Accepted Manuscript

Title: Highly sensitive and non-invasive electrochemical

immunosensor for salivary cortisol detection

Authors: Nidhi Dhull, Gurpreet Kaur, Vinay Gupta, Monika

Tomar

PII: S0925-4005(19)30708-7
DOI: https://doi.org/10.1016/j.snb.2019.05.020
Reference: SNB 26543

To appear in: Sensors and Actuators B

Received date: 24 December 2018

Revised date: 4 May 2019
Accepted date: 6 May 2019

Please cite this article as: Dhull N, Kaur G, Gupta V, Tomar M, Highly sensitive and
non-invasive electrochemical immunosensor for salivary cortisol detection, Sensors and
amp; Actuators: B. Chemical (2019), https://doi.org/10.1016/j.snb.2019.05.020

This is a PDF file of an unedited manuscript that has been accepted for publication.
As a service to our customers we are providing this early version of the manuscript.
The manuscript will undergo copyediting, typesetting, and review of the resulting proof
before it is published in its final form. Please note that during the production process
errors may be discovered which could affect the content, and all legal disclaimers that
apply to the journal pertain.
Highly sensitive and non-invasive electrochemical immunosensor for
salivary cortisol detection

Nidhi Dhull1, Gurpreet Kaur1, Vinay Gupta1, and Monika Tomar2,*

monikatomar@gmail.com

1
Department of Physics and Astrophysics, University of Delhi, Delhi-110007, India
2
Department of Physics, Miranda House, University of Delhi, Delhi-110007, India

PT
*
Corresponding author: Contact.: +91 9871346452

RI
Highlights

SC
 NiO thin film-based label-free electrochemical immunosensor has been devised using
rf magnetron sputtering

U
A detection limit as low as 0.32 pg/mL with high sensitivity in a broad range of linearity
N
from 1 pg/mL to 10 µg/mL has been achieved
 The fabricated immunoelectrode possesses high selectivity against other common
A
interferents
M

 The cortisol biosensor exhibits a stability of 9 weeks.

 Cortisol in actual saliva samples have been successfully assessed and validated using
ED

standard ELISA.
PT

Abstract

A non-invasive electrochemical immunosensor for detection of salivary cortisol has been

E
CC

proposed using RF sputtered NiO thin film as sensing platform. Cortisol, a steroid hormone,

plays a vital part in administrating various physiological activities and also functions to be a
A

stress biomarker. Thus, extremely sensitive and selective non-invasive mechanism for the

detection of salivary cortisol has been focused in the present work. Covalently immobilized

cortisol specific antibody on the NiO matrix has been used as the sensing platform. Cyclic

voltammetry (CV) and differential pulse voltammetry (DPV) studies have been implemented
to calibrate the electrode using standard cortisol solutions. The label-free immunoelectrode has

been successfully devised and used to achieve detection limit as low as 0.32 pg/mL with high

sensitivity in a broad range of linearity from 1 pg/mL to 10 µg/mL. It has been further assessed

to detect cortisol in actual saliva samples and the results exhibit a great scope to foster the

development of portable, integrated and efficient miniaturized non-invasive sensing devices

for cortisol determination.

PT
Keywords: Cortisol; Nickel oxide; Immunosensor; label-free; Saliva

RI
1. Introduction

SC
Cortisol, a steroid hormone, is produced by the adrenal glands positioned right over our

U
kidneys. An optimum level of cortisol plays a vital part in administrating various physiological
N
activities for instance maintaining blood pressure, glucose levels and carbohydrate metabolic
A
rate besides proper functioning of numerous segments of human anatomy such as immune,
M

endocrine, skeletal, cardiovascular and renal systems [1], [2,3]. Cortisol levels in a healthy

human being are at peak during the daybreak and drop down by night-time, following a day-
ED

night rhythmic pattern [4,5]. Taking into considerations the fluctuations in the cortisol levels

due to dietary habits and physical activities, a normal range has been defined to be between
PT

250 ng/mL to 20 ng/mL [3]. While excess of cortisol in the body may lead to Cushing’s
E

syndrome with obesity, fatigue and bone fragility as some of the symptoms [6]; deficiency of
CC

cortisol may result in Addison’s disease pronounced by extreme weight loss, weariness and

darkening of skin folds [7].

A

The psychological stress has the most prevailing effect on cortisol levels leading to its common

name “stress-hormone”. It has been established that cortisol can be used as a stress biomarker

[8]. Protracted experience of stress leads to gesticulation from brain to the adrenal cortex

resulting in release of cortisol. The complicated regime of living in this world of globalization
is the foremost reason of increasing psychological stress which results in numerous health

disparities and are a major issue of concern for most of the countries around the globe. Thus,

development of highly sensitive, reliable and efficient detection techniques is required to

quantify stress for prevention and timely diagnosis of resulting abnormalities [2].

Quantification of cortisol levels is also crucial in the field of sports to prevent the illicit use of

steroids to boost up performance on the field and also to monitor the health of astronauts before

PT
and after a space expedition [9], [10]. Military personnel are incessantly exposed to stress as

RI
the field requires high physical and mental tenacity thus making it crucial to monitor stress

levels closely [11]. Since cortisol secretion is highly variant depending upon the environmental

SC
and behavioral stimuluses, point-of-care detection becomes imperative.

U
Apart from the human serum, cortisol is present in measurable amounts in various body-fluids
N
such as urine, sweat, and saliva, all of these being accessible without the use of any surgical
A
device. Urinal cortisol levels can be misrepresented to a great extent by pregnancy and diuretics
M

medication and sweat has its own intrinsic shortcomings in collecting reliable and recurring

samples for cortisol level determination. Cortisol in the saliva occurs in relevant biologically
ED

active form and causes no distress to the person participating in the test [12]. Thus, due to ease

of collection and handling, concurrent monitoring of salivary cortisol is a promising area of

PT

research. However, the salivary cortisol levels are up to 100-folds lower than that in serum
E

ranging from 0.1 ng/mL to 10ng/mL [13]. Thus, highly sensitive and selective immunosensors
CC

are required with low limit of detection for proficient and non-invasive detection of cortisol.

