Reduced nonspecific protein adsorption by application of diethyldithiocarbamate in receptor layer of diphtheria toxoid electrochemical immunosensor

Journal Pre-proofs Reduced nonspecific protein adsorption by application of diethyldithiocarbamate in receptor layer of diphtheria toxoid electrochemical immunosensor Robert Zió łkowski, Adrianna Kaczmarek, Ilona Kośnik, Elżbieta Malinowska PII: DOI: Reference:

S1567-5394(19)30602-4 https://doi.org/10.1016/j.bioelechem.2019.107415 BIOJEC 107415

To appear in:

Bioelectrochemistry

Received Date: Revised Date: Accepted Date:

1 September 2019 12 November 2019 12 November 2019

Please cite this article as: R. Zió łkowski, A. Kaczmarek, I. Kośnik, E. Malinowska, Reduced nonspecific protein adsorption by application of diethyldithiocarbamate in receptor layer of diphtheria toxoid electrochemical immunosensor, Bioelectrochemistry (2019), doi: https://doi.org/10.1016/j.bioelechem.2019.107415

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© 2019 Published by Elsevier B.V.

Reduced nonspecific protein adsorption by application of diethyldithiocarbamate in receptor layer of diphtheria toxoid electrochemical immunosensor Robert Ziółkowski*1, Adrianna Kaczmarek1, Ilona Kośnik1, Elżbieta Malinowska1,2 1

Warsaw University of Technology, Faculty of Chemistry, The Chair of Medical Biotechnology, Noakowskiego 3, 00-664 Warsaw, Poland 2

CEZAMAT PW, Poleczki 19, 02-822 Warsaw, Poland

*Corresponding author. Tel.: +48 222347573, Fax: +48 226 282741. E-mail address: [email protected] (R. Ziółkowski)

1

Abstract

The immunoassay technology is of particular importance for both the environmental industry and clinical analysis. Biosensors, with the sensing layer based on antibodies or their fragments, offer high selectivity and short detection times. However, analytical devices where the electrochemical signal corresponds to changes in the interfacial region (sensing layer/electrode surface) are very susceptible to any nonspecific adsorption. Unfortunately, proteins (including antibodies) belong to the molecules showing high non-specific interactions with solid substrates. Herein, we propose diethyldithiocarbamate as a new antifouling and highly conductive agent. The investigations were conducted to evaluate its interaction with chosen proteins and the mechanism of its co-adsorption with biotinylated thiol (an anchor point for immune-sensing elements). The developed receptor layer is characterised by reduced nonspecific protein adsorption and high conductivity with the same preserved specificity of the antibodies (immobilised by the streptavidin/biotin bioaffinity technique). This allowed for selective detection of the diphtheria toxoid, an inactive toxin secreted by virulent strains of Corynebacterium diphtheria, at the level of 510-6 gml-1 (110-6 Lfml-1) and in the real-life sample.

Keywords: diethyldithiocarbamate, antifouling agent, electrochemical immunosensor, diphtheria toxoid

2

1. Introduction

The immunoassay technology is of particular importance for environmental, industrial, and clinical analysis. This concerns the determination of low concentration levels of emerging pollutants (e.g. pesticides, hormones, antibiotics) as well as drugs, proteins, hormones or bacterial toxins, generally called antigens (Ag) [1]. The most commonly used analytical immunoassay is the Enzyme-Linked Immunosorbent Assay (ELISA). In its simplest form, Ag is immobilised onto a solid support and a specific antibody (Ab) is added, which is directed against the Ag. If the Ab is linked to an enzyme, the addition of the solution containing the enzyme's substrate results in the generation of a detectable signal (e.g. a colour change) [2]. The broad application of the ELISA is attributed to its high selectivity provided by the antibodies. This technique, however, is used for specific analyses in central laboratories or facilities [2, 3]. To improve diagnostic efficiency, the so-called point-of-care testing (POCT) is being developed. The analysis can be performed directly at the patients’ bedside, at accident sites, or in operating outpatient clinics [4-6]. As the POCT devices are mobile and have compact dimensions, only miniature immunosensors can be used as detectors (which are usually electrochemical). In this case, the sensing is also achieved by the antigen-antibody recognition. However, the significantly shorter time of the analysis in contrast to the traditional ELISA method and similar detection limits is of undeniable advantage [6, 7-11]. This is achieved by registering the sensor signal (e.g. voltammetry or electrochemical impedance spectroscopy), which is proportional to the changes in the interfacial region (sensing layer/electrode surface) [9]. However, proteins show high non-specific adsorption at the solid substrates, which can significantly influence the interfacial properties. Similar negative effects can result from inappropriate antibodies immobilisation at the electrode surface [12]. To obtain the competitive working parameters of electrochemical immunosensor, it is indispensable that both of the following are present: (i) reduction of non-specific adsorption of proteins and (ii) preservation of antibodies’ specificity after their immobilisation. The first can be achieved by introduction into the receptor layer molecules with the appropriate terminal groups. Polyethylene glycol (PEG) has widely been considered as a gold standard of antifouling properties [13-15]. Nonetheless, its application in electrochemical sensors is limited because of the blocking of the electron transfer [16]. Another approach is to apply self-assembled monolayers (SAM) onto the electrode surface, which exhibit appropriate molecular-level characteristics: (i) are hydrophilic, (ii) include hydrogen-bond acceptors, (iii) do not include hydrogen-bond donors, and (iv) whose overall electrical charge is neutral. These requirements fulfil oligo (ethylene glycol), 3

