MWCNTs sensor for highly-sensitive cardiac troponin T detection

Journal Pre-proof A novel molecularly imprinted polymer PMB/MWCNTs sensor for highly-sensitive cardiac troponin T detection Kewarin Phonklam (Methodology) (Investigation) (Formal analysis) (Writing - original draft), Rodtichoti Wannapob (Methodology) (Visualization), Wilaiwan Sriwimol (Validation) (Resources), Panote Thavarungkul (Conceptualization)Writing reviewing and editing) (Supervision), Tonghathai Phairatana (Conceptualization)Writing - reviewing and editing) (Supervision) (Project administration)

PII:

S0925-4005(19)31829-5

DOI:

https://doi.org/10.1016/j.snb.2019.127630

Reference:

SNB 127630

To appear in:

Sensors and Actuators: B. Chemical

Received Date:

28 August 2019

Revised Date:

11 December 2019

Accepted Date:

23 December 2019

Please cite this article as: Phonklam K, Wannapob R, Sriwimol W, Thavarungkul P, Phairatana T, A novel molecularly imprinted polymer PMB/MWCNTs sensor for highly-sensitive cardiac troponin T detection, Sensors and Actuators: B. Chemical (2019), doi: https://doi.org/10.1016/j.snb.2019.127630

This is a PDF file of an article that has undergone enhancements after acceptance, such as the addition of a cover page and metadata, and formatting for readability, but it is not yet the definitive version of record. This version will undergo additional copyediting, typesetting and review before it is published in its final form, but we are providing this version to give early visibility of the article. 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. © 2019 Published by Elsevier.

A novel molecularly imprinted polymer PMB/MWCNTs sensor for highly-sensitive cardiac troponin T detection Kewarin Phonklam1, Rodtichoti Wannapob2,3, Wilaiwan Sriwimol4,

1

of

Panote Thavarungkul2,3** and Tonghathai Phairatana1,5* Institute of Biomedical Engineering, Faculty of Medicine, Prince of Songkla University, Hat

Center of Excellence for Trace Analysis and Biosensor, Prince of Songkla University, Hat Yai,

-p

2

ro

Yai, Songkhla, 90110, Thailand

Songkhla, 90112, Thailand

Department of Physics, Faculty of Science, Prince of Songkla University, Hat Yai, Songkhla,

re

3

4

Department of Pathology, Faculty of Medicine, Prince of Songkla University, Hat Yai,

ur na

Songkhla, 90110, Thailand 5

lP

90112, Thailand

Medical Science Research and Innovation Institute, Prince of Songkla University, Hat Yai,

Songkhla, 90112, Thailand

Jo

*Corresponding author: [email protected] ** Corresponding author: [email protected]

Highlights  A novel cTnT-molecularly imprinted polymer (MIP) sensor was fabricated  on a screen-printed carbon electrode. 1

    

Key parameters involved in cTnT-MIP sensor fabrication were optimized via a step-wise approach. The developed cTnT-MIP sensors exhibited excellent sensitivity, selectivity and binding affinity. The novel MIP sensor provides a high potential for the development of cTnT  point-of-care testing.

of

Abstract

ro

Cardiac troponin T (cTnT) is a cardiac biomarker introduced for early diagnosis, patient follow-

-p

up and guidelines for acute myocardial infarction treatment. In this work a novel cTnTmolecularly imprinted polymer (MIP) based electrochemical sensor using a screen printed

re

carbon electrode (SPCE) was developed. The MIP sensor employed an electrodeposited polymethylene blue (PMB) redox probe on multi-walled carbon nanotubes (MWCNTs) modified

lP

SPCE, with the electropolymerized polyaniline around the immobilized cTnT templates. The sensor response was performed using differential pulse voltammetry, wherein the decrease in the

ur na

PMB current correlated with an increase in the concentration of cTnT. Linearity was obtained in the range of 0.10-8.0 pg mL-1 with a detection limit of 0.040 pg mL-1. The MIP sensor exhibited an excellent binding affinity (KD = 2.810-13 mol L-1), high selectivity and good stability, whilst retaining more than 90% of the sensitivity after 6 weeks of storage at room temperature. The

Jo

determination of cTnT in human plasma using the developed sensor compared well with the gold standard electrochemiluminescense method (P > 0.05). This novel cTnT-MIP sensor provides a high potential for the development of cTnT point-of-care testing applications.

2

Keywords: Cardiac troponin T; Molecularly imprinted polymer; Electrochemical biosensor; Electropolymerized polyaniline methylene blue; Screen printed carbon electrode; Synthetic

Jo

ur na

lP

re

-p

ro

of

receptor

3

1. Introduction Cardiovascular disease is one of the leading causes of morbidity and mortality within today’s world population, especially acute myocardial infarction (AMI), which is a type of myocardial necrosis resulting from an ischemic syndrome. Among patients with AMI, the mortality rate is higher than 40% [1]. Currently the diagnosis of AMI, according to the World Health Organization, depends on the following criteria: the clinical history of chest discomfort,

of

electrocardiogram (ECG) changes and variations in the amount of biomarkers of myocardial

ro

infarction [2]. At present, the detection of biomarkers plays an important role because symptoms of chest pain and ECG changes may disappear or may not be specific to AMI [3,4]. Owing to the

-p

high sensitivity coupled with specificity, cardiac troponin T (cTnT) has been widely used as a

re

crucial cardiac biomarker for AMI diagnosis. This is due to its prominent release in the bloodstream during cardiac ischemia [5,6]. The cTnT level in healthy people is less than 10 pg

lP

mL-1, therefore its elevation, adjoined with likely pathology, can be utilized to predict the presence of AMI [7]. Hence, an early, rapid and accurate detection of this biomarker would be

ur na

beneficial for AMI diagnosis. As a consequent, this helps increase the survival rate, reduces healthcare cost coupled with time on the prognosis of the disease. Both commercial use and research on the quantitative cTnT detection are mostly based on its capture by natural antibodies, such as enzyme linked immunosorbent assay (ELISA) [7],

Jo

electrochemiluminescence immunoassay (ECLIA) [8], fluorescence immunoassay [9] and surface plasmon resonance (SPR) [10]. Because the use of antibodies contributes to their high costs and relatively low stability [11,12], one way to resolve these issues is via the use of molecularly imprinted polymers (MIPs), these being: synthetic materials with recognition sites possessing specific cavities that are designed for a target molecule, which offers a more cost-

4

effective approach and long-term stability [13–15]. MIPs are synthesized using polymerization of functional monomers in the presence of a target molecule acting as a template. Removal of the templates creates polymeric cavities with shape, size and chemical functionality for further selective rebinding to the target molecules. Attention on the fabrication of electrochemical MIPs biosensors is evidenced [16], due to their many benefits, such as high stability, sensitivity and

of

selectivity, reusability, inexpensiveness, and low limit of detection (LOD) [17]. In particular, when the MIPs is integrated with a small, portable electrochemical device for point-of-care

ro

diagnosis [18–20].

An interesting strategy for the development of protein-imprinted polymer is

-p

electropolymerization, which presents a simple and controllable approach for the direct

re

deposition of the MIP on an electrode [21–23]. The polymer thickness can be well-controlled by adjusting parameters that can affect the amount of charge passed for the MIP synthesis, leading easy

protein

removal

[24].

