Miniaturized Electrochemical Immunosensors for the Detection of Growth Hormone

by

Qi Li

A thesis submitted in conformity with the requirements for the degree of Master of Science

Department of Chemistry University of Toronto

© Copyright by Qi Li 2012

ii

Miniaturized Electrochemical Immunosensors for the Detection of

Growth Hormone

Qi Li

Master of Science

Department of Chemistry

University of Toronto

2012

Abstract

The first part of this research involves the development of a gold-nanoparticle based sandwich

type immunosensor to identify trace amounts of human chorionic gonadotropin hormone based

on the direct electrochemical detection of Au nanoparticles. The second part of this research is to

design a biosensor that can be easily handled, has higher specificity, sensitivity, low-cost, and

rapid response and has a better detection of growth hormone (GH). Current bioanalytical

techniques have reported the difficulty to detect GH doping. This research aims to address the

issue of measuring GH in small volumes, which has been challenging the limits of analytical

detection systems. The electrochemical measurements utilize the redox activity of

ferro/ferricyanide in cyclic voltammetry and impedance spectroscopy. The detection limit 10

pg/mL was observed for GH in 20 µL sample volume, which indicated that this versatile

platform can be easily adapted for decentralized electrochemical immunosensing of clinically

important proteins.

iii

Acknowledgments

First and foremost, I would like to give my sincerest appreciation to my supervisor, Prof. Kagan

Kerman for his continued support and guidance throughout my research project and in helping

me develop scientific thinking and research knowledge. I have not only gained invaluable

research experience under his enthusiastic approach to electrochemistry, but had the opportunity

to excel as an individual with his mentorship in all aspects of life.

Working in the Kerman Laboratory would not have been the same without its past and present

members – Anthony, Tiffiny, Vinci, Xavier, Yuki, Nan, Vlad, Justin, Julian, Hosay and Sharon –

whose group discussions and encouraging advice have made my tenure in the lab productive and

inspirational. It was a pleasure to work alongside a remarkable group of students having

developed long lasting friendships with all of them.

Moreover, I would like to give my thanks to Professor Eiichi Tamiya – for his kindly support

with Screen Printed Carbon Strips and Mini-potentiostat system. I am also grateful to Professor

Paul Le Tissier and Professor David R. Grattan - for their kindly support with Rat Growth

hormone (GH) antibody and mouse GH antigen. I would also like to thank Professor Heinz-

Bernhard Kraatz for serving as a second reader in reviewing my MSc thesis.

Finally I would like to give thanks to my family and friends for their continued encouragement.

To my parents, thank you for the confidence for allowing me to pursue my research studies with

unwavering support.

iv

Table of Contents Acknowledgments ........................................................................................................................ iii

Table of Contents ......................................................................................................................... iv

List of Figures ............................................................................................................................... vi

List of Abbreviation ...................................................................................................................... x

Chapter 1 Electrochemical Immunosensors ............................................................................... 1

1.1 Introduction ......................................................................................................................... 1

1.2 Electrochemistry ................................................................................................................. 2

1.2.1 Principles ................................................................................................................. 3

1.2.1.1 The Electrical Double Layer ................................................................................... 4

1.2.1.2 Mass Transport ........................................................................................................ 5

1.2.1.3 Reversible, Irreversible, Quasi-reversible Reactions .............................................. 8

1.2.2 Electrode Materials ................................................................................................. 8

1.2.2.1 Solid Electrodes ...................................................................................................... 9

1.2.2.2 Carbon Electrodes ................................................................................................. 10

1.2.2.3 Carbon Paste Electrodes ....................................................................................... 10

1.2.3 Electrochemical Techniques ................................................................................. 11

1.2.3.1 Potential Sweep Methods ...................................................................................... 11

1.2.3.2 Cyclic Voltammetry .............................................................................................. 11

1.2.3.3 Differential Pulse Voltammetry ............................................................................ 15

1.2.3.4 Square-Wave Voltammetry .................................................................................. 17

1.2.3.5 Electrochemical Impedance Spectroscopy ........................................................... 18

1.2.4 Application of Electrochemistry: Nanomaterials and Electrochemical

Biosensors ......................................................................................................................... 20

1.3 Ideal Immunosensor Properties ......................................................................................... 22

1.3.1 General Working Principle of Immunosensors ..................................................... 22

1.3.2 Characterization of Immunosensors in Clinical Analysis ..................................... 23

1.4 Antigen, Antibody, and Their Recognition Reaction ....................................................... 23

1.5 Applications of Electrochemical Immunosensors ............................................................ 25

1.5.1 Competitive Immunoassay Systems ..................................................................... 26

1.5.2 Non-Competitive Immunoassay Systems ............................................................. 28

1.5.3 Antibody Immobilization ...................................................................................... 29

1.5.3.1 Biotin-Streptavidin Interaction ............................................................................. 29

1.5.3.2 Antibody Binding Proteins ................................................................................... 31

1.5.4 Nanomaterials Based Immunosensors .................................................................. 33

1.5.4.1 Gold Nanoparticle Based Immunosensors ............................................................ 37

1.6 Objectives ......................................................................................................................... 38

Chapter 2 Gold nanoparticle based electrochemical detection of Human Chorionic

Gonadotropin.......................................................................................................................... 40

2.1 Introduction ....................................................................................................................... 40

2.2 Experimental ..................................................................................................................... 43

2.2.1 Instrument and Materials ...................................................................................... 43

2.2.2 Methods ................................................................................................................. 44

2.2.2.1 Immobilization of Primary Antibody onto Working Electrode Surface ............... 44

2.2.2.2 Preparation of Au Nanoparticle-Labelled hCG Antibody (Au-Mab-hCG) .......... 44

v

2.2.2.3 Immobilization of hCG and Secondary Antibody and Detection of the

Antigen-Antibody Reaction .................................................................................. 45

2.3 Results and Discussion ..................................................................................................... 46

2.4 Conclusion ........................................................................................................................ 48

Chapter 3 Label Free Electrochemical Detection of Growth Hormone (GH) ...................... 49

3.1 Introduction ....................................................................................................................... 49

3.2 Experimental ..................................................................................................................... 51

3.2.1 Instrument and Materials ...................................................................................... 51

3.2.2 Methods ................................................................................................................. 52

3.2.2.1 Immobilization of Antibody on Working Electrode Surface ................................ 52

3.2.2.2 Direct Redox-Based Detection of Antigen-Antibody Reaction ............................ 52

3.3 Results and Discussion ..................................................................................................... 53

3.3.1 Construction of the Immunosensor and its Characterization ................................ 53

3.3.2 Electrochemical Analysis of Antigen-Antibody Binding ..................................... 55

3.4 Conclusion ........................................................................................................................ 59

Chapter 4 Conclusion and Future Directions .......................................................................... 60

4.1 Conclusion ........................................................................................................................ 60

4.2 Future Directions .............................................................................................................. 61

References .................................................................................................................................... 62

vi

List of Figures

Fig. 1.1: Schematic representation of the electrical double layer. IHP=inner Helmholtz plane;

OHP=outer Helmholtz plane (Figure drawn by adaptation from [4]). ........................................... 5

Fig. 1.2: Three modes of mass transport (Figure drawn by adaptation from [5]). .......................... 6

Fig. 1.3: Cyclic Voltammetry. (A) Potential waveform. The potential sweeps between two values

(V1 and V2). The scan rate is determined by the slope of the line. (B) The resultant

voltammogram of current is plotted against the applied potential of ferri/ferrocyanide solution

(Figure drawn by adaptation from [8]). ........................................................................................ 12

Fig. 1.4: (A) Schematic waveform of pulses superimposed on a staircase to form differential

pulse voltammetry. (B) The typical diffierential pulse voltammogram of current is plotted against

the applied potential. (Figure drawn by adaptation from [4]) ....................................................... 15

Fig. 1.5: (A) Schematic waveform for square-wave voltammetry. (B) The typical square-wave

voltammogram of current is plotted against the applied potential. (Figure drawn by adaptation

from [5]) ........................................................................................................................................ 17

Fig. 1.6: (a) Randles cell schematic diagram (b) sample Nyquist plot, assuming Rs=20 Ω and

Rct=250 Ω. (Figure drawn by adaptation from [11])..................................................................... 19

Fig. 1.7: The biosensor process including biological recognition element, physicochemical

transducer and signal processing steps. ........................................................................................ 22

Fig. 1.8: A schematic illustrates the “Y”-shaped structure of an antibody. The region between the

heavy chain and the light chain is where antigen binding occurs. This open arm portion of the “Y”

shape is generally denoted as Fab, while the non-antigenic binding site in the base portion is

referred to as Fc. (Figure drawn by adaptation from [35]) ........................................................... 24

Fig. 1.9: Schematic representation of (a) non-competitive and (b) competitive immunoassay

formats. (Figure drawn by adaptation from [6]) ........................................................................... 26

vii

Fig. 1.10: A schematic illustrates a competitive immunoassay format used for the analysis of

BSA. The rabbit anti-bovine serum albumin (IgG fraction) attached to microbeads and the

cymantrene labelled BSA attached to IgG-coated agarose beads. (Figure drawn by adaptation

from [39]) ...................................................................................................................................... 27

Fig. 1.11: Schematic representation of biotin-streptavidin interaction where biotinylated antibody

attached to the streptavidin coated solid phase. (Figure drawn by adaptation from [48]) ............ 29

Fig. 1.12: Schematic representation of the immunoassay based on a two-step labelling procedure

using ProtA-GEB biocomposite as a transducer. (A) RIgG immobilisation on the surface of the

electrode based on its interaction with protein A. (B) Competitive immunoassay, using anti-RIgG

and biotinylated anti-RIgG. (C) Enzyme labelling using HRP-streptavidin. (D) Electrochemical

enzyme activity determination. (Figure drawn by adaptation from [54]) ..................................... 32

Fig. 1.13: Multiprotein electrical detection protocol based on different inorganic colloid

nanocrystal tracers. (A) Introduction of antibody-modified magnetic beads; (B) binding of the

antigens to the antibodies on the magnetic beads; (C) capture of the nanocrystal-labeled

secondary antibodies; (D) dissolution of nanocrystals and electrochemical stripping detection.

(Figure drawn by adaptation from [69]) ....................................................................................... 35

Fig. 2.1: (a) Schematic representation of and subunits of human chorionic gonadotropin

hormone; (b) Anti- –FSH antibody (Polyclonal anti-human -subunit of follicle-stimulating

hormone) as primary antibody; (c) Anti- hCG (anti-chorionic gonadotropin -subunit (ab 1))

antibody as secondary antibody. ................................................................................................... 40

Fig. 2.2: Screen Printed Carbon Strip (SPCS) chips ..................................................................... 42

Fig. 2.3: Schematic illustration of the disposable immunosensor system. (a) The primary

antibody was immobilized directly on the SPCS chips, and a series of sandwich type

immunoreactions took place on the electrode surface. (b) A high potential, at 1.2 V, was applied

for pre-oxidation of Au nanoparticles and then the voltammetric measurements was taken. ...... 43

Fig. 2.4: Differential pulse voltammograms of the Au-Mab-hCG on SPCS at 20 mV/s in 0.1 M

HCl. The concentration of hCG ranged from 0 pg/mL to 1 ng/mL. ............................................. 46

viii

Fig. 2.5: Corresponding relation between the peak current intensity of Au nanoparticles with

hCG concentrations. Error bars indicate the relative standard deviation of the three measurements

(n = 3) performed with three different samples. ........................................................................... 47

Fig. 3.1: The cyclic voltammograms of 10 mM K4[Fe(CN)6]/K3[Fe(CN)6] solution in 50 mM

phosphate buffer using autolab system (A) and mini-potentiostat system (B) at a bare carbon

electrode (a) and at the rat GH antibody-modified carbon electrode before (b) and after the

addition (c) of mouse GH antigen (200 pg/mL). ......................................................................... 53

Fig. 3.2: Nyquist plot of rat GH antibody + mouse GH antigen immobilized on chip with 10 mM

potassium ferri/ferrocyanide (ratio 1/1) dissolved in 50 mM PBS without NaCl buffer solutions.