tripropylenesulfoxide, maltose, or tertiary amine groups [17, 18]. However, in most cases, they also significantly reduce the electron transfer and increase the electrode resistivity. To further anchor the antibodies to such SAMs, it is necessary to use random crosslinking methods [12, 19-23]. This may lead to a significant loss in the antibody-antigen binding capability of the ready-to-use immunosensor’s receptor layer (forming linkages on active sites, toxic reagents, complicated chemistry, and a long-lasting procedure) [12, 22-25]. The exploitation of the bioaffinity immobilisation techniques existing in nature (streptavidin/biotin; protein A or G) can significantly improve the receptor layer quality [12, 19, 21, 22, 24, 26, 27]. Herein, we present an electrochemical immunosensor created with the use of a conductive, mixed self-assembled monolayer which significantly reduces unspecific protein adsorption. The antibody was immobilised by a biotin-streptavidin reaction. To secure the conductivity of such prepared receptor layer and limit the unspecific protein adsorption, for the first time diethyldithiocarbamate (DEDTC) was introduced into the immunosensing layer. Its antifouling capability was compared to other commonly-used electrode surface blocking agents. During the studies, two mechanisms of detection were investigated. It is shown that the developed immunosensor is capable of diphtheria toxoid quantitative detection in complex samples with high sensitivity, low detection limit, and increased antifouling ability.

4

2. Materials and methods

Equipment Cyclic voltammetry (CV), square-wave voltammetry (SWV), and electrochemical impedance spectroscopy were performed using a CHI 660A and CHI 1040A electrochemical workstation (CH Instruments, USA). The experiments were conducted with a three-electrode system consisting of a gold disk electrode (GDE) (CH Instruments, USA), an Ag/AgCl/1.0 mol·L-1 KCl reference electrode (Mineral, Poland) and a gold wire as an auxiliary electrode (Sigma Aldrich, Germany) at room temperature. EIS was carried out at a DC potential of 0.2V for potassium ferricyanide. With the AC amplitude of 5 mV for the frequency in the range from 1 Hz to 50000 Hz. Cyclic voltammetry was executed at 0.1 V/s scan rate, whereas square-wave voltammetry was performed at a pulse amplitude of 25 mV, an increment of 4 mV and frequency of 15 Hz. The single-potential amperometry experiments were performed at -0.1V versus Ag/AgCl electrode. The quartz-crystal microbalance experiments were performed with a CHI 440C electrochemical workstation equipped with the quartz crystal microbalance (QCM) module (CH Instruments, USA). As a transducer, the quartz crystal for QCM with a gold layer was used (CH Instruments, USA).

Reagents All chemicals and reagents were of analytical reagent grade and used as received, without further purification. Concentrated HCl (36.5–38.0%), NaOH, NaCl, NaH2PO4·H2O, Na2HPO4·7H2O, K2HPO4, cysteamine hydrochloride, 11-mercaptoundecanoic acid (MUA), 6mercapto-1-hexanol (MCH), 1-octanethiol, sodium diethyldithiocarbamate trihydrate (DETC), sodium citrate dihydrate, citric acid, 3,3′,5,5′-tetramethylbenzidine (TMB), H2O2, K4[Fe(CN)6]·3H2O, K3Fe(CN)6, H2SO4, HNO3 (65 %), ethanol (99.8 %), DMF, DMSO, streptavidin, bovine serum albumin (BSA), human serum albumin (HSA), lysozyme, immunoglobulin G (IgG), and immunoglobulin E (IgE), were purchased from Sigma-Aldrich, Germany. The HS-(CH2)11-NH-C(O)-Biotin (HS-L-Biotin) was purchased from ProChimia, Poland. The diphtheria toxoid (DT) was purchased from the National Institute for Biological Standards and Control (NIBSC), United Kingdom. The polyclonal antibodies against diphtheria toxoid, biotinylated (AbI) and labelled with horseradish peroxidase (AbII), were purchased from Abcam, United Kingdom. Human saliva used as a real sample was collected from a male

5

volunteer. The solutions were prepared with Milli-Q water. Milli-Q water and all aqueous buffer solutions were sterilised using an autoclave.

Solutions The solutions used in the experiments were as follows: 20 mM phosphate buffer (PBS, 20 mM NaH2PO4/Na2HPO4, 15 mM M NaCl, pH 7.4), 50 mM citrate buffer (citric acid 21.2 mM/sodium citrate 28.8 mM), 10 mg·ml-1 TMB in DMSO (as a stock solution), 0.1 mg·ml-1 TMB in citrate buffer, 2.5 mM K3Fe(CN)6/2.5 mM K4Fe(CN)6 in PBS, 0.5 M NaOH, 1 M and 0.1 M H2SO4, acidic piranha solution (H2SO4: H2O2 (3:1) (v/v)), basic piranha solution (H2O2:NH4OH:H2O (1:1:5) (v/v)). The pH was adjusted with 1 M NaOH or HCl water solution.