Polypyrrole

lP

to

[25],

polyaniline

(PANI),

poly(3,4-

ethylenedioxythiophene) [26] and polythiophene [27] are conducting polymers, that have been

ur na

extensively used in a micro- or nanostructured MIP electrosynthesis. Among these, PANI with its desirable features, such as easy synthesis, good chemical and mechanical stability and electrical conductivity [28,29], has exhibited good performance for a protein surface-imprinting sensor [30,31]. This is because of its protonated amine groups that can interact with the

Jo

carboxylic groups of protein.

There have been some recent studies on the electrosynthesis of MIPs for cTnT

recognition with the use of exterior redox probe (ferrocyanide/ferricyanide, [Fe(CN)6]3-/4-) to obtain the electrochemical signals that showed good affinity and selectivity [20,21]. However, for point-of-care applications, redox electroactive probe deposited on the modified electrode

5

surface is a more suitable alternative. Various electroactive polymer redox probes have been reported to be able to enhance the performance of electrochemical sensors [32–34], wherein polymethylene blue (PMB) has been presented as an excellent choice [35,36]. Additionally, the use of nanomaterials, that can enhance the specific surface area and electrical conductivity, can help improve the performance of MIP electrochemical sensors. Multi-walled carbon nanotubes (MWCNTs), in combination with electroactive polymers, have been widely used since they can

of

promote the electrical current along with mechanical properties of polymers [37].

ro

This work is the first report of a layer-by-layer cTnT-MIP sensor in which the electrochemical signals were obtained through the MWCNTs in combination with the directly

-p

electrodeposited polymethylene blue (PMB) and a controllable electropolymerized MIP layer.

re

This is a more suitable candidate for a point-of-care device. The MWCNTs increased the electrode surface area for the MIP coating as well as enhanced the electron transfer rate of PMB.

lP

In the fabrication of the MIPs, two layers of polyaniline (PANI) were controllably prepared by electropolymerization. The first was to provide an immobilization platform for the human cTnT

ur na

templates while the thickness of the second layer was fine-tuned by the electropolymerization procedure to cover the cTnT molecules. This approach offers simple, controllable preparation and a calculable thickness of the PANI, with good stability and low cost MIP materials.

Jo

Experimental parameters involved in both the MIP process and redox probe coating; such as, MIP film thickness were evaluated and optimized via a step-wise approach. Finally, the developed cTnT-MIP sensor was tested with human plasma samples, upon which it was then compared to the gold standard method.

2. Experimental section 6

2.1 Reagents and apparatus Cardiac Troponin T (cTnT, ≥95%), cardiac troponin I (cTnI, ≥60%, Calbiochem®, Germany), methylene blue (MB, Gurr Certistain, England), aniline (ANI, 99%, Sigma-Aldrich, USA), hydrochloric acid (HCl, 37%, Merck, Germany), dimethylformamide (DMF, Fisher, USA), glutaraldehyde (25% in H2O, Sigma-Aldrich) and lithium perchlorate (Sigma-Aldrich)

of

were used as received. Chemicals for interference testing and other substances were obtained from Sigma-Aldrich (Steinheim, Germany). Carboxylic functionalized multi-walled carbon

ro

nanotubes (f-MWCNT) were from Shenzhen Nano-Technologies Port Co., Ltd. (Shenzhen,

-p

China). Standard stock solution of cTnT, 50 µg mL-1, was prepared in 10 mmol L−1 phosphatebuffered saline solution (PBS) at pH 7.40 and stored at −20°C when not in use. The PBS

re

supporting electrolyte consisted of di-sodium hydrogen orthophosphate and sodium hydrogen orthophosphate (Ajax Finechem, USA), with 100 mmol L-1 potassium chloride (Merck,

lP

Germany) (100 mmol L-1, pH 7.00). All aqueous solutions were prepared using water purified with a Milli-Q purification system (resistivity ≥18 MΩ cm).

ur na

Electrochemical experiments were conducted using a computer-controlled potentiostat µAutolabIII/FRA2, monitored by the NOVA 1.11 software (Metrohm, Netherland). Screenprinted carbon electrodes (SPCEs) were from DropSens (DRP-C110DIEL). Each included a

Jo

carbon disk working electrode (4 mm in diameter), a silver pseudo-reference electrode and a carbon counter electrode. Scanning electron micrographs were obtained using a Field Emission Scanning Electron Microscope (FE-SEM, Apreo, FEI Company, USA). Atomic force microscope (AFM) images were received using SEIKO-SPA400 (Seiko instruments Inc., Japan). All experiments were carried out at room temperature (25°C).

7

2.2 cTnT-MIP sensor fabrication Preliminary cyclic voltammetric study of the PMB on a glassy carbon electrode with and without the f-MWCNTs showed a three times higher current signal of the PMB/f-MWCNT/GCE than that of the PMB/GCE (Supplementary data Fig.S1), i.e., the f-MWCNTs modified electrode

of

can contribute to the PMB oxidation signal. It has been reported that f-MWCNTs significantly

ro

increase the deposition amount of PMB [38], leading to the enhancement of the electron transfer rate and also the reduction of the PMB degradation [39]. Hence, f-MWCNTs were used to

-p

modify the SPCE surface before the electrodeposition of PMB in this study.

Fabrication steps of the developed cTnT-MIP sensor are schematically illustrated in Fig.

re

1A. A SPCE was pre-treated by applying a 1.2 V constant potential for 3 min in saturated

lP

Na2CO3 to remove excessive organic substances on the electrode surface [40], rinsed with deionized water and dried with nitrogen gas. The working surface was modified by 3.0 µL of fMWCNTs (5.0 mg mL-1 in DMF), and then allowed to dry in an incubator at 35°C for 30 min.

ur na

The electropolymerization of a PMB redox probe layer on the f-MWCNTs was then performed in 0.25 mmol L-1 MB in 100 mmol L-1 PBS, pH 7.00, by applying a fixed potential of 1.5 V for 900 s [41]. This was followed by applying a range of potential between -0.70 and 0.40 V for 30

Jo

cycles at a scan rate of 100 mV s-1. After this step, the PMB/f-MWCNTs/SPCE was washed with PBS and dried with nitrogen gas. For the MIP, the PMB/f-MWCNTs/SPCE was modified with two layers of polyaniline

(PANI). In general, the polymerization of aniline is performed in acidic media [42,43]. Thus, the first PANI layer was electropolymerized by 10 cycles of cyclic voltammetric scanning between 0.20 and +1.0 V at a scan rate of 50 mV s-1 in a simply prepared 5.0 mmol L-1 aniline in 200 8

mmol L-1 HCl. This layer acts as an anchor platform with amine groups, activated with 2.0% glutaraldehyde at room temperature for 30 min [44,45]. After washing with de-ionized water, a 5.0 µL of 50 µg mL-1 cTnT was drop-casted on the SPCE and left overnight at 4°C. After washing several times with de-ionized water, the other PANI layer was electrodeposited to form the MIP. To avoid the denaturation of cTnT templates in strong acid, the electropolymerization of the second PANI layer was performed in 10 mmol L-1 PBS (pH 5.80) containing 10 mmol L-1

of

aniline, with the addition of 5.0 mmol L-1 LiClO4 [46] for 20 scan cycles. The cTnT templates

ro

were removed by immersing the cTnT-PANI/PMB/f-MWCNTs/SPCE in 500 mmol L-1 acetic acid for a period of time that provided the maximal removal (see section 3.2), these were then

