach, Au/magnetic NP-CNT (Au/MNP-CNT) was proposed for the immobilization of a DNA probe

2.2.4. Surface Enhanced Raman Scattering-based DNA biosensors Surface Enhanced Raman Scattering (SERS) is a variation of standard Raman spectroscopy that provides an enhanced Raman signal (typically 106 to 1014 orders of magnitude) through electromagnetic interactions between the analyte and the metal nanostructures [64]. SER-based biosensors are well known for high sensitivity, rapid results, comparatively low cost, multiplexed detection, and portability [65–67]. In a typical SERS method, the nanomaterials are directly conjugated with specific Raman reporter molecules (e.g. organic dyes) and specific bioreceptor molecules (e.g., aptamers, oligonucleotides, mRNA, etc.) to form SERS tags to provide high Raman signals for sensing of target DNA [68]. For example, a silver nanorod-based SERS assay was reported for 5

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Table 1 Carbon-based nanomaterials for DNA biosensing. Order

Nanomaterial

Biosensing principle

DNA type

Detection limit

Linear range

Reference

1

Luminescent

ssDNA

Raw Information 40 pM

In fM 40,000

2 3

Amine-modified MWCNT CNT/ferrocene GCE/rGO/MWCNT

0.1–20 nM

[73]

Electrochemical Electrochemical

Hepatitis C virus genomic DNA BRCA1 5382insC mutation

– 3.5 × 10-19 M

– 0.00035

[80] [85]

0.79 8400 2.5 0.3

0.1 fM - 1 pM 1.0 × 10-18 1.0 × 10-10 M 10 - 107 fM 1 pM - 10 nM 10 fM - 0.1 nM 1 fM - 10 pM

4 5 6 7

MoS2/MWCNT Au/MNP-CNT WS2-MWCNT MWCNT-Ppy

Electrochemical Electrochemical Electrochemical Electrochemical

0.79 fM 8.4 pM 2.5 fM 0.3 fM

8 9 10 11 12 13 14 15 16 17 18 19

Au NP/TB-GO rGO/Au NR rGO-FET CeO2-rGO AgNC/GO GQD cCQD GD NS AuNC/GR AgNC/CNP NCD CNNS

Electrochemical Electrochemical Electrochemical Electrochemical Luminescent Luminescent Luminescent Luminescent Electrochemical Luminescent Luminescent Electrochemiluminescence

ssDNA Influenza virus DNA Hepatitis B virus genomic DNA rpoB gene of Mycobacterium tuberculosis MDR1 gene ssDNA peptide nucleic acid-DNA hybrid Aeromonas hydrophila DNA HIV DNA ssDNA DNA oligonucleotide ssDNA Exonuclease III HIV DNA GSU303 DNA Hemin-labeled ssDNA

2.95 × 10-12 M 3.5 × 10-17 M 100 fM 10-16 M 1.18 nM 75 pM 17.4 nM 25 × 10-12 M 0.057 fM 0.4 nM – 2 fM

2950 0.035 100 0.1 1,180,000 75,000 17,400,000 25,000 0.057 400,000 – 2

1.0 × 10-11 - 1.0 × 10-9 M 1.0 × 10-16 - 1.0 × 10-9 M – 1.0 × 10-15 - 1.0 × 10-8 M 10–100 nM 6.7–46.0 nM 0.04–400 nM 0 - 5.0 × 10-9 M 0.02 fM - 20 pM 1–50 nM – –

[103] [106] [108] [111] [113] [118] [121] [204] [115] [114] [205] [124]

[93] [4] [96] [77]

enzyme-based systems. Also, ferrocene has been used to evaluate the performance of interactions between large biomolecules and the ligands/receptors of biosensors [78,79]. In this respect, a CNT/ferrocene biosensor was used to directly detect hepatitis C virus genomic DNA in clinical isolates in the linear range of 0.1 fM – 1 pM [80]. Along similar lines, 1-pyrenebutyric acid-N-hydroxysuccinimide ester (PANHS) has often been adopted as a linker in the fabrication of CNT based electrochemical biosensing platforms [81]. The PANHS molecules strongly attract the base planes of carbon tubes and sheets via π stacking due to the presence of inherent hydrophobic π functionalities [82]. Moreover, the presence of a succinimide ester functionality makes PANHS molecules prime targets for nucleophilic attack by the amine groups present in protein systems [83]. Graphene is a single atom thin, two-dimensional sheet of sp2 hybridized carbon with very high surface area, great mobility, good electro-catalytic characteristics, and favorable biocompatibility [84]. A conjugated biosensing platform GCE/rGO/PANHS prepared on modified glassy carbon electrodes was demonstrated with good performance towards the detection of BRCA1 5382insC mutation (breast

through interactions between Au NPs and the thiol functionality in the DNA probe [4]. The Au/MNP-CNTs were successfully employed to quantify influenza virus DNA with satisfactory performance (detection limit of 8.4 pM with a linear range of 1 pM–10 nM) and excellent selectivity [4] (Table 1) (Fig. 6). Despite enhanced potential for biosensing, CNTs often suffer from agglomeration in polar solvents (e.g., water) due to van der Waals and hydrophobic interactions [74,75]. In this respect, the encapsulation of CNTs in conductive organic polymers has been touted as a possible solution as these polymers may enhance the electrical and mechanical properties of the material [76]. Adopting this approach, MWCNTs were hybridized with polypyrrole (Ppy) through an easy two-stage electrochemical patterning [77]. The resulting composite (WMCNT-Ppy) was effectively utilized as a transducer element to design the electrochemical DNA sensing of rpoB genes of mycobacterium tuberculosis (DL of 0.3 fM over 1 fM - 10 pM of a linear range) in highly complex matrices (e.g., PCR samples) [77] (Table 1). Ferrocene is known to be a promoter of organometallic redox markers and hence has been actively utilized as a redox mediator in

Table 2 Transition metal-based nanomaterials for DNA biosensing. Order

Nanomaterial

Biosensing principle

DNA type

Detection limit

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18

CoOOH CdTe QD CdTe QD/TiO2/CdS-Mn CdTe QD-ZnO Fe3O4/PDA Ag NC Ag NC Ag NC Au NP Au NP Au NP Au NP AuNTsA PANI/Au NP WS2 Mn2O3 MoS2 TMD NS