Blank signal was with nothing added on the chip. ........................................................................ 54

Fig. 3.3: (A) Corresponding relation between the GH concentrations (pg/mL) and the Rct values

(Kohm) from impedance spectra of rat GH antibody+mouse GH antigen immobilized on chip

with 10 mM potassium ferri/ferrocyanide (ratio 1/1). (B) Plot of the relationship between ratio of

RAb-RAb+GH/RAb and the negative logarithm value of mouse GH concentration from 10 pg/mL to

200 pg/mL to fit impedance data to Randles equivalent circuit. (n = 6, R² = 0.9953) ................. 55

Fig. 3.4: Corresponding relation between the GH concentrations (pg/mL) and the peak current I

(µA) (A) and bar graphs indicating the relation between the GH concentrations (pg/mL) and the

peak current I (µA) (B) using mini-potentiostat system from differential pulse voltammograms of

rat GH antibody+mouse GH antigen immobilized on chip at 10 mV/s with 10 mM potassium

ferri/ferrocyanide (ratio 1/1). ........................................................................................................ 56

Fig. 3.5: Corresponding relation between the GH concentrations (pg/mL) and the peak current I

(µA) using mini-potentiostat system (A) and bench-top potentiostat system (C); Bar graphs

indicating the relation between the GH concentrations (pg/mL) and the peak current I (µA) using

mini-potentiostat system (B) and bench-top potentiostat system (D) from cyclic voltammograms

of rat GH antibody+mouse GH antigen immobilized on chip at 50 mV/s (mini-potentiostat

system) and 100 mV/s (bench-top potentiostat system) with 10 mM potassium ferri/ferrocyanide

(ratio 1/1). ..................................................................................................................................... 57

ix

Fig. 3.6: Corresponding relation between the GH concentrations (pg/mL) and the peak potential

(V) using mini-potentiostat system (A) and bench-top potentiostat system (C); and corresponding

relation between the GH concentrations (pg/mL) and the change of peak potential ∆V(V) using

mini-potentiostat system (B) and bench-top potentiostat system (D) from Cyclic voltammograms

of rat GH antibody+mouse GH antigen immobilized on chip at 50 mV/s (mini-potentiostat

system) and 100 mV/s (bench-top potentiostat system) with 10 mM potassium ferri/ferrocyanide

(ratio 1/1). ..................................................................................................................................... 58

x

List of Abbreviation

AD Alzheimer’s disease

AFP Alpha Feta Protein

A-GEB Protein A- Graphite-Epoxy Biocomposite

AP Aminopyridine

ATP Adenosine triphosphate

AuNP Au nanoparticle

A Amyloid beta peptide

BBB Blood brain barrier

BSA Bovine serum albumin

CA 125 Carcinoma antigen 125

CEA Carcinoembryonic antigen

CNF Carbon nanofiber

CNT Carbon nanotube

CPE Carbon paste electrode

CSC Cathodic stripping voltammetry

CV Cyclic voltammetry

DDT Dichlorodiphenyltrichloroethane

DME Dropping mercury electrode

DMSO Dimethyl sulfoxide

DNA Deoxyribonucleic acid

DPP Differential pulse polarography

DPV Differential pulse voltammetry

EIS Electronic Impedance Spectroscopy

xi

ELISA Enzyme-linked immunosorbent assay

GCE Glassy carbon electrode

GEC Graphite-epoxy composites

GH Growth hormone

GHRF Growth hormone releasing factor

GO Glucose oxidase

GPES General purpose electrochemistry software

hCG Human chorionic gonadotropin

HIV Human immunodeficiency virus

HPLC High performance liquid chromatography

HRP Horseradish peroxidase

IEP Isoelectric point

Ig Immunoglobulin

IHP Inner Helmholtz plane

ITO Indium tin oxide

JS1 2-[(4-nitrophenyl)carbonyl]thieno[2,3-b]pyridine-3-amine

JS2 Thieno [2,3-b] pyridine, methanone derivative or methanone (3-

aminothieno[2,3-b] pyridine-2-yl) phenyl

JS3 3-aminothieno[2,3-b]pyridine-2-carbonitrile

JS4 3-aminothieno[2,3-b]pyridine-2-carboxamide

Mab-FSH Anti-FSH antibody

Mab-hCG Anti-hCG antibody

MFE Mercury film electrode

NA Nucleic acid

NHE Normal hydrogen electrode

xii

NPV Normal pulse voltammetry

OC Open circuit

ODN Oligodeoxynucleotide

OHP Outer Helmholtz plane

PAP Propargyl alcohol propoxylate

PBS Phosphate buffered saline

PCB Polychlorinated biphenyl

PCR Polymerase chain reaction

PEG Polyethylene glycol

PET Positron emission tomography

phGH Pituitary extracted human growth hormone

PQI p-quinone imine

PTK Protein tyrosine kinase

QD Quantum dot

RFU Relative fluorescence unit

rhGH Recombinant human growth hormone

RIA Radioimmunoassay

RIgG Rabbit IgG

SAE Solid amalgam electrodes

SAM Self assembled monolayer

SARA-CoV SARS-associated corona virus

SARS Severe acute respiratory syndrome

SC super coiled

SEM Scanning electron microscopy

xiii

SPCA Screen printed carbon array

SPCS Screen printed carbon strip

SPECT Single photon emission computed tomography

SWNT Single-wall carbon nanotube

SWV Square wave voltammetry

ThT Thioflavin T

Tyr Tyrosine

Vtg Vitellogenin

WADA World Anti-Doping Agency

1

Chapter 1 Electrochemical Immunosensors

1.1 Introduction

There is a continuing demand for fast and simple analytical methods for the determination of

many clinical, biochemical and environmental analytes. The requirement for immunologically

based biosensors is generally considered to be in the diagnostic field and particularly in the home

diagnostic field. Examples such as detection of contaminants in food or water are the most

obvious application. However, this general approach has limitations. Home diagnostic kits for

potentially epidemic infections frequently cannot legally be put on the open market. Thus kits

detecting bacteria causing venereal disease cannot be sold in countries where this remains a

‘notifiable’ disease and home testing kits for the human immunodeficiency virus (HIV) are still

banned in many countries. Projected applications in the clinical laboratory will have to prove

superiority either in detection limits or in cost effectiveness to gain acceptance and the same is

true for forensic or military use. With these caveats, it is possible to envisage a multitude of

biological applications such as personal detection alarms for allergy sufferers, probes for

detecting food additives or contaminants, and self monitoring systems for patients with chronic

autoimmune conditions to allow them to administer drugs only when appropriate in the same

manner as diabetic patients can monitor their current blood glucose level and take insulin only

when appropriate1. In this respect, immunoassays and immunosensors that rely on antibody-

antigen interactions provide a promising means of analysis due to their specificity and sensitivity.

High specificity of immunoassays and immunosensors is achieved mainly by the molecular

recognition of target analytes by antibodies or antigens to form stable immunocomplexes. On the

other hand, sensitivity depends on several factors, including the affinity of antibodies, the

amount of immobilized immunological recognition elements, and the choice of transducer and

signal probe. The affinity constant for antibody-antigen binding can span a wide range,

extending from below 105 L/mol to above 10

12 L/mol. Therefore, the improvement of

immunoassay and immunosensor performance mainly relies on the development of antibody

preparation techniques, the improvement of immobilization and tagging methods, and the

adoption of a high-performance transduction method2.

2

Moreover, electrochemical detection overcomes problems associated with other modes of

detection of immunoassays and immunosensors. For example, the short half-life of radioactive

agents, concerns of health hazards, and disposal problems are frequently raised in

radioimmunoassay, while limited sensitivity in the analysis of colored or turbid samples is

achieved in immunoassays coupled with optical detection. In contrast, electrochemical

immunoassays and immunosensors enable fast, simple, and economical detections that are free

of these problems. Furthermore, electrochemistry is an interfacial process in which the relevant

reactions take place at the electrode-solution interface, rather than in bulk solution. Therefore, in

conjunction with developments in micro- and nano-electrochemical sensors, electrochemistry

offers an added bonus of detecting analytes in very small volumes3.

1.2 Electrochemistry

Electrochemistry is the science concerned with the physical and chemical properties of ionic

conductors as well as with phenomena occurring at the interfaces between ionic conductors and

electronic conductors or semiconductors, or even insulators (including gases and vacuum). A

process of this kind can always be represented as a chemical reaction and is known generally as

an electrode process. Electroanalytical techniques are concerned with the relationship between

the measurements of electrical quantities, such as current, potential, or charge, and the chemical

parameters. Such electroanalytical measurements have been found to have a vast range of

applications, including biomedical analysis, industrial quality control, and environmental

monitoring. Instead of involving homogeneous bulk solutions, electrochemical processes take

place at the electrode-solution interface. Different types of electrical signal used for the

quantitation reflect the distinction between various electroanalytical techniques4.

The two principal types of electroanalytical measurements are potentiometric and potentiostatic.

And both types require at least two electrodes (conductors) and a contacting sample (electrolyte)

solution, which constitute the electrochemical cell. The electrode surface is thus a junction

between an ionic conductor and an electronic conductor. One of the two electrodes, which is

called working electrode, give response to the target analyte. The other electrode, called

reference electrode, has constant potential and is independent of the properties of the solution.

Potentiometric measurement is a static (zero current) technique, where the information of the

3

sample composition is obtained from measurement of the potential established across a

membrane. Various types of membrane materials have been developed to meet the needs from

different ion-recognition processes. The resulting potentiometric probes have been used to

monitor ionic species, such as calcium, fluoride, and potassium ions, within in complex

environment. Potentiostatic technique (controlled potential) is the study of charge-transfer

processes at the electrode-solution interface and deals with dynamic (no zero current) situations.

The electrode potential is applied to an electron-transfer reaction and the resultant current is

measured. Such potential can be considered as “electron pressure”, which drives the chemical

species to gain an electron (reduction) or to lose an electron (oxidation). The resulting current

shows the rate at which electrons move across the electrode-solution interface. Therefore,

potentiostatic technique can be used for any electroactive species. Nonelectroactive species may

also be detected through connection with indirect procedures. Potentiostatic technique has

several advantages, including high sensitivity and selectivity towards electroactive species, a

wide linear range, portable and low-cost instrumentation, speciation capability, and a wide range

of electrodes suitable for unusual environments4.

1.2.1 Principles

Electrochemistry is the study of interchange of chemical and electrical energy.

Oxidation/reduction involves the exchange of electrons from one chemical species to another.

The electrode can act as only a source (for reduction) or a sink (for oxidation) of electrons

transferred to or from species in solution, as in

O ne R [1-1]

where O and R are the oxidized and reduced species, respectively. Electrode reactions are

heterogeneous and take place in the interfacial region between electrode and solution, the region

where charge distribution differs from that of the bulk phases. Each has a standard electrode

potential, Eo, which is measured relative to the normal hydrogen electrode (NHE) with all

species at unit activity 1. For half reactions at equilibrium [1], the potential, E, can be used to

establish the concentration of the electroactive species at the surface [CO (0, t) and CR (0, t)]

through the Nernst equation5

4

° .

log ,

, [1-2]

where R is the universal gas constant (8.413 J K-1 mol-1), T is the Kelvin temperature, n is the

number of electrons transferred in the reaction, and F is the Faraday constant (96,487 Coulombs).

The resulting current from a change in oxidation state of the electroactive species is termed the

faradaic current because it obeys Faraday’s law, which is the reaction of 1 mole of substance

involves a change of n×96,487 Coulombs. The faradaic current is a direct measure of the rate of

the redox reaction. The resulting current-potential plot, known as the voltammogram, is a display

of current signal (vertical axis) versus the potential signal (horizontal axis). The exact shape and

magnitude of the voltammetric response is governed by the processes involved in the electrode

reaction. The total current is the summation of the faradaic currents for the sample and blank

solutions, as well as the nonfaradaic charging background current5.

1.2.1.1 The Electrical Double Layer

The electrical double layer is the array of charge particles and/or oriented dipoles that exists at

every material interface. In electrochemistry, such a layer reflects the ionic zones formed in the

solution to compensate for the excess of charge on the electrode. A positively charged electrode

thus attracts a layer of negative ions (vice versa).

5

Fig. 1.1: Schematic representation of the electrical double layer. IHP=inner Helmholtz plane;

OHP=outer Helmholtz plane (Figure drawn by adaptation from [4]).

The inner layer (closest to the electrode), known as the inner Helmholtz plane (IHP), contains

solvent molecules and specifically adsorbed ions, which are not fully solvated. The next layer,

the outer Helmholtz plane (OHP), reflects the imaginary plane passing through the center of

solvated ions at their closest approach to the surface (Fig. 1.1). The solvated ions are

nonspecifically adsorbed and are attracted to the surface by long-range coulomb forces. Both

Helmholtz layers represent the compact layer. However, the Helmholtz model does not take into

account the thermal motion of ions, which loosens them from the compact layer. The outer layer

(beyond the compact layer), referred to as the diffuse layer, is a three dimensional region of

scattered ions, which extends from the OHP into the bulk solution. Such an ionic distribution

reflects the counterbalance between ordering forces of the electrical field and the disorder caused

by a random thermal motion4.

1.2.1.2 Mass Transport

The rate of an electrode reaction is affected not only by the electrode itself but also by the

transport of species to and from bulk solution.

6

Mass transport can occur by three different modes (Fig. 1.2)5:

• Diffusion: the spontaneous movement under the influence of concentration gradient that

is from region of high concentration to region of lower concentration aimed at

minimizing concentration differences.

• Convection: transport to the electrode by a gross physical movement; such fluid flow

occurs with stirring or flow of the solution and with rotation or vibration of the electrode

(forced convection) or due to density gradients (natural convection).

• Migration: movement of charged particles along an electrical field.

Fig. 1.2: Three modes of mass transport (Figure drawn by adaptation from [5]).

The flux (J) is a common measure of the rate of mass transport at a fixed point. It is defined as

the number of molecules penetrating a unit area of an imaginary plane in a unit of time, and has

the units of mol cm-2 s-1. The flux to the electrode is described mathematically by a differential

equation, known as the Nernst-Planck equation4,

, !" #$,

#$! %&

#'$,

#$ (, ), [1-3]

Where D is the diffusion coefficient (cm-2s-1), ∂C(x, t)/∂x represents the concentration gradient at

distance x and time t, z is the charge, C is the concentration, ∂Φ(x, t)/∂x is the potential gradient

and V(x, t) is the hydrodynamic velocity. The current (i) is directly proportional to the flux4

* !+,- [1-4]

7

This equation can be greatly simplified by suppressing the electromigration or convection,

through the addition of excess inert salt or use of a quiescent solution, respectively. Under these

conditions, the movement of the electroactive species is limited by diffusion. The reaction

occurring at the surface of the electrode generates a concentration gradient adjacent to the

surface, which in turn gives rise to a diffusion flux. According to Fick’s first law, the rate of

diffusion (flux) is directly proportional to the slope of the concentration gradient4:

, !" #$,

#$ [1-5]

Combination of equations yields a general expression for the current response

* +,-" #$,

#$ [1-6]

The current (at any time) is proportional to the concentration gradient of the electroactive species.

The diffusion flux is time dependent. Such dependence is described by Fick’s second law (for

linear diffusion) 4:

#$,

# " #.$,

#$. [1-7]

After substitute Fick’s second law equation to, it leads to the well-known Cottrell equation:

* /&0

1&2/. [1-8]

That is, the current decreases in proportion to the square root of time, with (πDot) 1/2

corresponding to the diffusion layer thickness. The Cottrell equation can be further modified into

two components4:

* /&0

1&2/. /&

4 [1-9]

The current response of a spherical electrode following a potential step thus contains time

dependent and time-independent terms, reflecting the planar and spherical diffusion field

respectively4.