Gold disc electrode and QCM transducer cleaning Prior the surface modification, gold disk electrodes were incubated for 10 minutes in piranha solution, then washed with distilled water and cycled in 0.5 M NaOH between -0.5 and -1.4 V (scan rate 0.05 V·s-1, 10 cycles). The electrodes were then polished on microcloth pads (Buehler, USA) using Al2O3 slurries of the grain sizes 0.3 and 0.05 μm. After washing with distilled water, the electrodes were sonicated at room temperature in H2O: ethanol (1:1, v/v) for 10 minutes. Then, the electrodes were electrochemically cleaned using cyclic voltammetry in 1 M H2SO4 (scan rate 0.3 V·s-1, potential range from -0.3 to 1.7 V, 10 cycles) followed by scanning in 0.5 M H2SO4 containing 0.01 M KCl (scan rate 0.3 V·s-1, potential range from -0.3 to 1.7 V, 10 cycles). Finally, the purity of the gold surface was accounted for, referring to the gold oxide reduction peak recorded by cyclic voltammetry in 0.1 M H2SO4 (scan rate 0.3 V·s-1, potential range from -0.3 to 1.7 V, 2 cycles). The electrodes were then incubated in 5 mM K3Fe(CN)6/5 mM K4Fe(CN)6 and cyclic voltammetry, square-wave voltammetry, and impedance spectroscopy measurement were performed for a clean gold surface. Immediately after the measurement, the electrodes were subjected to modification. In the case of QCM, the gold transducers were cleaned with a base piranha solution treatment at 70 ◦C for 15 min. Then, the sensor was washed with enough water and absolute ethanol, and then dried under an argon atmosphere.

Immunosensor receptor layer preparation procedure First, the different thiol concentrations and times of immobilisation were investigated. This concerned the subsequent but separate immobilisation of DEDTC and HS-L-Biotin in various concentrations and times. Nonetheless, after this preliminary investigation for the best 6

electrochemical signal repeatability, the thiols were immobilised for 24 hours from their mixture in the milimolar proportion (1:1; 1:2 or 1:5) of DEDTC: HS-L-Biotin prepared in DMF. Thereafter, the electrodes were rinsed several times with DMF, ethanol, water, and finally with PBS. The protein components of the receptor layer were then immobilised. All were prepared in a PBS buffer. Firstly, the streptavidin, then the AbI. Such a prepared immunosensor was subjected to analysis.

7

3. Results and discussion

The signal registered in electrochemical immunosensors, which depend on the physicochemical changes in the interfacial region (sensing layer/electrode surface), can be greatly disturbed by the nonspecific adsorption of the sample components. From this point of view, antifouling and highly conductive monolayers are of the utmost importance [9, 13, 14, 16, 28].

Diethyldithiocarbamate as a conductive monolayer of the reduced protein fouling properties

In the present study, the protein antifouling characteristic was evaluated for the set of most popular electrode-blocking agents or molecules used for on-chip biological receptor layers’ formation in electrochemical biosensors (Table 1) [29-31 ].

Table 1. Characteristics of the molecules frequently used as electrode blocking agents (1, 4) or for onchip biological receptor layers formation (1-3, 6) and dithiocarbamate (5). The unmodified QCM transducer with bare gold (7) was added for RC% comparison. RC% - the relative change of QCM transducer frequency (vs. HS-L-Biotin) after interaction with 10 gml-1 streptavidin in PBS; “-“ not applicable.

No.

Structure

Name

HS

1

cysteamine

NH2

6-mercapto-1-hexanol

HS

2

OH

(MCH)

O

3

11-mercaptoundecanoic acid

HS CH3

HS (CH 2)3

HS

1-octanethiol

CH3

diethyldithiocarbamate

N

5 HS

(DEDTC)

CH3 H N

S

6

HS

NH (CH 2)9 O

7

(MUA)

OH

(CH2)7

4

Surface

N H

QCM transducer with bare gold

HS-(CH2)11-NH-C(O)-Biotin O

(HS-L-Biotin) bare gold

8

hydrophilic

hydrophilic

hydrophilic

Hydrogen bond donor / acceptor donor / acceptor donor / acceptor

pKa (bulk

RC%

Ref.

10.8

7.49%

[29, 32]

~17

48.8%

[29, 33]

4.7

15.74% [29, 31]

solution)

hydrophobic

-

-

5.24%

[29]

hydrophilic

acceptor

-

0.49%

[30]

100%

[29]

Stable complex used for binding probes

hydrophilic

acceptor

-

45.79% [29, 30]

In addition, diethyldithiocarbamate (DEDTC) was proposed as it potentially meets the requirements presented in [17] and mentioned in the introduction. The initial investigations of nonspecific protein adsorption were performed with quartz crystal microbalance (QCM) for streptavidin (MW 60 kDa, pI 5.5), the relatively acidic and slightly negatively charged protein under the conditions of the experiment (PBS, pH=7.4). This molecule was further used for the formation of the immunosensor’s sensing layer in the bioaffinity process with the biotinylated antibody (the biotin/streptavidin interaction is almost insensitive to pH and temperature changes, proteolysis, and denaturing agents) [22]. The HS-(CH2)11-NH-C(O)-Biotin (HS-LBiotin) was used as the intermediate binding molecule between the gold surface and the streptavidin (Table 1, Scheme 1).

(1)

(2)

i) antifouling agent CH3 CH3 N

CH3 CH3

CH3 CH3

CH3 CH3

CH3 CH3

N

N

N

N

SH SH

ii) HS-L-Biotin

Au

HS

S

S

S

Au

S

S

streptavidin

S

S

S

S

S

Au

Scheme 1. Immunosensor intermediate layer preparation steps: (1) co-adsorption of the antifouling agent and HS-L-Biotin, molecule dedicated to streptavidin attachment; (2) streptavidin immobilisation. The gold QCM transducer (modified or unmodified) frequency change during the interaction with 10 gml-1 streptavidin in PBS is presented in Fig. 1. The monolayer deposition was performed outside of the QCM flow cell by the incubation of the clean transducer in 2 mM solution (ethanolic or DMF as in the case of HS-L-Biotin or DEDTC) of the given compound for 24 hours. Then the transducer was washed with DMF or ethanol, respectively, and water and placed inside the thermostated QCM flow cell.

