85
Quantitative Analysis of H5N1 DNA Hybridization on Nanowell Array Electrode Min Seok Cha The Graduate School Yonsei University Department of Biomedical Engineering
Quantitative Analysis of
H5N1 DNA Hybridization
on Nanowell Array Electrode
Min Seok Cha
The Graduate School
Yonsei University
Department of Biomedical Engineering
Quantitative Analysis of
H5N1 DNA Hybridization
on Nanowell Array Electrode
A Dissertation
Submitted to the Department of Biomedical Engineering
and the Graduate School of Yonsei University
in partial fulfillment of the
requirements for the degree of
Doctor of Philosophy
Min Seok Cha
June 2013
2
This certifies that the dissertation of Min Seok Cha is approved.
___________________________
Thesis Supervisor: Young Ro Yoon
___________________________
Thesis Committee Member #1: Tae Min Shin
___________________________
Thesis Committee Member #2: Sang Woo Lee
___________________________
Thesis Committee Member #3: Dae Sung Yoon
___________________________
Thesis Committee Member #4: Sung Oh Hwang
The Graduate School
Yonsei University
June 2013
3
ACKNOWLEDGMENTS
먼저 저를 대학원으로 이끌어 주시고, 석-박사기간 동안 학문의 아버지로서 항상 바른
길을 제시해 주시며 많은 가르침과 질책을 아끼지 않으신 윤영로 교수님께 진심으로 감사의
말씀을 드립니다. 특히 바쁜 가운데에서도 저의 논문을 심사해 주시고 조언해주신 신태민
교수님, 이상우 교수님, 윤대성 교수님, 황성오 교수님께 감사의 말씀을 드립니다. 또한
의공학과를 다니면서 많은 가르침을 주신 윤형로 교수님, 이윤선 교수님, 이경중 교수님,
김동윤 교수님, 김영호 교수님께 감사 드립니다.
대학원 생활 동안 물심 양면으로 많은 도움을 주신 생체신호처리 연구실 선배님께 감사
드립니다. 항상 선배님들의 앞서 가시는 길을 보면서 많은 도전을 받았습니다. 그리고 제가
선배라는 이유만으로 잘 따라주고 도와준 연구실 후배들 특히, 심훈, 현철, 승환에게 고마운
마음을 전합니다.
직장 가운데에서 논문을 쓸 수 있도록 지원 해주신 김관식 사장님과 큰 도전의식을
심어주신 오병도 상무님께도 감사 드립니다. 또한 논문이 진행되도록 이끌어 주신 임선희
박사님께 진심으로 감사 드리고, 먼 미국에서도 흔쾌히 도와주신 이혜연 교수님께 감사
드립니다. 같이 고민하고 실험해 준 이주경 연구원에게도 제일 고마웠다고 말해주고 싶습니다.
연구 주제에 대해서 자기 일처럼 고민해주신 이창우 박사님께 감사 드립니다. 매주 말씀으로
기도로 영혼을 깨워 준 철준 형에게도 감사의 말을 전하고 싶습니다.
마지막으로 저를 있게 해주신 아버지와 어머니께 그리고 제 인생의 동반자를 낳아주신
장인 장모님께 이 작은 결실을 드립니다. 논문을 쓰도록 가정에서 사랑과 기도로 헌신해준
사랑하는 아내 수현에게 제일 고맙고, 두 아들 현빈과 현우는 이 논문을 완성하는데 큰 힘이
되었습니다. 끝으로 본 논문을 쓸 수 있도록 완벽하게 계획하시고 이끌어 주시고, 이 많은
사람들을 만나게 해주시고 능히 해낼 능력과 지혜를 더해 주신 살아계신 하나님께 모든 영광과
감사를 올려 드립니다.