-p

rinsed with PBS under stirring conditions. The cTnT-PANI/PMB/f-MWCNTs/SPCE is called:

re

‘MIP sensor’ in the following section. For comparison, a non-imprinted polymer (NIP) SPCE

lP

was constructed using the same steps, but without the cTnT template.

ur na

2.3 Electrochemical measurements

Electrode modification steps were evaluated by cyclic voltammetry between -0.60 and 0.10 V at a scan rate of 50 mVs-1 in 100 mmol L-1 PBS. The determination of the binding of cTnT, either to the MIP or NIP sensor, was by adding 40 µL of standard cTnT solution,

Jo

incubated for 30 min, before washing with PBS. The binding response was determined through the oxidation current of MB by differential pulse voltammetry (DPV) between –0.60 V and 0.10 V, with a pulse amplitude of 50 mV in 10 mL of 100 mmol L-1 PBS. The binding between the imprinted sites and the target cTnT impeded the electron transfer of the oxidation current of MB to the electrode surface (Fig. 1B). The response was taken as the change of current which related to the amount of cTnT molecules. 9

2.4 Optimization studies A set of parameters, which can influence the properties and performance of the MIP sensor, were optimized in the range of 0.10-100 pg mL-1 cTnT. The amount of deposited fMWCNTs was first studied with 5.0, 10, 15, 20 and 25 µg. For this f-MWCNTs experiment, PMB, the first PANI and second PANI layers were electropolymerized at 30, 10 and 20 scan

of

cycles, respectively. The thickness of PMB was then optimized by applying 30, 40, 50 and 60 electropolymerized scan cycles. Another factor that was taken into account was the thickness of

ro

the two PANI layers, as the first layer was optimized for the number of electropolymerization

-p

scans at 5, 6, 7, 8, 9 and 10 cycles, whereas the second layer was tested at 5, 10, 15, 20 and 25 cycles. The optimal condition was the one that obtained the highest sensitivity (slope of the

re

calibration curve). To ensure maximal removal of the cTnT templates, removal time was also

ur na

2.5 Selectivity

lP

optimized by observing the PMB DPV signal in 0.10 mol L-1 PBS (pH 7.00).

The MIP sensor was tested with human blood plasma samples, thus, its selectivity to cTnT was examined for possible blood interfering substances. Glucose (Glu), ascorbic acid (AA), creatinine (Cr) and uric acid (UA) as well as human serum albumin (HSA) were tested at

Jo

higher levels than their physiological concentrations, in comparison to 1.0 pg mL-1 cTnT. Cardiac troponin I (cTnI), another cardiac biomarker of AMI was also investigated since it can be found at the same level as cTnT in AMI blood plasma [3]. This was examined at 50 and 100 times higher than the tested cTnT concentration. These molecules were analyzed both separately as well as when mixing with cTnT.

10

2.6 Reproducibility, long-term and storage stabilities Six electrodes were fabricated under the same conditions, and used to determine cTnT in the range of 1.0-8.0 pg mL-1 for the reproducibility test. The long-term stability was evaluated by considering the sensitivity of one sensor by testing it once a week; these were kept in a closed, dry chamber at room temperature between each test. So as to take into consideration storage

of

stability, nine sensors were prepared at the same time, these were then stored in a closed, dry

ro

chamber at room temperature, and their sensitivities were examined one by one, each week.

-p

2.7 Real sample analysis

re

Blood plasma samples from Songklanagarind Hospital, Hat Yai, Thailand were analysed (approved by the local ethics committee (REC.61-143-25-2)). The plasma samples were simply

lP

pretreated by dilution in 10 mmol L-1 PBS (pH 7.40) before testing. The matrix effect was first evaluated by spiking a series of cTnT concentrations (before dilution). Both the slope of the

ur na

standard calibration curves as well as the spiked blood plasma curves were compared by twoway ANOVA, so as to clarify any matrix effects. Samples, with a 100-times dilution factor, were then tested by the MIP sensor. Results from the above were then compared with the values obtained from the gold standard electrochemiluminescence immunoassay (ECLIA), by the

Jo

Wilcoxon signed-rank test.

11

3. Results and discussion 3.1 Characterization of the MIP sensor 3.1.1 Surface morphology The surface of the modified SPCE was observed using FE-SEM. The fibrous structure fMWCNTs layer showed an average diameter of 53±4 nm (Fig. 2A). The PMB

of

electropolymerization made the fibrous-like structure thicker, with a larger diameter of 74±6 nm (Fig. 2B). When the PANI was electrodeposited (MIP layer), small beads were observed (Fig.

-p

ro

2C), similar to those reported earlier [47].

3.1.2 Electrochemical characterization

re

The stepwise modification of the developed MIP sensor was electrochemically characterized, wherein both the cyclic voltammograms (CVs) of the SPCE, (Fig. 2D, curve a)

lP

and after modifying with f-MWCNTs (curve b) showed no redox peaks. However, the coating of f-MWCNTs provided higher background current, indicating an increase in the conductivity of

ur na

the electrode surface. With a layer of PMB its oxidation and reduction peaks can be clearly observed at -0.20 V and -0.42 V, respectively (curve c). These current peaks decreased after the electropolymerization of the first PANI layer (curve d), due to the blocking of electron transfer by the polymer on the PMB surface. The PMB redox current peaks further decreased when

Jo

glutaraldehyde and cTnT molecules were added (curve e), this indicated the successful attachment of cTnT, thus, hindering the electron transfer to the electrode surface. When the second PANI film was deposited, the redox current signal of PMB disappeared, and revealed a higher background current (curve f). This confirmed the covering of the second PANI layer on the previous layers. After cTnT removal, the PMB redox current peaks increased, confirming the

12

successful fabrication of the MIP layer with cTnT cavities which in turn allowed more electron transfer to the electrode surface (curve g). The characterization of cTnT removal was also confirmed by DPV and electrochemical impedance spectroscopy (EIS). After dipping the modified electrode in 500 mmol L-1 acetic acid solution for 4 h [48] the oxidation peak of PMB increased (Supplementary data Fig. S2A). This suggests that the removal of cTnT molecules creates the cavities for charges to penetrate through

of

the MIP layer to reach the PMB layer. This is consistent with the EIS results, by considering the

ro

charge transfer resistance (Rct) obtained from the diameter of the Nyquist plot semicircle. The Rct visibly decreased from 97.3±4.1 Ω to 7.2±1.2 Ω (Supplementary data Fig. S2C) due to the cTnT

-p

cavities of the imprinted polymer film. In contrast, both DPV and EIS results, from before and

re

after cTnT removal were unchanged for the NIP sensor (Supplementary data Fig. S2B and Fig. S1D, respectively). This is as expected, since there were no binding cavities in the polymer film.

lP

Additionally, AFM was used to investigate the surface topography. The AFM image of the MIP revealed an observable rougher surface (Supplementary data Fig. S2E), with a root

ur na

mean square (RMS) surface roughness value of 4.22 nm compared to 2.34 nm of the NIP surface (Supplementary data Fig. S2F). This supported that the templates were removed from the MIP

Jo

film after washing.