9

Figure 1.

QCM experiment of streptavidin adsorption at the modified or unmodified gold QCM transducer. Through the QCM cell the PBS or 10 gml-1 streptavidin in PBS was flowing (0.2 ml/min.). The sharp peaks indicate the peristaltic pump stops and the changes in the medium from PBS into streptavidin/PBS and back to PBS. (1) bare gold electrode; (2) HS-L-Biotin; (3) MCH; (4) MUA; (5) cysteamine; (6) 1octanethiol; (7) DEDTC.

The QCM transducers, unmodified or modified with a particular compound, were subjected to streptavidin for at least 15 minutes. First, the PBS solution flowed through the QCM cell, then the peristaltic pump was stopped (the sharp peak in the graphs), and the solution was changed to the 10 gml-1 streptavidin in PBS. After the given time, the flow was stopped again (sharp peak) and the QCM transducer was again subjected to PBS only. The change in the QCM transducer frequency registered in PBS (F), before and after the streptavidin, is directly related to the mass deposited at the surface (in this case the amount of streptavidin). To compare the antifouling properties of the analysed compounds, it was assumed that the maximal surface coverage,

100%,

was

obtained

for

HS-L-Biotin

(Fig.

1,

(2))

where

the

FHS-L-Biotin was 53,04 Hz (strong and non-covalent interaction of biotin and streptavidin). For bare gold and other compounds, the deposition rate of streptavidin was defined according to the equation (1): 10

𝑅𝐶% = ∆𝐹

∆𝐹𝑥

𝐻𝑆−𝐿−𝐵𝑖𝑜𝑡𝑖𝑛

∙ 100%

(1)

where Fx – frequency change for respective transducer; FHS-L-Biotin - frequency change for transducer modified with HS-L-Biotin (53,04 Hz).

The calculated RC% values are shown in Table 1. For the unmodified transducer (QCM transducer with bare gold) or modified with 6-mercapto-1-hexanol (MCH, the classical electrode blocking agent of strong hydrogen bond donor/acceptor properties), very similar RC% results (45.79% and 48.86%, respectively) were obtained. As there are no target molecules for streptavidin with which it could form non-covalent bonds, it can be assumed that the registered high efficiency of protein adsorption is a process of an uncontrollable nature. This concerns the protein orientation and stability, which often results in denaturation and/or detachment of protein previously adsorbed [12]. A significantly smaller signal change (15.74%) was observed for 11-mercaptoundecanoic acid (MUA), which can be assumed as a hydrophilic compound possessing negative charge in the conditions of the experiment (pKa 4.7). For cysteamine (pKa 10.8, positive charge during the experiment, weaker than MCH hydrogen bond donor/acceptor properties) and 1-octanethiol (hydrophobic surface), comparable results were obtained; 7.49% and 5.24% respectively. The highest antifouling properties were registered for diethyldithiocarbamate (DEDTC), RC% = 0.49%. This may result from the tertiary amine groups placed at the surface of the transducer with hydrogen bond acceptor properties, creating a hydrophilic surface with an overall neutral electrical charge in the given conditions. Aside from reducing the unspecific adsorption of the proteins, the blocking agents in the electrochemical immunosensor should also maintain the electrode conductivity. The electrochemical responses (SWV and EIS) in 2.5 mM K3Fe(CN)6/K4Fe(CN)6 of the electrodes modified with the analysed compounds are presented in Fig. 2.

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A)

B)

Figure 2.

The electrochemical response of the bare gold disc electrode and electrode modified with respective thiol. A) Square-wave voltammetry (SWV), B) electrochemical impedance spectroscopy (EIS). The thiols were deposited from the 2 mM solution (ethanolic or DMF as in the case of HS-L-Biotin or DEDTC) for 24 hours. The experiments were conducted in the presence of 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH 7.4). (1) bare gold electrode; (2) HS-L-Biotin; (3) MCH; (4) MUA; (5) cysteamine; (6) 1-octanethiol; (7) DEDTC.

As shown in Fig. 2, the electrochemical response (SWV) of the electrode modified with HS-LBiotin, MCH, MUA or 1-octanethiol was significantly lower than that obtained for the clean gold disc electrode. Similar observations were made for EIS, where the electrode resistance significantly increased in comparison to the bare gold. Only for cysteamine and DEDTC the 12

opposite results were obtained. In the case of cysteamine, the increase in the electrode conductivity is related to the positive charge of this molecule at the electrode surface and the same attraction of the free redox couple. Moreover, such a charge accumulation may also result in nonspecific ionic interactions with complex proteins [31], whereas the significantly better electrode response (higher currents and lower electrode resistivity) in the case of DEDTC is related to delocalised electronic states (the DEDTC resonance structures are presented in the Fig. 5), and its proximity to the Fermi level of gold (a reduced charge injection barrier across the metal–molecule interface results in a drop in the contact resistance in about two orders of magnitude compared to thiols) [34]. The same can be concluded, that the DEDTC-modified electrode should not introduce any unspecific interactions with the redox marker used in the detection process of the ready-to-use electrochemical immunosensor nor with the proteins themselves. This, with the above results (Table 1, Fig. 1 and 2), are a clear indication of the possible advantages of DEDTC usage in the electrochemical immunosensor receptor layers as a blocking agent.

