Introduction

Scrub typhus is a zoonotic infection caused by intracytosolic, Gram negative bacterium Orientia tsutsugamushi, resulting in an acute febrile illness and indigenous to ‘tsutsugamushi triangle’ (Paris et al. 2013; Prakash 2017). Human is an incidental, dead-end host, transferred by the bite of infected parasitic larval stage of the mite Leptotrombidium deliense (Shivalli 2016). The disease is highly prevalent in rural areas of Southeast Asia including India, Thailand, Korea, Australia, Russia, The Pacific Islands and Japan (Saraswati et al. 2018). Scrub typhus cases accounts for up to 23% of all febrile illnesses, with an estimated case occurrence of one million each year, in endemic areas (Chunchanur 2018). The fatality rate of scrub typhus can reach 30% or even higher in the lack of appropriate treatment (Xu et al. 2017). Coinciding symptoms of scrub typhus with other co-endemic diseases such as leptospirosis, dengue and typhoid makes it difficult to distinguish (Bonell et al. 2017). Presence of eschar in mite biting site is a specific (98.9%) marker used for clinical diagnosis of scrub typhus; however, the presence of eschar is varied widely in patients from 7 to 97% (Lee et al. 2009). The laboratory-based diagnosis of scrub typhus relies on serological assays like weil-felix test, indirect immunofluorescence assay, indirect immunoperoxidase assay, ELISA and immunochromatographic tests (ICT) etc. Among all serological assays, IFA and ELISA-based method is most reliable for detection of scrub typhus (Xu et al. 2017). Immuno-based methods required paired sampling for confirmation of scrub typhus as standard cutoff value for antibody titer has not yet defined and varied among different studies conducted in endemic areas (Saraswati et al. 2019). Due to retrospective approach, paired sampling-based methods become unsuitable for early-stage diagnosis of scrub typhus. On the other hand, molecular assays rely on PCR-based diagnosis of scrub typhus with more specificity and sensitivity using eschar and blood samples. However, genetic variations among strains of O. tsutsugamushi and lower recovery of DNA from patients’ blood sample are the biggest challenge for its diagnosis using PCR. More than 20 antigenically distinct strains have been identified till now that includes 3 prototypic strains Karp, Kato and Gilliam. Antigenic variations among strains have been linked to highly diverse 56-kDa TSA immunodominant gene, specific to O. tsutsugamushi (Koh et al. 2010; Phetsouvanh et al. 2013; Kelly et al. 2009; Lin et al. 2011). It encodes a primary immunogen (type-specific antigen) located on the outer membrane surface of O. tsutsugamushi and responsible for eliciting neutralizing antibodies (Blacksell et al. 2008). Apart from 56 kDa TSA gene, few other genes have also been identified as genetic markers including 47-kDa high temperature requirement A (HtrA) and 16S rRNA gene etc. (Rodkvamtook et al. 2011; Yang et al. 2012). Among all 56 kDa TSA is most widely used genetic marker for detection of scrub typhus due to its well studied genetic variations among serotypes of O. tsutsugamushi (Kala et al. 2020; Premaratna et al. 2017; Biswal et al. 2018; Usha et al. 2016; Zheng et al. 2015; Zhang et al. 2015; Varghese et al. 2015; Usha et al. 2015).