Luminescent Luminescent Electrochemical Luminescent Luminescent Luminescent Luminescent Luminescent Electrochemiluminescence Electrochemical Electrochemical Luminescent Electrochemical Electrochemical Luminescent Electrochemical Luminescent Luminescent

ssDNA HIV DNA ssDNA smDNA ssDNA ssDNA ssDNA ssDNA HIV-1 DNA ssDNA Breast cancer gene BRCA1 Genetically modified maize genome Mycobacterium tuberculosis DNA dsDNA ssDNA ssDNA ssDNA ssDNA

Raw Information 0.5 nM 0.76 pM 27 aM 9.3 × 10-16 M 0.05 nM 7.3 nM 2.5 nM 25 nM 5.0 fM 1 nM 1.0 pM 1.5 aM 0.05 ng μL-1 0.01 fM 500 pM 120 × 10-21 M 500 pM 0.05 nM

6

In fM 500,000 760 0.027 0.93 50,000 7,300,000 2,500,000 25,000,000 5 1,000,000 1000 0.0015 – 0.01 500,000 0.00012 500,000 50,000

Linear range

Reference

1–50 nM 1 pM - 10 nM 50 aM- 50 pM 10-11 - 10-14 M 0.1–7.0 nM – 5–100 nM 0–2 μM 0.02–1.0 pM – 1 pM - 500 nM 0.01–1000 fM 0.01–100 ng μL-1 0.001–1000 pM 1–20 nM 10-1 - 10-11 μM 0–50 nM 0–5 nM

[128] [132] [135] [136] [130] [139] [141] [140] [206] [207] [208] [147] [209] [210] [150] [154] [156] [211]

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Table 3 Miscellaneous nanomaterials for DNA biosensing. Order

[A] Hairpin assemblybased biosensors 1

Nanomaterial

Biosensing principle

DNA type

Detection limit Raw Information

In fM

Linear range

Reference

Luminescent

p53 DNA

0.85 fM

0.85

2.0 fM - 1.0 nM

[156]

2

DNA machine based on hairpin assembly cyclic activation Au-NP-based biobarcode

Electrochemical

1 fM

1

–

[168]

3 4

MoS2 Au/Pt NP

Luminescent Electrochemical

Helicobacter pylori DNA ssDNA ssDNA

15 pM 0.3 aM

15,000 0.0003

[165] [169]

5 6 7 [B] MOF-based biosensors 1 2 3 4 5 6 7 8

Iridium(III) complex Au NP Au NP

Electrochemical Luminescent Electrochemical

ssDNA ssDNA ssDNA

9 fM 0.5 nM 0.2 fM

9 500,000 0.2

0–200 pM 1.0 × 10-18 1.0 × 10-7 M 0.025–100 pM – 1.0 fM - 1.0 nM

[212] [213] [170]

MIL-101 AuNP/Fe-MIL-88 MIL-88B NR H2dtoaCu ZIF-8 UiO-66-NH2 Fe-MIL-88NH2 {[Zn2(Cmdcp)(bipy)2(H2O)5] (NO3)2.3H2O}n {[Cu(dcbb)2(H2O)2].10H2O}n [Cu(phen)2]n {[Zn(Cbdcp)(H2O)3].H2O}n MIL-101 Cu-MOF {[Ag3(L)3].(H2O)(CF3SO3)3}n

Luminescent Luminescent Luminescent Luminescent Luminescent Luminescent Electrochemical Luminescent

ssDNA ssDNA ssDNA Multiplexed DNA ssDNA ssDNA ssDNA ssDNA

73 pM 11.4 nM 10.0 pM 0.22 nM 1.2 nM 20 fM 0.21 fM 7.4 nM

73,000 11,400,000 10,000 220,000 1,212,000 20 0.21 7,475,000

0.1–14 nM 30–150 nM 0–5 nM 1–10 nM 10–100 nM 50 fM – 2 nM 1 fM – 10 nM 0–60 nM

[189] [192] [214] [215] [129] [187] [191] [183]

Luminescent Electrochemical Luminescent Luminescent Luminescent Luminescent

dsDNA ssDNA HIV-1 ds DNA ssDNA HIV-1 ds DNA ssDNA

1.42 nM 1.99 × 10-13 M 10 pM 1 nM 196 0.0057 μM

1,434,297 199 10,100 1,000,000 197,973 5,700,000

1–120 nM 10-12 – 10-6 M 1–80 nM 1–20 nM 0–50 nM 0.03–0.09 μM

[184] [190] [186] [216] [185] [217]

TSDR-cruciform DNA crystal ScGFP HRCA-based molecular beacon SERS nanorattle

Electrochemical Luminescent Electrochemical Luminescent

0.21 pM 1 nM 8.9 aM 100 aM

210 1,000,000 0.0089 0.1

Luminescent Luminescent Luminescent

50 pM 0.09 nM 4.0 aM

50,000 90,000 0.004

1 pM - 100 nM 1–100 nM 0.01 fM - 10 pM 1.0 × 10-11 1.0 × 10-10 M 0–5 nM 0.5–40 nM 5–700 aM

[218] [219] [220] [221]

Ta2NiS5 Eu/BTC CP dumbbell-shaped DNA probe

HIV DNA HT-29 cancer DNA phi29 DNA Plasmodium falciparum DNA ssDNA ssDNA Exonuclease III

9 10 11 12 13 14 [C] Miscellaneous biosensors 1 2 3 4 5 6 7

[222] [223] [224]

proposed based on Photonic spin Hall effect (PSHE) for detection of analyte DNA using hybridization between probe DNA and complementary target DNA through simulation [86]. A correlation was established between the spin-dependent displacements and the refractive index variations of sensed solution (Phosphate buffered Saline, PBS) in their work. For the simulation of DNA molecules detection, the amplified transverse shifts in PBS solution after absorption of probe DNA/Target DNA hybrid structure were calculated using a mathematical relation between the refractive index of sensed solution and concentration of DNA molecules absorbed. It was seen that the refractive index of the analyte medium was altered with the adsorption of biomolecules by graphene-MoS2 heterostructure. Such alteration in refractive index then induced the shifts in PSHE which cannot be directly detected (e.g., generally nanometer orders of magnitude). Thus, a signal enhancement technology, known as weak measurement technique, was employed to detect the biorecogntion event. The use of weak measurement technique enhanced the original PSHE shift by around 4 orders of magnitude when using preselection to select the polarization and postselection to produce destructive interference [87]. In addition to signal amplification, the weak measurement technology also simplifies the components of optics involved while reducing the cost of device [88]. The work expounded the theoretical foundation of the proposed biosensor for detection of DNA molecules. The proposed sensor can also distinguish the association of DNA molecules. The results of this study hence confirmed the enhanced potential of PSHE for practical applications toward biological, clinical, and medical diagnostic fields. Metal dichalcogenides (e.g., VS2, WS2, MoS2, and SnS2) bearing a highly layered structure have attracted great interest in the