8

For a reaction involving the reduction of O to R, since the surface concentration of O is zero at

the new potential, a concentration gradient is established near the surface. The region within

which the solution is depleted of O is known as the diffusion layer, and its thickness is given by

δ. The concentration gradient is steep at first, and the diffusion layer is thin. As time goes by, the

diffusion layer expands and the concentration gradient decreases4.

1.2.1.3 Reversible, Irreversible, Quasi-reversible Reactions

If the oxidized and reduced species involved in an electrode reaction are in equilibrium at the

electrode surface, the Nernst equation can be applied. The electrode reaction is then known as a

reversible reaction since it obeys the condition of thermodynamic reversibility. The

concentrations of species at the interface depend on the mass transport of these species from bulk

solution, often described by the mass transfer coefficient kd. Within a reversible reaction, the

kinetics of the electrode reaction is much faster than the transport. The kinetics is expressed by a

standard rate constant, k0. For the reversible reaction, it is k0>>kd. By contrast, an irreversible

reaction is one where the electrode reaction cannot be reversed. It has to overcome a high kinetic

barrier, which is achieved by application of an extra potential (extra energy) called the

overpotential. For the irreversible reaction, it is k0<<kd. The third type reaction is called quasi-

reversible reactions, which exhibit intermediate behaviour between reversible and irreversible

reactions. In this case, the overpotential has a relatively small value and this extra potential

reaction can be reversed6.

1.2.2 Electrode Materials

The potentiostatic control contains three electrode system and a combination of operational

amplifiers and feedback loops. Here, the reference electrode is placed as close as possible to the

working electrode and is connected to the instrument through a high resistance circuit that draws

no current from it. There are two types of reference electrodes that are commonly used. The first

one is saturated calomel electrode, which is based on the reaction between elemental mercury

and mercury chloride. The aqueous phase in contact with the mercury and the mercury chloride

is a saturated solution of potassium chloride in water. The electrode is normally linked via a

porous frit to the solution in which the other electrode is immersed. And this porous frit is a salt

bridge. The second one is called standard hydrogen electrode, whose absolute electrode potential

9

is estimated to be 4.44 ± 0.02 V at 25 °C, but to form a basis for comparison with all other

electrode reactions, hydrogen's standard electrode potential (E0) is declared to be zero at all

temperatures. Because the flow cannot occur through the reference electrode, a current-carrying

auxiliary (counter) electrode is placed in the solution to complete the current path. Therefore, the

current flows through the solution between the working and the auxiliary electrodes. Symmetry

in the placement of these electrodes is important for the assumption that the current paths from

all points on the working electrode are equivalent5.

The open circuit (oc) potential is the potential of the working electrode relative to the reference

electrode when no potential or current is being applied to the cell. When a potential is applied

relative to oc, the system measures the open circuit potential before turning on the cell, then

applies the potential relative to that measurement7.

The choice of an electrode material depends on a great extent on the useful potential range of the

electrode in the particular solvent employed and the qualities and purity of the materials. The

usable potential range is limited by one or more of the following factors:

• Solvent decomposition.

• Decomposition of the supporting electrolyte.

• Electrode dissolution or formation of a layer of an insulating/semiconducting substance

on its surface.

Additionally, solid electrodes can be adversely affected by poisoning through contact with

solutions containing contaminants6.

1.2.2.1 Solid Electrodes

Accordingly, solid electrodes with extended anodic potential windows have attracted

considerable analytical interest. Of the many different solid materials that can be used as

working electrodes, the most often used are carbon, platinum, and gold. Silver, nickel, and

copper can also be used for specific applications. An important factor in using solid electrodes is

the dependence of the response on the surface state of the electrode. Therefore, the use of such

electrodes requires precise electrode pretreatment and polishing to obtain reproducible results.

10

The nature of these pretreatment steps depends on the materials involved. Mechanical polishing

and potential cycling are commonly used for metal electrodes, while various chemical,

electrochemical, or thermal surface procedures are added for activating carbon-based electrodes.

Unlike mercury electrodes, solid electrodes present a heterogeneous surface with respect to the

electrochemical activity4.

1.2.2.2 Carbon Electrodes

Solid electrodes based on carbon are currently in widespread use in electroanalysis, primarily

because of their broad potential window, low background current, rich surface chemistry, low

cost, chemical inertness, and suitability for various sensing and detection applications. In

contrast, electron transfer rates observed at carbon surfaces are often slower than those observed

at metal electrodes. While all carbon electrode materials share the basic structure of a six-

member aromatic ring and sp2 bonding, they differ in the relative density of the edge and basal

planes at their surfaces. The edge orientation is more reactive than the graphite basal plane

toward electron transfer and adsorption. A variety of electrode pretreatment procedures have

been proposed to increase the electron transfer rates. The type of carbon, as well as the

pretreatment method, thus has a profound effect upon the analytical performance. The most

popular carbon electrode materials are those involving glassy carbon, carbon paste, carbon fiber,

screen printed carbon strips4.

1.2.2.3 Carbon Paste Electrodes

A carbon-paste electrode (CPE) is made from a mixture of conducting graphite powder and a

pasting liquid. These electrodes are simple to make and offer an easily renewable surface for

electron exchange. In general, CPEs are popular because carbon pastes are easily obtainable at

minimal costs and are especially suitable for preparing an electrode material modified with

admixtures of other compounds thus giving the electrode certain pre-determined properties.

Electrodes made in this way are highly selective sensors for both inorganic and organic

electrochemistry. Despite their growing popularity, the exact behaviour of carbon-paste

electrodes is not fully understood. It is possible that some of the electrochemistry observed at

these electrodes involves permeation of the pasting liquid layer by the electroactive species. The

biggest disadvantage of CPEs, which limits their applicability in practical analysis, is that

11

success in working with carbon paste electrodes depends on the experience of the user. In

contrast to commercially available solid electrodes for which basic electrochemical

characteristics are comparable for almost all products from each manufacturer, each carbon paste

unit is an individual, where the physical, chemical and electrochemical properties may differ

from one preparation to another6.

1.2.3 Electrochemical Techniques

1.2.3.1 Potential Sweep Methods

Of all the methods available for studying electrode processes, potential sweep methods are

probably the most widely used, particularly by non-electrochemists. They consist in the

application of a continuously time-varying potential to the working electrode. These indicate the

occurrence of oxidation or reduction reactions of electroactive species in solution (faradaic

reactions), possibly adsorption of species according to the potential, and a capacitive current due

to double layer charging. In linear sweep voltammetry, the potential scan is done in only one

direction, stopping at a chosen value, Ef. The scan direction can be positive or negative and, in

principle, the sweep rate can have any value5.

1.2.3.2 Cyclic Voltammetry

Cyclic voltammetry (CV) is the most widely used technique for acquiring qualitative information

about electrochemical reactions. The power of CV results from its ability to rapidly provide

considerable information on the thermodynamics of redox processes, on the kinetics of

heterogeneous electron-transfer reactions, and on coupled chemical reactions or adsorption

processes. It is usually the first experiment performed in an electroanalytical study, since it offers

a rapid location of redox potentials of the electroactive species, and convenient evaluation of the

effect of media upon the redox process. It enables the electrode potential to be rapidly scanned in

search of redox couples. Once located, a couple can then be characterized from the potentials of

peaks on the cyclic voltammogram and from changes caused by variation of the scan rate5.

The repetitive triangular potential excitation signal for CV causes the potential of the working

electrode to sweep back and forth between two designated values (the switching potentials). To

12

obtain a cyclic voltammogram, the current at the working electrode is measured during the

potential scan (Fig. 1.3).

Fig. 1.3: Cyclic Voltammetry. (A) Potential waveform. The potential sweeps between two values

(V1 and V2). The scan rate is determined by the slope of the line. (B) The resultant

voltammogram of current is plotted against the applied potential of ferri/ferrocyanide solution

(Figure drawn by adaptation from [8]).

In the forward scan, the potential is scanned positively and a diffusion layer between the

electrode surface and the surrounding solvent ions is created from an increase in potential. When

the [Fe(CN)6]4- is continuously oxidized (Fig. 1.3), the anodic current is increased as a result of

the movement of electrons from a result of [Fe(CN)6]4- oxidizing to [Fe(CN)6]

3-. The current will

continue rise until the concentration of the species [Fe(CN)6]4- at the electrode surface is

13

substantially diminished and results in a current peak. And the current will further decrease as

the solution around the electrode is depleted of [Fe(CN)6]4- due to the electrolytic conversion to

[Fe(CN)6]3-. This occurs when the diffusion layer is produced and the flux of reactant has been

depleted from the electrode8.

During the scan in reverse direction, the potential is scanned negatively. Initially the potential is

still sufficiently positive to retain [Fe(CN)6]3- ions, so that a positive anodic current is observed,

though the potential is scanning in the negative direction. When the electrode is becoming a

sufficiently reductant, [Fe(CN)6]3- ions that accumulated on the electrode surface will undergo

reduction process causing a cathodic current. Once [Fe(CN)6]4- has accumulated on the electrode

surface, the current signal will read neutral. This completes a redox reaction where both anodic

and cathodic currents are produced and is defined as reversible reaction as discussed in section

1.2.1.3.8.

The important parameters in a cyclic voltammogram are the peak potentials (Epc, Epa) and peak

currents (ipc, ipa) of the cathodic and anodic peaks, respectively. For a reversible reaction, the

formal reduction potential E° is given by

E° 789:78;

[1-10]

The peak separation is

ΔE> ?E>@ ! E>A? 2.303RT/nF [1-11]

Thus, for a reversible redox reaction at 25°C with n electrons ΔE>should be 0.0592/n V or about

60 mV for one electron. However, practically this value is difficult to attain because of such

factors as cell resistance. Both the cathodic and anodic peak potentials are independent of the

scan rate. It is possible to relate the half-peak potential (Ep/2, where the current is half of the peak

current) to the polarographic half-wave potential, E1/2:

E>/ EG/ H .I

J V [1-12]

14

As the oldest technique in electrochemistry in 1900s, it is most well studied. It is also useful in

the study of absorption process within the liquid-electrode junction and the interfacial behaviours

of the analyte as it undergoes absorption and desorption processes9. Due to the rapid nature of

this technique, sensitivity is compromised. However, as a linear sweep waveform, background

charging current and current formed from the electroactive species cannot be differentiated.

15

1.2.3.3 Differential Pulse Voltammetry

Differential Pulse Voltammetry (DPV) is an extremely useful technique for measuring trace

levels of organic and inorganic species. This technique is comparable to normal pulse

voltammetry in that the potential is also scanned with a series of pulses. However, it differs from

Normal Pulse Voltammetry (NPV) because each potential pulse is fixed, of small amplitude (10

to 100 mV), and is superimposed on a slowly changing base potential (Fig 1.4).

Fig. 1.4: (A) Schematic waveform of pulses superimposed on a staircase to form differential

pulse voltammetry. (B) The typical diffierential pulse voltammogram of current is plotted against

the applied potential. (Figure drawn by adaptation from [4])

16

Current is measured at two points for each pulse, the first point just before the application of the

pulse and the second at the end of the pulse. These sampling points are selected to allow for the

decay of the nonfaradaic current. The difference between current measurements at these points

for each pulse is determined and plotted against the base potential. The resulting differential

pulse voltammogram consists of current peaks, the height of which is directly proportional to the

concentration of the corresponding analytes. Such quantitation methods depend not only on the

corresponding peak potentials but also on the widths of the peak. The peak shaped response of

differential pulse measurements resulted in improved resolution between two species with

similar redox potentials. The peak-shaped response, coupled with the flat background current,

makes the technique particularly useful for analysis of mixtures4.

The selection of the pulse amplitude and potential scan rate usually requires a trade-off among

sensitivity, resolution, and speed. For example, larger pulse amplitudes result in larger and

broader peaks. Pulse amplitudes of 25-50 mV, coupled with a 5 mV/s scan rate, are commonly

employed. Irreversible redox systems result in lower and broader current peaks compared with

those predicted for reversible systems10. In addition to improvements in sensitivity and resolution,

the technique can provide information about the chemical form in which the analyte appears,

such as oxidation states, complexation, etc.

17

1.2.3.4 Square-Wave Voltammetry

The excitation signal in Square-wave voltammetry (SWV) consists of a symmetrical square-

wave pulse superimposed on a staircase waveform, where the forward pulse of the square wave

coincides with the staircase step (Fig. 1.5). The net current is obtained by taking the difference

between the forward and reverse currents and centered on the redox potential. The peak height is

directly proportional to the concentration of the electroactive species and direct detection limits

as low as 10-8 M is possible.

Fig. 1.5: (A) Schematic waveform for square-wave voltammetry. (B) The typical square-wave

voltammogram of current is plotted against the applied potential. (Figure drawn by adaptation

from [5])

18

Square-wave voltammetry has several advantages, including excellent sensitivity and the

rejection of background currents. Another is its ability to scan the voltage range over one drop

during polarography with the DME. Applications of square-wave voltammetry include the study

of electrode kinetics with regard to preceding, following, or catalytic homogeneous chemical

reactions, determination of some species at trace levels, and its use with electrochemical

detection in HPLC5.

1.2.3.5 Electrochemical Impedance Spectroscopy

Electrochemical Impedance Spectroscopy (EIS) or AC impedance methods have seen

tremendous increase in popularity in recent years. Initially applied to the determination of the

double layer capacitance, they are now applied to the characterization of electrode processes and

complex interfaces. EIS studies the system response to the application of a periodic small

amplitude AC signal. These measurements are carried out at different AC frequencies and

analysis of the system response contains information about the interface, its structure and

reactions taking place there. However, EIS is a very sensitive technique and it must be used with

great care. Besides, it is not always well understood. It should be stressed that EIS cannot give all

the answers. It is a complementary technique and other methods must also be used to elucidate

the interfacial processes11.

