miniaturized electrochemical immunosensors for the ... · miniaturized electrochemical...
TRANSCRIPT
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
References
1. Campbell, A.M. Immunosensors. Monoclonal antibody and immunosensor technology:
the production and application of rodent and human monoclonal antibodies; Elsevier:
Amsterdam (Netherlands), 1991; pp 343-370.
2. Ju, H.; Zhang, X.; Wang, J. Nanomaterials for Immunosensors and Immunoassays.
Nanobiosensing: Principles, Development and Application; Springer: Berlin (Germany),
2011; pp 425-445.
3. Ronkainen-Matsuno, N. J.; Thomas, J. H.; Halsall, H. B.; Heineman, W. R. TrAC, Trends
Anal. Chem. 2002, 21, 213-225.
4. Wang, J. Analytical Electrochemistry, 2ed. Edition; Wiley-VCH: New York, 2000.
5. Brett, C.; Brett, A. Electrochemistry: Principles, Methods and Applications; Oxford
University Press: Oxford, 1993.
6. Brad, A. J.; Faulkner, L. R. Electrochemical Methods: Fundamentals and Applications, 2ed
Edition; John Wiley & Sons, Inc.: New York, 2001.
7. Pohanka, M.; Skladal, P. J. Appl. Biomed. 2008, 6, 57-64.
8. Osteryoung, J.; Osteryoung, R. Anal. Chem. 1985, 57, 101A.
9. Kissinger, P.; Heineman, W. J. Chem. Educ. 1983, 60, 702-706.
10. Diamond, D. Electroanalysis. 1993; 5, 795.
11. Grieshaber, D.; Mackenzie, R.; Voros, J.; Reimhult, E. Sensors, 2008, 8, 1400-1458.
12. Stoynov, Z. B.; Grafov, B. M.; Savova-Stoynova, B. S.; Elkin, V. V.; Damaskin, B. B.
Electrochemical Impedance;, Nauka: Moscow, 1991.
13. Covington, A. K. Introduction: Basic electrode type, classification, and selectivity
considerations. Ion-Selective Electrode Methodolog; CRC Press: Florida (USA), 1979; pp
15.
14. Yu, X.; Munge, B.; Patel, V.; Jensen, G.; Bhirde, A.; Gong, J. D.; Kim, S. N.; Gillespie, J.;
Gutkind, J. S.; Papadimitrakopoulos, F.; Rusling, J. F. J. Am. Chem. Soc. 2006, 128, 11199-
11205.
15. Okuno, J.; Maehashi, K.; Matsumoto, K.; Kerman, K.; Takamura, Y.; Tamiya, E.
Electrochem. Commun. 2007, 9, 13-18.
16. Khan, M.A.K.; Kerman, K.; Petryk, M.; Kraatz, H.-B. Ananl. Chem. 2008, 80, 2574-2582.
17. Malfoy, B.; Reynaud, J. A. J. Electroanal. Chem.1980, 114, 213-223.
18. Pumera, M.; Aldavert, M.; Mills, C.; Merkoci, A.; Alegret, S. Electrochimica Acta. 2005, 50,
3702-3707.
19. Kerman, K.; Saito, M.; Morita, Y.; Takamura, Y.; Ozsoz, M.; Tamiya, E. Anal. Chem. 2004,
76, 1877-1884.
20. Levent, A.; Yardim, Y.; Senturk, Z. Electrochimica Acta. 2009, 55, 190-195.
21. Meyerhoff, M. F.; Opdycke, W. N. Advances in Clinical Chemistry. 1986, 25, 1-47.
22. Stefan, R. I., Van Staden, J. F.; Aboul-Enein, H. Y. Fresenius J. Anal. Chem. 2000, 366,
659–668.
23. Luppa, P. B.; Sokoll, L. J.; Chan, D. W. Clin. Chim. Acta. 2000, 314, 1-26
24. D’Orazio, P. Clin. Chim. Acta. 2003, 334, 66-71.
63
25. Mallat, E.; Barceló, D.; Barzen, C.; Gauglitz, G.; Abuknesha, R. Trac-Trends Anal. Chem.
2001, 20, 124–132.