3.2 Optimization study Parameters affecting the performance of the MIP sensor were evaluated as follows.

3.2.1 Amount of f-MWCNTs

13

The amount of f-MWCNTs is an important parameter to achieve a large surface area for PMB formation, and a high capacity of the MIP binding site. As shown in Fig. 3A, the sensor sensitivity was higher with the increased amount of f-MWCNTs due to the larger surface area of f-MWCNTs on the electrode surface. The highest sensitivity was at 20 µg, at which point it became stable, indicating that this amount of f-MWCNTs was sufficient. Therefore, 20 µg f-

of

MWCNTs was chosen for further fabrication of the MIP sensors.

ro

3.2.2 PMB thickness

The thickness of PMB also plays an essential role, since it acts as a redox probe that

-p

generated the current signal, which is related to cTnT concentrations. Cyclic voltammograms

re

recorded during the electropolymerization of MB on f-MWCNTs showed the increase of electrochemical response of PMB with the number of scan. The PMB surface coverage (Γ) can

lP

be estimated from equation (1) [41]:

𝑄

𝛤 = 𝑛𝐹𝐴

(1)

ur na

where Q is the total charge involved in the electropolymerization process, n is the number of transferred electrons, which is equal to 2, F is Faraday's constant (96485 C mol-1) and A is the electrode surface area (0.090 cm2). The film thickness (d) of PMB was then calculated from Γ

Jo

by equation (2) [49]:

𝑑 =𝑣 ×𝛤

(2)

where v is the molecular volume of MB in the polymer, 400 cm3 mol-1 [50]. Γ of the 30, 40, 50 and 60 scans were 4.4, 5.6, 5.9 and 6.5 nmol cm-2, and the thickness of PMB films were 17.6, 22.5, 24.7 and 26.8 nm, respectively. From the results, excess PMB thickness caused a decrease in sensitivity after 40 scan cycles (Fig. 3B). This is mainly due to the blocking of electron

14

transfer by the thicker layer of MB at 50 and 60 scan cycles, resulting in a smaller MB response signal (Supplementary data Fig. S3A and S3B). In addition, the PMB oxidation currents tested before the MIP fabrication showed stable signals where the average of eight successive DPV peak currents had a RSD of 1.3% (Supplementary data Fig. S3C). Therefore, the optimal condition was chosen to be 40 scans because it gave a maximal current response adjoined with

of

the highest sensitivity.

ro

3.2.3 PANI layer

Thickness of the two PANI layers has a potential influence on the performance of the

-p

MIP sensor. The first layer provides amino groups for glutaraldehyde crosslinking with the

re

template, thus, affecting the amount of the immobilized cTnTs. Excessive PANI coating could reduce the electrochemical signal of PMB, the greatest sensitivity being displayed at 7 cycles

lP

(Fig. 3C).

The second PANI layer, formed after the immobilization of cTnT on the electrode

ur na

surface, needs to be thick enough to create the molecular cavities for cTnT detection. The thickness of the second PANI layer was estimated from the amount of charge under the electropolymerization process by equation (3) [51]: 𝑄𝑀 𝑛𝐹𝐴𝜌

(3)

Jo

𝑑=

where Q is the total charge in the electropolymerization, 𝑀 is molecular weight of aniline, n is number of electrons (0.5), F is Faraday's constant, A is electrode surface area and  is aniline density. PANI thickness of 5, 10, 15, 20 and 25 scanning cycles were 3.2, 3.8, 4.4, 4.9 and 5.1 nm, respectively. As shown in Fig. 3D, the highest sensitivity is presented at 10 scan cycles. At 5 scans, the thickness of the PANI layer might not be sufficient to surround the cTnT molecules, 15

thus, the imprinting cavity could not readily bind to the target. It was most likely that at 15, 20 and 25 scan cycles, the PANI layer might be too thick, hindering the electron transfer, and leading to the decrease of their sensitivities. This is confirmed by the decrease of peak current, and the shift of the peak potential of the baseline signal to a more positive value from -0.35 to 0.30, -0.25, -0.10, and -0.10 V, respectively (Supplementary data Fig. S4), indicating the hindering of electron transfer to the electrode surface from the PANI layer with higher number of

of

scanning cycles. In the following experiments, 7 and 10 cycles were applied as the optimal

ro

conditions for the first and second PANI layers, respectively.

-p

3.2.4 The removal time

re

The removal time of cTnT templates was tested between 1-6 h. Fig.3E shows stable

ur na

3.3 Analytical performances

lP

current signals after 4 h, i.e., the maximal removal of cTnTs was achieved at 4 h.

3.3.1 Linearity, limit of detection and binding isotherm The analysis of cTnT was validated under the optimum electrode modification conditions. Examples of the DPV response are shown in Fig. 4A where the oxidation current of PMB

Jo

decreased with the increase of cTnT concentrations. The plot of the decreased current (∆I) of the PMB oxidation signal, versus cTnT concentration showed a linear relationship in the range of 0.10 to 8.0 pg mL-1, with a correlation coefficient of 0.999. The limit of detection (LOD) was found to be 0.040 pg mL-1 (3S/N)[52]. On the other hand, the NIP sensor showed almost no change in current response for all concentrations of cTnT.

16

The relationship between the resulting current response of imprinted sites and the cTnT concentrations can be used to predict the affinity of the imprinted sites and the cTnT molecules in equilibrium by considering a binding isotherm. A Langmuir isotherm model was considered [53] as shown in equation (4) 𝐼𝑚𝑎𝑥 [𝐹]

(4)

𝐾𝐷 +[𝐹]

of

𝐼𝐹 =

where IF is the current response with the bound cTnTs, F is the concentration of the cTnT (mol L), Imax is the maximum current response when cTnT binding site is saturated and KD is the

ro

1

dissociated constant (mol L-1). The linearized Langmuir plotting of 1/IF and 1/[F] [54,55] showed

-p

a good fit, with a correlation coefficient of 0.999, providing a KD of 2.810-13 mol L-1

re

(Supplementary data Fig. S5). Using the graph-fitting GraphPad Prism 8 software, the resulting cTnT binding response of the MIP sensor also showed a relatively good fit to the Langmuir

1

lP

isotherm with an R-squared of 0.992 (Fig. 4B). The KD was also calculated to be 2.810-13 mol L. A lower KD indicates tighter binding (higher affinity) for cTnT molecules. Compared to other

ur na

studies (Table 1), the developed MIP sensor provided a higher affinity for cTnT.

3.3.2 Selectivity

Jo

The selectivity study was evaluated by determining possible interfering substances that can be found in human blood plasma. Both smaller (cTnI, Glu, AA, Cr and UA) along with larger (HSA) molecules than cTnT were considered, especially; cTnI, which has a similar molecular structure to cTnT [56]. There was concern that smaller molecules could move into the cavities of the MIP sensor, while HSA, the main protein within human blood plasma, may be adsorbed onto the electrode surface. The results showed no interfering effect from all tested 17

molecules (Fig. 4C). That is, the imprinted sites are very selective and rebind to analyte with specific sizes and shapes.

3.3.3 Reproducibility, long-term and storage stabilities Sensitivities of the six MIP sensors were determined to be 0.3857±0.0078,

of

0.3837±0.0092, 0.398±0.010, 0.3880±0.0084, 0.3949±0.0061 and 0.3867±0.0092 µAmL ng-1.

ro

The average sensitivity, with a RSD of 1.4%, indicated an acceptable reproducibility (Supplementary data Fig. S6A).