Fig. 5. Schematic illustration of the principle of DNA measurement on MWCNT-based lateral flow biosensor; reproduced with permission from Ref. [73].

cancer). A detection limit of 0.00035 fM was reported for a detection linear range of 1.0 × 10-18 - 1.0 × 10-10 M) [85] (Table 1). Recently, a novel optical biosensor built as graphene-MoS2 heterostructure was 7

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Fig. 6. Detection of influenza virus DNA by Au/MNPCNT biosensor. (a) Illustration of DNA detection process. (b) IV curve change due to hybridization. (c) Resistance difference (((R(pro+tar) -Rpro)/Rpro)) for the influenza virus DNA sensitivity test. (d) Selectivity test with different types of DNA; reproduced with permission from Ref. [4].

with NMs can be an effective solution, which can even enhance the selectivity and sensitivity of electrochemical biosensors [101]. In this respect, an Au NP/TB-GO nanocomposite biosensor was developed for the effective detection of the MDR1 gene (detection limit of 2.950 pM with a linear range of 1.0 × 10-11 - 1.0 × 10-9 M) [103] (Table 1) (Fig. 7). Although GO shows great potential for biosensing applications, GObased nanocomposites have also been observed to form aggregates [104]. As a consequence, the reduced graphene oxide (rGO) coated with a suitable polymer (e.g., negative charge bearing poly(sodium-pstyrene sulfonate)) is often used in place of GO to overcome their shortcomings [105]. On this trend, Au nanorods (Au NRs) were combined with rGO to form composites rGO/Au NR to effectively detect DNA (detection limit of 0.035 fM with a linear range of 1.0 × 10-16 1.0 × 10-9 M) [106] (Table 1). The label-free FET biosensors do not require electrochemical or fluorescent tags. However, it is a highly

area of electrochemical biosensing due to their very high surface area and electrical conductivity [89,90]. Among these, MoS2 (two layers of sulfur atoms sandwiching a single layer of Mo metal and held together by van der Waals forces) has been extensively explored. The MoS2 framework can facilitate the planar transport of electrical charges through electron-electron interactions among Mo atoms [91,92]. Despite such advantages, the pristine MoS2 also has some disadvantages like lower conductivity (e.g., relative to graphene-based materials) [90]. This shortcoming can be effectively overcome by taking advantage of the synergistic effects of MoS2 composites [92]. For instance, a MoS2/MWCNT nanocomposite biosensor was synthesized via a hydrothermal approach on which glucose oxidase (GOD) was immobilized to act as a redox enzyme [93]. The observed detection limit for the MoS2/MWCNT nanocomposite biosensor was 0.79 fM with a linear range of 10–107 fM (Table 1). Likewise, WS2 (hexagonal W layer sandwiched between double chalcogen layers) also holds great promise due to its several merits like facile large-scale production (as compared to graphene), high dispersibility in water, superior electrical conductivity, and better oxidative/thermal stability than MoS2 [94,95]. A biosensor based on a WS2-MWCNT hybrid (with enhanced mechanical and electrical properties) was also proposed. The hybrid material was synthesized via a hydrothermal approach. The above biosensor was applied successfully for the sensing of hepatitis B virus genomic DNA (detection limit of 2.5 fM with a linear range of 10 fM - 0.1 nM) [96] (Table 1). Apart from graphene, its water-soluble variant (i.e., graphene oxide (GO)) has also been successfully utilized as a biosensor surface due to its many advantages (e.g., abundant oxygen functionalities on the surface, high biocompatibility, and good dispersion capabilities) [97,98]. GO is of great interest in electrochemical biosensing applications as it offers highly favorable mobility for electrons while acting as electrode surfaces to accommodate active groups for sensing purposes [99,100]. Au NPs are actively combined with different NMs and provide relatively easy immobilization through Au-S or Au-NH bonds. Such composites facilitate better transport of electrons between electrodes and target biomolecules [98]. Toluidine blue (TB) has been widely utilized as a redox probe for various electrochemical detection applications due to its high electrical conductivity and chemical stability [101]. However, the practical application of redox mediators is often limited due to leakage from the sensing platform [102]. However, the covalent linking of mediators

Fig. 7. Fabrication and detection process of the Au NP/TB-GO biosensor for DNA sensing; reproduced with permission from Ref. [103]. 8

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attractive option because it can provide high selectivity and sensitivity [107]. An rGO field-effect transistor (FET) biosensor (rGO-FET) was developed by a drop-casting approach with immobilized peptide nucleic acid (PNA) for the effective detection of peptide nucleic acid-DNA hybrids [108]. The rGO-FET biosensor displayed good performance (detection limit of 100 fM) with enhanced reusability and selectivity [108] (Table 1). Ceria (CeO2) is a very attractive NP from an environmental perspective due to its biocompatibility and high catalytic activity [109,110]. Ceria has been used for DNA probe immobilization through π-π interactions (non-covalent bonding) between the DNA bases and the conjugated interface of ceria [110]. Due to the positive synergistic effects, CeO2-rGO biosensors showed good detection capability (detection limit of 0.1 fM with a linear range of 1.0 × 10-15 - 1.0 × 10-8 M) for aeromonas hydrophila DNA [111] (Table 1). Similarly, silver nano clusters (AgNCs) have been recognized as good alternative fluorophores for DNA sensing as compared to quantum dots and organic dyes due to their superior photostability and high specific surface that can promote complete DNA-label molecule hybridization [13,112]. AgNCs are also considered to be a preferable option due to their innocuous and biocompatible properties [13]. However, the AgNC/GO nanocomposite could not be satisfactorily utilized for the optical sensing of HIV-DNA (limit of detection of 1,180,000 fM with a linear range of 10–100 nM) [113] (Table 1). A FRET sensor has been proposed utilizing carbon nanoparticle (CNP) oxides as a quencher for the fluorescence of DNA-functionalized AgNC. This sensor also offered sensing of HIV-DNA strands but with a slightly better limit of detection of 0.4 nM within a linear range of 1–50 nM [114] (Table 1). Similar to AgNC/GO, a nanocomposite consisting of gold nano clusters and graphene (AuNC/GR) was prepared [115]. This detector was utilized successfully for the sensing of Exonuclease III with good performance (limit of detection of 0.057 fM with a linear range of 0.02 fM - 20 pM) (Table 1). Graphene quantum dots (GQDs) have also been known to overcome the drawbacks of organic dyes and semiconducting quantum dots for DNA detection in terms of photostability, photo bleaching, and potential toxicity [116,117]. A GQD-based FRET system was employed to detect DNA samples. The performance of the sensor was not very promising as it only offered a limit of detection of 75,000 fM with a linear range of 6.7–46.0 nM [118] (Table 1). Although FRET sensors may be highly useful in principle, their practical applications are often limited due to the requirement of a very selective pairing of quencher and fluorophore for optimum operation (apart from the potential cytotoxicity of many potential NMs) [119,120]. In this respect, carbon quantum dots (CQDs) have proven to be useful due to their easy synthesis, capability for functionalization, high stability, and low toxicity [116]. Carboxylic functionalized CQDs (cCQDs) were employed for the successful detection of DNA, although the observed performance was relatively poor and therefore should require further refinement in the sensor design (detection limit of 17,400,000 fM with a linear range of 0.04–400 nM) [121] (Table 1). Recent investigations have shown that carbon nitride nanosheets (CNNS) can emit cathodic electrochemiluminescence (ECL) [122]. However, in the case of DNA biosensing, ECL is only observed when oxygen molecules are present to endogenously produce hydrogen peroxide as a co-reactant on the surface of the electrode [123]. Using the principle of ECL, a CNN-based detector was prepared to efficiently detect Hemin-labeled ssDNA with a detection limit of 2 fM [124] (Table 1) (Fig. 8).