The capacitance is only due to the working electrode, whilst the resistance includes the resistive

components of the electrode process, of the solution. In some cases a combination of resistance

and capacitance in parallel has also been used. However, a disadvantage of this type of technique

is that the impedance of the whole cell is measured, whereas in the investigation of electrode

processes one is interested in the properties of one of the electrodes. It is possible to reduce the

contribution of the unwanted components by using an auxiliary electrode with an area large

relative to that of the electrode being studied11.

19

The Randles cell is one of the simplest cell models. It includes a solution resistance (Rs), a

double layer capacitor (Cdl) and a charge transfer (Rct) or polarization resistance (Rp)11. In

addition to be a useful model in its own right, the Randles model is the starting point for other

more complex models. The equivalent circuit for the Randles cells is shown in Fig. 1.6. The

double layer capacity is parallel with the impedance due to the charge transfer reaction. Fig is the

sample Nyquist plot for a typical Randles cell. The Nyquist plot for a Randles cell is always a

semicircle. The solution resistance can be found by reading the real axis value at the high

frequency intercept, which is the intercept near the origin of the plot. In this case the solution

resistance is 20 Ω. The real axis value at the other intercept (low frequency) is the sum of the

charge transfer resistance and the solution resistance. The diameter of the semicircle is therefore

equal to the charge transfer resistance. In this case the Rct is 250 Ω.

(a) (b)

Fig. 1.6: (a) Randles cell schematic diagram (b) sample Nyquist plot, assuming Rs=20 Ω and

Rct=250 Ω. (Figure drawn by adaptation from [11])

EIS has become a mature and well understood technique. It is now possible to acquire, validate

and quantitatively interpret the experimental impedances. However, the most difficult problem in

EIS is modeling of the electrode processes. There is almost an infinite variety of different

reactions and interfaces that can be studied (corrosion, coating, conducting polymers, batteries

and fuel cells, etc.) and the main effort is now applied to understand and analyze these

processes11.

20

1.2.4 Application of Electrochemistry: Nanomaterials and Electrochemical Biosensors

Biosensors can be applied to a large variety of samples including body fluids, food samples, and

cell cultures and be used to analyze environmental samples. Designed for the purpose, biosensors

are generally highly selective due to the possibility to tailor the specific interaction of

compounds by immobilizing biological recognition elements on the sensor substrate that have a

specific binding affinity to the desired molecule12. Typical recognition elements used in

biosensors are: enzyme, nucleic acids, antibodies, whole cells, and receptors. Of these, enzymes

are among the most common. Today, a multitude of instruments referred to as biosensors can be

found in labs around the world and there is a growing number of biosensors being used as

diagnostic tools in point-of-care testing, but the realization of cheap handheld devices is almost

limited to one well-known example: the glucose sensor. In many cases the main limitation in

realizing point-of-care testing/sensing devices is the ability to miniaturize the transduction

principle and the lack of a cost-effective production method. Thus, they have to be confirmed to

expert users of high-cost equipment in a lab environment and cannot be used by patients

themselves or doctors in the field. Other inherent advantages of electrochemical biosensors are

their robustness, easy miniaturization, excellent detection limits, also with small analyte volumes,

and ability to be used in turbid biofluids with optically absorbing and fluorescing compounds12.

Several electrochemical biosensors associated with nanomaterials have been developed to fulfill

their practical needs. For example, the combination of electrochemical immunosensors using

single-wall carbon nanotube (SWNT) forest platforms with multi-label secondary antibody-

nanotube bioconjugates was described for highly sensitive detection of a cancer biomarker in

serum and tissue lysates13. Also, a SWNT-arrayed microelectrode chip has been reported to

detect the electroacitive amino acids: L-Tyrosine, L-Cysteine and L-Tryptophan14. Although,

carbon nanotubes (CNTs) have emerged as a novel class of nanomaterials and consequently

receive considerable interest in a plethora of areas, metal impurities in CNTs have been

confirmed to be responsible for the electrocatalysis problem of signal detection15. In addition, the

electrochemical properties of amino acids containing no sulfur atoms have been investigated

using stationary or rotating solid electrodes such as Au and vitreous carbon, due to the fact that

among 20 amino acids, only tryptophan and tyrosine are specifically oxidizable at a gold,

21

platinum or carbon electrode16. On the other hand, a voltammetric method for a direct

determination of gold nanoparticles using graphite-epoxy composite electrode has been

described17. Moreover, the electrochemical signal of the monobase modified colloidal gold

nanoparticles can be used to monitor the electrochemical coding of single-nucleotide

polymorphisms18. At last, a sensitive immunosensor based on the direct electrical detection of

Au nanoparticles have been reported to detect human chorionic gonadotropin hormone,

pregnancy marker19.

Electrochemical biosensors have existed for nearly fifty years and seem to possess great potential

for the future. This technology gains practical usefulness from a combination of selective

biochemical recognition with the high sensitivity of electrochemical detection. With the

development of technology, such biosensors profit from miniaturized electrochemical

instrumentation and are thus very advantageous for some sophisticated applications requiring

portability, rapid measurement and use with a small volume of samples. Numerous commercial

applications confirm the attractive advantages of electrochemical biosensors12.

22

1.3 Ideal Immunosensor Properties

1.3.1 General Working Principle of Immunosensors

The general working principle of the immunosensors is based on the fact that the specific

immunochemical recognition of antibodies (antigens) immobilized on a transducer to antigens

(antibodies) in the sample media can produce analytical signals dynamically varying with the

concentrations of analytes of interest20. The general immunosensor design consists of three

individual parts in close contact: a biological recognition element, a physicochemical transducer,

and an electronic part. Antibodies or antibody derivatives (antigens or haptens) are usually

served as the biological recognition elements, which are either integrated within or intimately

associated with a physicochemical transducer (Fig. 1.7). This recognition reaction defines the

high selectivity and sensitivity of the transducer device. The electronic part is used to amplify

and digitalize the physicochemical output signal from the transducer devices such as

electrochemical (potentiometry, conductometry, capacitive, impedance, amperometry), optical

(fluorescence, luminescence, refractive index), and microgravimetric devices.

Fig. 1.7: The biosensor process including biological recognition element, physicochemical

transducer and signal processing steps.

It has been suggested that an ideal immunosensor design should posses the following

specifications: the ability to detect and quantify the antigens (antibodies), the capacity to

transform the binding event without externally added reagents, the ability to repeat the

measurement on the same device, and the capacity to detect the specific binding of the antigens

(antibodies) in real samples. All of these specifications have been the main issues to pursue in

developing immunosensors applied in various fields.

23

1.3.2 Characterization of Immunosensors in Clinical Analysis

As an important branch of immunoassay techniques, immunosensors possess all essential

performance characteristics of immunoassays. They have been the subject of expanding interest

in the immunochemical studies with enormous potential in clinical diagnosis21-23

, environmental

analysis24-25

, and biological process monitoring27. As for the medical diagnosis of some diseases,

herein considerable efforts have been devoted to the development of precise, rapid, sensitive, and

selective immunosensors by measurement of the markers or pathogenic microorganisms

responsible for the diseases, such as proteins, enzymes, viruses, bacteria, and hormones21, 27-28

.

For example, Chagas disease, an American trypanosomiasis caused by the hemoflagellate

Trypanosoma cruzi, is one of them. It has been reported to probe the presence of antibodies

against T. cruzi, in blood donors and also to follow the antibody decay during treatment of

chagasic patients with the available drugs through amperometric immunosensor29. A

piezoelectric immunosensor was developed for the on-line detection of severe acute respiratory

syndrome (SARS)-associated corona virus (SARA-CoV) in sputum in the gas phase with a

relatively fast speed and low cost30. In addition, the analysis of some tumour markers plays an

important role in diagnosing, screening, and determining the prognosis of a cancer disease.

Wilson proposed an electrochemical immunosensor for the simultaneous detection of two

tumour markers of CEA and AFP31. Although there are still problems associated with the assay

of analytes in real sample, there is an increasing number of utilization of immunosensors for

diagnosing infectious disease.

1.4 Antigen, Antibody, and Their Recognition Reaction

The immune response can be defined as any mechanism of identifying “nonself” substance from

‘self’ substances in an organism, which usually results in a more rapid destruction of those

substances identified as “noself”32. Some degree of this ability to identify and respond to foreign

substances has been found in very simple life forms, such as microbes. In higher forms of life,

particularly in mammals, the immune system is a complex mechanism in which identification

and communication take place in the blood and lymph.

When a foreign substance enters the body of an advanced animal, certain proteins are

synthesized to identify the invader and to prohibit its harmful effects. Antibodies are biologically

24

defined as the proteins that are formed when an animal is immunized with an antigen (nonself

substance). Antibodies show very high specificity and binding constants toward their

corresponding antigens. An antigen has been defined as “any agent that gives rise to antibody

formation specific for that agent when transferred to a living cell system containing cells of the

immunologically competent type”33. The natural antigens may be such macromolecule

substances as proteins and nucleic acids.

Antibodies are a family of glycoprotein known as immunoglobulin (Ig). There are generally five

distinct classes of glycoprotein (IgA, IgG, IgM, IgD, and IgE) with IgG being the most abundant

class (approximately 70%) and the most often used in immunoanalytical techniques34. As shown

in Fig. 1.8, IgG is a “Y”-shaped molecule based upon two distinct types of polypeptide chains.

The molecular weight of the smaller (light) chain is approximately 25000 Da, while that of the

larger (heavy) chain is approximately 50000 Da. In each IgG molecule, there are two light and

two heavy chains held together by disulfide linkages34. The variable and hypervariable regions of

Fab create an active portion that recognized a specific area of the antigen. The singular segment

at the other end of Y shape is known as Fc fragment, which cannot bind with antigen but has the

ability to affix the cell surface and to pass through the placenta35.

Fig. 1.8: A schematic illustrates the “Y”-shaped structure of an antibody. The region between the

heavy chain and the light chain is where antigen binding occurs. This open arm portion of the “Y”

shape is generally denoted as Fab, while the non-antigenic binding site in the base portion is

referred to as Fc. (Figure drawn by adaptation from [35])

25

Many different types of antibodies exist in the serum of animals immunized with specific

antigens. The mixture of these different types of antibodies is a so-called polyclonal antibody.

Because these antibodies arise from the clones of a number of separate “B” cells, they are

heterogeneous, and different antibodies in this mixture react with different antigenic

determinants. With the ever-increasing sophisticated genetic techniques, the production and use

of monoclonal antibodies have attracted more interest since this technique was first developed by

Kohler and Milstein in 1975, who won the Nobel Prize in 198436. Monoclonal antibodies are

produced by the fusion of myeloma (tumour) cells cultivated in vitro with mouse spleen B-

lymphocytes immunized with a specific antigen. Because a monoclonal antibody reacts only

with one specific antigenic determinant, it shows a higher sensitivity and better specificity than

the conventional polyclonal antibodies for immunoassays.

The specific binding between antigen and antibody is a collection of noncovalent forces,

including electrostatic forces, hydrophobic attractions, hydrogen bonding, and van der Waals

interactions. The interaction between antigen and antibody is quite strong, as indicated by the

large association constant of 105 – 10

12 /M

37. Therefore, the antibody-antigen complex does not

dissociate so readily unless some harsh solutions such as buffers at pH higher than 10 or lower

than 3, organic solvents, and saline solutions at high concentration are used to regenerate it38.

1.5 Applications of Electrochemical Immunosensors

Immunoassays are the quantitative methods o f analysis where antibodies are the primary

binding agents for the antigen (which is often the analyte) of interest. The net results of an

immunoassay are thus often the investigation of the binding between an antibody and its antigen

and the differentiation between bound and unbound antigen. In other words, all immunoassays

depend on measuring the fractional occupancy of the recognition sites. However, such a

measurement can rely on either the assessment of occupied sites or, indirectly, on measuring

unoccupied sites. This leads to the development of either a “competitive” or a “non-competitive”

immunoassay format, as described below.

26

1.5.1 Competitive Immunoassay Systems

In a competitive immunoassay, the sample analyte is mixed with labelled analyte, both of which

compete for a limited number of antibody-binding sites. This is schematically depicted in Fig.

1.9(b). In electrochemical immunoassays, an enzyme label or an electroactive label is commonly

used. Quantitative analysis can be achieved by determining the amount of labelled analyte that

interacted at the binding sites. Therefore, with a fixed number of antibody sites, a smaller signal

is expected when the ratio between the quantities of sample to labelled analyte is large. In

contrast, a larger signal is obtained when there is a small quantity ratio. Hence, the signal

produced by the bound labelled analyte is usually inversely proportional to the amount of sample

analyte.

Fig. 1.9: Schematic representation of (a) non-competitive and (b) competitive immunoassay

formats. (Figure drawn by adaptation from [6])

27

Competitive immunoassays can be used to quantify the extremely low concentrations of analytes

contained in body fluids (serum and urine). An electrochemical immunosensor based on the use

of a novel organometallic (η5-cyclopentadienyl)-tricarbonylmanganese redox label (cymantrene)

bound to bovine serum albumin (BSA) 39. A schematic diagram of this immunoassay is depicted

in Fig. 1.10. In their system, BSA was first labelled by cymantrene entities (redox marker). Then

the rabbit anti-bovine serum albumin (IgG fraction) was coupled to microbeads. The IgG-coated

agarose beads were put directly into the electrochemical cell containing labelled BSA molecules.

No separation step was needed. The electrochemical detection was based on the impedance

measurements of a one-electron reversible reduction of the organometallic probe in the

frequency range 1 Hz to 100 kHz. Cymantrene bound to BSA yielded the reversible redox

exchange at -1.84 V against the Ag/AgCl reference electrode. The relationship between the

electrochemical signal and the concentration was linear in a reasonably wide concentration range

(0.1 – 1.0 µM) 39.