26. Marco, M. P.; Gee, S.; Hammock, B. D. Trends Anal. Chem. 1995, 14, 341-350.
27. Sadana, A.; Vo-Dinh, T. Biotechnol. Prog.1998, 14, 782-790.
28. Turner, A. F. P.; Wilson, G. S. Biosensors: Fundamentals and Applications; Oxford
University Press: Oxford, 1987; pp 5-8.
29. Morgan, C.L.; Newman, D.J.; Price, C.P. Clin, Chem. 1996, 42, 193-209.
30. Ferreira, A.A.P.; Colli, W.; Costac, P.I.; Yamanaka, H. Biosens. Bioelectron. 2005, 21, 175-
181.
31. Zuo, B.; Li, S.; Guo, Z.; Zhang, J.; Chen, C. Anal. Chem. 2004, 76, 3536-3540.
32. Wilson, S.M. Anal. Chem. 2005, 77. 1496-1502.
33. Ferenicik, M. Handbook of Immunochemistry; Chapman and Hall: New York, 1993.
34. Clausen, J. Immunochemical Techniques for the Identification and Estimation of
Macromolecules; North Holland: London, 1972.
35. Roit, I.; Brostoff, J.; Male, D. Immunology; Mosby International Ltd: London, 1998.
36. Stryer, L. Biochemistry; Freeman: San Francisco, 1981.
37. Köhler, G.; Milstein, C. Nature. 1975, 256, 495-497.
38. Harlow, E.; Lane, D. Antibodies: A Laboratory Manual; Cold Spring Harbor Laboratory:
Cold Spring Harbor, NY, 1988.
39. Fu, Z.F.; Hao, C.; Fei, X.Q.; et al. J. Immunol. Meth. 2006, 312, 61–67.
40. Hromadova, M.; Salmain, M.; Fischer-Durand, N.; Pospisil, L.; Jaouen, G. Langmuir. 2006,
22, 506-511.
41. Zhou, Y.M.; Wu, Z.Y.; Shen, G.L.; Yu, R.Q. Sensors and Actuators, B: Chemical. 2003, 89,
292-298.
42. Zhong, T.S.; Liu, G. Anal. Sci. 2004, 20, 537-541.
43. Ball, L.; Jones, A.; Boogaard, P.; Will, W.; Aston, P. Biomarkers. 2005, 10, 127-137.
44. Thomas, J.H.; Kim, S.K.; Hesketh, P.J.; Halsall, H.B.; Heineman, W.R. Analy. Biochem.
2004, 328, 113-122.
45. Chemburu, S.; Wilkins, E.; Abdel-Hamid, I. Biosens. Bioelectro. 2005, 21, 491-499.
46. Singh, C.; Agarwal, G.S.; Rai, G.P.; Singh, L.; Rao, V.K. Electroanalysis. 2005, 17, 2062-
2067.
47. Ionescu, R.E.; Gondran, C.; Consnier, S.; Gheber, L.A.; Marks, R.S. Talanta, 2005, 66, 15-
20.
48. Viswanathan, S.; Wu, L.C.; Huang, M.R.; Ho, J.A.A. Anal. Chem. 2006, 78, 1115-1121.
49. Zacco, E.; Pividori, M.I.; Alegret, S. Biosens. Bioelectron. 2006, 21, 1291-1301.
50. Gitlin, G.; Bayer, E.A.; Wilchek, M. J. Biochem. 1987, 242, 923-926.
51. Jones, M.L.; Kurzban, G.P. Biochemistry. 1995, 34, 11750-11756.
52. Diaz-Gonzalez, M.; Gonzalez-Garcia, M.B.; Costa-Garcia, A. Biosens. Bioelectron. 2005,
20, 2035-2043.
53. Langone, J.J. Adv. Immunol. 1982, 32, 157-252.
54. Bjoerck, L.; Kronvall, G. J. Immunol. 1984, 133, 969-974.
55. Zacco, E.; Pividori, M.I.; Liopis, X.; Del-Valle, M,; Alegret, S. J. Immunol. Methods. 2004,
286, 35-46.