Fig. 8. Schematic illustration of CNN-based ECL sensing platform for DNA detection; reproduced with permission from Ref. [124].

oxyhydroxide (CoOOH) can be a good option due to its quick large scale synthesis, and appreciable catalytic activity [127]. Therefore, CoOOH nanoflakes were used for sensing DNA based on a mix-and-detect approach (detection limit of 500 nM with a linear range of 1–50 nM) [128] (Table 2). Essentially, ssDNA labeled with 6-carboxyfluorescein (FAM) was set to adsorb on the surface of CoOOH (primarily through electrostatic forces) with the fluorescence quenching of FAM [129]. However, the fluorescent properties were retained upon the hybridization of the probe with the target DNA. The observed restoration of fluorescence might be ascribable to weak interactions between CoOOH and the formed dsDNA [128]. Along similar lines, core-shell structured Fe3O4 polydopamine (Fe3O4-PDA, responsive to external magnetic field with great biocompatibility) NPs loaded with FAM were designed [130]. In the presence of the target DNA, a hybridization chain reaction (enzyme-less amplification) starts to quench the fluorescence (detection limit of 50,000 fM with a linear range of 0.1–7.0 nM) [130] (Table 2). In recent years, quantum dots (QDs) have emerged as a promising NM for the development of lateral flow biosensing assays due to their advantageous properties (e.g., good photostability, extended lifetime of fluorescence, and high specific surface (hence large quantities of binding sites)) [116,131]. CdTe QDs bearing hairpin DNA (as the investigator probe) were synthesized via a hydrothermal approach [132]. The CdTe QD displayed good performance towards the sensing of HIV DNA (detection limit of 760 fM over a linear range of 1 pM–10 nM) using an enzyme-free toehold-based amplification approach [132] (Table 2) (Fig. 9). Interestingly, the photo electrochemical-based biosensing approach is a rapidly growing area of research due to the very low background noise along with easy, economical, and portable device fabrication [133,134]. To enhance the sensitivity of the photo electrochemical biosensors, the system should have a high intensity photocurrent while discouraging the recombination of electrons and holes [133]. In this regard, CdTe QD/TiO2/CdS-Mn was used as a DNA assay in which CdTe QDs helped in the amplification of the signal (detection limit of 0.027 fM with a linear range of 50 aM- 50 pM) [135] (Table 2). Upon hybridization with the target DNA, the DNA probe underwent a change to affect the photocurrent with a suitable DNA detection mechanism [134]. In a similar approach, (3-aminopropyl)triethoxysilane was used as a linker to covalently bind CdTe QDs with ZnO nanosheets [136]. The CdTe QD-ZnO hybrid showed highly favorable band alignment and porosity, resulting in enhancements in photo electrochemical characteristics. Also, CdTe QD-ZnO displayed very high absorbency of visible light through the utilization of hydrogen peroxide as an electron acceptor (resulting in a highly sensitive photocurrent response towards H2O2 consumption/formation) [136]. In addition, the transport of various species (e.g., reactants and products) is easily facilitated by the

3.2. Transition metal/oxide-based nanomaterials NMs consisting of transition metals or their oxides have proven to be highly effective and attractive options for the detection of a wide range of target analytes due to their various merits (e.g., very high surface to volume ratio, eco-friendliness, availability, good redox activity, and great abundance in nature) [125,126]. For instance, cobalt 9

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Fig. 9. HIV DNA biosensing by CdTE QD. (a) Detail of lateral flow assay strip. (b) Schematic of strand displacement amplification and positive or negative expression on strips; reproduced with permission from Ref. [132].

two dimensional and highly interconnected framework of ZnO nanosheets. The high porosity and connected framework greatly favor the capture of excitation light through multiple scatterings. The CdTe QDZnO hybrid showed a detection limit of 0.93 fM with a linear range of 10-11 - 10-14 M (Table 2).