Fig. 1.10: A schematic illustrates a competitive immunoassay format used for the analysis of

BSA. The rabbit anti-bovine serum albumin (IgG fraction) attached to microbeads and the

cymantrene labelled BSA attached to IgG-coated agarose beads. (Figure drawn by adaptation

from [39])

There are many other examples of competitive electrochemical immunoassays and

immunosensors for detecting clinically important analytes40-42

. Despite simplicity, a

disadvantage of competitive immunoassay is that labelling the analyte may reduce, or totally

remove, its binding affinity for antibody. This would occur if the analyte were labelled at a site

that is closely associated with an epitope.

28

1.5.2 Non-Competitive Immunoassay Systems

In a non-competitive immunoassay (also known as a “sandwich” immunoassay), the sample

analyte is captured by an excess of a capture antibody, separating it from the bulk sample. The

captured analyte is then exposed to an excess of second signal antibody, which will only bind to

the existing capture antibody-analyte complex. As shown schematically in Fig. 1.9(a), this

structure is a classic two-site immunoassay complex in which the analyte is sandwiched between

two antibodies. In this system, the signal antibody is often conjugated to either an enzyme label

or an electroactive label that produces a signal proportional to the amount of bound analyte.

In an ideal non-competitive immunoassay, no signal would be produced in the absence of any

analyte because there are no appropriate sites available for binding to the signal antibody.

However, in practice, this is not the case due to non-specific interactions between the signal

antibody and other components of the immunoassay. Therefore, it is always desirable to use a

blocking reagent to reduce these non-specific interactions. Non-specific adsorption also needs to

be considered when determining the quantity of signal antibody for use in a system. Although

this immunoassay format often offers superior specificity, it can only be used for the

quantification of analytes with two antigenic determinants that can be simultaneously recognized.

Heineman et al. proposed an enzyme-labelled sandwich immunoassay on paramagnetic

microbeads with mouse IgG as the analyte and β-D-galactosidase as the enzyme label43. β-

Galactosidase converted p-aminophenyl β-D-galactopyranoside to p-aminophenol (PAP). This

enzyme reaction was measured continuously by positioning the microbeads near the electrode

surface with a magnet. Electrochemical recycling occurred with PAP oxidation to p-quinone

imine (PQI) at +290 mV followed by PQI reduction to PAP at 300 mV vs. Ag/AgCl. A

calibration curve of PAP concentration vs. anodic current was linear between 10-4 and 10

-6 M

with a detection limit of 3.5×10-15 mol mouse IgG

43.

Another example is a non-competitive immunoassay system developed for the detection of

pathogenic Listeria monocytogenes in food samples using horseradish peroxidise (HRP)-labelled

signal antibody44. There are also examples of non-competitive assays in the literature for

analyzing different clinically important species45-47

.

29

1.5.3 Antibody Immobilization

The manner in which a capture antibody is immobilized on a solid phase is a critical aspect that

requires careful consideration in the design of an immunoassay system, whether it is competitive

or non-competitive. A desirable feature of the chosen method is that it results in an immobilized

capture antibody that is oriented with minimal steric hindrance to interact favourably with its

target antigen. Equally important, it is highly desirable to immobilize the antibody without a

significant change in its ability to bind its antigen. Clearly, all these features have a direct

bearing on the level of sensitivity and dynamic range achievable by an immunosystem. There are

several strategies for immobilizing a capture antibody on a solid phase including covalent

attachment, physical adsorption or electrostatic/physical entrapment in a polymer matrix.

1.5.3.1 Biotin-Streptavidin Interaction

Specific affinity interactions for antibody immobilization have been widely used in

immunoassay systems in recent years. The streptavidin-biotin interaction is one of the examples.

This technique may be used to immobilize various types of biomolecules such as nucleic acids,

polysaccharides, and proteins, including the capture antibody in immunoassay system48. This

technique usually involves biotinylating the capture antibody and coating a solid phase with

either avidin or streptavidin (Fig. 1.11).

Fig. 1.11: Schematic representation of biotin-streptavidin interaction where biotinylated

antibody attached to the streptavidin coated solid phase. (Figure drawn by adaptation from [48])

30

The dissociation constants of biotin-avidin and biotin-streptavidin interactions are of the order of

10-15 mol/L and are some of the largest free energies of association yet observed for non-covalent

interactions49. The complexes also withstand high temperatures, pH variations, and are resistant

to dissociation when exposed to chemicals such as detergents and protein denaturants50. Equally

important, the use of this immobilization technique maintains the biological function of the

immobilized antibody48. In some cases, neutravidin, which is an almost neutrally charged (pI of

6.3) variation of avidin, is used to minimize any non-specific binding by charged species to

maintain high binding affinity for biotin.

An electrochemical immunosensor for the detection of Mycobacterium tuberculosis, based on the

immobilization of a capture antibody using the biotin-streptavidin interaction, has been

reported51. In this system, biotinylated anti-M. tuberculosis antibody was immobilized on the

surface of a streptavidin-modified Screen Printed Carbon Electrode (SPCE). Incubation between

antigen M. tuberculosis and monoclonal mouse anti- M. tuberculosis was carried out remotely,

and then introduced to the sensor surface for capture by the immobilized capture antibody. The

immunosensor structure was completed by introducing Aminopyridine (AP)-labelled rabbit anti-

mouse antibody. The substrate 3-indoxyl phosphate was then introduced and converted to its

Indigo product by AP. Indigo was converted to hydrosoluble indigo carmine and the analytical

signal was produced by either cyclic or square-wave voltammetry. A detection limit of 1 ng/mL

M. tuberculosis was achieved by this immunoassay. The results were compared to those of a

similar assay, which relied on the passive adsorption of monoclonal rabbit anti-mouse antibody

directly onto the surface of a pre-treated SPCE. This assay format yielded a detection limit of 40

ng/mL, indicating that the biotin-streptavidin interaction used to immobilize capture antibody is

a suitable support for electrochemical immunosensing51.

31

1.5.3.2 Antibody Binding Proteins

Another commonly used affinity-based immobilization technique for capture antibodies in

immunoassay systems involves a bacterial antibody-binding protein. The two most common of

which are protein A and protein G. These proteins bind specifically to antibodies through their

non-antigenic (Fc), which allow the antigen binding sites of the immobilized antibody to be

oriented away from the solid phase and be available to bind the target analyte. As these proteins

interact directly with the Fc region of antibodies, there is no need for antibody biotinylation.

Protein A has a molecular weight of approximately 42 kDa and was originally isolated from the

cell wall of Staphylococcus aureus52. It is known to contain five Fc binding domains located

towards its –NH2 terminal. However, the building capacity of Protein A is limited to three human

IgG subclasses (IgG 1, 2 and 4)53. Also, protein A will not bind to goat and rat IgG, and only

weakly to mouse IgG (30). The second bacterial antibody binding protein, protein G, is a cell

surface protein of group C and G streptococci with three Fc binding domains located near its C-

terminal, and has specificity for subclasses of antibodies from many species53.

32

Zacco et al. reported a rigid material for use as a scaffold in electrochemical immunosensing that

is based a protein A bulk-modified graphite-epoxy biocomposite (Protein A-GEB) 54

. This

biocomposite not only provides a means to securely immobilize the capture antibody, but also

acts as the transducer for the electrochemical signal. First of all, rabbit antibody (RIgG) was

introduced to the layer and allowed to interact with protein A. Biotinylated anti-RIgG was then

introduced to bind to the immobilized RIgG. Streptavidin-labelled HRP was then introduced to

bind to the bound anti-RIgG before processing the immunoassay by introducing the substrate

H2O2 (Fig. 1.12).

Fig. 1.12: Schematic representation of the immunoassay based on a two-step labelling procedure

using ProtA-GEB biocomposite as a transducer. (A) RIgG immobilisation on the surface of the

electrode based on its interaction with protein A. (B) Competitive immunoassay, using anti-RIgG

and biotinylated anti-RIgG. (C) Enzyme labelling using HRP-streptavidin. (D) Electrochemical

enzyme activity determination. (Figure drawn by adaptation from [54])

33

This assay could distinguish between 2 pmol and 10 pmol of anti-RIgG. Furthermore, they have

also shown that the Protein A-GEB layer can be regenerated by polishing with abrasive and

alumina papers, which yields a smooth mirror finish containing freshly exposed protein A that

may be reused in subsequent assays. By applying a solution with the appropriate pH and ionic

strength, the interaction between proteins A or protein G and the antibody can be reversed,

enabling easy renewal of sensing surfaces55, 56

. This has been demonstrated by Yakovleva et al.

who have developed a renewable microfluidic immunosensor using protein G as the

immobilization aid56.

1.5.4 Nanomaterials Based Immunosensors

The unique properties of nanoscale materials offer excellent prospects for designing highly

sensitive and selective bioassays of nucleic acids and proteins. The creation of such designer

nanomaterials for specific biosensing and bioassay applications greatly benefits from being able

to vary the size, composition, and shape of the materials and hence tailor their physical and

chemical properties. Due to the tiny size of nanomaterials, their properties are strongly

influenced by the binding of target biomolecules. Nanoparticles of different compositions and

dimension have been widely used in recent years as versatile and sensitive tracers for the

electronic, optical, and microgravimetric transduction of different biomolecular recognition

events63-67

. The enormous signal enhancement associated with using nanoparticle amplifying

labels and with forming nanoparticle-biomolecule assemblies provides the basis for ultrasensitive

optical and electrical detection. Such protocols couple the amplification features of nanoparticle-

biomolecule assemblies with highly sensitive optical or electrochemical transduction schemes.

Multi-amplification protocols, combining several nanomaterials-based amplification units and

processes, can also be designed for addressing the high sensitivity demand of modern bioassays.

The unique catalytic properties of metal nanoparticles stimulate their enlargement by the same

metal or another one to offer substantial signal amplification. It is also possible to dramatically

increase the number of tags per binding event and achieve enormous signal amplification by

encapsulating numerous signal-generating molecules within a nanoparticle host. These

nanomaterials-based biosensing and bioassays can be combined with additional amplification

processes, such as surface preconcentration or enzymatic recycling.

34

Over the last two decades, considerable attention has been paid to the development of new

biocompatible nanomaterials with suitable hydrophilicity, high porosity, and large surface area

for immunological recognition element immobilization.

Carbon nanofiber (CNF) has been recognized as a very promising material based on its

nanostructure and properties. The oxidation of CNF with nitric acid can produce carboxyl groups

without degradation of the structural integrity of its backbone. Compared to CNT, CNF has a

much larger functional surface area and higher ratio of surface active groups to volume.

Therefore, it can be used for covalent binding of proteins and mediators with the help of cross-

linking reagent. Ju’s group68 used soluble CNF to construct an immunosensor for a rapid

separation-free immunoassay for carcinoma antigen 125 (CA 125). The acidic oxidation of the

CNF provided its solubility and wettability for the convenient preparation of a porous CNF

membrane and a larger number of active sites for covalent binding of CA 125 and thionine as the

electron-transfer mediator. The covalent attachment of proteins to the surface overcame the

problems of instability and inactivation. With a competitive mechanism, the CNF-based

immunosensor was able to detect CA 125 concentrations from 2 to 75 U/mL68.

35

Fig. 1.13: Multiprotein electrical detection protocol based on different inorganic colloid

nanocrystal tracers. (A) Introduction of antibody-modified magnetic beads; (B) binding of the

antigens to the antibodies on the magnetic beads; (C) capture of the nanocrystal-labeled

secondary antibodies; (D) dissolution of nanocrystals and electrochemical stripping detection.

(Figure drawn by adaptation from [69])

An electrochemical immunoassay protocol for the simultaneous measurements of proteins, based

on the use of different inorganic nanocrystal tracers (Fig 1.13), was reported by Liu et al.69. The

multi protein electrical detection capability is coupled to the amplification feature of

electrochemical stripping transduction and with an efficient magnetic separation. The

multianalyte electrical sandwich immunoassay involves a dual binding event, based on

antibodies linked to the nanocrystal tags and magnetic beads. Carbamate linkage is used for

conjugating the hydroxyl-terminated nanocrystals with the secondary antibodies. Each

biorecognition event yields a distinct voltammetric peak, whose position and size reflects the

identity and level, respectively, of the corresponding antigen. These nanocrystal labels exhibit

similar sensitivity. Such electrochemical coding could be readily multiplexed and scaled up in

36

multiwall microtiter plates to allow simultaneous parallel detection of numerous proteins or

samples and is expected to open new opportunities for protein diagnostics and biosecurity.

To enhance the sensitivity of the nanoparticle label-based electrochemical immunosensors,

Wang et al. developed a novel electrochemical immunosensor based on poly (guanine)-

functionalized silica nanoparticle (NP) labels and mediator-generated catalytic reaction70.

Biotinylated primary antibodies are first immobilized on an avidin-modified electrode and mouse

IgG then bound onto the antibody, followed by interaction with mouse IgG specific antibody-

silica NPs covered with poly [G], which introduces a large amount of guanine residues on the

electrode surface. Guanines on silica NPs catalyze the oxidation of Ru (bpy) 32+. The amplitude

of the oxidation current depends on the amount of guanine, which is related to the concentrations

of sample solutions. The amplification of the catalytic signals is attributed to the attachment of a

large number of guanine markers per antibody-antigen-antibody complex formed. This

immunobiosensor is very sensitive, and the limit of detection was found to be 0.02 ng/mL. An

attractive feature of this method is it makes it feasible to develop a cheap, sensitive, and portable

device for multiplexed diagnoses of different proteins70.