Receptor layer composition As was already mentioned (Scheme 1), the immunosensor intermediate layer was prepared by co-adsorption of HS-L-Biotin and DEDTC. However, this procedure was the result of initial experiments where the compounds were deposited from separate solutions subsequently. The different compound concentrations and different times of their immobilisation were analysed (results are not shown). Nonetheless, the obtained electrochemical signal was of low reproducibility. Significantly better repeatability was observed when the compounds were coadsorbed for 24 hours. During the studies, the three different milimolar ratios of DEDTC:HSL-Biotin were tested: 1:1, 1:2, and 1:5, respectively. The electrochemical responses of such prepared gold disc electrodes are shown in Fig. 3.

13

A)

B)

Figure 3.

The electrochemical response of the gold disc electrode modified with a mix of thiols in given millimolar ratios. The thiols were deposited from the DMF solution in 24 hours. The experiments were conducted in the presence of 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH 7.4). (1) bare electrode; (2) DEDTC; (3) HS-L-Biotin; (4) 1:1 ratio; (5) 1:2 ratio; (6) 1:5 ratio.

It was observed that together with the HS-L-Biotin concentration increase in the deposited solution, the electrode resistivity was also increasing (Fig. 3 B), from about 500  for bare gold electrode to 1 k for 1:2 and 70 k for 1:5 thiol solutions (for pure HS-L-Biotin it was calculated at the level of 200 k, Fig. 2 and 3). However, in the case of 1:1 ratio, the electrode

14

resistivity decreased from 500 to 200  (10  for GDE modified only with DEDTC, Fig. 2 and 3). Respective changes after electrode modification can also be observed for the SWV experiments (Fig. 3 A). As the developed immunosensor will use electrochemical detection techniques, only 1:1 and 1:2 mixtures were further investigated. Both abovementioned mixed monolayers were subjected to QCM investigations (according to the steps shown in Scheme 2). It concerned the evaluation of the efficiency of the receptor layer formation (steps 1-2), the preservation of its activity (steps 3 and 4), and selectivity (IgG, BSA, HSA).

Scheme 2. The schematically presented receptor layer formation, its interaction with proteins, and the attachment of the secondary antibody (Ab II) to diphtheria toxoid. The step number is given in circles, which corresponds with the QCM results presented in Fig. 6. The QCM results for both analysed mixed monolayers after each of the steps (the numbers in brackets) presented in Scheme 2 are shown in Fig. 4.

A)

(3) (1)

(4) (2)

15

B)

(3)

(1) (4) (2)

Figure 4.

QCM experiment for subsequent transducer modifications described in the text and Scheme 2. The investigations were performed for two different thiol layers prepared from A) 1:1 and B) 1:2 mixture of DEDTC:HS-L-Biotin. Through the QCM cell the PBS or a 10 gml-1 solution of a given protein in PBS was flowing (0.2 ml/min.). Only the diphtheria toxoid concentration was 0.1 lf (about 0.5 gml-1). The sharp peaks indicate the peristaltic pump stops required for medium change.

Before any QCM experiments, the transducer was cleaned and modified by immersion (for 24 hours outside of the QCM flow cell) in DMF solution of DEDTC and HS-L-Biotin in 1:1 or 1:2 milimolar ratio. Afterwards, it was rinsed with DMF and water and installed in the thermostated (21 oC) flow-cell where the PBS buffer was flowing. The conditioning was carried out until the constant baseline was obtained. Then, the solutions of streptavidin (1), primary antibody – AbI (2), IgG, BSA HSA, diphtheria toxoid (3) and secondary antibody – AbII (4) in PBS were respectively introduced into the QCM flow-cell (Fig. 4). The similar changes in resonance frequencies (F) were obtained for both tested mixed thiols layers. The highest F was registered for AbI (step 2), almost twice as high as for streptavidin (step 1). This correlates with the molecular weight of both proteins (52.8 kDa streptavidin (tetramer) and 144 kDa (the average weight of the goat polyclonal anti-diphtheria toxin antibody). The initial selectivity studies with IgG, BSA and HAS (step 3) proved that such prepared receptor layer corresponds almost exclusively to the diphtheria toxoid (diphtheria toxin 70 kDa). According to the assumption of quantitative proteins interaction based on their masses and registered F, it can

16

be concluded that a single toxoid molecule does not fall on every molecule of AbI. Further reduction in this proportion was also observed for the last step (4), where the interaction with the secondary antibody, AbII (goat polyclonal anti-diphtheria toxin antibody, 144 kDa, conjugated with horseradish peroxidase, 44 kDa), was investigated. However, the results presented in Fig. 4 graphs show the receptor layer formation, the preservation of the antibodies specificity after its immobilisation, and the selectivity of the sensing layer towards diphtheria toxoid with a very slight response only in comparison to IgG when the 1mM:1mM proportion of thiols were used. To evaluate whether the observed slight unspecific response for IgG results from the interactions with DEDTC, electrochemical measurements were taken (Fig. 5). After GDE modification with the DEDTC (black line), it was subsequently incubated for 30 minutes in the PBS solutions containing one of the following proteins (1gml-1): streptavidin, AbI, AbII, diphtheria toxoid (0.2 Lfml-1), IgG, BSA, HSA, IgE, or lysozyme. Between the incubations, the electrode was washed with PBS and the experiments in 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couples (in 20 mM PBS, pH 7.4) was performed. As the electrode response was very similar for GDE modified with DEDTC and all the abovementioned proteins, in Fig. 5 only the graphs registered for DEDTC (green), streptavidin (blue), AbI (red), diphtheria toxoid (light blue) and IgG (brown) are shown and compared to the bare gold electrode (black).