3.2.1. Plasmonic nanoparticles The measurement of the fluorescence intensity at two distinct wavelengths (ratiometric fluorescence) is favorable over the single steadystate fluorescence technique because the former method can effectively reduce the environmental noise and enhance the accuracy of the detection method [137,138]. However, ratiometric fluorescence-based biosensors often encounter shortcomings like bleaching of fluorescent dyes, recombination of acceptor-donor pairs, and the need for labels [137]. In this respect, Ag NCs have been proposed for the sensing of DNA, wherein they do not need any labeling as the cytosine base (C base) itself has a great affinity for Ag+ ions. Hence, the generation of Ag NCs took place through the reduction of Ag+ by oligonucleotides rich in C base [139]. Moreover, the fluorescent characteristics of Ag NCs bearing immobilized DNA mainly depend upon the DNA structure and sequence [140]. As a consequence, a single-nucleotide mutation can be easily detected in a hybridized DNA duplex scaffold due to the formation of Ag NCs (Fig. 10). The best performance of this DNA sensing system was observed at a detection limit of 2,500,000 fM over a linear range of 5–100 nM [141] (Table 2). Similar to the wide application of Ag NCs, Au NPs have also been extensively utilized for DNA detection, as summarized in Table 2. Au NPs boast a very high surface plasmon resonance, size/distance-based optical characteristics, and large sorption extinction coefficient [142]. Au NPs functionalized with oligonucleotides have been used as signaling probes for sensing DNA via hybridization of nucleic acid or primer extension with the support of PCR [143]. In general, Au NPs

Fig. 10. Schematic illustration of label-free ratiometric fluorescence strategy utilizing DNA-regulated Ag NCs for DNA detection; reproduced with permission from Ref. [139].

show an unordered agglomeration state after the reaction and detection of the target DNA, which can be effectively exploited to generate sensing signals for the target DNA [144]. In many cases, the agglomerated Au NPs showed a color shift, which can be easily monitored both by UV–vis spectrophotometer and by the naked eye [144]. In electrochemical approaches, gold electrodes (functionalized with ssDNA) were used to immobilize Au NPs due to the interactions between nitrogen bearing bases and Au, which subsequently facilitated electron transfer to the electrode [145]. In the presence of the target DNA in the system, ssDNA was hybridized to dsDNA to resist sorption by Au NPs (onto the electrode surface). This is primarily due to the phosphate backbones of DNA, which bear a negative charge that enhances resistance to charge transfer for the detection of the target DNA [145,146]. The best 10

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performance by an Au NP was observed for the detection of a genetically modified maize genome (detection limit of 0.0015 fM with a linear range of 0.01–1000 fM) [147] (Table 2). 3.2.2. Miscellaneous The electrochemical detection capabilities of DNA by dichalcogenide composites with carbon-based NMs were discussed in Section 3.1. WS2 is known to behave as a nanoquencher for fluorophore-labeled DNA probes, similar to carbon-based NMs such as GO [148]. Although the fluorescent behavior of GO is highly dependent on its degree of oxidation; this is not the case for dichalcogenides [149]. A peptide nucleic acid (PNA)-based probe was immobilized onto WS2 nanosheets for the effective detection of DNA (detection limit of 500,000 fM with a linear range of 1–20 nM) [150] (Table 2). Note that PNA is a neutral molecule behaving as a DNA mimic with better hybridization characteristics [151]. In recent years, semiconducting NMs have been utilized extensively for the amplification of the output signal [152]. Essentially, metal-oxide NMs with a low band-gap showed FET-like behavior, which can be tailored to suit the desired purpose through actuation [153]. Interestingly, the catalytic characteristics of many semiconducting NMs greatly lowers the over-potential, which improves the electron transfer properties resulting in enhanced performance of electrochemical biosensing [152]. Similarly, Mn2O3 (showing semiconducting behavior with low band gap) nanofibers were synthesized via electrospinning for the effective detection of dengue virus DNA (detection limit of 0.00012 fM with a linear range of 10-1 - 10-11 μM) [154] (Table 2).

Fig. 11. Schematic illustrating the principle of hairpin assembly-triggered cyclic activation of a DNA machine for ultrasensitive chemiluminescence detection of target p53 DNA; reproduced with permission from Ref. [156].

oxidation of a suitable chemiluminescent species (e.g., luminol)) [156,164] (Fig. 11). A MoS2-based DNA sensor built through the abovementioned working principle showed a detection limit of 15,000 fM with a linear range of 0–200 pM [165] (Table 3). Interestingly, the DNA used in hairpins (hpDNA) exhibits a stemloop-based endogenous base pairing pattern of molecules. These pairing patterns may take place when two areas of the same strand (having a complementary sequence of nucleotides as observed in an opposite direction) interact to make a double helix ending in an unpaired loop [166,167]. The double helix architecture of hpDNA can intercalate with a dsDNA-based intercalator [166]. Employing this approach, a bio barcode based on hpDNA (containing 15 base pairs) along with a DNA probe (complementary to about 50% of the target DNA sequence) was fixed onto Au NPs (hpDNA/Au NP/rpDNA) [168]. The other 50% of the target DNA was complementary to a DNA capture probe immobilized onto Au electrodes. The hpDNA/Au NP/rpDNA was able to effectively detect Helicobacter pylori DNA (infecting approximately 15% of the world's entire population) with a detection limit of 1 fM [168] (Table 3). Using the hairpin approach, Au/Pt NPs based on a core-shell structure were prepared as nano-bimetallic catalysts and functionalized with DNA [169]. The H1 was fixed onto Au electrodes functionalized with Au NPs. Interestingly, the H1 was designed so that it undergoes self-hybridization into a stem-loop framework when the target DNA (tDNA) is absent [166]. In contrast, the stem-loop opens up in the presence of tDNA due to the tDNA bonding and the formation of a double strand bearing product with 21 bases [167]. Consequently, H2 opens and results in the formation of dsDNA with 39 base hybridization along with the release and displacement of tDNA [169]. The process keeps on repeating as explained earlier (Fig. 11). The Au/Pt NP showed good performance towards the sensitive detection of tDNA (detection limit of 0.0003 fM with a linear range of 1.0 × 10-18 - 1.0 × 10-7 M) [169] (Table 3). In a similar approach, a DNA capture probe (C-DNA, containing exonuclease III) was immobilized onto an Au NP-based electrode [170]. When tDNA was present in the system, C-DNA hybridizes with tDNA to give rise to a duplex region, revealing the initiator (5′ complementary sequence) [171]. The exonuclease III contains hydroxyl functionalities that selectively consume the duplex region, leaving the remaining initiator behind [172]. Subsequently, the unimpaired tDNA was detached from the framework due to the catalysis of exonuclease III to ensure the next hybridization phenomenon [173]. The hybridization chain reaction was then triggered by the remnant initiator molecules and two hairpin-based signal probes