-p

The long-term stability of one cTnT-MIP sensor was examined from its relative

re

sensitivity. During the first 3 weeks, it was found to be more than 90% with a 2.7% RSD (Supplementary data Fig. S6B). The decline in the relative response may be owing to the loss in

lP

the properties of some modified layers on the SPCE, especially PMB. It was noticed that on the 4th week the initial PMB oxidation peaks were reduced (Supplementary data Fig. S6C).

ur na

In regards to storage stability, the relative sensitivity of the sensor was > 90% up to 6 weeks, with an average of 96.2±4.2 % (RSD = 4.4%) (Supplementary data Fig. S6D). At the 8th week, 83.2 % of the sensitivity still remained. That is the developed MIP sensors exhibited

Jo

excellent storage stability, even at room temperature.

3.3.4 Comparing to other sensors The analytical performance of the MIP/PMB/f-MWCNTs/SPCE as a cTnT sensor was

compared with other MIP sensors for cTnT detection, as shown in Table 1 . Our developed MIP sensor exhibited the lowest limit of detection, two orders of magnitude lower than the best one in these cited reports (Table 1). Although the linear range was not as wide as some similar work 18

[21], it was however, more than sufficient for the cTnT detection, even when tested with the 100 times diluted human blood plasma samples in patients with AMI. In the case of stability, this sensor offered better performance than other cited reports. In addition, the developed MIP sensor showed a relatively low KD value, indicating a better affinity to cTnT binding, than that of

of

sensors assessed in the cited works.

3.3.5 Analysis of human blood plasma samples

ro

To show the applicability of the developed MIP sensor, human blood plasma samples were tested. AMI patients usually have an elevated cTnT level of >10 pg mL-1 [7]. In this study,

-p

the standard curve of the MIP sensor exhibited a linear range of 0.10-8.0 pg mL-1, therefore

re

human blood plasma samples were diluted 100 times before testing. The study of the matrix effect was first performed by comparing a calibration plot of standard cTnT solution and spiked

lP

sample solutions at 100 times dilution. The sensitivity of the spiked samples presented no significant difference to that of standard cTnT (P > 0.05 two-way ANOVA). Hence, there was no

ur na

matrix effect at this dilution. The cTnT values of the diluted samples were then calculated from the calibration equation of the cTnT standard curve. The cTnT values of eight ‘real’ samples obtained from the MIP sensor and the gold standard electrochemiluminescence immunoassay (ECLIA) in clinical use (from the hospital) were in good agreement (P > 0.05) (Fig. 5).

Jo

The MIP sensor was also validated by analyzing recovery from spiking cTnT (100, 200,

400, 600 and 800 pg mL-1) into three human blood plasma samples, so as to obtain the final concentrations after 100 times dilution (1.0, 2.0, 4.0, 6.0 and 8.0 pg mL-1). Recoveries were in the range of 91–112% (Supplementary data Table.1), which fall into the acceptable range (40120%) [57]. These results indicate that the MIP sensor has high potential for cTnT detection.

19

4. Conclusions This study presents the successful fabrication of cTnT-MIP based electrochemical biosensor, using a PMB redox probe modified on SPCE. The developed MIP sensors showed excellent sensitivity, selectivity and binding affinity for the cTnT detection. Without an exterior

of

redox probe, the sensor achieved a wide linear range of 0.10-8.0 pg mL-1 with limit of detection of 0.040 pg mL-1. This low LOD allowed the sensor to be simply applied for the direct detection

ro

of cTnT in a diluted complex, ‘real’ sample. This biomimetic sensor technique, integrated with

-p

an inexpensive SPCE has high potential for cTnT point-of-care testing.

re

Credit Author Statements

ur na

lP

Kewarin Phonklam: Methodology, Investigation, Formal Analysis, Writing-Original draft preparation. Rodtichoti Wannapob: Methodology, Visualization. Wilaiwan Sriwimol: Validation, Resources. Panote Thavarungkul: Conceptualization, Writing-Reviewing & Editing, Supervision. Tonghathai Phairatana: Conceptualization, Writing-Reviewing & Editing, Supervision, Project Administration.

Declaration of interests The authors declare that they have no known competing financial interests or personal

Jo

relationships that could have appeared to influence the work reported in this paper.

Acknowledgements This work was supported by the grant from Faculty of Medicine, Prince of Songkla University (MED6104081S). We gratefully acknowledged the student scholarship from Faculty of Medicine and the Graduate School, Prince of Songkla University awarded to Ms. Kewarin Phonklam. We

20

would also like to thank the Center of Excellence for Trace Analysis and Biosensor, and Medical Science Research and Innovation Institute at Prince of Songkla University, Thailand. Thanks also to the International Affairs Office of the Faculty of Medicine, Prince of Songkla University

of

for editing the English of the manuscript.

[1]

ro

References

A. Moore, H. Goerne, P. Rajiah, Y. Tanabe, S. Saboo, S. Abbara, Acute Myocardial

S. Mendis, K. Thygesen, K. Kuulasmaa, S. Giampaoli, M. Mahonen, K.N. Blackett, L.

re

[2]

-p

Infarct, Radiol. Clin. North Am. 57 (2019) 45–55. doi:10.1016/j.rcl.2018.08.006.

Lisheng, World Health Organization definition of myocardial infarction: 2008-09

L. Babuin, A.S. Jaffe, Troponin: The biomarker of choice for the detection of cardiac

ur na

[3]

lP

revision, Int. J. Epidemiol. 40 (2011) 139–146. doi:10.1093/ije/dyq165.

injury, Cmaj. 173 (2005) 1191–1202. doi:10.1503/cmaj/051291. [4]

A.S. Jaffe, J. Ravkilde, H. Katus, R. Roberts, M. Galvani, U. Naslund, F.S. Apple, It’s Time for a Change to a Troponin Standard, Circulation. (2000) 1216–1220.

Jo

doi:10.1161/01.cir.102.11.1216.

[5]

A.A. Mohammed, J.L. Januzzi, Clinical applications of highly sensitive troponin assays, Cardiol. Rev. 18 (2010) 12–19. doi:10.1097/CRD.0b013e3181c42f96.

[6]

F.S. Apple, A.H.B. Wu, A.S. Jaffe, European Society of Cardiology and American College of Cardiology guidelines for redefinition of myocardial infarction: How to use 21

existing assays clinically and for clinical trials, Am. Heart J. 144 (2002) 981–986. doi:10.1067/MHJ.2002.124048. [7]

R. Twerenbold, A. Jaffe, T. Reichlin, M. Reiter, C. Mueller, High-sensitive troponin T measurements: What do we gain and what are the challenges?, Eur. Heart J. 33 (2012) 579–586. doi:10.1093/eurheartj/ehr492. T. Lenderink, E. Boersma, C. Heeschen, A. Vahanian, M.J. De Boer, V. Umans, M.J.B.M.

of

[8]

ro

Van den Brand, C.W. Hamm, M.L. Simoons, Elevated troponin T and C-reactive protein predict impaired outcome for 4 years in patients with refractory unstable angina, and

-p

troponin T predicts benefit of treatment with abciximab in combination with PTCA, Eur.