A)

17

B)

Figure 5.

The electrochemical response of the bare GDE (black) and for the same gold electrode modified with DEDTC and after its consecutive 30 minutes incubation in 1 gml-1 solutions of streptavidin, AbI, diphtheria toxoid, and IgG. The DEDTC was deposited from the DMF solution for 24 hours. The CV and EIS experiments (graph A and B, respectively) were conducted in the presence of 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH=7.4). The DEDTC resonance structures are provided on the base of [34].

According to the results presented in Fig. 5, it can be concluded that the electrode modified with DEDTC exhibits more reversible electrochemical reactions that the clean gold electrode. The peak separation potential for cyclic voltammetry was reduced from 107 mV for the bare electrode to 73 mV for the electrode modified with DEDTC. The electrode resistivity was also decreased from 300 to approximately 10 . What is more, after the consecutive electrode incubation (each for 30 minutes) in the PBS solutions with different proteins (each in 1 gml1

concentration), no significant changes in the above given values were observed. This may

confirm that a slight frequency change for IgG observed in QCM graph (Fig. 4), in the case of 1mM DEDTC: 1mM HS-L-Biotin thiols layer, did not result from the protein-unspecific adsorption on the DEDTC-covered gold. The receptor layer formation with the use of both mixed SAM layers was also investigated with electrochemical impedance spectroscopy (Fig. 6).

18

A)

B)

Figure 6.

The EIS characteristics for receptor layer prepared on A) 1mM DEDTC: 1mM HSL-Biotin and B) 1mM DEDTC:2mM HS-L-Biotin. The experiments were conducted in the presence of 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH 7.4); bare gold electrode (dark yellow); mix of thiols (green); streptavidin (blue); AbI (red); DT (black); AbII (brown).

Depending on the thiols ratio (DEDTC:HS-L-Biotin) at the electrode surface, significantly different impedance values were obtained after each step of the receptor layer preparation. For the 1mM: 1mM ratio, the electrode resistivity was about five times smaller than 1mM:2mM. This probably resulted from the significantly smaller density of the HS-L-Biotin deposited at 19

the gold surface. This, in turn, influenced the amount of each of consecutive receptor layer component (streptavidin, AbI, DT and AbII). As the observed (Fig. 6 A) impedance changes for the monolayer deposited from 1mM: 1mM ratio were rather scarce (200 ), it can be concluded that protein immobilisation did not occur or that only single molecules were adsorbed. On the contrary, for 1mM: 2mM ratio the classical impedance increased after each of the sensing layer preparation steps were registered (Fig. 6 B). Such a change (5.5 k) is related rather to the formation of steric or electrostatic hindrances for redox markers by immobilised proteins, which indicates its successful immobilisation.

Analysis of the quantitative electrochemical immunosensor response and the selectivity studies

After the confirmation of the sensing layer formation, the electrochemical immunosensor response compared to different concentrations of diphtheria toxoid (DT), and its selectivity compared to IgG, BSA, HSA, IgE and lysozyme, were investigated. In the present study, two classical mechanisms of detection were investigated (Scheme 3).

Scheme 3. Two investigated mechanisms of detection are schematically presented. The first mechanism of detection is based on the changes in the redox mediator, K3Fe(CN)6/K4Fe(CN)6, and the interaction with the receptor layer and gold surface after toxoid recognition. The second is based on the amperometric reduction current of the enzymatically oxidised 3,3',5,5'-tetramethylbenzidine (TMB) with the use of horseradish peroxidase conjugated to the secondary anti-diphtheria toxin antibody (AbII).

20

The first (the results presented in Fig. 7 A and B) is based on the change in the SWV response for K3Fe(CN)6/K4Fe(CN)6 redox couple after binding the analyte. The sensor signal was related to the SWV current before (I0) and after toxoid detection (I) and calculated according to the equation: (𝑰𝟎 −𝑰) 𝑰𝟎

.

(1)

A)

B)

Figure 7.

A) Square wave voltammetry response of the ready to use immunosensor I0 primary anti-diphtheria toxoid antibody (AbI) and I - after its interaction with diphtheria toxoid (DT); B) the relationship between the immunosensor response, calculated with equation 1, and the diphtheria toxoid concentration; B inset) the selectivity studies of the ready to use immunosensor. The experiments were 21

conducted in the presence of 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH 7.4). The analytical information in the case of the second detection mechanism (Fig. 8) is the amperometric reduction current of the enzymatically oxidised 3,3',5,5'-tetramethylbenzidine (TMB) [35]. This should occur only when the secondary anti-diphtheria toxin antibody conjugated with horseradish peroxidase (AbII) is present in the sensing layer (Scheme 3).

A)

B)

Figure 8.