2013 년 6 월
차민석 올림
i
Table of Contents
List of Figures .......................................................................................................................... iii
List of Tables ............................................................................................................................. v
Abstract .................................................................................................................................... vi
1. Introduction ....................................................................................................................... 1
2. Basic Theory ...................................................................................................................... 4
2.1. Biosensors ............................................................................................................................................. 4
2.1.1. Biosensors in Pathogen Detection ...................................................................................... 5
2.2. Electrochemical Impedance Spectroscopy ................................................................................ 7
2.2.1. Impedance Data Representation (Nyquist, Bode) ........................................................ 8
2.2.2. Circuit Modeling and Physical Analysis ........................................................................ 10
2.2.3. Electrical Double Layer ...................................................................................................... 13
2.2.4. Charge Transfer Resistance ............................................................................................... 14
2.3. Nanowell Electrode ........................................................................................................................ 16
2.3.1. Benefits of Nanoscale Electrodes ...................................................................................... 16
2.3.2. Fabrication Methods of Nanowell Array Electrode .................................................. 18
2.4. DNA Immobilization ...................................................................................................................... 21
2.4.1. Chemisorption ........................................................................................................................ 21
2.4.2. Covalent Attachment on Functionalized Surfaces ..................................................... 22
2.4.3. Streptavidin-Biotin Interactions ....................................................................................... 23
3. Materials and Methods .................................................................................................... 24
3.1. Materials ........................................................................................................................................ 24
3.2. Methods ........................................................................................................................................ 25
3.2.1. Nanowell Array Electrode Fabrication .......................................................................... 25
3.2.2. Immobilization of Probe DNA on the Nanowell Coated with Streptavidin ....... 28
3.2.3. Physical Characterization of DNA Hybridization ...................................................... 29
3.2.4. Electrochemical Characterization of Target DNA Hybridization ........................ 31
ii
4. Results and Discussion..................................................................................................... 34
4.1. Determination of Probe DNA Concentration........................................................................ 34
4.2. Surface Morphologies and Characterization of DNA Hybridization by AFM ......... 39
4.3. Characterization of Nanowell Electrode Sensitivity ........................................................... 41
4.4. Development of Equivalent Circuit Model ............................................................................ 45
4.5. Hybridization of Probe DNA and Non-Complementary DNA ....................................... 54
4.6. Hybridization of Probe DNA and Target DNA .................................................................... 59
5. Conclusions ...................................................................................................................... 66
References and Notes .............................................................................................................. 67
국문 초록 ................................................................................................................................ 73
iii
List of Figures
Figure 1. Biosensor system ................................................................................................. 4
Figure 2. Nyquist plot ......................................................................................................... 8
Figure 3. Bode plot (Top: Magnitude plot, Bottom: Phase plot) ......................................... 9
Figure 4. Double layer capacitance and circuit modeling ................................................. 12
Figure 5. Advantage of nanowell array electrode ............................................................. 16
Figure 6. Nanowell array using nanolithography .............................................................. 18
Figure 7. Molded nanowell electrode using soft lithography ............................................ 19
Figure 8. Nanowell fabrication method using UV nanoimprinting ................................... 20
Figure 9. Schematic presentation of DNA immobilization using EDC coupling.............. 22
Figure 10. Fabrication of nanowell-based electrochemical sensor ................................... 25
Figure 11. Photograph of fabricated samples and SEM image of nanowell electrode ...... 26
Figure 12. Schematic procedure of probe/target DNA hybridization assay on nanowell
array ........................................................................................................................... 28
Figure 13. Counter, reference, working electrode and resorvoir ....................................... 31
Figure 14. Schematic design of electrode holder .............................................................. 32
Figure 15. EIS measurement in 3-electrode system .......................................................... 33
Figure 16. Nyquist plot of probe DNA immobilization on nanowell electrode ................ 36
Figure 17. Surface morphology characterization of 1 µM probe DNA immobilization on
nanowell array electrode using AFM ......................................................................... 37
Figure 18. Nyquist plot for DNA concentration on a macro electrode ............................. 38
Figure 19. Surface morphology characterization of nanowell array electrode using AFM
for bare nanowell array electrode .............................................................................. 39
Figure 20. Surface morphology characterization of nanowell array electrode using AFM
for hybridization of probe/target DNA (1 pM) .......................................................... 40
Figure 21. Characterization of nanowell electrode sensitivity by EIS .............................. 41
Figure 22. Characterization of nanowell electrode sensitivity by EIS .............................. 43
Figure 23. Randles circuit model ...................................................................................... 45
Figure 24. Nyquist plot indicating impedance of the Randles circuit ............................... 45
Figure 25. Nyquist and Bode plot of fitting results with equivalent circuit ...................... 47
Figure 26. Nyquist and Bode plot of equivalent circuit-based fitting results .................... 49
iv
Figure 27. Nyquist and Bode plot of Randles circuit-based fitting results ....................... 52
Figure 28. Nyquist and Bode plot showing Randles circuit-based fitting results ............. 53
Figure 29. Nyquist plot of probe DNA and non-complementary DNA hybridization ...... 54
Figure 30. Bode Plot of probe DNA and non-complementary DNA hybridization .......... 56
Figure 31. Nyquist plot of probe DNA and target DNA hybridization ............................. 61
Figure 32. Bode plot of probe DNA and target DNA hybridization ................................. 62
Figure 33. ΔRct based quantitative analysis of probe/target DNA hybridization. .............. 65
v
List of Tables
Table 1. Sequences of 18-mer oligonucleotide ...................................................... 24
Table 2. Elemental parameters values obtained from fitting of various equivalent
circuit models .................................................................................................. 53
Table 3. Elemental parameters values obtained from equivalent circuit fitting .... 58
Table 4. Elemental parameters values obtained from equivalent circuit fitting .... 64
vi
Abstract
Quantitative Analysis of H5N1 DNA Hybridization
on Nanowell Array Electrode
Min Seok Cha
Department of Biomedical Engineering
The Graduate School
Yonsei University
A nanowell array electrode-based electrochemical quantitative system without
amplification was developed and applied for the detection of H5N1 target DNA.