64
56. Quinn, J.; Patel, P.; Fitzpatrick, B.; Manning, B.; Dillon, P.; Daly, S.; Okennedy, R.; Alcocer,
M.; Lee, H.; Morgan, M.; Lang, K. Biosens. Bioelectron. 1999, 14, 587-595.
57. Yakovleva, J.; Davidsson, R.; Bengtsson, M.; Laurell, T.; Emneus, J. Biosens. Bioelectron.
2003, 19, 21-34.
58. Ramanaviciene, A.; Ramanavicius, A. Anal. Chem. 2002, 32, 245-252.
59. Darain, F.; Park, D.S.; Park, J.S.; Chang, S.C.; Shim, Y.B. Biosens. Bioelectron. 2005, 20,
1780-1878.
60. Gooding, J.J.; Mearns, F.; Yang, W.; Liu, J. Electroanalysis. 2003, 15, 81-96.
61. Love, J.C.; Estroff, L.A.; Kriebel, J.K. Nuzzo, R.G.; Whitesides, G.M. Chem. Rev. 2005, 105,
1103-1169.
62. Chen, S.; Liu, L.; Zhou, J.; Jiang, S. Langmuir. 2003, 19, 2859-2864.
63. Akram, M.; Stuart, M.C.; Wong, D.K.Y. Anal. Chim. Acta. 2004, 504, 243-251.
64. Niemeyer, C.M. Angew. Chem. 2001, 113, 4254-4287.
65. Alivisatos, P. Nat. Biotechnol. 2004, 22, 47-52.
66. Katz, E.; Willner, I.; Wang, J. Electroanal. 2004, 16, 19-44.
67. Willner, I.; Katz, E. Angew. Chem. Int. Ed. 2004, 43, 6042-6108.
68. Rosi, N.L.; Mirkin, C.A. Chem. Rev. 2005, 105, 1547-1562.
69. Wu, L.N.; Yan, F.; Ju, H.X. J. Immunol. Meth. 2007, 322, 12–19.
70. Liu, G.; Wang, J.; Kim, J.; Jan, M.R.; Collins, G.E. Anal. Chem. 2004, 76, 7126-7130.
71. Kerman, K.; Endo, T.; Tsukamoto, M.; Chikae, M.; Yamamura, S.; Tamiya, E. Talanta,
2007, 71, 1494-1499.
72. Wang, J. Anal. Chim. Acta. 2002, 469, 63-71
73. Rashid, M.H.; Bhattacharjee, R.R.; Kotal, A.; Mandal, T.K. Langmuir. 2006, 22, 7141-7149.
74. Frens, G. Nature (Phys Sci.). 1973, 241, 20-25.
75. Ullman, A. Chem. Rev. 1996, 96, 1533-1539.
76. Petit, C.; Lixon, P.; Pileni, M. J. Phys. Chem. B. 1993, 97, 12974-12979.
77. Suslick, K.S.; Fang, M.; Hyeon, T. J. Am. Chem. Soc. 1996, 118, 11960-11967.
78. Wang, J. Analyst. 2005, 130, 421-431.
79. Liu, G.; Lin, Y. J. Nanosci. Nanotech. 2005, 5, 1060-1065.
80. Kerman, K.; Kraatz, H-B. Biosens. Bioelectron. 2009, 24, 1484-1489.
81. Pumera, M.; Castaneda, M.T.; Pividori, M.I.; Eritja, R.; Merkoci, A.; Alegret, S. Langmuir,
2005, 21, 9625-9629.
82. Chikae, M.; Fukuda, T.; Kerman, K.; Idegami, K.; Miura, Y.; Tamiya, E. Bioelectrochemistry,
2008, 74, 118-123.
83. Idegami, K.; Chikae, M.; Kerman, K.; Nagatani, N.; Yuhi, T.; Endo, T.; Tamiya, E.
Electroanalysis, 2008, 20(1), 14-21.
84. Pingarron, J.M.; Yanez-Sedeno, P.; Gonzales-Cortes, A. Electrochimi. Acta. 2008, 53,
5848-5866.
85. Hiep, H.M.; Kerman, K.; Endo, T.; Saito, M.; Tamiya, E. Anal. Chim. Acta. 2010, 661, 111-
116.