Currently employed techniques for ascertaining cortisol levels include RIA (radio
A

immunoassay), chromatography and ELISA (Enzyme Linked Immunosorbent Assay). Each of

them poses issues related to multifaceted complexities at each step starting from the sample

collection and purification. Chromatographic techniques are based on the adsorption of analyte

induced by mass transfer. Due to involvement of several preprocessing procedures, such

techniques are limited for application at point of care (POC). Radioimmunoassay (RIA)

involved labeling by radioisotopes and are barely used today due to handling issues [14,15].

ELISA is a very extensively employed method but has limitations in the terms of need of

monoclonal antibodies as matched pairs which are costly. Also, the enzyme-substrate reaction

is short term which needs immediate reading of the microwells [16]. Biosensors based on

optical techniques such as surface plasmon resonance (SPR) are difficult to miniaturize and

PT
suffer from the issues of integration and portability of the complex optical components. Also,

RI
most of the laser instruments are difficult to use in the untreated or unpurified samples. Thus,

these techniques involve collection of biological samples from the human subjects in

SC
centralized laboratories where the samples are processed using sophisticated instrumentations.

In spite of recent advances, the availability of point-of-care devices in the remote areas of
U
military operations still lacks. In view of overcoming these shortcomings, electrochemical
N
immunosensing has surfaced to be the most appropriate method to develop highly sensitive,
A

specific and reliable non-invasive POC devices for real time cortisol level monitoring. They
M

are appropriate for microfabrication and integration of electronics with biology leading to the
ED

promise of easy-to-use analytical devices. Being based on the use of highly specific antibodies,

the greatest advantage offered by the electrochemical biosensors is the combination of

PT

specificity and selectivity towards the target analyte.

E

Sun et al (2008) reported micro-fabricated Au electrodes using alkaline phosphate enzyme for
CC

recognition of salivary cortisol by cyclic voltammetry studies, achieving 0.76 nmol/L as the

limit of detection [17]. Interdigitated µ-electrodes modified by DTSP self-assembled

A

monolayer (SAM) have been employed by Arya et al (2010) to quantify cortisol levels ranging

from 0.36 pg/mL-0.36 ng/mL in saliva via electrochemical impedance spectroscopy (EIS) [18].

M. Yamaguchietal et al (2013) used glucose oxidase labelled cortisol conjugate for cortisol

detection in the range 0.1-10 ng/mL [19]. ZnO has been used by Munje et al (2016) for attaining
detection limit of 1 pg/mL [20]. Most recently Tuteja et al (2018) achieved a detection limit of

0.1 ng/mL via electrochemical sensing of cortisol using Graphene embedded screen-printed

electrodes [21]. The major issue common in most of these reports is the use of labels for

detection of the analyte. Labelling not only adds to the increased number of steps in realization

of the biosensor but can also potentially alter the intrinsic properties of the antibody. Another

issue with the currently proposed biosensor designs is low stability of the matrix. The design

PT
of matrix is the most vital part of developing electrochemical immunosensors. Metal oxides

RI
are highly suitable for this purpose due to their nontoxicity, good biocompatibility, chemical

stability and high electrocatalytic activity. Nickel Oxide (NiO) is one such material that has

SC
been reported to be used in the form of thin films and nanostructures for recognition of a large

range of biomolecules like glucose, urea, DNA, LDL and cholesterol [22–26]. These sensors
U
have exhibited high sensitivity, low detection limits, excellent stability, specificity and
N
reproducibility. The charge transfer through a matrix occurs either via molecule to molecule
A

transfer or diffusion across its surface. The transfer of charge through lattice of the matrix
M

results in enhanced sensing response. NiO as a sensing platform or matrix, having a higher hole
ED

mobility and good oxygen ion conductivity, gives an enhanced electrochemical response and

also facilitates optimal loading of immobilized antibodies in a favorable orientation on its

PT

surface. Possessing a high isoelectric point (IEP) value of 10.5, NiO is also suitable for

enhanced adsorption of proteins having lower IEP due to greater electrostatic affinity. Also,
E
CC

since NiO has a redox couple of its own, it encourages the development of a reagentless

electrochemical immunosensor devoid of any external mediator in solution which on

A

integration with suitable electronics can lead to the subsistence of POC devices for on-site

monitoring of cortisol levels. The electrochemical immunosensors based on NiO have not

been explored for cortisol detection yet and especially for non-invasive detection using saliva.
Thus, in the present work NiO thin film has been used as the matrix for the determination of

salivary cortisol levels. Cortisol specific antibody has been covalently immobilized on the

matrix prepared by RF sputtering technique. The label-free immunoelectrode has been

successfully devised and used to achieve a low detection limit in a wide linear range of

detection with a high sensitivity. Cyclic voltammetry (CV) and differential pulse voltammetry

(DPV) analyses have been implemented to calibrate the electrode using standard cortisol

PT
solutions and have been further assessed to sense salivary cortisol in real samples. The sensing

RI
results have been validated using a standard ELISA assay.

SC
2. Experimental

2.1. Chemicals and reagents

U
Anti-Cortisol antibody derived from rabbit (Ab) and cortisol solution of concentration 1mg/mL
N
in methanol were purchased from Sigma-Aldrich (USA). Further dilution (10µg/mL) of the
A
antibody was prepared in a pH 7.5 Tris buffer (Tris(hydroxymethyl) Aminomethane)
M

consisting of 150mM of NaCl. Working cortisol dilutions (10 µg/mL-1 pg/mL) were prepared

in PBS (pH 7.5) and stored at 4oC. (3-Aminopropyl) triethoxysilane (APTES) was purchased
ED

from Alfa-Aesar (UK). Phosphate buffer saline (PBS) was prepared by mixing monobasic and
PT

dibasic sodium phosphate in an appropriate ratio along with 0.9% NaCl in de-ionised (DI)

water. Monobasic sodium phosphate, dibasic sodium phosphate and Tris buffer acquired from
E

SRL Pvt. Ltd (India) were used to prepare buffer solutions. N-(3-Dimethylaminopropyl)-N’-
CC

ethylcarbodiimide hydrochloride (EDC) and N-Hydroxysuccinimide (NHS) were purchased

from Sigma-Aldrich (USA). Rest other reagents utilized were of analytical grade and need not
A

required any further refinement. A standard ELISA kit from Cayman Chemicals Co. was used

for validation of real salivary cortisol detection results. The recommended assay protocols

were followed and experiments were performed using the 96-well plate.
 2.2. Preparation of immunoelectrode

NiO thin film (NiO/ITO) fabricated by the technique of RF magnetron sputtering using the

parameters reported elsewhere was used as working electrode [27]. Electrically conducting

substrate of ITO coated corning glass (size: 2 cm x 1 cm) was used. The substrate was masked

in half region (1 cm x 1 cm) to make electrical contacts for the electrochemical measurements.