3.3. Hairpin assembly-based biosensors As elucidated in previous sections, PCR is the most widely used conventional technique for DNA detection. However, in spite of its high sensitivity, PCR involves multiple step processing for the effective detection of DNA; hence, the increased complexity coupled with decreased selectivity may cause false positives [155]. In recent years, the advances in biotechnology, biomedical engineering, and chemistry have enabled the development of alternate amplification strategies to realize the highly sensitive and efficient sensing of DNA (e.g., loopbased isothermal amplification, rolling circle amplification, strand displacement polymerase reaction, ligase chain reaction, hybridization chain reaction, use of biocatalytic amplifiers such as nucleases, and bio barcodes) [156,157]. However, most of the methodologies mentioned above require chemical alteration of the oligonucleotide probes for the purpose of labeling electro active/optical moieties. These alterations may lead to increased cost and time for sensor probe preparation [156]. Also, in cases when the amount of target DNA is limited, it is desirable for the sensing system to have higher specificity and sensitivity. The usage of advanced NMs as typical electrochemical/luminescent probes with diverse operating mechanisms was discussed in previous sections. Further, the NM-based sensors for DNA can be designed to include cyclic activation through the triggering of a hairpin assembly as a possible option to improve performance wherever applicable [158,159]. A generalized hairpin-based DNA detection assembly typically consists of two hairpin structures (H1 and H2) (Fig. 11). As a first step of the reaction, the target DNA attaches itself to H1 and opens it [156]. In the second step, hybridization between H1 and H2 takes place to displace the target DNA, which triggers a new cycle of hybridization between H1 and H2 [160,161]. These hybridization cycles can result in the formation of large amounts of H1-H2 complexes (Fig. 11). The produced H1-H2 complexes undergo activation with specific enzymes (e.g., polymerase or a suitable nicking endonuclease) to produce large quantities of DNA fragments rich in guanine [162,163]. The guanineenriched DNA fragments then undergo further interaction with suitable reactants (e.g., hemin) to form activated species (DNAzymes) that can produce powerful chemiluminescence signals (e.g., through the 11

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virus RNA sequences along with ds HIV-1 DNA with limits of detection of 79 and 196 pM, respectively [185] (Table 3). In another report, a highly sensitive detection of ds HIV-1 DNA was demonstrated using a Zn-based 2-D MOF showing a very low limit of detection of 10 pM [186] (Table 3). Recently, the real time detection of DNA has been achieved using a water stable silver-based 2-D MOF [184]. In addition, the quenching ability of MOF-based nanoprobes has been explored for multiplexed detection of several analytes [187]. For the development of probes, the porous MOFs were used as a scaffold for loading of different dyes capped with DNA molecules (hairpin structure). The target DNA introduction resulted in competitive displacement-based reactions that released fluorophores from the MOF pores. As a result, an elevated luminescent signal was observed due to the large MOF loading capacity. The developed sensor enabled highly specific (even single base mismatch) and sensitive detection of target DNA (LOD = 20 fM). A nano-MOF (MIL-101) was utilized to develop a fluorescence anisotropy (FA)-based assay for the label-free sensing of the gene sequences pertaining to the respiratory syncytial virus (RSV) [188]. The difference in strengths of affinity between MOF and different types of DNA (e.g., positively charged MIL-101 towards negatively charged ssDNA (relatively strong) and dsDNA (relatively weak)) was exploited to assess the detection process of the sensing system. A decrease in the anisotropy value was observed after hybridization of the target DNA with ss probe DNA (labeled with 6-carboxyfluorescein) bound to the MOF, and this depended on the concentration of analyte DNA sequence. The developed sensor showed a detection limit of 1 pM towards target RSV DNA [188,189] (Table 3). In addition to optical sensors, a few MOF-based electrochemical sensors have also been reported for nucleic acid sensing. A carbon electrode modified with a nanocomposite of [Cu(phen)2]2þ (where phen = 1,10-phenanthroline) and graphene (GR) was reported for electrochemical detection of DNA [190]. In the assay, the MOF acted as both an electrochemical indicator for the hybridization process as well as an anchor for immobilization of probe DNA (through advanced intercalation among the aromatic functionalities (vertical groups) of phen and the partial double helix structure of probe DNA). The graphene (containing large amounts of surface docking sites) enhanced the electrical conductivity of the electrode through the immobilization of probe DNA. Upon introduction of target DNA, the square-wave voltammetry (SWV) technique was utilized to detect the hybridization process between the probe and complementary target DNA sequences. The results indicated that the oxidation peak currents of [Cu(phen)2]2] were linear with the logarithm of the concentrations of the complementary target DNA in the range of 10-12 - 10-6 M. The sensor showed good specificity and sensitivity with a detection limit of 0.199 pM [190] (Table 3). Also, a voltammetric DNA sensor using Fe-MOF nanocomposite was proposed recently for the detection of β1 adrenergic receptor gene (ADRB1) as a hypertension biomarker [191]. The nanocomposite consisted of platinum and hemin NPs in a Cu(II) ion-loaded amino-modified MOF (Fe-MIL-88-NH2). In the development of this biosensor, a glassy carbon electrode (GCE) was utilized to immobilize the nanocomposite. Next, the variation in electrocatalytic activity of the nanocomposite towards H2O2 was measured after addition of a target ADRB1 gene. An obvious variation in the electroactivity of H2O2 was observed after probe DNA hybridization with a target gene concentration ranging from 1 fM to 10 nM. The sensor showed a limit of detection of 0.21 fM with good specificity, having applications in precise clinical prognosis and personalized medicine [191] (Table 3). The catalytic peroxidase-like activity of Fe-MOF has also been exploited for the development of colorimetric DNA sensors [192]. In this work, Au NPs were immobilized on the Fe-MIL-88 surface to form a hybrid material. The prominent sorption characteristic of AuNPs towards dsDNA and ssDNA along with high peroxidase-like activity of MOF aided in the development of an efficient sensing platform using