These days DNA-based biosensors have become new advanced diagnostic tools and are highly sensitive and specific for their target sites (Singh et al. 2017a, b; Kaushal et al. 2016). They can become future of disease diagnosis especially for the diseases like scrub typhus, which is having issues of specificity and sensitivity. Furthermore, sensitivity of biosensors can be improved by modification with nanomaterials and nanocomposite that improves sensitivity of sensor and provide more surface area for immobilization of bioreceptors. Recently several advanced nanostructured carbon materials like MWCNT, SWCNT, graphene, and carbon nanofibers (CNF) have been employed for construction of biosensors due to their better electrocatalytic activity than traditional carbon materials (Li et al. 2007, Wu et al. 2010, Baby et al. 2010, Zhou et al. 2008). Among all nanostructured carbon materials, carbon nanofibers (CNF) have received much attention due to their larger surface area, thermal and mechanical stability with low ohmic resistance (Wang et al. 2012). The CNFs and CNTs are similar in terms of several properties like mechanical strength and electric properties but the advantage of CNFs is that its graphite ordering can be controlled and stacking of various size graphene sheets in its structure produces more edge plane that may facilitate the electron transfer of electroactive species. CNFs have a unique property that its whole surface area can be activated, which provides a much larger functionalized surface area. These properties make CNFs an ideal platform for immobilization of biomolecules and a better transducer (Huang et al. 2010). Unique properties of gold nanoparticles (AuNPs) such as good conductivity and nano-scaled dimension effect received much interest for preparing composite with other nanomaterials in the fabrication of sensors (Song et al. 2011; Hu et al. 2011; Wang et al. 2010; Ulianas et al. 2014). The biocompatibility of AuNPs aid to maintain the bioactivity of immobilized biomolecules and its higher surface-area-to-volume ratio enhances the electron transfer kinetics on the transducer surface (Lin et al. 2015). Gold nanoparticles in combination with other carbon nanomaterials have already been deployed for construction of biosensors for detection of analytes in biological samples (Table 1). However, only few studies have reported the use of AuNPs modified carbon nanofiber-based sensing platform for detection of biological components and revealed their potential application in the designing and fabrication of electrochemical sensors (Mohanaraj et al. 2019; Wang et al. 2012).

Table 1 Comparison of different types of carbon nanomaterials and nanocomposites based DNA sensor
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In this work, Methylene blue was used as a redox indicator dye to monitor the hybridization event on the transducer surface (working electrode) and convert them into valuable signals during electrochemical studies. The change in Ip values was monitors to analyze the samples (ssGDNA of O. tsutsugamushi) on the working electrode surface (AuNPs/CNF) modified with ssDNA probe complementary to the targeted gene (56 kDa tsa). Immobilization of the ssDNA probe onto the working electrode (transducer) surface is an essential step in the fabrication of electrochemical DNA sensor. A strong binding and a good orientation are required for immobilization of DNA probe onto the transducer surface for its stable response and higher reactivity. Several methods have been employed for the immobilization of DNA probes such as covalent bonding, adsorption, and Avidin biotin interaction (Yuce and Kurt 2017). Among all methods, covalent bonding is very strong and used in most of the studies for immobilization of DNA probes onto the transducer surface (Asal et al. 2019). So the present method used the carbodiimide cross-linking (EDC-NHS) chemistry, mostly used covalent bonding method for immobilization of DNA probe. The 1-ethyl-3-(3 dimethyl aminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS) are zero-length cross linkers, that make covalent attachment between the carboxyl group and amines (NH2) terminated DNA probe (Rashid and Yusof 2017). Our aim is to develop an approach which is highly specific and sensitive for detection of scrub typhus with less response time and cost. Present method describes use of gold nanoparticles modified carbon nanofibers (AuNPs/CNF)-based screen printed electrode (working electrode) modified with ssDNA probe using carbodiimide cross-linking (EDC-NHS) chemistry for the early and rapid detection of scrub typhus infection using patient samples (Fig. 1).

Fig. 1
figure 1

Schematic representation of electrochemical DNA biosensor fabrication process along with their characterization methods at each step of electrode modification

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Experimental methods

Materials

N-Hydroxysuccinimide (NHS), 1‐ethyl‐3‐(3‐dimethylaminopropyl) carbodiimide (EDC), methylene blue (MB), were obtained from Sigma Aldrich, USA. Ethanol, hydrochloric acid, sodium chloride, Potassium di-hydrogen phosphate, di‐sodium hydrogen orthophosphate, tris (hydroxymethyl) aminomethane (Tris), ethylenediamine‐tetraaceticacid (EDTA) and other chemicals were received from Qualigens, India. 5′-Amine modified ssDNA probe 20 mer (5′NH2-GCAATGTCTGCGTTGTCGTT-3′) of 56 kDa TSA gene was synthesized from Eurofins Genomics India Pvt. Ltd.