The substrate was prepared by successively cleaning in trichloroethylene, acetone and

PT
isopropyl alcohol in an ultrasonic bath for 10 min each prior to deposition to remove all sorts

RI
of impurities. The step-by-step process flow for preparation of the immunoelectrode is shown

in figure 1. Hydroxylation and silanization processes were carried out to covalently immobilize

SC
the antibody (Ab) on the surface of the NiO matrix via EDC-NHS chemistry [23]. The NiO

U
thin film (NiO/ITO) was hydroxylated (to attain -O-H bonds on its surface) by submerging it
N
in equal volumes of (25 %) ammonia solution and 30 % H2O2 further diluted in de-ionized (DI)
A
water for 30 min at 80oC. The electrodes were then rinsed with DI water and dried prior to
M

silanization by submerging the electrode in APTES (1% solution in toluene) for the night at

room temperature. This was followed by washing the electrode with toluene to eliminate
ED

unbound silanes from the surface. After drying the electrode, it was functionalized with 0.4 M

EDC and 0.1 M NHS solution by providing a 30 min incubation at room temperature. 10 µL
PT

of the Cortisol antibody (Ab) solution of concentration 10 µg/mL was then drop casted on the
E

surface of the electrode and incubated at 27 oC for 3 hours. In the present sensing protocol, in
CC

order to attain a low limit of detection, it is imperative to use a minimal antibody concentration

(to avoid stearic hindrance) without compromising with the signal strength excessively. The
A

devised immunoelectrode (Ab/NiO/ITO) was rinsed comprehensively using buffer solution

(PBS) after incubation to eradicate any unbound or physically adsorbed antibody molecules. A

set of 8 similar electrodes was prepared under identical conditions to study the variation in the
sensing response. The immunoelectrode was incubated at 27 oC for 2 hours with different

concentrations of cortisol for electrochemical studies stored at 4 oC when not being used.

2.3. Measurements and apparatus

The NiO thin films were analyzed for structural characterization. by a Rigaku Ultima X-Ray

diffraction (XRD) instrument. The surface morphological studies of the sensing layer were

PT
accomplished via TESCAN MIRA3 scanning electron microscope (SEM) and Park NX10

Atomic Force Microscopy (AFM). The effective immobilization of the anti-cortisol Ab on NiO

RI
matrix was validated by PerkinElmer Fourier transform infrared (FTIR) spectrophotometer

SC
(Frontier). Cyclic voltammetry (CV) and differential pulse voltammetry (DPV) studies were

performed utilizing Gamry Interface 1000 Potentiostat in a conventional three-electrode set-up

U
where the devised Ab/NiO/ITO immunoelectrode was employed as the working, Ag/AgCl as
N
the reference and platinum foil as the counter electrodes. PBS solution (pH 7) supplemented
A
with an external redox mediator in the form of 5mM [Fe(CN)6]3-/4- was used as the electrolyte
M

for electrochemical analysis.

ED

2.4. Saliva sample collection

Being non-invasive, it is easier to collect saliva samples from the subjects voluntarily without
PT

any medical help. For the present study whole saliva samples of five subjects were collected in
E

vials after thoroughly rinsing the oral cavity with drinking water. About 1 mL of saliva sample
CC

collected from each subject was stored at 4 oC before measurements. All measurements were

performed the same day to avoid degradation of unstable analytes in saliva and avoid freeze-
A

thaw cycles after storing at -20 oC (“Saliva Collection Handbook – Salimetrics,”). The samples

were directly used without any further purification to test the cortisol detection proficiency of

the devised sensing platform and correlate the electrochemical sensing results with the ELISA

results.
3. Results and discussion

3.1. Structural and morphological characterization

The XRD spectra presented in the supplementary figure 1 affirms the growth of polycrystalline

NiO thin film. The XRD peaks at 37.3o and 43.2o 2θ values are well in correspondence to (111)

and (200) planes of the cubic phase of NiO respectively [JCPDS file #471049]. The (111)

surface of NiO being highly polar enhances the antibody immobilization covalently on matrix

PT
surface [29]. The nickel and oxygen terminations are effective for the functionalization of

RI
APTES linker molecules [30]. The SEM and AFM images of the NiO thin film as shown in

supplementary figures 2(a) and 3 (a) respectively depict a porous and nanostructured

SC
morphology with uniformly dispersed nano sized particles. The nanostructured film enhances

U
the efficient bonding of antibody on its surface by providing a greater surface to volume
N
proportion [31]. The evident cracks in the NiO thin film may be attributed to the deposition
A
process which was carried out in 100 % oxygen ambience (Zhao et al., 2014). The Ab/NiO/ITO
M

immunoelectrode surface, as evident in supplementary figures 2(b) and 3 (b), is well-decorated

with globular structures signifying large antibody molecules bound to the NiO thin film in a
ED

mesh-like structure after covalent linkage. The large globular structures can be reasoned due

to the possible agglomeration of the immobilized antibody molecules on the nanostructured

PT

NiO matrix surface. The interaction of the cortisol molecules with the immunoelectrode is
E

depicted in supplementary figures 2 (c) and 3 (c). The cortisol molecules tend to uniformly
CC

bind with the covalently immobilized antibody molecules on the NiO matrix.

3.2. FTIR Studies

A

The NiO thin film and Ab/NiO were also deposited over silicon substrates (Si being transparent

to IR) under identical condition for FTIR studies. Supplementary figure 4 represents the FTIR

spectra of both of these samples. The adsorption bands at 420 cm-1 and 516 cm-1 in the NiO
thin film spectra correspond to the characteristic stretching vibration of Ni-O bond. The strong

band at 614 cm-1 is associated with Ni-O-H stretching bond [33,34]. The bands at 890 cm-1 and

1108 cm-1 show typical bending and stretching of Si-O bond due to presence of an interfacial

layer of SiO2 on the Si substrate [35]. All of these bands also appear after immobilization of

the antibody with diminished intensity and slight shift in frequency (Supplementary figure 3).