86. Shibata, K.; Katsuyama, I.; Izoe, H.; Matsui, M.; Muramatsu, H. J. Heterocycl. Chem. 1993,
277. 121-125.
87. Wang, J.; Yuan, R.; Chai, Y.; Yin, B.; Xu, Y.; Guan, S. Electroanalysis, 2008, 21, 707-714.
65
88. Vestergaard, M.; Kim, D.-K.; Kerman, K.; Hiep, H.-M.; Tamiya, E. Talanta, 2008, 74, 1038-
1042.
89. Gong, H.; Zhong, T.; Gao, L.; Li, X.; Bi, L.; Kraatz, H.-B. Anal. Chem. 2009, 81, 8639-8643.
90. Vestergaard, M.; Kerman, K.; Tamiya, E. Sensors, 2007, 7, 3442-3458.
91. Kerman, K.; Saito, M.; Tamiya, E. Anal. Bioanal. Chem. 2008, 391, 2759-2767.
92. Wang, J.; Xu, D.; Kawde, A.N.; Polsky, R. Anal. Chem. 2001, 73, 5576-5583.
93. Luo, X.; Morrin, A.; Killard, A.J.; Smyth, M.R. Electroanalysis, 2006, 18, 319-325.
94. Liang, K.Z.; Mu, W.J. Anal. Chim. Acta. 2006, 580, 128-135.
95. Li, M.; Lin, Y.C.; Su, K.C.; Wang, Y.T.; Chang, T.C. Lin, H.P. Sens. Actuators B. 2006, 117,
451-456.
96. Tang, D.Q.; Zhang, D.J.; Tang, D.Y.; Ai, H. J. Immunol. Meth. 2006, 316, 144-150.
97. Yin, T.; Wei, W.; Yang, L.; Gao, X.; Gao, Y. Sens. Actuators B. 2006, 117, 286-292.
98. Tang, D.; Yuan, R.; Chai, Y. Anal. Chim. Acta. 2006, 564, 158-164.
99. Li, Z.P.; Wang, Y.C.; Liu, C.H.; Li, Y.K. Anal. Chim. Acta. 2005, 551, 85-91.
100. Kerman, K.; Vestergaard, M. Tamiya, E. Anal. Lett. 2008, 41, 2077-2087.
101. Chetcuti, A.F.; Wong, D.K.; Stuart, M.C. Anal. Chem. 1999, 71, 4088-4095.
102. Kerman, K.; Saito, M.; Tamiya, E.; Yamamura, S.; Takamura, Y. TrAC-Trends Anal. Chem.
2008, 27, 585-592.
103. Le Roith, D.; Bondy, C.; Yakar, S.; Liu, J.L.; Butler, A. Endocr Rev. 2001, 22, 53-74.
104. Li. C.H.; Dixon, J.S. Arch Biochem Bio-phys. 1971, 233-236.
105. Doessing, S.M. Scand. J. Med. Sci. Sports. 2005, 15, 202-210.
106. Johannsson, G.; Bengtsson, B.A.; Andersson, B.; Isgaard, J.; Caidahl, K. Clin. Endocrinol.
1996, 45, 305-314.
107. Bidlingmaier, M.; Wu, Z. et al. Baillieres Best Pract Res Clin Endocrinol Metab. 2000, 14,
99-109.
108. Waite, E.; Lafont, C.; Carmignac, D.; Chauvet, N. Coutry, N.; Christian, H.; Robinson, I.;
Mollard, P.; Le Tissier, P. Endocrinology. 2010, 151, 234-243.
109. Castrique, E.; Fernandez-Fuente, M.; Le Tissier, P.; Herman, A.; Levy, A. J. Endocrinol.
2010, 205, 49-60.
110. Hashimoto, Y.; Ikeda, I.; Ikeda, M.; Takahashi, Y.; Hosaka, M.; Uchida, H.; Kono, N.; Fukui,
H.; Makino, T.; Honjo, M. J. Immunol. Methods. 1998, 221, 77.
111. Higham, C.E.; Trainer, P.J. Exp. Physiol. 2008, 93, 1157.
112. Bains, R.K.; Wells, S.E.; Flavell, D.M.; Fairhall. K.M.; Strom, M.; Le Tissier, P.; Robinson,
I.C. Endocrinology. 2004, 145, 2666-2679.