Instruments used

Potentiostat/Galvanostat (Product: Palmsens, model: Palmsens 4, manufactured by Palmsens, The Netherland). Screen printed Gold nanoparticles modified carbon nanofiber-based screen printed electrodes (Order no: 110CNF-GNP) were purchased from Dropsens, Spain and working electrode surface was modified at Amity University, Haryana. Field-Emission Scanning electron microscope (FE-SEM) (Make: JFEI company Of USA (S.E.A.) PTE LTD, Model: Nova Nano SEM-450). Nanodrop Spectrophotometer (Make: QIAGEN, Germany, model: QIAXPERT), Fourier transform infrared spectrometer (Perkin Elmer, Frontier), Raman spectrophotometer (ANDOR, SR-500i-B2) of 656 nm wavelength laser was used. Double distilled water (DW) for reagent preparation and milli Q water for DNA was used throughout the experimental studies.

Isolation of DNA from patient samples

The genomic DNA was isolated from patient’s Blood samples using QIAamp DNA Mini Kit. The different concentrations of isolated DNA were prepared and denatured to ssDNA at 95 °C for 5 min followed by hybridization with ssDNA probe on working electrode surface (AuNPs/CNF/ssDNA (probe)) for 10 min at room temperature (25 ± 3 °C). Genomic DNA (G-DNA) from other bacterial cultures (Salmonella enterica, Klebsiella pnuemoniae and Leptospira interrogans) was also isolated to check specificity of fabricated DNA biosensor.

Construction and response measurement of DNA biosensor for detection of Scrub Typhus

AuNPs/CNF-based DNA biosensor was constructed using ssDNA probe specific to 56 kDa TSA gene of O. tsutsugamushi. Screen printed electrode consist of AuNPs/CNF as working electrode, carbon as counter electrode and silver as a reference electrode. The working electrode was modified with equimolar amount (10 mM each) of EDC: NHS (1:1, v/v) in 1xPBS buffer (0.137 M NaCl, 0.0027 M KCl, Na2HPO4, 0.01 M, 0.0018 KH2PO4), pH 7.4 for 1.3 h, to activate the –COOH groups on the surface. After activation, electrode surface was washed with PBS buffer (pH 7.4) to remove excess of the reagents and air dried at room temperature (25 ± 3 °C).

5′-NH2 modified ssDNA Probe (10 pmol) in TE buffer (10 mM Tris, 1 mM EDTA), pH 8 was added onto the activated AuNPs/CNF working electrode surface for overnight to form the amide bond between –COOH group of AuNPs/CNF and –NH2 group of ssDNA probe results in construction of AuNPs/CNF/ssDNA(probe)-based working electrode. The immobilization step was followed by washing with TE buffer (pH 8.0) to remove unbound ssDNA probe and dried at room temperature (25 ± 3 °C). Different dilutions of single stranded genomic DNA (ssG-DNA) 0.04–2.6 ng was prepared in TE buffer (10 mM Tris, 1 mM EDTA), pH 8 and hybridized with AuNPs/CNF/ssDNA (probe)-based electrode for 10 min at room temperature ((25 ± 3 °C). in moisturized chamber. The hybridization step was followed by washing step using TE buffer (10 mM Tris, 1 mM EDTA), pH 8. Cyclic voltammetry (CV) and electrochemical impedance (EI) was recorded in Potentiostat/Galvanostat including 50 µl of methylene blue and 1 mM [K3 Fe(CN)6] in pH 7.4, PBS buffer.

Characterization of DNA biosensor

The electrochemical DNA biosensor was characterized using Field-Emission Scanning electron microscopy (FE-SEM), Fourier transform infrared spectroscopy (FTIR) and Raman spectroscopy. Characterization studies were performed to compare the fabrication steps of electrochemical DNA biosensor including bare electrode, immobilized ssDNA probe and hybridized cDNA of O. tsutsugamushi. The FE-SEM study was performed to monitor the morphological changes on electrode surface after immobilization of ssDNA probe and hybridization of cDNA. FTIR and Raman study reveals about functional groups imparted onto the surface of AuNPs/CNF electrode after immobilization of DNA probe.

Optimization of DNA biosensor

Optimization studies of electrochemical DNA biosensor was performed to find out the optimum conditions (pH, incubation time and analyte concentration) required for higher output of sensor. The pH was varied in between pH 5–9 at an interval of pH 1.0 to determine the effect of pH on hybridization efficiency using sodium phosphate buffer (pH 5–7), Tris–HCL EDTA buffer (pH 8–9). Similarly optimum hybridization time was calculated at different time intervals (1–15 min) and linearity of biosensor response was checked with different concentration of G-DNA (0.04–10 ng).