This apparent frequency shift may be due to variation in population of the chemical species

PT
after covalent binding of the antibody (“Positional Fluctuation of IR Absorption Peaks:

RI
Frequency Shift of a Single Band or Relative Intensity Changes of Overlapped Bands? |

American Laboratory”).

SC
The intense peak at 1078 cm-1 and the adsorption band at 820 cm-1 assign to the stretching of

U
C-O bond due to amide band in proteins [37]; [38]. The set of bands at 1230 cm-1, 1570cm-1
N
and 1650 cm-1 are assigned to the C-N stretching and N-H bending of Amide III and Amide II,
A
and C=O stretching of Amide I peptide linkages respectively [39]. An additional band at 1440
M

corresponds to C-N axial deformation of amine group band [40]. The significant band at 1712

cm-1 represents vibrational mode of C=O bond [41]. These results clearly support the existence
ED

of protein molecules on the NiO thin film surface and successful immobilization of the anti-

cortisol Ab on the NiO matrix.

PT

3.3. Cyclic Voltammetry (CV) studies

E

Figure 2 shows the CV curves for bare ITO, NiO/ITO and Ab/NiO/ITO in PBS comprising of
CC

[Fe(CN)6]3-/4- mediator and recorded at a 100 mV/s scan rate with applied potential varying
A

from -0.3 to 0.8 V. Distinct redox peaks can be identified in all the CV curves analogous to the

oxidation and reduction potentials of the mediator in the buffer. Though NiO itself possesses a

redox couple of (Ni2+ and Ni3+) its own, but the peak current values due to transition of Ni3+ to
Ni2+ are negligible compared to those due to the transition of Fe3- between Fe4- via subsequent

reversible reaction:

[Fe(CN)6 ]3− + e− → [Fe(CN)6 ]4−

The magnitude of peak oxidation current (Ipo) was very low in the case of bare ITO (520 µA)

which significantly increased for NiO/ITO (798 µA) because of an accelerated rate of electron

PT
transfer which in turn is due to excellent electro catalytic activity of the NiO matrix. Following

the successful immobilization of the cortisol specific antibody (Ab/NiO/ITO), the peak

RI
oxidation current (Ipo) decreased slightly to 691 µA from 798 µA implying decline in the

SC
electron transfer rate through the NiO matrix due to obstruction caused by the insulating

antibody molecules present on its surface. The inset of figure 2 represents the cyclic

U
voltammetry curves for NiO/ITO for 10 continuous cycles. Though, NiO undergoes a redox
N
reaction between the two states Ni2+/Ni3+, the surface of the electrode remains stable after
A
multiple redox cycles for application as a matrix for an immunosensor.
M

Supplementary figure 5 represents the scan rate (υ) studies of the (Ab/NiO/ITO)
ED

immunoelectrode performed from 60 mV/s to 150 mV/s. The peak oxidation current (Ipo)

constantly increases while the peak reduction current (Ipr) decreases with the increase in the
PT

scan rate. Both the peak currents show a linear variation with square root of the scan rate (Inset

of supplementary figure 3) affirming that current in the process is controlled by the quasi-
E
CC

reversible electron diffusion mechanism taking place at the surface of immunoelectrode [42].

The equations of the linear regression for the peak redox currents, Ipo and Ipr are as follows:
A

Ipo = 0.066ν1/2 + 0.034 with R = 0.998 and SD = 0.001

Ipr = −0.038ν1/2 − 0.114 with R = -0.994 and SD = 0.001

Where R is the Regression coefficient and SD is the standard deviation. The value of R close

to 1 and the SD <<1 imply the linear relation of peak redox currents with the square root of

scan rate. The surface concentration of the ionic elements (Г) on the immunoelectrode surface

was calculated using the following equation:

F2ГAn2ν
Ipo =
4RT

PT
Where, F is the Faraday’s constant (96485 C mol-1), ν is the scan rate, R is the gas constant

(8.314 J mol-1 K-1), T is the temperature, n is the number of charge carriers participating in the

RI
redox reaction and A is the area of the electrode [43]. Since it is hard to ascertain the precise

SC
value of n, in accordance with 2014 IUPAC recommendation it can be replaced by unity [44].

The value of Г for the prepared immunoelectrode was found to be 7.2 x 10-7 mol cm-2. A high

U
value of surface coverage is an implication of greater concentration of ionic elements on the
N
electrode surface.
A
M

3.4. Biosensing response studies

Biosensing studies were performed using the Cyclic voltammetry technique by varying the
ED

cortisol concentration from 10 µg/mL to 1 pg/mL (dilutions prepared in PBS of pH=7.5). The
PT

prepared immunoelectrodes were washed with PBS prior to the introduction of cortisol and

incubated for 2 hours with cortisol, providing sufficient time for the antibody -antigen
E

interactions. The corresponding CV curves are shown in the figure 3(a). It can be observed that
CC

the Ipo decreases continuously with increase in the cortisol concentration, the current (Ipo) being

highest for 1 pg/mL. The cortisol binds with the specific antibody present on immunoelectrode,
A

forming an immunocomplex over the surface which inhibits electron transport from the redox

probe to the matrix. Further, the immunocomplex formation leads to decreased active electrode

area for the redox reaction with redox couples in the solution. The difference in peak oxidation

current with varying concentrations of cortisol is directly correlated to the concentration being
measured. However, the difference does not remain appreciable as we approach higher

concentrations as shown in the inset of figure 3(a) where peak oxidation current with cortisol

concentration is given (calibration curve). Therefore, differential pulse voltammetry studies

with applied potential ranging from 0.1 V to 0.5 V with a step size of 2 mV were employed for

a better understanding of the sensing characteristics of the immunoelectrode towards cortisol.