(loaded with methylene blue) catalyzing the oligonucleotide polymerization into the long nicked dsDNA [170]. This hairpin-based approach showed a detection limit of 0.2 fM with a linear range of 1.0 fM 1.0 nM for the effective detection of DNA [170] (Table 3). 3.4. Metal-organic framework-based biosensors Metal–organic frameworks (MOFs) have emerged as advanced crystalline porous materials consisting of metal ions linked to organic ligands through coordinate bonds [174]. MOFs possess excellent properties like ultrahigh porosity, structural tunability, organic functionality, high surface areas, and robust thermal/mechanical stability [175,176]. Due to these fascinating properties, they have been widely used for various applications including storage and separation of gases, catalysis, drug delivery, imaging, and sensor development [177,178]. The MOFs have been used for sensing of several cations, anions, gases and other small molecules [179]. With respect to their application in sensing of nucleic acids, several works have been reported in literature in recent years as discussed in this section. The fluorescence quenching property of MOFs has been widely utilized for the development of nucleic acid biosensors. The hierarchical structure of the MOF helps in attaining good stability, reproducibility, and sensitivity in the developed sensor, while the biomolecule (ssDNA) probes conjugated to the MOFs provide specificity to the sensor [179,180]. The principle of such sensors is based on the differential binding ability of MOFs to single-stranded and double-stranded DNA molecules [179]. The fluorophore labeled ssDNA molecules bind to the MOF ligands through π-stacking and hydrophobic interactions. This may result in the quenching of its fluorescence through a photoinduced electron transfer (PET) mechanism. Further, when a complementary target DNA sequence is added to the complex probe (DNA/MOF), hybridization occurs between the target DNA and probe DNA molecules. This hybridization of the probe DNA alters its configuration with the MOF to allow it to detach from the MOF surface. Hence, the fluorescence of the fluorophore labelled probes is restored. The schematic of above sensing strategy is described in Fig. 12. Using such a sensing strategy, several authors have demonstrated the detection of HIV DNA using different MOFs (based on Cu, Fe, Zr, and Cr metal ions) with good sensitivity and selectivity [129,181] (Table 3). Similarly, an NMOF (UiO-66) was reported for detection of miRNA in living cancers cells [182]. Very recently, a highly sensitive ZIF-8-based detector was utilized for sensing of HIV-1 DNA with a limit of detection of 1.2 nmol L-1 [129]. Further, MOFs have been extended for sensing of HIV-1 double stranded DNA (dsDNA) [183,184]. In these works, the MOF conjugated FAM-labelled ssDNA probes interacted with the major grove of dsDNA present in HIV-1 through Hoogsteen hydrogen bonding to produce a triplex structure. This leads to either complete or partial recovery of FAM fluorescence-based on the analyte concentration. The reported ds DNA sensing assays based on zwitterionic Zn and Cu-based MOFs showed detection limits of 7.4 nM and 1.42 nM, respectively [183,184] (Table 3). Also, a 3-D Cu-MOF was reported for the detection of Sudan

Fig. 12. Schematic of fluorescence quenching-based nucleic acid sensors using MOFs; reproduced with permission from Ref. [56]. 12

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based electrochemical approaches might perform better for sensitive and practical DNA detection and may hold great potential for realworld applications. In this regard, the luminescent biosensing approaches for DNA sensing require intuitive and in-depth analysis so that their performance can be increased further to enable real applications.

the hybrid material. The sensing principle was based on the switchable peroxidase-like activity of the hybrid. Its peroxidase activity was observed to decrease as ss probe DNA molecules were immobilized over the hybrid surface. However, the introduction of complementary DNA in the analyte mixture facilitated the release of the probe DNA from the NP surfaces, thus restoring the peroxidase action of the hybrid material. The detector had a limit of detection of 11.4 nM over a dynamic range of 30–150 nM with good target specificity [192] (Table 3).

5. Challenges in nanomaterial technology for DNA sensing The applications of nanomaterials in biosensing technology has significantly enhanced the detection capability of sensors in clinical, diagnostic, food, and environmental sectors. The unique properties of nanomaterials have enabled the development of rapid, sensitive, specific, portable, and miniaturized biosensors. Further, multiplexed analysis can be performed using nanomaterial-based biosensors, which paves the way for high throughput multiarray sensing devices. However, the emerging field of nanosensors is restricted by certain challenges associated with them. One of the challenges in using nanomaterials for biosensor development is the inability to elucidate the structure and function of nanomaterials and biomolecules for the development of multifunctional nanocomposites, nanofilms, and nanoelectrodes [193]. Also, there is a lack of understanding of the detailed mechanisms involved in the conjugation of biomolecules (e.g. antibodies, nucleic acids, and enzymes) to nanomaterials, which may affect the stability of biomolecules/nanomaterials during the process of nanoprobe development for detection of specific analytes [193]. Further, some of the electrochemical or optical DNA biosensors require the development of homogeneous films of nanoprobes on substrates (such as carbon, silica, gold or graphite), which is a big challenge for obtaining maximum sensor reproducibility [194]. In general, the generation of stable and reproducible signals often requires a highly robust and uniform transducer surface to offering enhanced chemical inertness and easy accessibility to the target analyte. However, DNA biosensors often struggle in terms of quality assurance parameters (e.g., reproducibility). It is found that some biosensors (such as carbon dotmodified gold electrodes and MOF-based detectors) have enhanced reproducibly under laboratory conditions. Nonetheless, further improvement is required to realize applications for real-world matrices [40,179,180]. Notably, a sequence-selective DNA biosensor must possess large storage stability if its routine application is to be envisioned. However, most hybridization biosensors operating on the electrochemical principles showcase limited storage stability. In some studies, the usage of suitable alkanethiols have been suggested to enhance attractive forces between the self-assembly systems and substrate via the formation of robust bonds (in light of altered surface chemistry or increased amount of attachment points) [195,196]. However, as little efforts have been paid on the subject of storage stability for biosensors, further research is required to fabricate a truly robust DNA biosensor capable of operating under practical conditions. In addition, as complicated labelling steps (with metal tags or enzymes) are often required in the DNA detection procedure, actual diagnosis tends to proceed over an extended assay time. For the reduction of assay time, the applied DNA detection protocol needs to be simplified to a large extent. To this end, some authors have proposed the usage of optimized incubations steps and amplification strategies so as to minimize the assay time [197]. Additionally, DNA biosensors which do not require an amplification step, if developed with facile fabrication, may help in the further advancement of this highly challenging area of research. Notably, most of the developed nano-biosensors have been implemented on samples under laboratory conditions. Thus, the effect of environmental (or field) conditions and interferents present in real samples has not been evaluated for their use at commercial scale. Another difficulty in using nanomaterials for biosensing applications is the toxicity issues associated with these materials, which raises concern about public health [194]. Although DNA biosensors have showcased great possibilities in the laboratory atmosphere, their practical applicability or