A) The electrochemical response, amperometry, of the immunosensor; A inset) 3,3A,5,5A-tetramethylbenzidine electrochemical response with the indicated potential used in further amperometric experiments; B) the relationship between the immunosensor response diphtheria toxoid concentration; B inset) the selectivity 22

studies of the ready to use immunosensor. The experiments were conducted in the citrate buffer (pH 5.0) with 0,1 mgml-1 TMB and 1 mM H2O2. Remarkably, for both mechanisms of detection, considerable results concerning biosensor responses compared to different diphtheria toxoid concentrations were obtained only for the mixed monolayer coadsorbed from 1mM DEDTC:2mM HS-L-Biotin solution (Fig. 6 B). Because of the above, the selectivity studies were performed only for such a thiol ratio (insets in Fig.7B and Fig.8B). Such substantially different quantitative results obtained for two investigated ratios of thiols in solutions from which its deposition took place was very surprising, especially considering the QCM results (Fig. 4) of sensing layer formation and DT detection. In that case, no significant differences in transducer response (F) for both analysed thiol ratios were observed. Slightly higher overall F values for 1mM DEDTC:1mM HS-LBiotin should not necessarily be ascribed to the same area of the transducer’s gold surface subjected to the passing solutions in the flow-cell. This is in strict opposition to similar experiments performed with electrochemical techniques of detection (Fig. 6). As already mentioned, significantly different electrode resistivity characteristics for 1:1 and 1:2 milomolar ratios were observed. The inconsistency in the results obtained with QCM and electrochemical techniques probably resulted from the applied gold layer cleaning procedures and the differences in the speed and energies of thiols and dialkyldithiocarbamate formation on gold. As demonstrated, this is of particular importance in the case of co-adsorption of the mixture of both, HS-L-Biotin (thiol, modified with biotin and used as an anchor point for receptor layer preparation), and DEDTC (dialkyldithiocarbamate, used as an protein antifouling agent) [36, 37]. As proved by Raigoza et al. [36], after equimolar co-adsorption of dialkyldithiocarbamate with other thiol, almost only dialkyldithiocarbamate is present at the gold surface (most probably because DEDC has a higher binding energy than thiol, 1.5 eV compared to 1.3 eV) [36]. Despite thiols adsorbing approximately four times faster than DEDTC, the exposure of a surface to equimolar thiol and dithiocarbamate in solution must result in the formation of disordered or low-density thiolate phases, which are then displaced by DEDTC. This explains that any reasonably electrochemical immunosensor response was obtained only for 1:2 of thiols ratio (DEDTC:HS-L-Biotin) (Fig. 7 and 8). However, this does not explain the discrepancies between QCM (Fig. 4), where the specific transducer response was registered for both thiol ratios, and electrochemical results (Fig. 6), where only 1:2 thiols ratio allows for specific biological interactions. In both cases, the procedures of receptor layer preparation was the same. The only differences were the different quality of gold and the sensor cleaning procedure. In 23

[38], the authors proved that the oxidised gold surface significantly enhances the stability of thiol–gold contacts. The oxidised top layer of gold, in the present study, could be expected in QCM transducers which were subjected only to the hot-base piranha solution during the whole cleaning process. Because of the above, it can be concluded that the increased thiols stability on such prepared surfaces together with the thermodynamic preferential thiol adsorption (four alkanethiol molecules occupy the same amount of area as three DEDTC) [36], may finally result in the presence of HS-L-Biotin at the gold surface after co-adsorption, even from the 1:1 milimolar ratio solution with DEDTC. In Fig. 7 and Fig. 8, the qualitative and quantitative results obtained for prepared immunosensing layers are presented. The diphtheria toxoid concentration was calculated according to the assumption that 1Lfml-1 corresponds to 5 μgml-1 [39]. Based on the obtained results, it can be concluded that the mechanism of detection which is based on the differential measurements (Fig. 7) allows to detect as low as 510-6 gml-1, which corresponds to 110-6 Lfml-1 (with the linear range of 510-5 to 510-3 gml-1). Moreover, the response of the prepared immunosensors towards 1 gml-1 diphtheria toxoid (0.896±0.026; n=4) was significantly higher than any other interfering compound in the 10-fold higher concentration (sensor response at the level of 0.23). What is more, the immunosensor response was at the same level for the pure experimental buffer (PBS) as for PBS containing any of the interfering proteins (IgG, BSA, HSA, IgE or lysozyme). Such response may result from the reorientation of the protein (primary antibody) during the sensing layer incubation and measurements in 2.5 mM K3Fe(CN)6/K4Fe(CN)6 redox couple (in 20 mM PBS, pH 7.4). The preferred orientation of a protein is strictly related to its free energy minimum, resulting from attractive coulomb and van-der-Waals interactions, hydrogen bonds, and the entropy gain of solvent molecules or counter-ion release [31, 40]. Proteins are typically large and complex molecules and exhibit different affinities toward sensor substrates in different regions of their surface. As these affinities may change along with the environment in which the measurements are taken [31, 40] (the measuring buffer change), it will influence the orientation of the proteins. This, in turn, will also result in observed electrode responses for immunosensor subjected only to a measuring buffer. However, the observed signal change is repeatable and did not influence the overall immunosensor response. In the case of the second mechanism of detection (Fig. 8), it was possible to detect as low as 510-4 gml-1 (110-4 Lfml-1) with the linear range to 510-2 gml-1. An almost two orders of magnitude higher detection limit probably resulted from the fact that the measurements were conducted in the 10 ml volume container where the enzymatically-oxidised TMB could freely

24

diffuse. Nonetheless, the selectivity studies revealed that there was no immunosensor response obtained for 10 gml-1 of the interfering proteins. The additional experiments with significantly higher concentrations resulted in signal changes (the concentrations of the interfering proteins are given in Fig. 8B inset).