Quantum Dots (QDs) are the most eye-catching fluorophores developed for fluorescent images

and bioconjugates in the past two decades. They exhibit some important differences compared to

traditional fluorophores, such as organic fluorescent dyes and naturally fluorescent proteins. QDs

are nanometre-scale atom clusters, containing from a few hundred to a few thousand atoms of

semiconductor material such a CaSe or CaTe, which sometimes have been coated with an

additional semiconductor shell such as ZnS to improve their optical properties. Besides their

excellent quantum efficiency, QDs also show several other advantages over conventional organic

dyes. Their emission spectra are symmetric, narrow, and tunable according to their size and

chemical composition, permitting close spacing of different probes without substantial spectral

overlap. They exhibit excellent photostability, tolerating long-time excitation. QDs also display

broad adsorption spectra, making it possible to minimize sample autofluorescence by choosing

an appropriate excitation wavelength. Thus, QDs have attracted increasing interest as tags for

immunoassays besides their application in bioimage in the past decade71.

37

1.5.4.1 Gold Nanoparticle Based Immunosensors

Nanosized particles of noble metals, especially gold nanoparticles (AuNPs), have received great

attention due to their attractive electronic, optical, and thermal as well as catalytic properties and

potential applications in the fields of physics, chemistry, biology, medicine, and material science

and their different interdisciplinary fields72 and therefore the synthesis and characterization of

AuNPs have attracted considerable attention from a fundamental and practical point of view.

The preparation of AuNPs generally involves the chemical reduction of gold salt in the aqueous

organic phase or in two phases73. However, the high surface energy of AuNPs makes them

extremely reactive, this causes aggregation if their surfaces are not protected or passivated. Thus

special precautions have to be taken to avoid aggregation or precipitation. Typically, AuNPs are

prepared by chemical reduction of the corresponding metal salts in the presence of a stabilizer

that binds to their surface to impact high stability and rich linking chemistry and to provide the

desired charge and solubility properties. Some of the methods commonly used for surface

passivation include protection by self-assembled monolayers, the most popular being citrate73

and thiol-functionalized organics74; encapsulation in the H2O pools of reverse microemulsions75;

and dispersion in polymeric matrixes76.

From an electroanalytical point of view, more attention has been paid to AuNPs because of their

good biological compatibility, excellent conducting capability, and high surface-to-volume ratio.

These features provide excellent prospects for interfacing biological recognition events with

electronic signal transduction and make AuNPs extremely suitable for developing novel and

improved electrochemical sensing and biosensing systems77.

AuNPs have surprising electrostatic adsorption ability for proteins, and they have been extremely

used as an immobilized matrix for retaining the bioactivity of antigens and antibodies and for

promoting the direct electron transfer of the immobilized proteins. There characters allow

AuNPs to be widely used for the fabrication of various types of electrochemical immunosensors,

especially amperometric immunosensors. For example, a particle-based renewable

electrochemical magnetic immunosensor was developed by Wang et al. by using magnetic beads

and gold nanoparticle labels78. Anti-IgG antibody-modified magnetic beads were attached to a

renewable carbon paste transducer surface by a magnetic that was fixed inside the sensor. Gold

38

nanoparticle labels were capsulated to the surface of magnetic beads with a sandwich

immunoassay. Highly sensitive electrochemical stripping analysis offers a simple and fast

method to quantify the captured gold nanoparticle tracers and avoid the dissolution step and the

use of an enzyme label and substrate. The stripping signal of gold nanoparticles is related to the

concentration of target IgG in the sample solution. The detection limit of 0.02 µg/mL of IgG was

obtained under optimum experimental conditions78.

Furthermore, the interaction between saccharide and protein interactions was studied using a

couple of sialic acid derivatives and Alzheimer’s amyloid-beta (Aβ) 81

. Firstly, the Au

nanoparticles were electrochemically deposited on a screen printed strip, followed by SAM

formation of the acetylenyl group on AuNPs, and then the azide-terminated sialic acid was

immobilized on the AuNP modified strip through cycloaddition. The attachment of Aβ peptides

to the sialic acid layer was further confirmed from electrochemistry and atomic force microscopy

imaging. The intrinsic oxidation signal of the captured Aβ (1-40) and (1-42) peptides, containing

a single tyrosine (Tyr) residues, was monitored at a peak potential of 0.6 V (with respect to

Ag/AgCl reference electrode) using DPV. From this glycoside cluster effect, the immobilization

of the saccharides as the biorecognition materials on carbon electrodes provides new routes for

analysis of saccharide-protein interactions and electrochemical biosensor development. The

presence of both AuNP (for immobilization of biomolecules) and bare carbon (for

electrochemical detection on the electrode) enable the electrochemical sensing with easy

fabrication and low cost.

1.6 Objectives

The main goal of this research is to design a biosensor that can be easily handled, has higher

specificity, sensitivity, low-cost, and rapid response and have a better detection of desired

analytes.

1. The first part of this research involves the development of a gold-nanoparticle based

sandwich type immunosensor to identify trace amounts of human chorionic gonadotropin

hormone based on the direct electrochemical detection of Au nanoparticles. The human

chorionic gonadotropin hormone (hCG) is produced right after conception when implantation

39

of fertilized egg occurs in the uterine lining. As a result, increased levels of hCG can be

observed in the urine of pregnant females82. In this study, hCG concentrations were detected

based on the direct oxidation and reduction peak currents of the Au nanoparticles that serve

as a biosensor scheme. The properties of gold nanoparticles provide a greater advantage in

enhanced electrochemical detection when compared to conventional electrochemical

biosensor devices owing from its ability to enhance electron transfer between electroactive

constituents to the electrode, increase surface-to-volume ratio making them more sensitive to

adsorbed surfaces and its stability while maintaining the bioactivity of the immobilized

biomolecule83.

2. The second part of this research involves a miniaturized potentiostat in connection with

disposable screen-printed carbon strips (SPCS) for the point-of-care detection of proteins.

The performance of the miniaturized potentiostat was studied using growth hormone (GH) as

the target protein. GH has been considered to play an essential role in a variety of biological

processes84. Either excess or deficiency of GH produces significantly problems at various

ages. Despite its natural function, GH has also been banned by the World Anti-Doping

Agency (WADA) for its abuse. Current bioanalytical techniques have reported the difficulty

to detect GH doping, because recombinant GH is indistinguishable analytically from

endogenous GH. This research aims to address the issue of measuring GH in small volumes,

which has been challenging the limits of analytical detection systems. The electrochemical

measurements utilize the redox activity of ferri/ferrocyanide in cyclic voltammetry and

Impedance spectroscopy. Furthermore, comparison between miniaturized potentiostat and a

conventional electrochemical analyzer was studied. Therefore, this study indicated that this

versatile platform could be easily adapted for decentralized electrochemical immunosensing

of clinically important proteins.

40

Chapter 2 Gold nanoparticle based electrochemical detection of

Human Chorionic Gonadotropin

2.1 Introduction

Human chorionic gonadotropin (also known as beta-hCG), a 37 kDa glycoprotein hormone, is

normally produced by the syncytiotrophoblastic cells of the placenta and is elevated in

pregnancy86. In most normal pregnancies with hCG levels below 1200 mIU/mL, the hCG usually

doubles every 48-72 hours and increases by at least 60% every two days. Between 1200 and

6000 mIU/mL serum hCG levels in early pregnancy, the hCG usually takes 72-96 hours to

double. Above 6000 mIU/mL, the hCG often takes over four or more days to double86. Briefly,

hCG consists of two subunits, designated and , which are noncovalently bonded and are

synthesized separately (Fig. 2.1). The -subunit unit of hCG is identical to the -subunit of other

pituitary glycoprotein hormones, but the biological activity of hCG is conferred by the -subunit.

Both - and -subunit are species specific. It’s most important uses as a tumour marker are in

gestational trophoblastic disease and germ cell tumours. All gestational trophoblastic tumours

produce hCG. Therefore, the detection of hCG in diagnostic processes will provide an indication

of the effectiveness of a particular tumour treatmet87.

Fig. 2.1: (a) Schematic representation of and subunits of human chorionic gonadotropin

hormone; (b) Anti- –FSH antibody (Polyclonal anti-human -subunit of follicle-stimulating

hormone) as primary antibody; (c) Anti- hCG (anti-chorionic gonadotropin -subunit (ab 1))

antibody as secondary antibody.

41

Various commonly available methods were developed for the detection of hCG, such as

fluorescence labelled antibody immunoassay (detection limit of 2 mIU/mL), radioimmunoassay

(RIA; detection limit of 100 mIU/mL) 86, enzyme-linked immunosorbent assay (ELISA;

detection limit of 17 mIU/mL) 87, and colloidal gold-labelled test paper card assay (detection

limit of 50 mIU/mL) 88. Although the operation of colloidal gold labelled test paper card assay is

simple, its sensitivity is low. All other three methods are sufficiently sensitive and precise, but

these conventional immunoassay methods require a radioisotope, enzyme, or fluorescence

labelled antibody/antigen and may suffer from draw backs of required skilled personnel, time-

consuming procedures, and expensive chemicals. Thus development of a new method with high

sensitivity and specificity for direct detection of hCG is highly desirable87.

Electrochemistry immunoassay offers good possibilities for sensitive detection of unlabeled

protein because it is highly sensitive, low cost, low power requirement, and has high

compatibility with advance micromachining technologies. A variety of electrochemical

biosensing schemes involving enzymes88-90

and colloidal metal nanoparticles91 as labels have

been reported. In particular, due to its excellent conductivity and catalytic properties, metal

nanoparticle can act as “electronic wire” and promote the communication between the redox

centers in protein and electrode surfaces92. The catalytic activity of metal nanoparticles to

amplify the electrochemical reactions gives them a significantly priority in the design of

electrochemical biosensors. The metal nanoparticles are highly stable and less vulnerable to

degradation/denaturation caused by the solution matrix than their enzymatic counterparts.

Moreover, metal nanoparticles are suitable for the multiplexed detection schemes in a more cost-

effective way than the enzyme-based ones.

Recently, it has become abundant with the nanomaterials-based biosensors for the detection of

proteins. Liang and Mu93 modified screen printed electrodes with Au nanoparticles and

performed a flow-injection immunobioassay for the detection of interleukin-6 in humans. Li et

al.94 developed a real-time immunoassay utilizing an electroimmunosensing microchip and Au

nanoparticles for the capacitive immunosensing of transferring. The amplification of the antigen-

antibody interactions on quartz crystal microbalance immunosensors was performed by Tang et

al95. Yin et al.96 prepared ultrathin alumina sol-gel derived films containing Au nanoparticles for

the capacitive immunosensing of transferring. Tang et al.97 prepared a thionine and Au

42

nanoparticle-modified carbon paste interface for the electrochemical immunoassay of carcinoma

antigen 125 (CA 125). Chemiluminescence detection of Au nanoparticles in biological

conjugates has been utilized for the development of highly sensitive immunoassays by Li et al.98

For the voltammetric detection of Au nanoparticles, various techniques; such as direct oxidation

peak current detection, direct reduction peak current after pre-oxidation, and silver enhancement

were demonstrated. An electrochemical biosensor for the detection of DNA hybridization, in

which the oxidation and reduction steps of Au nanoparticles took place on a single surface, has

been reported by Alegret and co-workers80. Moreover, previous studies about an electrochemical

immunosensor for detection of hCG based on Au nanoparticles was reported by Idegami and co-

workers. The electrochemical reduction process involves the following reaction.

-L(MN 3O P -L 4(M

In our laboratory, we utilized a gold-nanoparticle based sandwich type immunosensor to identify

trace amounts of human chorionic gonadotropin based on the direct electrochemical detection of

Au nanoparticles.

Fig. 2.2: Screen Printed Carbon Strip (SPCS) chips.

This disposable sensor system is based on the three-electrode type of SPCSs with the strong

advantage of fabricating a large number of near identical electrodes at a low-cost (Fig. 2.2).

43

Followed by the direct immobilization of primary antibody on the working electrode, a

sandwich-type immunosensor was built up (Fig. 2.3).

Fig. 2.3: Schematic illustration of the disposable immunosensor system. (a) The primary

antibody was immobilized directly on the SPCS chips, and a series of sandwich type

immunoreactions took place on the electrode surface. (b) A high potential, at 1.2 V, was applied

for pre-oxidation of Au nanoparticles and then the voltammetric measurements was taken.

2.2 Experimental

2.2.1 Instrument and Materials

Polyclonal anti-human -subunit of follicle-stimulating hormone (Mab-FSH) with an affinity

constant of 2.8×109 M

-1 was purchased from abcam. The anti-chorionic gonadotropin -subunit

(ab 1) antibody (Mab-hCG) with an affinity constant of 4.9×109 M

-1, the human chorionic

gonadotropin (hCG) hormone with potency 10,000 IU/mg and the colloidal solution of Au

nanoparticles with diameter 20 nm were purchased from Sigma-Aldrich (Oakville, ON). All

44

chemicals, bovine serum albumin (BSA), sodium azide (NaN3), HCl, Na2HPO4, NaH2PO4,

Polyethylene glycol (PEG), K2HPO4, and KH2PO4 were purchased from Sigma-Aldrich (Oakville,

ON). All solutions were prepared with ultra-pure water using a Cascada LS (Pall Co., NY) water

purification system at 18.2 MΩ.

DPV was performed using a µAutolab-III electrochemical analyzer (Metrohm, Switzerland)

operated in conjunction with its general-purpose electrochemistry software (GPES). The planar

SPCS electrodes are consisted of a carbon electrode with geometric working area of 2.64 mm2, a

carbon counter-electrode, and the Ag/AgCl reference electrode. . All measurements were taken

at room temperature (24 ± 1˚C).

2.2.2 Methods

2.2.2.1 Immobilization of Primary Antibody onto Working Electrode Surface

For immobilizing the primary polyclonal antibody onto the carbon electrode surface, 2 µL of

Mab-FSH solution at 100 µg/mL in 50 mM phosphate buffered saline (PBS, pH 7.4) was

dropped onto the surface. After incubation at 4 °C for 18 h, excess antibodies were rinsed with

PBS. For the suppression of non-specific adsorption, 2 µL of blocking solution (1% BSA in PBS)

was incubated on the electrode surface at 4 °C for 24 h. This incubation was performed at a

controlled temperature in order to avoid undesired denaturation of BSA during the process.