A similar behavior in Ipo was observed in DPV towards varying cortisol concentrations. As can

PT
be observed in figure 3 (a) and (b), the change in peak oxidation current was linear in DPV

RI
than in the CV studies in the concentration varying from 10 µg/mL to 1 pg/mL. The

physiological levels of cortisol in saliva lie well within this range. Thus, realization of non-

SC
invasive cortisol immunosensor is feasible via the proposed protocol. On fitting the peak

oxidation current (Ipo) values from DPV curves of various concentrations

U
(Cortisol/AB/NiO/ITO) into the calibration curve (Inset of figure 3(b)) the linear regression
N
equation for determining the cortisol concentration was found to be
A
M

Ipo = -0.016 log [cortisol concentration] + 0.164

with a correlation coefficient (R2) of 0.99. The lower limit of detection (LOD) was calculated
ED

to be 0.32 pg/mL using the equation [45]

PT

y(LOD) = y(blank) + 3SD(blank)

E

where the average of Ipo values from the DPV studies of the immunoelectrode (Ab/NiO/ITO)
CC

in absence of cortisol was used as y(blank) and the standard deviation of the measurement was

used as SD(blank). Further, to study the matrix effect, the measurement of cortisol concentrations
A

in absence of the receptor antibodies were performed by incubating cortisol directly on the NiO

matrix and the results are shown in the inset of figure 3(b) (Cortisol/NiO/ITO). It can be

observed that exposing the NiO matrix directly to the varying cortisol concentrations (without

immobilizing the antibody) show no significant variation in the peak oxidation current (Ipo).
The Ipo observed in this case is same as observed in case of the blank NiO matrix. Table 1

shows the assessment of the analytical performance of the devised cortisol immunosensor in

the present work compared with other cortisol biosensors previously reported in the literature.

The limit of detection and linear range are better in the present work as compared to

corresponding reported results (Table 1). The NiO thin film has not been previously employed

for cortisol detection especially for non-invasive application. NiO possessing a redox couple

PT
of its own, can be further optimized for development of reagentless non-invasive cortisol

RI
immunosensors.

3.5. Selectivity, Stability and Reproducibility

SC
The selectivity and specificity of the biosensor was studied against common interferents that

U
are commonly present in human body along with a normal concentration of cortisol (10
N
ng/mL). The concentration of these interferents, i.e. LDL, Ascorbic acid, lactic acid, uric acid,
A
urea and glucose was kept at 72 mg/dL, 1 mg/dL, 1 mmol/L, 5 mg/dL, 10 mg/dL and 80 mg/dL
M

respectively which are similar or higher than the levels in saliva of a health person [46–50]. As

depicted in figure 4(a), the peak oxidation current in the presence of interferents show less than
ED

5% variation from that of cortisol. Thus, the prepared immunoelectrode is highly selective and
PT

specific towards cortisol.

Reproducibility of the prepared sensor was tested by fabricating a set of 8 electrodes

E

(Ab/NiO/ITO) under identical conditions and studying their response towards a cortisol
CC

concentration of 1 µg/mL (Cortisol/Ab/NiO/ITO). The variation in peak oxidation current

A

values of individual electrodes and their response towards cortisol lies within 5% deviation

from the average value (Figure 4(b)). Also, the shelf-life of the biosensor was studied by

repeating the measurements for a fixed concentration of cortisol (1 µg/mL) for several weeks.

The resulting plot of peak oxidation current from the DPV studies (Figure 4 (c)) shows that the
electrodes exhibited a stability of up to 90% of its response for a period of 9 weeks. The

repeated measurements of the immunoelectrode with cortisol reveal that the electrodes are

stable for more than 20 continuous measurement cycles (Figure 4 (d)). This also ensures the

measurement-to-measurement reproducibility of the devised sensing mechanism. The

numerous measurements depicted in figure 4 (c) and figure 4 (d) were performed by

regenerating a single immunoelectrode (Ab/NiO/ITO) for repetitive use by washing with a pH

PT
2.5 glycine-HCl buffer solution. The glycine buffer tends to disrupt the binding between the

RI
antibody and antigen and the electrode showed the same electrochemical response as prior to

the introduction of the cortisol solution. The covalent linkage of the antibody to the matrix

SC
accounts for the high chemical stability and highly repeatable antigen bindings to the surface

of the immunoelectrode after regeneration.

U
N
A
3.6. Salivary Cortisol Measurements
M

Cyclic and differential pulse voltammetry studies were performed by incubating the
ED

immunoelectrodes for 2 hours after application of one of the saliva samples and its dilutions in

PBS solution up to 20-fold dilution. As can be seen in the figure 5(a), the sensor did not show
PT

any significant response to up to 4-fold saliva dilution. The obtained sensing response was

considerable in case of 20-fold dilution (Figure 5(a)). This may be because human saliva is
E

extremely concentrated and viscous due to presence of large amounts of mucins and proteins.
CC

These components in saliva might have hindered its binding to the specific antibody on the
A

immunoelectrode [51]. Therefore, 20-fold dilution of each saliva sample was used for

measurement of cortisol using the prepared sensor and the same was analyzed using ELISA

assay. A good correlation can be observed from the figure 5 (b), between the electrochemical

measurements and ELISA results. Thus, the prepared biosensor is capable of detecting cortisol
levels directly in the saliva without any pre-treatment or filtration, paving a way towards

realization of non-invasive cortisol biosensors for on-site application in case of critical

illnesses.

4. Conclusion

The biosensing results clearly demonstrate that RF sputtered NiO thin film is capable of

PT
development of non-invasive cortisol immunosensors for point of care applications. A high

sensitivity in a wide linear range of 1pg/mL to 10µg/mL and a low limit of detection of 0.32

RI
pg/mL together with the capability of cortisol detection directly in real saliva samples are

SC
extremely promising. The proposed biosensor also exhibited high stability, specificity and

reproducibility. The cortisol levels in real saliva samples were accurately determined and

U
validated using standard ELISA assay. Thus, the present sensing protocol has a great scope to
N
foster the development of portable, integrated and efficient miniaturized non-invasive sensing
A
devices for cortisol determination.
M

Acknowledgement
ED

The authors gratefully acknowledge the Department of Science and Technology (DST),

Ministry of Science and Technology and the Department of Physics and Astrophysics,
PT

University of Delhi for the technical and financial support.

E
CC
A
References

[1] E.R. De Kloet, M. Joëls, F. Holsboer, Stress and the brain: from adaptation to disease,
Nat. Rev. Neurosci. 6 (2005) 463–475.
[2] R. Gatti, G. Antonelli, M. Prearo, P. Spinella, E. Cappellin, F. Elio, Cortisol assays and
diagnostic laboratory procedures in human biological fluids, Clin. Biochem. 42 (2009)
1205–1217.