113. Mejri, M.B.; Baccar, G.; Baldrich, E.; Del Campo, F.J.; Helali, S.; Ktari, T.; Simonian, A.;
Aouni, M.; Abdelghan, A. Biosens. Bioelectron. 2010, 26, 1261-1267.
114. Steyn, F.J.; Huang, J.; Ngo, S.T.; Leong, J.W.; Tan, H.Y.; Xie, T.Y.; Parlow, A.F.; Veldhius,
J.D.; Waters, M.J.; Chen, C. Endocrinology. 2011, 152, 3165-3171.
115. Powrie, J.K.; Bassett, E.E.; Rosen, T.; Jorgensen, J.O.; Napoli, R.; Sacca, L.; Christiansen,
J.S.; Bengtsson, B.A.; Sonksen, P.H. Growth Hormone & IGF Research. 2007, 17, 220-226.
116. Nelson, A.e.; Ho, K.K. Asian J Androl. 2008, 10, 416-425.
117. Holt, R.I.; Sonksen, P.H. Br J Pharmacol. 2008, 154, 542-556.
66
118. Travas-Sejdic, J.; Peng, H.; Cooney, R.P.; Bowmaker, A.G.; Cannell, M.B.; Soeller, C. Curr.
Appl. Phys. 2006, 6, 562-566.
119. Okuno, J.; Maehashi, K.; Kerman, K.; Takamura, Y.; Matsumoto, K.; Tamiya, E. Biosens.
Bioelectron. 2007, 22, 2377-2381.
120. Kerman, K.; Nagatani, N.; Chikae, M.; Yuhi, T.; Takamura, Y.; Tamiya, E. Anal. Chem.,
2006, 78, 5612-5616.
121. Minkstimiene, A.K.; Ramanaviciene, A.; Ramanavicius, A. Analyst, 2009, 134, 2051-2057.
122. Endo, T.; Kerman, K.; Nagatani, N.; Heip, H.M.; Kim, D.-K.; Yonezawa, Y.; Nakano, K.;
Tamiya, E. Anal. Chem. 2006, 78, 6465-6475.
123. Chang, C.Y.; Silverman, D. Expert. Rev. Mol. Diagn. 2004, 4, 63-69.
124. Rauk, A. Chem. Soc. Rev. 2009, 38, 2698–2715.
125. Selkoe, D.J. Nature. 2003, 426, 900-904.
126. Shastry, B.S. J. Mol. Biol. 2003, 43, 1-7.
127. Chiti, F.; Dobson, C.M. Annu. Rev. Biochem. 2006, 75, 333-366.
128. Martin, O.S.; Serpell, L.C. Fibre Diffraction Rev. 2004, 12, 29-35.
129. Jimenez, J.L.; Guijarro, J.I.; Orlova, E.; Zurdo, J.; Dobson, C.M.; Sunde, M.; Saibil. H.R.
EMBO J. 1999, 18, 815-821.
130. Joseph, J.; Shukitt-Hale, B.; Denisova, N.A.; Martin, A.; Perry, G.; Smith, M.A. Neurobiol.
Aging. 2001, 22, 131-146.
131. Robinson, S.R.; Bishop, G.M. Neurobiol. Aging. 2002, 23, 1051-1072.
132. Lashuel, H.A.; Hartley, D.; Petre, B.M.; Walz, T.; Lansbury Jr, P.T. Nature. 2002, 418, 219.
133. Volles, M.J.; Lansbury Jr, P.T. Biochemistry. 2003, 42, 7871-7878.
134. Mirzabekov, T.A.; Lin, M.C.; Kagan, B.L. J. Biol. Chem. 1996, 271, 1988-1992.
135. Takuma, K.; Yan, S.S.; Stern, D.M.; Yamada, K. J. Pharmacol. Sci. 2005, 97, 312-316.
136. William, R.; Markesbery, M.D. Arch. Neurol. 1999, 56, 1449-1452.
137. Aliev, G.; Smith, M.A.; Sevidove, D.; Neal, M.L.; Lamb, B.T.; Nunomura, A.; Gasimov, E.K.;
Vinters, H.V.; Perry, G.; LaManna, J.C.; Friedland, R.P. Brain Pathol. 2002, 12, 21-35.