Specificity study

Specificity of developed electrochemical DNA biosensor was checked with isolated G-DNA of O. tsutsugamushi and other similar bacteria causing monsoon disease (L. interrogans, S. enterica, K. pnuemoniae). The qualitative and quantitative estimation of isolated DNA samples were performed using nanodrop spectrophotometer. The isolated DNA samples were heated at 95 °C for 5 min and hybridize with AuNPs/CNF/ssDNA (probe) working electrode. The response of DNA biosensor for different bacterial DNA samples (same concentration) was recorded using cyclic voltammetry and change in peak current (Ip) with respect to probe (control) was measured to determine the specificity of DNA biosensor.

Selectivity study

Selectivity of electrochemical DNA biosensor against complementary DNA sequence was evaluated using different mismatched bases in DNA sequence (Table 2). Selectivity of the sensor was calculated based on percentage of CV current (% Ip ncDNA) using the following equation:

Table 2 Oligonucleotide sequence of DNA probe and cDNA containing different number of mismatched bases used in present study
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$${\text{\% }}I_{\text{p}} nc{\text{DNA}} = \left( {{\raise0.7ex\hbox{${I_{\text{p}} nc{\text{DNA}}}$} \!\mathord{\left/ {\vphantom {{I_{\text{p}} nc{\text{DNA}}} {I_{\text{p}} c{\text{DNA}}}}}\right.\kern-0pt} \!\lower0.7ex\hbox{${I_{\text{p}} c{\text{DNA}}}$}}} \right) \times 100{\text{\% ,}}$$
(1)

where Ip ncDNA and Ip cDNA are peak currents of hybridized DNA probe with mismatched DNA base and complementary DNA, respectively.

Regeneration of electrochemical DNA biosensor

Regeneration capability of electrochemical DNA biosensor was evaluated using 0.1 mM HCl solution. Acidic HCl solution breaks down the hydrogen bonds between DNA probe-cDNA and converts them in single stranded form. However, prolonged exposure to HCl solution causes break down of covalent bonds in between electrode surface and DNA probe, provides a way to reuse an electrode for multiple experiments. DNA probe-cDNA hybridization was performed at room temperature (25 ± 3 °C) for 10 min and 1 mM HCl treatment to the electrode was given for different time period of 60–300 s. The reproducibility of the sensor was evaluated by comparing Ip value of bare electrode with other values obtained after treatment of electrode with 1 mM HCl for different time interval.

Validation study

The developed electrochemical DNA biosensor was validated with GDNA isolated from patient’s sample. All collected patient blood samples (n = 9) was confirmed with ELISA and further subjected to DNA isolation for their analysis by electrochemical DNA biosensor. The Ip value of AuNPs/CNF/ssDNA (probe) was compared with values obtained from positive control, negative control and sample DNA for validation. All voltammetric readings were taken in triplicate and average value of respective samples and controls were considered for final validation.

Results and discussion

DNA Hybridization response

Figure 2 shows cyclic voltammogram of AuNPs/CNF bare electrode, AuNPs/CNF/ssDNA (probe) and its hybridization with different concentration of cDNA of O. tsutsugamushi. The Ip of AuNPs/CNF/double stranded DNA (dsDNA) was higher than AuNPs/CNF/ssDNA (probe) and concurrently increased with higher concentration of complementary G-DNA. Methylene blue was used as redox indicator dye which monitors the hybridization event on the transducer surface and converts them into valuable signals during electrochemical studies. The change in Ip values is due to the electrostatic interaction of methylene blue (MB) with exposed nitrogenous bases (guanine bases), which is minimum in bare electrode than AuNPs/CNF/ssDNA(probe) and AuNPs/CNF/dsDNA. The ssDNA probe contains limited number of exposed nitrogenous bases as compare to the exposed unhybridized single stranded genomic DNA (ssG-DNA) onto the transducer surface after probe-cDNA hybridization. Unhybridized ssDNA sequence contains numerous nitrogenous bases which binds more MB molecules through electrostatic binding (Lin et al. 2014, 2015) leading to increase in the Ip value. The average of three Ip values was plotted against ssG‐DNA concentrations with respect to probe (zero), that forms hyperbolic curve and relative current values also show linearity from 0.04–2.6 ng ssG-DNA with a regression coefficient of (R2) 0.90. The sensitivity (S) of DNA sensor was calculated using the formula S = m/A, where m is the slope of the linear equation, and A is the area of the working electrode (0.126 cm2). The LOD of the DNA sensor was calculated using the formula LOD = 3 (σ/S), where σ is the standard deviation and S is the sensitivity. The sensitivity of the electrochemical DNA biosensor was 7849 (µA/cm2)/ng DNA and LOD was calculated as 0.02 ng/µl of ssDNA.