[9]

re

Heart J. 24 (2003) 77–85. doi:10.1016/S0195-668X(02)00322-6. S. Mayilo, M.A. Kloster, M. Wunderlich, A. Lutich, T.A. Klar, A. Nichtl, K. Kürzinger,

lP

F.D. Stefani, J. Feldmann, Long-range fluorescence quenching by gold nanoparticles in a sandwich immunoassay for cardiac troponin T, Nano Lett. 9 (2009) 4558–4563.

ur na

doi:10.1021/nl903178n.

[10] R.F. Dutra, L.T. Kubota, An SPR immunosensor for human cardiac troponin T using specific binding avidin to biotin at carboxymethyldextran-modified gold chip, Clin. Chim.

Jo

Acta. 376 (2007) 114–120. doi:10.1016/j.cca.2006.07.029. [11] Q. Sheng, X. Qiao, M. Zhou, J. Zheng, Recent progress in electrochemical sensing of cardiac troponin by using nanomaterial-induced signal amplification, Microchim. Acta. 184 (2017) 1573–1585. doi:10.1007/s00604-017-2219-y. [12] F.S. Apple, Y. Sandoval, A.S. Jaffe, J. Ordonez-Llanos, Cardiac troponin assays: Guide to understanding analytical characteristics and their impact on clinical care, Clin. Chem. 63 22

(2017) 73–81. doi:10.1373/clinchem.2016.255109. [13] M. Cieplak, W. Kutner, Artificial Biosensors: How Can Molecular Imprinting Mimic Biorecognition?, Trends Biotechnol. 34 (2016) 922–941. doi:10.1016/j.tibtech.2016.05.011. [14] L. Uzun, A.P.F. Turner, Molecularly-imprinted polymer sensors: Realising their potential,

of

Biosens. Bioelectron. 76 (2016) 131–144. doi:10.1016/j.bios.2015.07.013.

ro

[15] G. Ertürk, B. Mattiasson, Molecular imprinting techniques used for the preparation of biosensors, Sensors (Switzerland). 17 (2017) 1–17. doi:10.3390/s17020288.

-p

[16] R. Gui, H. Jin, H. Guo, Z. Wang, Recent advances and future prospects in molecularly

re

imprinted polymers-based electrochemical biosensors, Biosens. Bioelectron. 100 (2018) 56–70. doi:10.1016/j.bios.2017.08.058.

lP

[17] M.C. Blanco-López, S. Gutiérrez-Fernández, M.J. Lobo-Castañón, A.J. Miranda-Ordieres, P. Tuñón-Blanco, Electrochemical sensing with electrodes modified with molecularly

ur na

imprinted polymer films, Anal. Bioanal. Chem. 378 (2004) 1922–1928. doi:10.1007/s00216-003-2330-2.

[18] J.A. Ribeiro, C.M. Pereira, A.F. Silva, M.G.F. Sales, Electrochemical detection of cardiac

Jo

biomarker myoglobin using polyphenol as imprinted polymer receptor, Anal. Chim. Acta. 981 (2017) 41–52. doi:10.1016/j.aca.2017.05.017.

[19] J.G. Pacheco, M.S.V. Silva, M. Freitas, H.P.A. Nouws, C. Delerue-Matos, Molecularly imprinted electrochemical sensor for the point-of-care detection of a breast cancer biomarker (CA 15-3), Sensors Actuators, B Chem. 256 (2018) 905–912.

23

doi:10.1016/j.snb.2017.10.027. [20] B.V.M. Silva, B.A.G. Rodríguez, G.F. Sales, M.D.P.T. Sotomayor, R.F. Dutra, An ultrasensitive human cardiac troponin T graphene screen-printed electrode based on electropolymerized-molecularly imprinted conducting polymer, Biosens. Bioelectron. 77 (2016) 978–985. doi:10.1016/j.bios.2015.10.068.

of

[21] N. Karimian, M. Vagin, M.H.A. Zavar, M. Chamsaz, A.P.F. Turner, A. Tiwari, An

50 (2013) 492–498. doi:10.1016/j.bios.2013.07.013.

ro

ultrasensitive molecularly-imprinted human cardiac troponin sensor, Biosens. Bioelectron.

-p

[22] D. Cai, L. Ren, H. Zhao, C. Xu, L. Zhang, Y. Yu, H. Wang, Y. Lan, M.F. Roberts, J.H.

re

Chuang, M.J. Naughton, Z. Ren, T.C. Chiles, A molecular-imprint nanosensor for ultrasensitive detection of proteins, Nat. Nanotechnol. 5 (2010) 597–601.

lP

doi:10.1038/nnano.2010.114.

[23] H. Yang, L. Li, Y. Ding, D. Ye, Y. Wang, S. Cui, L. Liao, Molecularly imprinted

ur na

electrochemical sensor based on bioinspired Au microflowers for ultra-trace cholesterol assay, Biosens. Bioelectron. 92 (2017) 748–754. doi:10.1016/j.bios.2016.09.081. [24] E. Júlia, V. Horváth, A. Yarman, F.W. Scheller, R.E. Gyurcsányi, Electrosynthesized

Jo

molecularly imprinted polymers for protein recognition, TrAC - Trends Anal. Chem. 79 (2016) 179–190. doi:10.1016/j.trac.2015.12.018.

[25] Z. Wang, F. Li, J. Xia, L. Xia, F. Zhang, S. Bi, G. Shi, Y. Xia, J. Liu, Y. Li, L. Xia, Biosensors and Bioelectronics An ionic liquid-modi fi ed graphene based molecular imprinting electrochemical sensor for sensitive detection of bovine hemoglobin, Biosens. Bioelectron. 61 (2014) 391–396. doi:10.1016/j.bios.2014.05.043. 24

[26] B.A. Menaker, V. Syritski, J. Reut, O. Andres, E. Gyurcsa, V. Horva, Electrosynthesized Surface-Imprinted Conducting Polymer Microrods for Selective Protein Recognition, Adv. Mater. 21(2009) 2271–2275. doi:10.1002/adma.200803597. [27] T. Huynh, A. Pietrzyk-le, C.B. KC, K.R. Noworyta, J.W. Sobczak, P. Sindhu, F.D. Souza, W. Kutner, Electrochemically synthesized molecularly imprinted polymer of thiophene

of

derivatives for flow-injection analysis determination of adenosine-5′-triphosphate (ATP),

ro

Biosens. Bioelectron. 41 (2013) 634–641. doi:10.1016/j.bios.2012.09.038.

[28] S.E.E. Profile, Electropolymerization of aniline ( Review ), Sci. Int. 20 (2008) 97–106.

-p

[29] A.K. Roy, V.S. Nisha, C. Dhand, B.D. Malhotra, Molecularly imprinted polyaniline film

re

for ascorbic acid detection., J. Mol. Recognit. 24 (2011) 700–706. doi:10.1002/jmr.1104. [30] H.J. Chen, Z.H. Zhang, D. Xie, R. Cai, X. Chen, Y.N. Liu, S.Z. Yao, Surface-imprinting

lP

sensor based on carbon nanotubes/graphene composite for determination of bovine serum albumin, Electroanalysis. 24 (2012) 2109–2116. doi:10.1002/elan.201200375.

ur na

[31] J. Luo, J. Huang, Y. Wu, J. Sun, W. Wei, X. Liu, Synthesis of hydrophilic and conductive molecularly imprinted polyaniline particles for the sensitive and selective protein detection, Biosens. Bioelectron. 94 (2017) 39–46. doi:10.1016/j.bios.2017.02.035.