138. Klunk, W.E. Neurobiol. Aging. 1998, 19, 145-147.
139. Klunk, W.E.; Wang, Y.; Huang, G.F.;Debnath, M.L.;Holt, D.P.; Mathis, C.A. Life Sci. 2001,
69, 1471-1484.
140. Begley, D.J. Curr. Phar. Des. 2004, 10, 1295-1312.
141. Bacskai, B.J.; Hickey, G.A.; Skoch, J.; Kajdasz, S.T.; Wang, Y.; Huang, G.F.; Mathis, C.A.;
Klunk, W.E.; Hyman, B.T. Proc. Nal. Acad. Sci. 2003, 100, 12462-12467.
142. Mathis, C.A.; Bacskai, B.J.; Kajdasz, S.T.; McLellan, M.E.; Frosch, M.P.; Hyman, B.T.; Holt,
D.P.; Wang , Y.; Huang, G.F.; Debnath, M.L.; Klunk, W.E. Bioorg. Med. Lett. 2002, 12,
295-298.
143. LeVine, H. Protein Sci. 1993, 2, 404-410.
144. Nilsson, M.R. Methods. 2004, 34, 151-160.
145. Hawe, A.; Sutter, M; Jiskoot, W. Pharm. Res. 2008, 25, 1487-1499.
146. Krebs, M.R.; Bromley, E.H.; Donald, A.M. J. Struct Biol. 2005, 149, 30-37.
147. Biancalana, M.; Makabe, K.; Koide, A.; Koide, S. J. Mol. Biol. 2008, 385, 1052-1063.
148. Stsiapura, V.I.; Maskevich, A.A.; Kuzmitsky, V.A.; Turoverov, K.K.; Kuznetsova, I.M. J.
Phys. Chem. A. 2007, 111, 4829-4835.
67
149. Bolder, S.G.; Sagis, L.M.; Venema, P.; van der Linden, E. Langmuir. 2007, 23, 4144-4147.
150. Nesterov, E.E.; Skoch, J.S.; Hyman, B.T.; Klunk, W.E.; Bacskai, B.J.; Swager, T.M. Angew.
Chem. Int. 2005, 44, 5452-5456.
151. Lukevics, E. Khim. Geterotsikl. Soedin., 1995, 723 [Chem.Heterocycl. Compd., 1995 (Engl.
Transl.)].
152. Selimov, F. A.; Dzhemilev, U. M.; Ptashko, O. A. Metallokompleksnyi kataliz v sinteze
piridinovykh osnovanii [Metal_Complex Catalysis in Synthesis of Pyridine Bases]; Khimiya:
Moscow, 2003 (in Russian).
153. Lakowicz, J.R. Principles of fluorescence spectroscopy, 3rd ed.; Springer: New York, 2006.
154. Perkampus, H.-H.; Grinter, H.C.; Threfall, T.L. UV-Vis Spectroscopy and its applications,
Springer: New York, 1992.
155. Bendich, A. The Nucleic Acids, Eds.: Chargaff, e.; Davidson, J.N.; Vol. 1; Academic Press:
New York, 1955, 119-120.
156. Palecek, E. Naturwissenschaften. 1958, 45, 186-187.
157. Palecek, E. Nature. 1960, 188, 656-657.
158. Palecek, E. Biochim. Biophys. Acta. 1961, 51, 1-8.
159. Palecek, E.; Janik, B. Arch. Biochem. Biophys. 1962, 98, 527.
160. Palecek, E. Topics in Bioelectrochemistry and Bioenergetics, Ed.: Milazzo, G.; Vol. 5; John
Wiley: Chichester, 1983, 65-155.
161. Palecek, E. Electroanalysis. 1996, 8, 7-14.
162. Palecek, E. Talanta. 2002, 56, 809-819.
163. Palecek, E.; Fojta, M.; Jelen, F.; Vetterl, V. The Encyclopedia of Electrochemistry, Vol. 9,
Bioelectrochemistry, Eds.: Bard, A.J.; Stratsmann, M.; Wiley-VCH: Weinheim, 2002, 365-
429.