Fig. 2
figure 2

Cyclic voltammogram of (A) AuNPs/CNF electrode, (B) AuNPs/CNF/ssDNA(probe) and (C–I) hybridization with different concentration of ssG-DNA of Orientia tsutsugamushi. The experiment was conducted using 1 mM methylene blue as a redox indicator with a potential range of − 0.6 V to 0.2 V and scan rate of 30 mV/s

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The reason behind the shifting of the potential in the CV plots of the present study is that the current itself is dependent on the rate at which material gets from the bulk of solution to the electrode, known as mass transport. Diffusion is the random movement of molecules from a region of high concentration to regions of lower concentration. At a low concentration of the analyte, diffusion occurs in a satisfactory way, but at high concentration, the diffusion process is disrupted. In this condition, mass transport will be in trouble, so, to compensate for this problem and re-establishment of mass transport, the electrochemical system applies more potential; therefore, the peak potential will be shifted.

The DNA biosensor was also characterized by electrochemical impedance spectroscopy in a frequency range of 10−2–105 Hz using 1 mM [K3Fe(CN)6] in PBS buffer, pH-7.4. The Rct value of bare AuNPs/CNF electrode was lower than AuNPs/CNF/ssDNA(probe) that indicates faster electron exchange kinetic of [Fe(CN)6]3−/4− with bare AuNPs/CNF electrode. Increase in Rct value of AuNPs/CNF/ssDNA(probe) resulted from the immobilization of ssDNA probe containing electronegative phosphate group, which prevents [Fe (CN)6]3−/4− ions to reach the AuNPs/CNF electrode surface (Fig. 3). Furthermore, AuNPs/CNF/ssDNA(probe) was hybridized with different concentrations of complementary ssG-DNA and increase in Rct value were obtained which can be attributed to the increase in negative charges of extra deoxyribose backbone of DNA leading to increase in repulsion with [Fe(CN)6]3−/4− ions. Impedance was in increasing order from bare electrode to immobilized probe and hybridized complementary DNA that shows the success of electrode fabrication process and DNA hybridization event.

Fig. 3
figure 3

Characterization of electrochemical DNA chip fabrication steps including (A) Bare electrode surface (AuNPs/CNF), (B) AuNPs/CNF/ssDNA (probe), (C–H) Hybridization with different concentration of ssG-DNA of Orientia tsutsugamushi, using electrochemical impedance spectroscopy. Electrochemical impedance studies were performed using 1 mM potassium ferricyanide (in Phosphate buffer saline, pH 8.0) solution. The Impedance values were in increasing order from the bare working electrode (A) to immobilized ssDNA probe modified electrode (B) and subsequently increased with concentrations (C–H) of complementary ss GDNA of Orientia tsutsugamushi

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Characterization studies

Electrochemical DNA biosensor was characterized at each step of fabrication using FE-SEM, FTIR and Raman spectroscopy.

The FE-SEM images of the bare AuNPs/CNF electrode, AuNPs/CNF/ssDNA(probe) and AuNPs/CNF/ssDNA(probe) hybridized with complementary G-DNA of O. tsutsugamushi is shown in Fig. 4a–c, respectively. All FE-SEM images were taken in magnification of 25,000 × using the accelerating voltage of 500 kV. The morphological changes at each step of electrode modification could be observed clearly from the FE-SEM images. In Fig. 4a of bare AuNPs/CNF electrode, we can observe the fibrous structure of CNFs decorated with AuNPs. In the case of AuNPs/CNF/ssDNA (probe), the covalently immobilized ssDNA probe formed a thin film (Fig. 4b) which changed into a dense surface morphology (Fig. 4c) after hybridization with ssGDNA in the case of AuNPs/CNF/dsDNA.