PT
[3] A. Levine, O. Zagoory-Sharon, R. Feldman, J.G. Lewis, A. Weller, Measuring cortisol
in human psychobiological studies, Physiol. Behav. 90 (2007) 43–53.

RI
[4] D. Corbalán-Tutau, J.A. Madrid, F. Nicolás, M. Garaulet, Daily profile in two circadian
markers “melatonin and cortisol” and associations with metabolic syndrome
components, Physiol. Behav. 123 (2014) 231–235.

SC
[5] N.A. Nicolson, Measurement of cortisol, Handb. Physiol. Res. Methods Heal. Psychol.
1 (2008) 37–74.
B.S. McEwen, Cortisol, Cushing’s syndrome, and a shrinking brain—new evidence for
[6]
U
reversibility, J. Clin. Endocrinol. Metab. 87 (2002) 1947–1948.
N
[7] O. Edwards, J.M. Galley, R.J. Courtenay-Evans, J. Hunter, A. Tait, Changes in cortisol
metabolism following rifampicin therapy, Lancet. 304 (1974) 549–551.
A
[8] F. Holsboer, M. Ising, Stress hormone regulation: biological role and translation into
M

therapy, Annu. Rev. Psychol. 61 (2010) 81–109.

[9] M. Frasconi, M. Mazzarino, F. Botrè, F. Mazzei, Surface plasmon resonance
immunosensor for cortisol and cortisone determination, Anal. Bioanal. Chem. 394
ED

(2009) 2151–2159.
[10] C. Leach, An overview of the endocrine and metabolic changes in manned space flight,
Acta Astronaut. 8 (1981) 977–986.
PT

[11] A. Omar, J. Rashid, A. Latif, K. Karim, O. Bakar, W. Yunus, Development of cortisol

immunosensor based reduced graphene oxide (rGO) for future application in monitoring
stress levels among military personnel, Def. S T Tech. Bull. 10 (2017) 142–148.
E

[12] E. Van Caenegem, K. Wierckx, T. Fiers, H. Segers, E. Vandersypt, J.M. Kaufman, G.

CC

T’Sjoen, Salivary cortisol and testosterone: a comparison of salivary sample collection

methods in healthy controls, in: Endocr Abstr, 2011: p. 355.
[13] E. Aardal, A.-C. Holm, Cortisol in Saliva - Reference Ranges and Relation to Cortisol
A

in Serum, Clin. Chem. Lab. Med. 33 (1995) 927–932.

[14] G.E. Abraham, J.E. Buster, L.A. Lucas, P.C. Corrales, R.C. Teller, Chromatographic
separation of steroid hormones for use in radioimmunoassay, Anal. Lett. 5 (1972) 509–
517.
[15] M. Kao, S. Voina, A. Nichols, R. Horton, Parallel radioimmunoassay for plasma cortisol
and 11-deoxycortisol, Clin. Chem. 21 (1975) 1644–1647.
[16] A. Lajtha, G.B. Baker, S. Dunn, A. Holt, Handbook of neurochemistry and molecular
neurobiology. Practical neurochemistry methods, Springer, 2007.
[17] K. Sun, N. Ramgir, S. Bhansali, An immunoelectrochemical sensor for salivary cortisol
measurement, Sensors Actuators B Chem. 133 (2008) 533–537.
[18] S.K. Arya, G. Chornokur, M. Venugopal, S. Bhansali, Dithiobis (succinimidyl
propionate) modified gold microarray electrode based electrochemical immunosensor
for ultrasensitive detection of cortisol, Biosens. Bioelectron. 25 (2010) 2296–2301.
[19] M. Yamaguchi, Y. Matsuda, S. Sasaki, M. Sasaki, Y. Kadoma, Y. Imai, D. Niwa, V.
Shetty, Immunosensor with fluid control mechanism for salivary cortisol analysis,

PT
Biosens. Bioelectron. 41 (2013) 186–191.
[20] R.D. Munje, S. Muthukumar, A. Panneer Selvam, S. Prasad, Flexible nanoporous
tunable electrical double layer biosensors for sweat diagnostics, Sci. Rep. 5 (2015)

RI
14586. doi:10.1038/srep14586.
[21] S.K. Tuteja, C. Ormsby, S. Neethirajan, Noninvasive Label-Free Detection of Cortisol

SC
and Lactate Using Graphene Embedded Screen-Printed Electrode, Nano-Micro Lett. 10
(2018) 41. doi:10.1007/s40820-018-0193-5.
[22] C. Li, Y. Liu, L. Li, Z. Du, S. Xu, M. Zhang, X. Yin, T. Wang, A novel amperometric

U
biosensor based on NiO hollow nanospheres for biosensing glucose, Talanta. 77 (2008)
455–459. doi:10.1016/j.talanta.2008.06.048.
N
[23] M. Tyagi, M. Tomar, V. Gupta, NiO nanoparticle-based urea biosensor, Biosens.
A
Bioelectron. 41 (2013) 110–115.
[24] S. Mohan, P. Srivastava, S.N. Maheshwari, S. Sundar, R. Prakash, Nano-structured
M

nickel oxide based DNA biosensor for detection of visceral leishmaniasis (Kala-azar),
Analyst. 136 (2011) 2845–2851.
ED

[25] S.K. Arya, M. Datta, B.D. Malhotra, Recent advances in cholesterol biosensor, Biosens.
Bioelectron. 23 (2008) 1083–1100.
[26] G. Kaur, M. Tomar, V. Gupta, Nanostructured NiO-based reagentless biosensor for total
cholesterol and low density lipoprotein detection, Anal. Bioanal. Chem. 409 (2017)
PT

1995–2005.
[27] N. Dhull, G. Kaur, V. Gupta, M. Tomar, Development of nanostructured nickel oxide
E

thin film matrix by rf sputtering technique for the realization of efficient bioelectrode,
Vacuum. 158 (2018) 68–74. doi:10.1016/J.VACUUM.2018.09.034.
CC