164. Palecek, E.; Lukasova, E.; Jelen, F.; Vojtiskova, M. Bioelectrochem. Bioenerg. 1981, 8,
497-506.
165. Palecek, E.; Postbieglova, I. J. Electroanal. Chem. 1986, 214, 359-371.
166. Palecek, E. Anal. Biochem. 1988, 170, 421-431.
167. Fojta, M.; Vetterl, V.; Tomschik, M.; Jelen, F.; Nielsen, P.; Wang, J.; Palecek, E. Biophys. J.
1972, 72, 2285-2293.
168. Wang, J.; Bollo, S.; Paz, J.L.L.; Sahlin, E.; Mukherjee, B. Anal. Chem. 1999, 71, 1910-1913.
169. Trnkova, L. Talanta. 2002, 56, 887-894.
170. Vetterl, V.; Papadopoulos, N.; Drazan, V.; Strasak, L.; Hason, S.; Dvorak, J. Electrochim.
Acta. 2000, 45, 2961-2971.
171. Dryhurst, G.; Elving, P.J. J. Electrochem. Soc. 1968, 115, 1014–1022.
172. Goyal, R.N.; Sangal, A. J. Electroanal. Chem. 2002, 521, 72–80.
173. Oliveira-Brett, A.M.; Brett, C.M.A.; Silva, L.A. Bioelectrochemistry. 2002, 56, 33–35.
174. Brett, A.M.O.; Matysik, F.M. J. Electroanal. Chem. 1997, 429,95–99.
175. Oliveira-Brett, A.M.; Piedade, J.A.P.; Silva, L.A.; Diculescu, V.C. Anal. Biochem. 2004, 332,
321–329.
176. Palecek, E.; Jelen, F.; Hung, M.; Lasovsky, J.; Bioelectrochem. Bioenerg. 1981, 8, 621-631.
177. Palecek, E.; Osteryound, J. Osteryoung, R.A. Anal. Chem. 1982, 54, 1389-1394.
178. Trnkova, L.; Friml, J. Dracka, O. Bioelectrochemistry. 2001, 54, 131-136.
68
179. Jelen, F.; Yosypchuk, B.; Kourilova, A.; Novotny, L.; Palecek, E. Anal. Chem. 2002, 74,
4788-4793.
180. Erdem, A.; Pividori, M.I.; del Valle, M.; Alegret, S. J. Electroanal. Chem, 2004, 567, 29–37.
181. Dryhurst, G. Anal. Chim. Acta. 1971, 57, 137-149.
182. Fojta, M.; Bowater, R.P.; Stankova, V.; Havran, L. Lilley, D.M.J.; Paleck, E. Biochemistry,
1998, 37, 4853-4862.
183. Fojta, M.; Teijeiro, C.; Palecek, M. Bioelectrochem. Bioenerg. 1994, 34, 69-76.
184. Oliveira-Brett, A.M.; Matysik, F.-M. J.Electroanal. Chem. 1997, 429, 95-99.
185. Oliveira-Brett, A.M.; Matysik, F.-M. Bioelectrochem. Bioenerg. 1997, 42, 111-116.
186. Bin, X.; Kraatz, H.-B. Analyst, 2009, 134, 1309-1313.
187. Halliwell, B.; Gutteridge, M.C. Free Radicals in Biology and Medicine, 3rd edition; Oxford
University Press: New York, 1999.
188. Diakowsko, P.M.; Kraatz, H.-B. Chem. Commun. 2009, 1189-1191.
189. Lyne, P.D. Drug Discovery Today, 2002, 7, 1047-1053.
190. Mahadevan, S.; Palaniandavar, M. Inorg Chim Acta. 1997, 254, 291-295.
191. Kumar, C.V.; Punzalan, E.H.A.; Tan, W.B. Tetrahedron. 2000, 56, 7027-7033.
192. Erdem, M.; Ozsoz, Turk. J. Chem. 2001, 25, 469-173.
193. Palecek, E.; Fojta, M. Anal. Chem. 1994, 66, 1566-1560.
194. Palecek, E.; Fojta, M. Anal. Chem. 2001, 73, 74A.