Jo

[32] Y. Umasankar, S.M. Chen, Separation and concentration effect of f-MWCNTs on electrocatalytic responses of ascorbic acid, dopamine and uric acid at f-MWCNTs incorporated with poly (neutral red) composite films, Electrochim. Acta. 52 (2007) 5985– 5996. doi:10.1016/j.electacta.2007.03.047. [33] S.M. Chen, S.A. Kumar, Electroanalysis of NADH Using Conducting and Redox Active

25

Polymer/Carbon Nanotubes Modified Electrodes-A Review, Sensors. 8 (2008) 739–766. doi.org/10.3390/s8020739 [34] M.C. Kum, K.A. Joshi, W. Chen, N. V. Myung, A. Mulchandani, Biomolecules-carbon nanotubes doped conducting polymer nanocomposites and their sensor application, Talanta. 74 (2007) 370–375. doi:10.1016/j.talanta.2007.08.047.

of

[35] M. Veerapandian, R. Hunter, S. Neethirajan, Dual immunosensor based on methylene

ro

blue-electroadsorbed graphene oxide for rapid detection of the influenza A virus antigen, Talanta. 155 (2016) 250–257. doi:10.1016/j.talanta.2016.04.047.

-p

[36] N. Ajami, A. Ehsani, F. Babaei, R. Safari, Electrochemical properties, optical modeling

re

and electrocatalytic activity of pulse-electropolymerized ternary nanocomposite of poly (methylene blue) in aqueous solution, J. Mol. Liq. 215 (2016) 24–30.

lP

doi:10.1016/j.molliq.2015.12.023.

[37] M.M. Barsan, M.E. Ghica, C.M.A. Brett, Electrochemical sensors and biosensors based

ur na

on redox polymer/carbon nanotube modified electrodes: A review, Anal. Chim. Acta. 881 (2015) 1–23. doi:10.1016/j.aca.2015.02.059. [38] J.Y. Wang, P.C. Nien, C.H. Chen, L.C. Chen, K.C. Ho, A glucose bio-battery prototype

Jo

based on a GDH/poly(methylene blue) bioanode and a graphite cathode with an iodide/triiodide redox couple, Bioresour. Technol. 116 (2012) 502–506. doi:10.1016/j.biortech.2012.03.083.

[39] Y. Umasankar, S.M. Chen, Multi-walled carbon nanotubes with poly(methylene blue) composite film for the enhancement and separation of electroanalytical responses of catecholamine and ascorbic acid, Sensors Actuators, B Chem. 130 (2008) 739–749. 26

doi:10.1016/j.snb.2007.10.040. [40] G. Cui, J. H. Yoo, J. S. Lee, J. Yoo, J. H. Uhm, G. S. Cha, H. Nam, Effect of pretreatment on the surface and electrochemical properties of screen-printed carbon paste electrodes, Analyst. 126 (2001) 1399–1403. doi:10.1039/b102934g. [41] M.I.C. Marinho, M.F. Cabral, L.H. Mazo, Is the poly (methylene blue)-modified glassy

ro

based molecules?, J. Electroanal. Chem. 685 (2012) 8–14.

of

carbon electrode an adequate electrode for the simple detection of thiols and amino acid-

doi:10.1016/j.jelechem.2012.08.023.

-p

[42] J. Zang, Y. Wang, X. Zhao, G. Xin, S. Sun, X. Qu, S. Ren, Electrochemical synthesis of

re

polyaniline on nanodiamond powder, Int. J. Electrochem. Sci. 7 (2012) 1677–1687. [43] A. Kellenberger, D. Ambros, N. Plesu, Scan rate dependent morphology of polyaniline

lP

films electrochemically deposited on nickel, Int. J. Electrochem. Sci. 9 (2014) 6821–6833. [44] N. Prabhakar, K. Arora, H. Singh, B.D. Malhotra, Polyaniline based nucleic acid sensor, J.

ur na

Phys. Chem. B. 112 (2008) 4808–4816. doi:10.1021/jp711853q. [45] M.P. Khesuoe, F.O. Okumu, M.C. Matoetoe, Development of a silver functionalised polyaniline electrochemical immunosensor for polychlorinated biphenyls, Anal. Methods.

Jo

8 (2016) 7087–7095. doi:10.1039/c6ay01733a. [46] J. Camalet, J. Lacroix, T.D. Nguyen, S. Aeiyach, M.C. Pham, Aniline electropolymerization on platinum and mild steel from neutral aqueous media, J. Electroanal. Chem. 485 (2000) 13–20. doi.org/10.1016/S0022-0728(00)00080-2 [47] Y. Song, Z. Guo, Z. Hu, J. Wang, S. Jiao, Electrochemical self-assembly of nano-

27

polyaniline film by forced convection and its capacitive performance, RSC Adv. 7 (2017) 3879–3887. doi:10.1039/c6ra24707e. [48] X. Shen, Y. Ma, Q. Zeng, J. Tao, J. Huang, L. Wang, Molecularly imprinted electrochemical sensor for advanced diagnosis of alpha-fetoprotein, Anal. Methods. 8 (2016) 7361–7368. doi:10.1039/c6ay01922f.

of

[49] D.D. Schlereth, A.A. Karyakin, Electropolymerization of phenothiazine, phenoxazine and

ro

phenazine derivatives: Characterization of the polymers by UV-visible difference

spectroelectrochemistry and Fourier transform IR spectroscopy, J. Electroanal. Chem. 395

-p

(1995) 221–232. doi:10.1016/0022-0728(95)04127-A.

re

[50] L. Tan, Q. Zie, S. Yao, Electrochemical and spectroelectrochemical studies on pyridoxine hydrochloride using a poly(methylene blue) modified electrode, Electroanalysis. 16

lP

(2004) 1592–1597. doi:10.1002/elan.200302993.

[51] Q. Qin, Y. Guo, Preparation and Characterization of Nano-Polyaniline Film on ITO

ur na

Conductive Glass by Electrochemical Polymerization, J. Nanomater. 2012 (2012) 1–6. doi:10.1155/2012/519674.

[52] B. Magnusson, U. Örnemark (Eds.), The Fitness for Purpose of Analytical Methods, in:

Jo

Eurachem Guid., 1998: pp. 1–61. doi:978-91-87461-59-0. [53] R.J. Ansell, Characterization of the binding properties of molecularly imprinted üpolymers, in: Adv. Biochem. Eng. Biotechnol., 2015: pp. 51–93. doi:10.1007/10_2015_316. [54] A. Saadaoui, C. Sanglar, R. Medimagh, A. Bonhomme, R. Baudot, S. Chatti, S. Marque,

28

D. Prim, M.S. Zina, H. Casabianca, New biosourced chiral molecularly imprinted polymer: Synthesis, characterization, and evaluation of the recognition capacity of methyltestosterone, J. Mol. Recognit. 30 (2017) 1–9. doi:10.1002/jmr.2594. [55] T.C. Cheng, Y.T. Huang, C.Y. Chang, K.S. Yao, C.C. Hwang, Molecular recognition of sulfaquinoxaline and sulfapyridine with molecularly imprinted polymer, J. Chil. Chem.

of

Soc. 54 (2009) 295–298. doi:10.4067/S0717-97072009000300019.

ro

[56] T. Keller, T. Zeller, D. Peetz, S. Tzikas, A. Roth, E. Czyz, C. Bickel, S. Baldus, A.