[28] Saliva Collection Handbook – Salimetrics, (n.d.). https://www.salimetrics.com/saliva-

collection-handbook/ (accessed August 23, 2018).
A

[29] P.M. Oliver, G.W. Watson, S.C. Parker, Molecular-dynamics simulations of nickel
oxide surfaces, Phys. Rev. B. 52 (1995) 5323–5329. doi:10.1103/PhysRevB.52.5323.
[30] B. Sasi, K.G. Gopchandran, Nanostructured mesoporous nickel oxide thin films,
Nanotechnology. 18 (2007) 115613.
[31] M.M. Rahman, A.J. Ahammad, J.-H. Jin, S.J. Ahn, J.-J. Lee, A comprehensive review
of glucose biosensors based on nanostructured metal-oxides, Sensors. 10 (2010) 4855–
4886.
[32] Y. Zhao, H. Wang, C. Wu, Z.F. Shi, F.B. Gao, W.C. Li, G.G. Wu, B.L. Zhang, G.T. Du,
Structures, electrical and optical properties of nickel oxide films by radio frequency
magnetron sputtering, Vacuum. 103 (2014) 14–16. doi:10.1016/j.vacuum.2013.11.009.
[33] X.H. Xia, J.P. Tu, J. Zhang, X.L. Wang, W.K. Zhang, H. Huang, Electrochromic
properties of porous NiO thin films prepared by a chemical bath deposition, Sol. Energy
Mater. Sol. Cells. 92 (2008) 628–633.
[34] A. Rahdar, M. Aliahmad, Y. Azizi, NiO nanoparticles: synthesis and characterization,
J. Nanostructures. 5 (2015) 145–151.
[35] E. San Andrés, A. Del Prado, F.L. Martınez, I. Mártil, D. Bravo, F.J. López, Rapid

PT
thermal annealing effects on the structural properties and density of defects in SiO 2 and
SiN x: H films deposited by electron cyclotron resonance, J. Appl. Phys. 87 (2000)
1187–1192.

RI
[36] Positional Fluctuation of IR Absorption Peaks: Frequency Shift of a Single Band or
Relative Intensity Changes of Overlapped Bands? | American Laboratory, (n.d.).
https://www.americanlaboratory.com/913-Technical-Articles/1244-Positional-

SC
Fluctuation-of-IR-Absorption-Peaks-Frequency-Shift-of-a-Single-Band-or-Relative-
Intensity-Changes-of-Overlapped-Bands/ (accessed June 28, 2018).
[37] S. Sudhasree, A. Shakila Banu, P. Brindha, G.A. Kurian, Synthesis of nickel
U
nanoparticles by chemical and green route and their comparison in respect to biological
effect and toxicity, Toxicol. Environ. Chem. 96 (2014) 743–754.
N
[38] J. Coates, Interpretation of infrared spectra, a practical approach, Encycl. Anal. Chem.
A
(2000).
M

[39] J. Kong, S. Yu, Fourier transform infrared spectroscopic analysis of protein secondary
structures, Acta Biochim. Biophys. Sin. (Shanghai). 39 (2007) 549–559.
[40] Z. Matharu, G. Sumana, S.K. Arya, S.P. Singh, V. Gupta, B.D. Malhotra, Polyaniline
ED

Langmuir−Blodgett Film Based Cholesterol Biosensor, Langmuir. 23 (2007) 13188–

13192. doi:10.1021/la702123a.
[41] A. Barth, Infrared spectroscopy of proteins, Biochim. Biophys. Acta (BBA)-
PT

Bioenergetics. 1767 (2007) 1073–1101.

[42] A.J. Bard, L.R. Faulkner, Electrochemical Methods, 2nd, J. Wiley. New York. (2001)
669.
E

[43] R. Guidelli, R.G. Compton, J.M. Feliu, E. Gileadi, J. Lipkowski, W. Schmickler, S.

CC

Trasatti, Defining the transfer coefficient in electrochemistry: An assessment (IUPAC

Technical Report), Pure Appl. Chem. 86 (2014) 245–258.
[44] E. Laviron, Surface linear potential sweep voltammetry: Equation of the peaks for a
A

reversible reaction when interactions between the adsorbed molecules are taken into
account, J. Electroanal. Chem. Interfacial Electrochem. 52 (1974) 395–402.
[45] J.N. (James N.. Miller, J.C. (Jane C. Miller, Statistics and chemometrics for analytical
chemistry, Pearson/Prentice Hall, 2005.
https://books.google.co.in/books/about/Statistics_and_Chemometrics_for_Analytic.ht
ml?id=Efx77dxOC3sC (accessed August 25, 2018).
[46] E. Mäkilä, P. Kirveskari, A study of ascorbic acid in human saliva, Arch. Oral Biol. 14
 (1969) 1285–1292. doi:10.1016/0003-9969(69)90201-5.
[47] V. Kalburgi, S. Leburu, S. Warad, Saliva as a surrogate to explore the association
between lipid profiles and chronic periodontitis: A case-control study., Dent. Res. J.
(Isfahan). 11 (2014) 619–23. http://www.ncbi.nlm.nih.gov/pubmed/25540654 (accessed
October 19, 2018).
[48] R. Segura, C. Javierre, J.L. Ventura, M.A. Lizarraga, B. Campos, E. Garrido, A new
approach to the assessment of anaerobic metabolism: measurement of lactate in saliva.,
Br. J. Sports Med. 30 (1996) 305–9. http://www.ncbi.nlm.nih.gov/pubmed/9015591
(accessed October 19, 2018).

PT
[49] Y. Xia, C. Peng, Z. Zhou, P. Cheng, L. Sun, Y. Peng, P. Xiao, [Clinical significance of
saliva urea, creatinine, and uric acid levels in patients with chronic kidney disease].,
Zhong Nan Da Xue Xue Bao. Yi Xue Ban. 37 (2012) 1171–6. doi:10.3969/j.issn.1672-
7347.2012.11.016.

RI
[50] C. Jurysta, N. Bulur, B. Oguzhan, I. Satman, T.M. Yilmaz, W.J. Malaisse, A. Sener,
Salivary glucose concentration and excretion in normal and diabetic subjects., J.