Therefore, BSA adsorbed on the uncoated parts of the working electrode surface to help prevent

the further adsorption of interfering biomolecules on these vulnerable sites. Finally, the blocking

solution was rinsed with PBS. Mab-FSH-immobilized immunosensor was stored at 4 °C until

use.

2.2.2.2 Preparation of Au Nanoparticle-Labelled hCG Antibody (Au-Mab-hCG)

For the preparation of Au-Mab-hCG, an aliquot 200 µL of Mab-hCG solution (50 µg/mL in 5

mM KH2PO4, pH 7.5) was mixed with 1.8 mL of 10% Au nanoparticle solution, and kept for 10

min at room temperature. Then, 100 µL of 1% PEG in 50 mM KH2PO4 (pH 7.5) and 200 µL of

10% BSA in 50 mM KH2PO4 (pH 9.0) were added to block the uncoated surface on Au

45

nanoparticles. After the immobilization and blocking procedures, Au nanoparticle-conjugated

Mab-hCG (Au-Mab-hCG) was collected through centrifuge (8000 g for 15 min at 4 °C). Au-

Mab-hCGs were suspended in 2 mL of the preservation solution (1% BSA, 0.05% PEG 1000,

0.1% NaN3 and 150 mM NaCl in 20 mM Tris-HCl buffer, pH 8.2), and collected again through

centrifuge as before. For the stock solution, Au-Mab-hCGs were suspended in the preservation

solution.

2.2.2.3 Immobilization of hCG and Secondary Antibody and Detection of the Antigen-Antibody Reaction

Different concentrations (between 0 and 1 ng/mL) of hCG were prepared by diluting with 1%

BSA in PBS. For the detection of the antigen and antibody reaction, 2 µL of these sample

solution were applied onto the Mab-FSH-immobilized immunosensor for 30 min at room

temperature with moderate shaking. In particular, 0 ng/mL of hCG was prepared by applying

only 2 µL of 1% BSA onto the working electrode surface. After rinsing with PBS, 2 µL of Au-

Mab-hCG solution was applied onto the electrode surface for another 30 min at room

temperature with moderate shaking, and rinsed with blank PBS.

Then, the direct redox reaction was performed using 30 µL 0.1 M HCl covering the entire three-

electrode zone at room temperature. The preoxidation of Au nanoparticles was performed at 1.2

V for 40 s, followed by differential pulse voltammetry measurement from 1 V to 0 V with a step

potential of 5 mV, pulse amplitude of 50 mV, and a pulse period of 0.5 s. The potentials were

recorded against the reference electrode (Ag/AgCl) printed within the SPCS.

46

2.3 Results and Discussion

This technique is served as a way to analyze the concentration of hCG based on current intensity.

As reported by Idegami and co-workers, the overnight incubation of BSA was sufficient for an

effective blocking and incubation with BSA for longer periods of time did not show significantly

change82. Particularly, the incubation of BSA at 4 °C was necessary, since non-specific

adsorption of the undesired biomolecules was observed under room temperature incubation82.

The reduction signal of Au nanoparticle was observed approximately at +0.65 V and the peak

current intensity increased in proportion with the increasing hCG concentration (Fig. 2.4).

Although the Au nanoparticle signal was observed to be 0.4 V under DPV, this signal shifting

could be due to the pH changes. As different layers of antibody and antigen were added on the

electrode surface, the pH of the solution, within which the redox reaction of Au nanoparticle

took place, could change. Also saturation between antigen and antibody binding reaction was

observed above 800 pg/mL hCG. This could be due to the maximum binding between antigen

and antibody was reached within limited space on the working electrode surface.

Fig. 2.4: Differential pulse voltammograms of the Au-Mab-hCG on SPCS at 20 mV/s in 0.1 M

HCl. The concentration of hCG ranged from 0 pg/mL to 1 ng/mL.

47

The analytical range and sensitivity of the immunosensor were investigated by measuring

various concentrations of hCG between 0 and 1 ng/mL (Fig. 2.5). The reduction peak current

signals depended linearly on the concentration of hCG, and the correlation coefficient was 0.987.

The detection limit was found to be 100 pg/mL. This value was lower than the previously

reported ones for the detection of hCG based on electrochemical principles99-101

. The high

sensitivity of our immunosensor was attributed to the combination of the high performance of

our SPCSs with high affinity antibodies and direct redox detection of Au nanoparticles. For the

enzyme labelled detection systems, the electrode surface is covered with the immune-complexes

and the blocking agents (BSA). These biomolecules will remain on the surface during

measurement, which will disturb the performance of the detection system. However, for our

system, the preoxidation of Au nanoparticles was at a high potential, and the denaturation of the

biomolecules was in highly acidic condition; both of were carried out simultaneously. In this

way, the detachment of the possible blocking molecules from the surface provided us a larger

electro-active area for the pre-oxidized Au ions to reduce back efficiently during the DPV scan.

In addition, the loss of oxidized Au ions through diffusion effect could be reduced by the

negative charge of the chelated compounds with the high concentration of chloride ions in the

acidic electrolyte. The constant application of highly positive voltage attracted the negatively

charged Au chelates and promoted their electrodeposition on the electrode surface.

Fig. 2.5: Corresponding relation between the peak current intensity of Au nanoparticles with

hCG concentrations. Error bars indicate the relative standard deviation of the three

measurements (n = 3) performed with three different samples.

y = 0.0005x + 0.821

R² = 0.987

0.75

0.85

0.95

1.05

1.15

1.25

1.35

1.45

100 300 500 700 900

pe

ak

curr

en

t (μ

A)

[hCG] (pg/mL)

48

2.4 Conclusion

The health of the pregnancy can be monitored by the level of hCG being produced where the

amount of hCG produced and secreted increases every day. Since current signal is linearly

correlated to the concentration of hCG, the level of hCG can be served as a tool to detect how far

along a woman is into her pregnancy. In high risk pregnancy cases, a diagnostic tool is used to

monitor the level of -hCG and if it does not rise fast enough, this may signal abnormal growth

or even abortion. This device can also be made reusable with the replacement and one time use

of these electrode strips.

This report demonstrates a technique to provide sensitive detection of the hCG hormone.

Utilizing disposable SPCS, this immunosensor technique allows for increased sensitivity while

providing the portability, use of small sample volume (2 µL), and cost effectiveness in a potential

biosensing device.

49

Chapter 3 Label Free Electrochemical Detection of Growth Hormone

(GH)

3.1 Introduction

GH is naturally produced by the somatotroph cells of the anterior pituitary gland and is released

in a pulsatile pattern. It is sometimes considered a prohormone because after its genetic code is

translated, hGH undergoes several modifications in the pituitary gland and circulation102

. The

stimuli for the secretion of hGH from the anterior pituitary gland requires the release of growth

hormone releasing factor (GHRF) from the hypothalamus.103,104 Human growth hormone has a

short half-life of about 20 min105

with concentrations returning to baseline 8-16 h after

intramuscular injection and 11-20 h after subcutaneous injection106

. GH plays an essential role in

a variety of biological processes, including somatogenesis, lactation, activation of macrophages,

and insulin-like and diabetogenic effects.107

An excess of the GH causes gigantism,108

hyperinsulinemia, impaired glucose tolerance, insulin resistance, and finally diabetes.109 A

deficiency of GH produces significantly different problems at various ages, including

hypoglycimia for newborn infants, growth failure in childhood, and a number of physical and

psychological symptoms, including poor memory, social withdrawal, and even depression for

adults.110

A number of factors are known to affect hGH release, such as aging, which can affect

hormone metabolism, diet, exercise, illness, endocrine pathologies, which can be congenital

orgenetic defects and stress, that may come from trauma or allergic reaction. Also some external

factors, such as pesticides Dichlorodiphenyltrichloroethane (DDT), Polychlorinated biphenyl

(PCB), a substance outside of the body may cause problems to our hormone circulation. Genes

that is passed from parent to child contain the instructions for the production of proteins and

some mutations or damage may also cause problems to our hormone production.

Despite its natural function, GH has also been considered as an ergogenic drug and is banned by

the world anti-doping agency (WADA)111

. Despite its ban, it has been reported to be in

widespread use in sports dating back to the late 1980s. The reasons hGH has become so popular

are because it is effective in benefiting athletic performance, has relatively mild side effects as

compared to anabolic steroids although substantial side effects accompany its use and it is

50

particularly attractive alternative to anabolic steroid use in female athletes because of the

relatively lower risk of undesired androgenic side effects104

.

In the past, a lot of studies have been done to develop a method that can effectively and

efficiently detect GH doping, such as luminescence immunoassay (detection limit: 0.05

ng/mL)100

, chemiluminescence immunoassay (detection limit < 0.05 ng/mL)101

, liquid

chromatography mass spectrometry (detection limit: 0.2 – 1 ng/mL)102

, enzyme-linked

immunosorbent assay (detection limit: 0.1 ng/mL)103, and surface plasmon resonance biosensor

(detection limit: 0.4 µg/mL)104

. However, detection of GH doping has been difficult because of

specific physiological and physicochemical properties of the hormone, including pulsatile release,

variable concentrations in normal subjects, amino acid and physicochemical similarity between

pituitary extracts (phGH) and recombinant preparations (rhGH), and short half-life in

circulation.112,113 For all mentioned reasons, an analytical procedure with rapid, quantitative,

robust, in-field capabilities for point-of-care GH detection is critical.

The detection of proteins is commonly accomplished via the antibody-based immunoassays,

where a binding event takes place between the target protein (antigen) and the primary antibody,

which is immobilized on a solid surface114. There are many immunoassay methods used in

practice such as enzyme-linked immunosorbent assay, radioimmunoassay, fluorescence or other

classical immunochemical techniques. Even atomic force microscopy methods could have been

applied to detect antigen-antibody complex formation115. Although these methods can provide

the desired sensitivity, specificity and selectivity, most of these methods are highly

disadvantageous because they imply the labelling of antigen or antibody, long analysis times and

extensive sample handling116

. Along with these disadvantages, it is quite difficult to fully

automate them. An excellent alternative to the traditional immunoassay method is

immunosensors117 that can be based on electrochemical114 or optical detection techniques118.

During the past years, the research in this field has evolved quickly with the aim of improving

the performance of the biosensors (specificity, stability, sensitivity, detection limit, etc.)119. In

this approach, recent studies have focused on the miniaturization of system by the use of

microelectrodes to develop microsensors and nanosensors.

Electrochemical impedance spectroscopy is a powerful electrochemical technique capable of

51

detecting small changes occurring at the solution-electrode surface. Accordingly, EIS has been

extensively exploited for the characterisation of materials and surface modification procedures,

and well as for the monitoring of binding events.120

In this study, we adopted the immunosensor system developed from Professor Tamiya’s

laboratory and fabricated a disposable sensor systems based on the three-electrode type of screen

printed carbon strips114

. The antibody was immobilized directly on the working electrode surface,

followed by the antigen binding. Then the voltammetric and impedance measurements were

performed. And, we also report for the first time on the electrochemical detection of GH using

the mini-potentiostat system. This equipment is compact, does not contain mobile parts, and is

easy to miniaturise, suggesting that mini-potentiostat could be easily used in-field with minimal

requirements. Furthermore, it combines rapid response, low detection limits, cost-effectiveness,

and the possibility of performing real-time monitoring of the samples, in contrast to established

strategies for GH detection.

3.2 Experimental

3.2.1 Instrument and Materials

Rat GH antibody and mouse GH antigen were kindly donated by Professor England. All

chemicals, bovine serum albumin (BSA), K2HPO4, KH2PO4, HCl were purchased from Sigma-

Aldrich (Oakville, ON). All solutions were prepared with ultra-pure water using a Cascada LS

(Pall Co., NY) water purification system at 18.2 MΩ.

CV and DPV were performed using a mini-potential stat system (BDT miniSTAT 100), which

was kindly donated by Professor Eiichi Tamiya (Osaka University, Japan) and Biodevice

Technology Ltd. (Ishikawa, Japan). Additionally, CV and electrochemical impedance

spectroscopy were performed using a µAutolab-III electrochemical analyzer (Eco Chemie,

Kanaawleg, The Netherlands) operated in conjunction with its general-purpose electrochemistry

software and frequency response analyser (GPES and FRA respectively). The planar screen-

printed carbon strip electrodes are consisted of a carbon electrode with geometric working area

of 2.64 mm2, a carbon counter-electrode, and the Ag/AgCl reference electrode (imprinted on the

electrode surface). All measurements were performed within room temperature (24±1 °C).

52

3.2.2 Methods

3.2.2.1 Immobilization of Antibody on Working Electrode Surface

For the immobilization of antibody on the carbon electrode surface, 2 µL of rat GH antibody at

50 µg/mL in 5 mM KH2PO4 (pH 7.5) was added onto the surface, followed by 18h incubation at

4 °C. Then excess antibodies were washed with 50 mM phosphate buffered saline (PBS, pH 7.4).

For the prevention of non-specific adsorption, 2 µL of 1% BSA as a blocking solution was

incubated on the working electrode surface at 4 °C for 24 h, followed by washing with PBS. This

incubation was performed at a controlled temperature to avoid undesired denaturation of BSA

during the process. Therefore, BSA adsorbed on the uncoated surface of the working electrode

surface and blocked the adsorption of unwanted target components on those vulnerable sites.

Prior to addition of antigen, the voltammetric and impedance measurements were performed to

assure the successful immobilization of Rat GH antibody on the working electrode surface.