Fig. 4
figure 4

FE-SEM micrograph of different steps of electrode modifications including a bare screen printed AuNPs/CNF electrode, b ssDNA Probe modified AuNPs/CNF electrode, and c hybridization with complementary ssDNA

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FTIR analysis was performed to ensure the immobilization of ssDNA probe onto the surface of AuNPs/CNF modified screen printed electrode (Fig. 5). The FTIR spectra in Fig. 5a show transmittance of the bare AuNPs-CNF electrode. In FTIR spectrum transmittance peaks around 777, 1680 and 2512 cm−1 indicative of aromatic C–H bending, C=O and O–H stretching. In addition, Fig. 5b shows FTIR spectra of AuNPs-CNF-ssDNA probe exhibits transmittance peaks at 638,734, 1044 and 1182 cm−1 corresponding to the C2=O stretch of thymine, C=O stretch of guanine, in plane vibration of cytosine and C7=N vibration of adenine, respectively, ensures the presence of 4 nucleotide bases of DNA on the surface. Peaks at 2983 and 1238 cm−1 correspond to C–H and PO2− stretch of deoxyribose backbone, respectively. The peaks at 1410 and 1558 cm−1 are corresponding to C–O–H and CO–NH (amide bond), respectively. The peaks show the carbodiimide cross-linking of NH2 group of the probe with –COOH group on the electrode surface.

Fig. 5
figure 5

FTIR spectra of a AuNPs-CNF and b AuNPs-CNF-ssDNA at frequency range of 500–3500 cm−1 ensures the immobilization of ssDNA probe onto the surface of AuNPs-CNF electrode

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Furthermore, the electrode surface was characterized by Raman spectroscopy (Suppl. Fig. S1) to ensure the DNA immobilization process. The Raman shift was in agreement with the FTIR spectra and shows the peaks at around 355 & 547 cm−1 (vibration of adenine), 668 cm−1 (vibration of T; G; A), 780 cm−1 (vibration of cytosine) and 1073 cm−1 (vibration of PO2−) that ensure the immobilization of the ssDNA probe onto the bare electrode surface. The graphitization pattern also appeared on bare working electrode surface as the peak appears at 1277, 1566, 2660 cm−1 corresponds to D, G and 2D bands of graphene-based materials.

Optimization of electrochemical DNA biosensor response

The developed nanosensor was optimized at different pH conditions, hybridization time and analyte concentrations. The highly acidic and alkaline solution can damage the electrode surface as well as can reduce DNA solubility and hybridization rate. In the present study the biosensor shows the maximum response for DNA hybridization at pH 8.0. Furthermore, the incubation time for DNA hybridization was optimized as 10 min as maximum Ip value was obtained till this point and become constant in subsequent incubation time periods. The incubation for ssDNA-Probe hybridization was performed at room temperature (25 ± 3 °C) and no effect on Ip values was observed by changing incubation temperature from 7 to 28 °C. Hence the electrochemical biosensor response was found linear in the range of 0.04–2.6 ng ssDNA and the current shows saturation after this value as current becomes constant after this point.

Performance of electrochemical DNA biosensor

The performance of electrochemical DNA biosensor was evaluated in context of specificity, selectivity and restoration competency.

The specificity of the developed electrochemical DNA biosensor with O. tsutsugamushi and other similar bacteria causing monsoon disease (L. interrogans, S. enterica, K. pnuemoniae,) is shown in Fig. 6. After hybridization increase in Ip value with respect to probe was observed only with O. tsutsugamushi, whereas other samples show inconsiderable changes in Ip value after hybridization. The sensor specificity was evaluated at 10 ng/µl DNA concentration for all organisms. The increase in Ip value only in case of O. tsutsugamushi confirms the specificity of the sensor towards the targeted organism and overcome the specificity issues of other diagnosis methods used for the detection of scrub typhus. Our study confirms that the electrochemical DNA biosensor is highly specific to 56 kDa TSA gene of O. tsutsugamushi and shows higher efficiency of binding to the complementary site only, as no binding efficiency was observed with other organism’s DNA samples. The electrochemical DNA biosensor shows a significant limit of detection (0.02 ng/µl), sensitivity [7849 (µA/cm2)/ng DNA] and specificity for O. tsutsugamushi.