Real-life sample analysis

Based on the obtained results, the possibility of use of the prepared immunosensor in a real-life sample was also investigated. The developed immunosensor is dedicated to the detection of diphtheria toxoid, an inactive form of the diphtheria toxin, secreted by a virulent form of Corynebacterium diphtheriae. The disease, diphtheria, develops rapidly and without the administration of the diphtheria antitoxin, usually leads to death. If the infection is suspected, doctors usually use a swab from the back of the patient’s throat and test it for the Corynebacterium diphtheriae and the diphtheria toxin. Accordingly, to evaluate the possibility of using such a prepared sensor in diphtheria toxin (toxoid) detection, saliva from a male volunteer was diluted 10-fold and spiked with the diphtheria toxoid of 0.01 Lfml-1 (0.05 gml-1) concentration. The lethal dose for humans is about 0.1 μg of diphtheria toxin per kg of body weight [41]. A signal change observed after incubation with the saliva sample reached 0.86±0.05 for the first mechanism of detection (this value lies with the signal response of 0.82±0.02 and 0.89±0.03 for toxoid concentration of 0.01 and 1 gml-1 respectively, presented in Fig. 7B). The immunosensor response for the second mechanism of detection was at the level of -0.051±0.004A (presented in Fig. 8B, signal response for 0.5 and 0.05 gml-1 was -0.088±0.008A and -0.035±0.004A, respectively). The above results may indicate that such an immunosensor construction, where the antifouling agent (proposed by us) is used, provides a strong motivation for further investigations of the development of the point-of-care device dedicated to real-life applications.

25

4. Conclusions

The present study is a dedicated to further studies on the electrochemical immunosensor for diphtheria toxoid development. As in our previous publication [39], a significant sensor response was also observed for interfering compounds; in the present study, we describe how to

reduce

this

disadvantage.

We

propose

a

new

antifouling

agent,

namely

diethyldithiocarbamate (DEDTC), which exhibits four molecular-level characteristics: (i) is hydrophilic; (ii) includes hydrogen-bond acceptors; (iii) does not include hydrogen-bond donors; and (iv) whose overall electrical charge is neutral. It was found that DEDTC characterises with almost no affinity to any of the analysed proteins (streptavidin, AbI, AbII, diphtheria toxoid IgG, BSA, HSA, IgE, lysozyme) and influences neither the proteins’ stability nor their activity. Moreover, because of its delocalised electronic states and with its proximity to the Fermi level of gold, the electrodes modified with DEDTC showed more reversible responses to redox markers than to the unmodified one. The neutral charge of this compound does not influence the electrostatic interactions with complex proteins, as can be observed for thiols possessing additional functional groups characterised with the appropriate pKa value [31, 40]. During the sensing layer preparation, the DEDTC was co-adsorbed at the electrode surface with HS-L-Biotin (an anchoring point for sensing elements of receptor layer). It was shown that to obtain a satisfactory ratio of HS-L-Biotin:DEDTC at the gold electrode surface (cleaned chemically and scanned electrochemically before the above compounds deposition), it is crucial to use a solution in which the thiol molar ratio is higher than dithiocarbamate. It was also shown that the gold substrate cleaned only in a hot-base piranha solution (the QCM experiments in the presented studies) significantly increases the stability of thiols, which allows for their co-adsorption from the solution of equimolar ratio with dithiocarbamate. After the electrochemical immunosensor preparation, two mechanisms of detection were used for diphtheria toxoid detection. The first, based on the differential analysis of SWV currents, allowed to obtain a lower detection limit at the level of 510-6 gml-1, which corresponds to 110-6 Lfml-1. On the other hand, the amperometric reduction of the enzymatically-oxidised 3,3A,5,5A-tetramethylbenzidine resulted in the almost two orders of magnitude higher detection limit of 510-4 gml-1 (110-4 Lfml-1). However, the obtained working parameters of the ready-to-use biosensor are similar to other, already developed electrochemical immunosensors e.g. towards human immunodeficiency virus p24 (LOD 1.510-

26

7

gml-1 with a linear range of 510-78.510-3 gml-1 [42]), hepatitis B virus surface antigen

(LOD 1.910-7 gml-1 with a linear range of 310-7110-3 gml-1 [43]), glial fibrillary acidic protein (LOD 110-6 gml-1 with a linear range of 110-6110-2 gml-1 [44]) and others [45]. In the case of both analysed mechanisms of detection, the biosensor response was selective, and any nonspecific sensor response was obtained only when the significantly higher interfering protein concentration was used. The high selectivity of the obtained immunosensor was also confirmed during the studies with the real sample; human saliva diluted 10-fold and spiked with diphtheria toxoid of 0.01 Lfml-1. In both detection mechanisms, the obtained signals lie within the expected regions on the respective calibration curves. This is of particular importance, considering the ultimate application of any of the biosensors which are planned to be used as detection elements in more sophisticated analytical POC devices.

Acknowledgements This work was financially supported by the National Centre for Research and Development, Poland (LIDER/35/0041/L-7/15/NCBR/2016) and by the Warsaw University of Technology.

Conflict of Interest The authors declare that there is no conflict of interest regarding the publication of this article.

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Highlights

 New highly conductive and protein antifouling agent was proposed  Reduction in nonspecific protein adsorption at the surface of electrochemical sensor  Preserving specificity of the immobilized antibodies  Low detection limits of diphtheria toxoid at the level of 510-6 gml-1

31

Conflict of Interest Authors declare that there is no conflict of interest regarding the publication of this article.

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