SC
Biomed. Biotechnol. 2009 (2009) 430426. doi:10.1155/2009/430426.
[51] K.R. Atkinson, K.R. Lo, S.R. Payne, J.S. Mitchell, J.R. Ingram, Rapid saliva processing
techniques for near real‐time analysis of salivary steroids and protein, J. Clin. Lab. Anal.
22 (2008) 395–402.
U
N
[52] V. Pinto, P. Sousa, S.O. Catarino, M. Correia-Neves, G. Minas, Microfluidic
immunosensor for rapid and highly-sensitive salivary cortisol quantification, Biosens.
A
Bioelectron. 90 (2017) 308–313. doi:10.1016/J.BIOS.2016.11.067.
M

[53] H. Wu, H. Ohnuki, S. Ota, M. Murata, Y. Yoshiura, H. Endo, New approach for
monitoring fish stress: A novel enzyme-functionalized label-free immunosensor system
for detecting cortisol levels in fish, Biosens. Bioelectron. 93 (2017) 57–64.
doi:10.1016/J.BIOS.2016.10.001.
ED

[54] J. Jeon, S. Uthaman, J. Lee, H. Hwang, G. Kim, P.J. Yoo, B.D. Hammock, C.S. Kim,
Y.-S. Park, I.-K. Park, In-direct localized surface plasmon resonance (LSPR)-based
nanosensors for highly sensitive and rapid detection of cortisol, Sensors Actuators B
PT

Chem. 266 (2018) 710–716. doi:10.1016/J.SNB.2018.03.167.

E
CC
A
 Figures

PT
RI
SC
U
N
Figure 1: Schematic for fabrication of the immunoelectrode (Ab/NiO/ITO)
A
and interaction of cortisol with the antibody.
M
ED
E PT
CC
A
 PT
RI
SC
U
N
A
M

Figure 2: Cyclic Voltammetry curves for bare ITO, NiO/ITO and Ab/NiO/ITO electrodes.
Inset shows the cyclic voltammetry curves for NiO/ITO for 10 multiple cycles.
ED
E PT
CC
A
 3 (a)
A
CC
EPT
ED
M
A
N
U
SC
RI
PT
 0.24 Ipo = - 0.016 log [cortisol concentration] + 0.164
0.26
Ab/NiO/ITO

Peak oxidation Current (mA)

0.24
1pg/mL
0.22
10pg/mL
0.20 0.20
100pg/mL
0.18 1ng/mL
0.16 10ng/mL
Current (mA)

0.14 100ng/mL
0.12
1ug/mL
0.16 0.10
Cortisol/NiO/ITO
Cortisol/Ab/NiO/ITO
Linear Fit
10ug/mL
0.08
1E-4 0.001 0.01 0.1 1 10 100 1000 10000 100000
Cortisol concentration (ng/mL)

PT
0.12

RI
0.08

SC
0.04

U
N
0.00
A
0.0 0.1 0.2 0.3 0.4
Voltage (V)
M

3 (b)
Figure 3: Biosensing response studies (a) CV (b) DPV of the Ab/NiO/ITO immunoelectrode
ED

by varying cortisol concentrations (Inset: Corresponding calibration curve)

E PT
CC
A
Figure 4: (a) Selectivity against common interferents and (b) reproducibility studies of the
prepared immunoelectrode (c) Shelf-life studies of the cortisol immunoelectrode
(Ab/NiO/ITO) (d) Stability of the immunoelectrode with repeated measurements of cortisol

PT
RI
SC
U
N
4 (a)
A
M
ED
E PT
CC
A

4 (b)
 PT
RI
SC
4 (c)
U
N
A
M
ED
E PT
CC
A

4 (d)
Figure 4: (a) Selectivity against common interferents and (b) reproducibility studies of the
prepared immunoelectrode (c) Shelf-life studies of the cortisol immunoelectrode
(Ab/NiO/ITO) (d) Stability of the immunoelectrode with repeated measurements of cortisol
 PT
RI
SC
U
N
5 (a)
A
Subject 5
M

Subject 4
ED

Subject 3
PT

Subject 2
ELISA
Sensor
E

Subject 1
CC

0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5

Cortisol Concentration (ng/mL)
A

5 (b)

Figure 5: Salivary cortisol measurements: (a) response of the immunoelectrode towards

varying dilutions of real saliva samples. Inset in the figure shows the corresponding cyclic
voltammetry curves. (b) Real saliva sample measurements using prepared sensor and
standard ELISA assay.
A
CC
EPT
ED
M
A
N
U
SC
RI
PT
Table 1: Assessment of the analytical performance of the cortisol biosensor devised in the
present work compared to those previously reported in the literature.

Sensing platform Technique Detection Linear Analysis Reference

(Interaction Limit Range Time
Type)

PT
Au electrodes Electrochemical 0.76 10 min [17]
functionalized by (Ab-Ag) nmol/L
alkaline phosphate

RI
enzyme
MUA, HC80 coating Optical (Ab-Ag) 4 µg/L 9 to 132 20 min [9]
µg/L

SC
interdigitated u- Electrochemical 1 pM 1 pM to 30 min [18]
electrodes (Ab-Ag) 10 nM
functionalized by
dithiobis(succinimidyl
propionate) U
N
Glucose oxidase Electrochemical 0.1 to 10 35 min [19]
labelled cortisol (Ab-Ag) ng/mL
A
conjugate
ZnO thin film on Electrochemical 1 pg/mL 10 to 200 15 min [20]
M

nanoporous (Ab-Ag) ng/mL

polyamdie membrane
Colorimetric detection Optical (Ab- 18 pg/mL 0.01 – 20 1 Hr [52]
ED

using CMOS HRP labelled ng/mL

photodiode cortisol)
Enzyme-labelled Electrochemical 1.01 1.25 to 25 min [53]
SAM (Enzyme) ng/mL 200
PT

ng/mL
LSPR-based cuvette- Optical (Ab-Ag) 8 ng/mL 1 to 104 1 Hr [54]
type sensor ng/mL
E

Graphene embedded Electrochemical 0.1 ng/mL 0.1 to 200 2 min [21]

screen-printed (Ab-Ag) ng/mL
CC

electrode
Nanostructured NiO Electrochemical 0.32 1pg/mL 2 Hrs Present
thin film (Ab-Ag) pg/mL to work
A

(0.89 pM) 10µg/mL

(2.75 pM
to 27.5
mM)