4. Performance comparison of nanomaterial-based biosensors for DNA The development of advanced NM-based sensing platforms is a critical step for overcoming the drawbacks and shortcomings associated with the conventional methodologies for DNA detection. The availability of a huge number of NMs along with highly favorable characteristics makes them very attractive candidates for the efficient detection of various DNA sequences. In this respect, the NM-based biosensing approach (both luminescent and electrochemical) is the most attractive and commonly utilized strategy for the detection of DNA. Therefore, it is highly beneficial to compare the performances of diverse NM-based sensors to recommend ways to refine and enhance their capabilities. In this section, the performances of DNA detection between different systems have been assessed based on some specific performance metrics, e.g., detection limit and the linear range. The performance can be assessed based on the data compiled in Tables 1–3. Among all the tested NMs, the highest sensitivities (e.g., the best DNA detection limit) were recorded for electrospun Mn2O3 nanofibers (DL of 0.00012 fM and a linear range of 10-1 - 10-11 μM) for the detection of dengue virus DNA [154] (Table 2). This observation indicates that low band-gap semiconducting NMs with FET-like tendencies can be effectively utilized to amplify the output signal for highly sensitive detection of DNA. Also, further exploration of such semiconducting NMs is imperative, as their high catalytic activity may greatly lower the over potential. This may improve electron transfer characteristics to help achieve superior electrochemical-based biosensing strategies. In contrast, the poorest DNA detection limit among NMs was observed for an Ag NC sensing system, which showed a detection limit of 25 nM over a linear range of 0–2 μM [140] (Table 2). However, this detection limit may still be meaningful for practical applications, indicating the generally improved sensitivities of NM-based sensors. In fact, Ag NCs exhibit some favorable characteristics for DNA sensing (e.g., great affinity of Ag+ ions towards C base as well as no requirement of labelling in some cases). However, an in-depth analysis and understanding of the working principle and mechanism of the Ag NCs is crucial. The acquisition of such information will help make their applications practical (Table 2). Interestingly, both the best and poorest DNA detection limits were found for the NMs based on transition metals and their oxides, highlighting the incredible diversity and possibilities associated with these materials. For the carbon-based NMs, the best performance was observed for GCE/rGO/PANHS, which efficiently detected breast cancer mutations with a detection limit of 0.00035 fM over a linear range of 10-18 - 1010 M [85] (Table 1). The good performance of GCE/rGO/PANHS was attributed to the excellent synergy between PANHS molecules and the sheets of GO. In the case of MOFs, the best performance was observed for Fe-MIL-88NH2, which displayed a DNA detection limit of 0.21 fM in the linear range of 1 fM – 10 nM for a hypertension biomarker [191] (Table 3). As such, the great sensing potential of MOFs has been realized towards a multitude of environmental pollutants and important chemicals. Nonetheless, their application towards the effective and practical biosensing of DNA needs to be improved manifold as most MOFs showed relatively poor detection limits compared to other NMs (Table 3). Interestingly, a survey of Tables 1–3 indicates that the electrochemical-based DNA biosensing approaches have superior performance over methods based on luminescence. This indicates that NM13

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sensitive, economical, swift, mobile, and robust natures. Such developments may also assist in the development of effective sensors for the detection of other potentially important biological and environmental target analytes.

commercialization has been severely limited to date. The market for biosensors has typically very narrow opening as it is commonly subject to a slower commercialization rate (due primarily to the high cost and difficulties associated with quality assurance (e.g., sensitivity and stability)) [198–200]. Additionally, the advantages offered by novel nanomaterial-based DNA sensors are at present not lucrative enough relative to the already existing technologies in a viewpoint of commercialization [201]. Nevertheless, the rapid advancements in the application of nanotechnology towards biotechnological systems are expected to stimulate the transition of the DNA biosensing tools to a more advanced forms (e.g., chip-based platform) so as to realistically overcome the requirement of conventional approach (e.g., PCR) [202,203]. In addition, future research endeavors should be directed toward the proper evaluation of performance between different DNA biosensors against clinically relevant real-world samples so as to further upgrade their practical applications and commercialization [201].

Acknowledgements This research acknowledges the support made by the R&D Center for Green Patrol Technologies through the R&D for Global Top Environmental Technologies funded by the Ministry of Environment (Grant No: 2018001850001) as well as by a grant from the National Research Foundation of Korea (NRF) funded by the Ministry of Science, ICT & Future Planning (Grant No: 2016R1E1A1A01940995). This work was also supported by “Cooperative Research Program for Agriculture Science & Technology Development (Project No. PJ014297)” Rural Development Administration, Republic of Korea. AD also acknowledges the financial grant from the Department of Biotechnology, India through project BT/PR18868/BCE/8/1370/2016.

6. Conclusions and future scope

References

This review article was designed to provide an overview of the current NM-based biosensing strategies for effective and sensitive detection of DNA. Among a variety of available sensing techniques, the biosensing approach has been regarded as the most promising for effective detection of DNA due to its efficiency, simplicity, and easy operation. The biosensors are generally built with a recognition element (of biological origin) along with a transducer element, which generates detectable signals. Consequently, a highly sensitive, robust, and chemically stable transducer element ensures high sensitivity with great stability and reproducibility. In this respect, NMs have been touted as highly promising candidates for the real-world and practical detection of DNA due to their manifold advantages (e.g., high surface area, biocompatibility, innocuous tendency, appreciable conductivity, chemical stability, and good catalytic activity). The DNA sensing potential of various NMs (e.g., carbon-based, transition metal/oxide-based, MOFs, and others) was also evaluated in terms of some important performance metrics (such as the limit of detection and the available linear range) (Tables 1–3). The beneficial properties of available NMs (e.g., their versatility, easy applicability, and specific tenability) hold great promise for the advancement of this highly challenging area of research. Interestingly, electrochemical biosensors based on semiconducting NMs with low band-gap have shown exceptionally good performance towards DNA detection; these warrant further investigation for real-world applications. In this review, an entire section was also dedicated to comparing the performance of various NMs to help readers build the effective sensing methods for DNA. Most NMs demonstrated very good performance in general. However, some specific biosensors (e.g., some MOF-based sensors) may need extensive refinement to further upgrade the present performance levels. An analysis of the currently available NM-based DNA biosensing methods suggests the need to overcome a list of shortcomings: (i) lack of knowledge and literature on the application of nanocomposites such as nano films for DNA sensing (they have excellent sensing capabilities for other target analytes), (ii) lack of understanding and studies on the bio conjugation of enzymes, nucleic acids, and antibodies onto the NM surface as they may affect the stability of the DNA detection probe, (iii) limited studies on real-world samples as they may contain potentially interfering species that may greatly affect the performance of the sensor, and (iv) the potential toxicity of some of the NMs raise concerns on their application and management for real-world samples. An interactive interdisciplinary collaboration among scientists from different research areas may help overcome the present technical hurdles while extending the present understanding of NM-based DNA sensors (e.g., with respect to their sensitivity and overall performance). The combined advancement of diverse research fields will pave the way for the dawn of a new era for practical biosensing systems with highly

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