Warnholtz, M. Fröhlich, C.R. Sinning, M.S. Eleftheriadis, P.S. Wild, R.B. Schnabel, E.

-p

Lubos, N. Jachmann, S. Genth-Zotz, F. Post, V. Nicaud, L. Tiret, K.J. Lackner, T.F. Münzel, S. Blankenberg, Sensitive Troponin I Assay in Early Diagnosis of Acute

lP

doi:10.1056/nejmoa0903515.

re

Myocardial Infarction, N. Engl. J. Med. 361 (2009) 868–877.

[57] AOAC, Guidelines for Standard Method Performance Requirements (Appendix F),

ur na

AOAC Off. Methods Anal. (2012) 1–17.

[58] F.T.C. Moreira, R.A.F. Dutra, J.P.C. Noronha, A.L. Cunha, M.G.F. Sales, Artificial antibodies for troponin T by its imprinting on the surface of multiwalled carbon nanotubes: Its use as sensory surfaces, Biosens. Bioelectron. 28 (2011) 243–250.

Jo

doi:10.1016/j.bios.2011.07.026.

[59] P. Palladino, M. Minunni, S. Scarano, Cardiac Troponin T capture and detection in realtime via epitope-imprinted polymer and optical biosensing, Biosens. Bioelectron. 106 (2018) 93–98. doi:10.1016/j.bios.2018.01.068.

29

30

of

ro

-p

re

lP

ur na

Jo

Figure legends Fig. 1 Schematic illustration showing (A) the fabrication of the MIP sensor and (B) the binding

ur na

lP

re

-p

ro

of

detection.

Fig. 2 SEM images of the modified SPCE: (A) f-MWCNTs, (B) PMB/f-MWCNTs and (C) MIP coated PMB/f-MWCNTs. (D) CVs of the SPCE modification steps in 100 mmol L-1 PBS, pH

Jo

7.00 at a scan rate of 50 mV s-1: (a) SPCE, (b) f-MWCNTs/SPCE, (c) PMB/f-MWCNTs/SPCE, (d) PANI/PMB/f-MWCNTs/SPCE, (e) cTnT/GA/PANI/PMB/f-MWCNTs/SPCE, (f) after electropolymerization of the second PANI layer and (g) after removal.

31

of ro -p re lP ur na

Fig. 3 Influences on the sensitivity of the MIP sensor to cTnT of: (A) amount of f-MWCNTs, (B) the number of scans for the MB electropolymerization, (C) number of electropolymerization scans for the first PANI layer, and (D) the second PANI layer. These sensitivities were received

Jo

by performing DPVs at different cTnT concentrations in 100 mmol L-1 PBS (pH 7.00). (E) The current signals of PMB after various cTnT removal time.

32

33

of

ro

-p

re

lP

ur na

Jo

Fig. 4 Sensor performance: (A) differential pulse voltammograms (DPVs) of the developed MIP sensor obtained with 0.0-40 pg mL-1 of cTnT in 0.10 mol L-1 PBS (pH 7.00). (B) plot of the change in current response (∆I) versus concentration of cTnT of the MIP and NIP sensors fitted with the Langmuir isotherm using GraphPad Prism 8 and the inset showing the linear range (n=3). (C) effects of interfering substances on the response of the cTnT-MIP sensors after the

of

addition of cTnI, HSA, Glu, AA, Cr and UA in a solution containing 1.0 pg mL-1 cTnT and the different concentrations of cTnI, HSA, Glu, AA, Cr and UA were tested using the developed

Jo

ur na

lP

re

-p

ro

MIP sensor.

34

35

of

ro

-p

re

lP

ur na

Jo

Fig. 5 Comparison of the analytical results between the developed MIP sensor and the gold standard electrochemiluminescence immunoassay (ECLIA) (from the hospital) for cTnT

Jo

ur na

lP

re

-p

ro

of

determination in human plasma sample.

36

of

Table 1 A comparison of the performance of MIP sensors for cTnT detection

Modification types

Detection Method

Mediator

detection

Linearity

Potentiometric

NR

0.16 µg mL-1

Electrochemical

Fe(CN)6]3-/4-

cTnT imprinted polypyrrole-pyrroleElectrochemical 3-carboxylic acid/RGO/SPCE

chip

surface plasmon

Fe(CN)6]3-/4-

ur na

cTnT imprinted polydopamine/Au

-1

0.009 ng mL

lP

phenylenediamine)/AuE

1.41–20.8 µg mL-1

re

cTnT imprinted poly (o-

-p

MWCNTs-cTnT imprinted polyacrylamine/AuNWs

KD

Incubation

(mol L-1)

time (min)

Stability

References

NR

30

NR

[58]

2.4  10-12

10

NR

[21]

7.3  10-13

30

7 days*

[20]

ro

Limit of

-1

0.006 ng mL

-1

-1

0.009-0.80 ng mL

-1

0.01-0.10 ng mL

-1

NR

0.57 µg mL

0.66-10.7 µg mL

NR

2

NR

[59]

Methylene blue

0.04 pg mL-1

0.10-8.0 pg mL-1

2.8  10-13

30

3 weeks

this study

resonance

cTnT imprinted polyaniline

Electrochemical

/PMB/MWCNTs/SPCE

Jo

KD; equilibrium dissociation constant, NR; not reported *No further experiment after 7 days`

37

of

Author Biographies Kewarin Phonklam received her B.Sc. in Medical Technology (2016) from Walailak

ro

University, Thailand. She is a M.Sc. student in the Institute of Biomedical Engineering at

-p

Prince of Songkla University, Thailand. Her research covers the development of chemical sensors and biosensors for medical applications.

re

Rodtichoti Wannapob obtained his Ph.D. in Chemistry (2016) from Prince of Songkla University. He is currently working as a post-doctoral researcher for the development of

lP

DNA sensors in the Bioengineering Laboratory at RIKEN, Japan. His research interests focus on electrochemical sensors, biosensors, and conducting polymer.

ur na

Wilaiwan Sriwimol obtained her Ph.D. in Medical Technology from Mahidol University, Thailand. She is currently a lecturer at Department of Pathology, Faculty of Medicine, Prince of Songkla University, Thailand. Her research interests include toxicology and clinical chemistry.

Jo

Panote Thavarungkul is an Associate Professor at the Department of Physics, Prince of Songkla University, Thailand. She is a member of the Center of Excellence for Trace Analysis and Biosensor, Prince of Songkla University. Her research interests include biosensors and chemical sensors for medical, environmental, and industrial applications. 38

of

Tonghathai Phairatana is a lecturer and a researcher at the Institute of Biomedical Engineering, Faculty of Medicine, Prince of Songkla University, Thailand. She received her

ro

M.Sc. and Ph.D. in Biomedical Engineering at Imperial College London, United Kingdom.

-p

Her current research interests focus on the development of clinical diagnostic systems

Jo

ur na

lP

re

including physiological monitoring, electrochemical biosensors and microfluidics.

39