3.2.2.2 Direct Redox-Based Detection of Antigen-Antibody Reaction

Various concentrations of mouse GH were prepared by diluting with 1% BSA in PBS. For the

detection of antigen-antibody reaction, 2 µL of these sample solutions were applied onto the

antibody-immobilized immunosensor for 30 min at room temperature with moderate shaking.

After washing with PBS, the electrochemical signals were measured from mini-potentiostat

system using CV from -0.5 V to 0.5 V with scan rate at 50 mV/s and DPV from -0.25 V to 0.25

V with a step potential of 5 mV, a pulse amplitude of 50 mV, and a pulse period of 0.5 s. We

have also collected the electrochemical signals of CV from -0.5 V to 0.9 V with scan rate at 100

mV/s and impedance measurements in the frequency range from 100 mHz to 100 kHz with

alternating voltage of 5.0 mV using our µAutolab-III system. The potentials were recorded

against the reference electrode (Ag/AgCl) and these measurements were carried out within 20 µL

of 10 mM ferri/ferrocyanide K4 [Fe (CN) 6]/K3 [Fe (CN) 6] (1:1, v/v)] solution in PBS at pH 7.

53

3.3 Results and Discussion

3.3.1 Construction of the Immunosensor and its Characterization

The electrochemical characterizations of rat GH antibody were carried out using CV and

electrochemical impedance measurements.

Fig. 3.1: The cyclic voltammograms of 10 mM K4[Fe(CN)6]/K3[Fe(CN)6] solution in 50 mM

phosphate buffer using autolab system (A) and mini-potentiostat system (B) at a bare carbon

electrode (a) and at the rat GH antibody-modified carbon electrode before (b) and after the

addition (c) of mouse GH antigen (200 pg/mL).

As shown in Fig. 3.1A-a and Fig. 3.1B-a, cyclic voltammograms on the bare carbon electrode

were collected from our autolab and mini-potentiostat systems respectively. A reversible redox

wave of 10 mM K4 [Fe(CN)6]/K3[Fe (CN)6] was observed, with oxidation and reduction peaks

around 0.25 V and -0.15 V respectively. As recorded, random physisorption is the easiest and

fastest strategy for biomolecule immobilization onto physical substrates. Additionally,

physisorption does not require biocomponent biotinylation, chemical modification, or the

utilisation of cross-linkers, and does not depend on multi-step and long experimental procedures.

Thus, the antibody was immobilized onto electrode surface through physisorption. Consequently,

the immobilization of rat GH antibody on carbon electrodes dramatically decreased the oxidation

peak from 8.75 × 10-5 A to 2.66 × 10-5 A (Fig. 3.1A-a and b). This also indicates the faradaic

54

response from this electrochemical immunosensor is getting smaller and charge transferred

towards the working electrode surface is getting smaller when different layer of antibody and

antigen are immobilized on the electrode surface. Similar results were also observed from mini-

potentiostat system (Fig. 3.1B-a and b). Since rat GH antibody contains its specific binding site

for mouse GH antigen, it will bind with each other non-covalently.16 As shown in Fig. 3.1A-b

and c, the binding of the mouse GH antigen to the antibody-modified carbon electrode

dramatically decreased the reversible redox peak from 2.66 × 10-5 A to 1.11×10

-5 A. Similar

results were also observed from mini-potentiostat system (Fig. 3.1B-b and c). The two cyclic

voltammograms obtained from autolab and mini-potentiostat systems are nearly

indistinguishable, which indicates that the two immunosensor detection systems exhibit

equivalent functionality.

Fig. 3.2: Nyquist plot of rat GH antibody + mouse GH antigen immobilized on chip with 10 mM

potassium ferri/ferrocyanide (ratio 1/1) dissolved in 50 mM PBS without NaCl buffer solutions.

Blank signal was with nothing added on the chip.

Fig. 3.2 shows the impedance spectra recorded with different interfaces. As shown in Fig. 3.2,

the immobilization of rat GH antibody on carbon electrode induced an increase in diameter of

the semicircle component of the Nyquist plot, which is consistent with the

K4[Fe(CN)6]/K3[Fe(CN)6] redox reactions from 7.73 kohm to 16.48 kohm. These results of CV

55

and electrochemical impedance measurements indicate that a highly compact rat GH antibody

has been immobilized on the carbon electrode surface and the resulting rat GH antibody layer

blocks the electrode surface, which impedes the electron transfer between the redox probes of

K4[Fe(CN)6]/K3[Fe(CN)6] and the electrode surface.

In addition, the electron transfer resistance of K4[Fe(CN)6]/K3[Fe(CN)6] redox reactions was

attenuated from 16.48 kohm to 24.78 kohm as shown in Fig. These results indicate that the

binding of the mouse GH antigen to the rat GH antibody further blocks the electrode transfer

barrier, which are also responsible for both the decreased redox current in CV and the increased

electron transfer resistance in electrochemical impedance measurements.

3.3.2 Electrochemical Analysis of Antigen-Antibody Binding

Fig. 3.3: (A) Corresponding relation between the GH concentrations (pg/mL) and the Rct values

(Kohm) from impedance spectra of rat GH antibody+mouse GH antigen immobilized on chip

with 10 mM potassium ferri/ferrocyanide (ratio 1/1). (B) Plot of the relationship between ratio of

RAb-RAb+GH/RAb and the negative logarithm value of mouse GH concentration from 10 pg/mL to

200 pg/mL to fit impedance data to Randles equivalent circuit. (n = 6, R² = 0.9953)

In order to quantify the immunosensor response, the calibration curves corresponding to the

variation of the charge transfer resistance Rct vs. concentration of mouse GH, is presented in Fig.

3.3. It shows that the charge transfer resistance is linearly correlated to the concentration of GH

ranging from 10 pg/mL to 200 pg/mL, with a low detection limit of 10 pg/mL (n = 4, R2 =

56

0.9937). In addition, the ratio between the charge transfer resistance resulting from the antibody

modified carbon electrode to antibody-antigen binding modified carbon electrode is linearly

related to the negative logarithm of mouse GH concentration ranging from 10 pg/mL to 200

pg/mL, with a low detection limit of 10 pg/mL (n = 4, R2 = 0.9953).

Fig. 3.4: Corresponding relation between the GH concentrations (pg/mL) and the peak current I

(µA) (A) and bar graphs indicating the relation between the GH concentrations (pg/mL) and the

peak current I (µA) (B) using mini-potentiostat system from differential pulse voltammograms of

rat GH antibody+mouse GH antigen immobilized on chip at 10 mV/s with 10 mM potassium

ferri/ferrocyanide (ratio 1/1).

Additionally, mini-potentiostat system was also utilized to investigate the analytical range and

sensitivity of the immunosensor. As shown in Fig.3.4, the oxidation peak current signals

depended linearly on the concentration of mouse GH ranging from 10 pg/mL to 200 pg/mL, with

a low detection limit of 10 pg/mL (n = 4, R2 = 0.9969). This was due to the fact that when

different layers of antibody and antigen were incubated on the electrode surface; the diffusion

layer thickness was increased, which induced the decrease of peak current signals.

57

Fig. 3.5: Corresponding relation between the GH concentrations (pg/mL) and the peak current I

(µA) using mini-potentiostat system (A) and bench-top potentiostat system (C); Bar graphs

indicating the relation between the GH concentrations (pg/mL) and the peak current I (µA) using

mini-potentiostat system (B) and bench-top potentiostat system (D) from cyclic voltammograms

of rat GH antibody+mouse GH antigen immobilized on chip at 50 mV/s (mini-potentiostat

system) and 100 mV/s (bench-top potentiostat system) with 10 mM potassium ferri/ferrocyanide

(ratio 1/1).

Fig. 3.5 shows that by using CV techniques from both mini-potentiostat system and bench-top

potentiostat system, the oxidation peak current signals depended linearly on the concentration of

mouse GH ranging from 10 pg/mL to 200 pg/mL, with a low detection limit of 10 pg/mL. This is

similar to the DPV results previously discussed by using our mini-potentiostat system (Fig. 3.4).

58

Fig. 3.6: Corresponding relation between the GH concentrations (pg/mL) and the peak potential

(V) using mini-potentiostat system (A) and bench-top potentiostat system (C); and corresponding

relation between the GH concentrations (pg/mL) and the change of peak potential ∆V(V) using

mini-potentiostat system (B) and bench-top potentiostat system (D) from Cyclic voltammograms

of rat GH antibody+mouse GH antigen immobilized on chip at 50 mV/s (mini-potentiostat

system) and 100 mV/s (bench-top potentiostat system) with 10 mM potassium ferri/ferrocyanide

(ratio 1/1).

As shown in Fig 3.6, peak potential increased with increasing of GH concentration. This was due

to the fact that when different layers of antibody and antigen were incubated on the electrode

surface; the diffusion layer thickness was increased, which induced the increase of peak potential

signals.

Based on the similar low detection limit, equivalent functionality is further confirmed between

our autolab system and mini-potentiostat system. It is known that bulky electrochemical

instruments should be miniaturized for future on-site measurement applications. This thus

highlights the potential of our mini-potentiostat system for the portable and decentralized

electrochemical immunosensor detection system with high specificity, stability, selectivity and

59

sensitivity, and low detection limit.

The high sensitivity of our immunosensor was attributed to the combination of the high

performance of our SPCS with high-affinity antibody-antigen binding and the direct redox

detection of ferri/ferrocyanide couple. Although the sensitivity with the present biosensing

system is impressive, sensitivity and detection limit might be further enhanced by covalently

modifying the carbon working electrode surface to increase the binding efficiency between

antibody and antigen.

3.4 Conclusion

GH has been reported to have varying degree of metabolic and immune-modulatory effects on

mammalian species. GH-deficient population are characterized by delayed puberty and impaired

fertility rather than complete reproductive function. Most importantly, recombinant hGH is

abused in sports, but adequate routine doping tests are lacking. To date, detection of exogenously

administered hGH has not been possible, since hGH is naturally synthesized by the body, it is

difficult to distinguish endogenously produced hGH from exogenously administered hGH.

Additionally, because hGH responds markedly in a pulsatile manner to stress, including nutrition,

sleep, emotion, and exercise, it is difficult to determine supraphysiologic levels, indicative of

doping. Moreover, hGH has a short half life in the blood and low concentration in the urine

being 100 to 1000 times less than in blood121. Thus simply quantifying the amount of hGH is not

sufficient to detect exogenous rhGH.

However, we have successfully fabricated a highly sensitive electrochemical immunosensor

system by using the direct detection of antibody-antigen binding on the disposable SPCS. The

required antigen sample was 2 µL. The measured dynamic range was from 10 pg/mL to 200

pg/mL and the detection limit was found to be 10 pg/mL for both autolab and mini-potentiostat

systems. Despite the fact that the mini-potentiostat system will possibly open up interesting

applications in the miniaturization of electrochemistry, our present approach provides a useful

tool for future research in the field of electrochemical immunosensor detection.

60

Chapter 4 Conclusion and Future Directions

4.1 Conclusion

Electrochemistry immunoassay offers good possibilities for sensitive detection of unlabeled

protein because it is highly sensitive, low cost, low power requirement, and has high

compatibility with advance micromachining technologies. In particular, due to its excellent

conductivity and catalytic properties, metal nanoparticle can act as “electronic wire” and

promote the communication between the redox centers in protein and electrode surfaces. The

catalytic activity of metal nanoparticles to amplify the electrochemical reactions gives them a

significantly priority in the design of electrochemical biosensors. The electrode system involves

SPCS which provide several advantages over conventional methods: 1) the small size affords the

use of small sample volumes to be used, 2) it is relatively inexpensive and readily disposable

which eliminates contamination over measurements and removes the process of regenerating the

electrode surface from irreversible oxidation processes, and 3) requires no modification or

pretreatment for measurements which allows for quick protein analysis.

Our studies demonstrate a technique to provide sensitive detection of the hCG hormone through

Au nanoparticle based sandwich type immunosensor. Utilizing disposable SPCS, this

immunosensor technique allows for increased sensitivity while providing the portability, use of

small sample volume (2 µL), and cost effectiveness in a potential biosensing device. In addition,

we have successfully fabricated a highly sensitive electrochemical immunosensor system using

the direct detection of antibody-antigen binding on the disposable SPCS. This label free

technique measured dynamic range of GH concentration from 10 pg/mL to 200 pg/mL and the

detection limit was found to be 10 pg/mL for both autolab and mini-potentiostat systems. Despite

the fact that the mini-potentiostat system will possibly open up interesting applications in the

miniaturization of electrochemistry, our present approach provides a useful tool for future

research in the field of electrochemical immunosensor detection.

61

4.2 Future Directions

Although immunoassay techniques emerged over two decades ago, there are still vigorous

research efforts and tremendous progress in the development of electrochemical immunoassays

and immunosensors. An extraordinary feature of these immunosystems is their specificity. There

are continuing studies on examining various strategies that will aid in aligning antibodies on a

solid phase in an optimal direction with minimal steric hindrance. Development in this area will

undoubtedly further enhance the degree of sensitivity achievable in analyses involving

immunoassays and immunosensors. In conjunction with electrochemical detection, these systems

will offer sensitive and selective analyses that are faster, simpler, and more economical. There is

also continuing interest in developing and applying suitable labels for electrochemical

immunoassays such that a more direct signal generation scheme can be used115

. The application

of electrochemical impedance spectroscopy has started to facilitate a label-free scheme, and this

is definitely an attractive, simpler alternative to others involving the required sensitivity and

dynamic range obtainable in amperometric detection. With the future direction in the

manufacture of miniaturized immunoassay devices, this will open up opportunities for

developing hand-held tools for instant on-site pharmaceutical and clinical diagnosis, particularly

in response to gradual shift towards home-based diagnosis125. Another challenge in this area is

the desire to integrate immunosensors in an array format to perform simultaneous analysis of

multiple analytes. Therefore, new immunosensor technologies are anticipated in the near future

in response to these exciting opportunities.

62

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