Fig. 6
figure 6

Specificity of electrochemical DNA chip with GDNA of O. tsutsugamushi and other similar bacteria causing monsoon disease (L. interrogans, S. enterica, K. pnuemoniae)

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In selectivity study percentage CV peak current of Probe hybridization with mismatched DNA sequences were compared with complementary DNA. The selectivity of electrochemical DNA biosensor was found higher for complementary DNA sequence as compared to mismatched base DNA sequences (Table 3). The CV peak current (Ip) percentages of probe hybridization with one, two and three-base mismatched DNAs were 59.5%, 51.4% and 43.3%, respectively, and probe hybridization with multiple base mismatched DNA shows a downfall in peak current to 14.3% relative to cDNA (Suppl. Fig. S2). The biosensor response towards the mismatched DNA sequences was found declining with increased mismatched bases and found equivalent to non-complementary DNA sequences with multiple mismatch bases.

Table 3 Percent CV peak current of Probe hybridization with complementary DNA and DNAs containing different number of mismatched bases
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Restoration competency of electrochemical DNA sensor was evaluated by treating the ss GDNA hybridized AuNPs/CNF/ssDNA(probe) electrode with 1 mM HCl solution. The Ip value was measured before and after treatment of the electrode with 1 mM HCl solution at different dipping time (60–300 s). The biosensor response was declined after treating electrode with 1 mM HCl due to the denaturation of ssDNA-Probe complex. The biosensor response after treatment of 1 mM HCl from 60 to 300 s with a time interval of 60 s were 17.7, 9.64, 9.54, 8.4 and 4.64 µA, respectively, with respect to initial biosensor response (Suppl. Fig. S3).

Validation of AuNPs/CNF-based sensor

The electrochemical DNA biosensor was validated using ELISA positive patient blood samples collected from MMU, Solan, India (Table 4). The DNA was isolated from patients sample using QIAamp DNA Mini Kit and further subjected to electrochemical study for validation. The change in CV peak current (Ip) with respect to DNA probe, negative control (Human GDNA) and positive control (GDNA of O. tsutsugamushi) analysed to validate the patient samples. The patient samples (6 µl ssDNA) were hybridized with AuNPs/CNF/ssDNA probe and CV peak current (Ip) was recorded and compared with the Ip of negative control and positive control for sample validation (Fig. 7). Total 9 ELISA positive blood samples were analysed and all were found positive using electrochemical DNA biosensor. Our biosensor shows a better correlation with the ELISA method and able to detect the infection with more sensitivity and in less time. So the developed sensor is efficient than ELISA and PCR in context of specificity, sensitivity and response time. The biosensor is able to detect the infection at early stages that is not possible with current available diagnostic methods like ELISA, IFA and ICT.

Table 4 Details of scrub typhus patients samples used for validation of DNA biosensor
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Fig. 7
figure 7

Validation of the electrochemical DNA chip with 9 ELISA positive scrub typhus patient blood DNA samples using positive control (PC) and negative control (NC)

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Conclusion

An electrochemical DNA biosensor for detection of scrub typhus based on AuNPs/CNF electrode modified with DNA probe specific for 56 kDa TSA gene was successfully fabricated. The biosensor shows better performance in terms of faster response time (60 s), wider linearity (0.04–2.6 ng), and lower limit of detection (0.02 ng) with higher specificity. The AuNPs/CNF-based sensor shows excellent sensitivity as the sensor is able to sense even lower concentrations of O. tsutsugamushi G-DNA. The higher selectivity and affinity of bioreceptors (DNA probe) towards the 56 KDA TSA gene of O. tsutsugamushi improves the LOD of the developed sensor. The sensitivity of the DNA biosensor is further enhanced by the use of AuNPs/CNF nanocomposite modified transducer (working electrode), which provides the larger surface area for immobilization of bioreceptors and larger electrochemical active area to work on a lower current range that will be beneficial for detection of the patients samples with lower recovery of O. tsutsugamushi G-DNA. Furthermore, the developed sensor is a portable device which can be used in remote areas for better monitoring of the infection without any requirement of trained staff and sophisticated laboratory facility. The developed sensor overcomes the sensitivity, specificity and selectivity issues of other diagnostic methods. No such biosensor has been reported till now and can become a future point of care system for diagnosis of scrub typhus.