An 18-mer probe was immobilized on a nanowell array electrode with a diameter
of 500 nm, which was coated with streptavidin and a self-assembly monolayer
(SAM). The surface properties of probe DNA hybridization with complementary
target DNA were characterized using atomic force microscopy (AFM) and
electrochemical impedance spectroscopy (EIS). The AFM image shows that the
depth of nanowell was reduced from 200 nm to 15 nm due to the formation of a
DNA hybridization complex on the streptavidin/SAM structure. Differences in
charge transfer resistance (ΔRct) in EIS upon hybridization of the probe DNA with
complementary target DNA were analyzed and used for the quantitation of H5N1
vii
DNA. This approach shows that the quantitative analysis of H5N1 DNA ranging
from 1 pM to 1 µM DNA is possible on a nanowell array electrode.
.
Keywords: Nanowell array electrode; H5N1 DNA; electrochemical impedance
spectroscopy (EIS); charge transfer resistance (Rct)
1
1. Introduction
In recent years, the interest in a DNA detection system based on the
hybridization of probe and complementary target DNA using radiochemical,
electrochemical, colorimetric, and chemiluminescent methods has been growing
due to its wide applicability in the different fields of pathogen diagnostics,
forensic analysis, and environmental monitoring [1, 2].
In the field of pathogen diagnostics, rapid and accurate detection of candidate
pandemic influenza virus strains currently circulating in humans is absolutely
necessary for the surveillance and control of pandemic influenza [3, 4]. Incorrect
diagnosis of apandemic strain, as a currently circulating influenza virus, would
delay intervention and reduce the likelihood of successful containment. The
incorrect diagnosis of a candidate pandemic strain, in a case of non-pandemic
influenza, would also have serious consequences for the country involved [3].
In case of avian influenza virus (AIV) subtype H5N1, it was first discovered
in the 1990s and since then its emergence has become a likely source of a global
pandemic and economic loss [5]. Currently accepted gold standard methods of
influenza detection, viral culture and rRT-PCR, are time consuming, expensive
and require special training and laboratory facilities which is not compatible with
field setting [5]. Those that can be used in a field setting are limited to
determining the presence or absence of a target nucleic acid fragment and do not
provide quantitative information.
2
The electrochemical detection method has been used widely for the detection
of pathogen DNA owing to its simplicity, portability, sensitivity, and fast response
[2, 6-8]. The main principle of electrochemical DNA biosensors is based on the
conversion of hybridization events to analytical signals via a transducer [9]. There
are many electrochemical methods of DNA hybridization [10]. For instance, the
most easy and rapid way is the direct detection of DNA oxidation signal through
voltammetry techniques (Topkaya et al., 2010) [9]. Specific DNA hybridization
events can also be monitored through indirect DNA detection using selective
redox indicators (Jelen et al., 2002) [11], using nanomaterial (Tamiya et al., 2008)
[12] or enzyme tags (Rochelet-Dequaire et al., 2009) for signal amplification.
Furthermore, many researchers have sought a paradigm for nanobiosensors that
can be miniaturized and integrated into analysis systems in numerous
biotechnology applications [13-15].
The benefits of the use of smaller electrodes are sensitive detection and faster
mass transport due to radial diffusion, which takes place predominantly on the
electrode surface, enabling kinetic measurements in a steady-state and making
electrochemical reactions faster at high speed [14].
A newly designed nanowell array electrode for the electrochemical detection
of a H5N1 target DNA has been developed. This sensor facilitated the fabrication
of a round nanowell structure, allowing rapid access of the target DNA to the
probe for detection and enabling the system to exhibit higher sensitivities than
conventional electrodes in electrochemical measurement. EIS is known as one of
3
the most powerful tools for examining the features of surface-modified electrodes
[2, 16, 17] and has been used to detect DNA hybridization by monitoring
electrical surface properties such as conductance, resistance, and capacitance of
the electrode/electrolyte interface [18-20]. To detect hybridization of the H5N1
target DNA with the probe DNA, EIS was performed due to determine the
features of the interaction between the electrode and immobilized DNA [2, 21, 22].
In this paper, we describe the development of a H5N1 target DNA detection
system based on a nanowell array electrode using EIS technology. The 18-mer
probe was immobilized on a streptavidin/SAM coated nanowell array electrode
with a diameter of 500 nm. The surface properties of probe DNA hybridization
with the complementary target DNA were characterized using atomic force
microscopy (AFM) and EIS. To quantify the hybridization efficiency between the
probe and target DNA, difference of charge transfer resistance (ΔRct) in EIS was
analyzed and plotted as a function of the amount of DNA. This approach
demonstrates that the quantitative analysis of H5N1 DNA at concentrations
ranging from 1 pM to 1 µM is possible on nanowell array electrodes.
4
2. Basic Theory
2.1. Biosensors
Biosensors are chemical sensors which converts a biological response into
electrical signals. A biosensor consists of two components:
The receptor is composes of biological recognition element such as enzyme,
antibody, cell, nucleic acid. The recognition element is immobilized on a support
material or a biointerface [23].
Figure 1. Biosensor system
The detector or the transducer (electrochemical, piezoelectric and optical)
which translates the biochemical signal produced by the immunological reaction
(antigen combines with antibody) into electrical signals [24]. Transducer,
electrical interface (electrode) and bioreaction are the key part of a biosensor.
Biosensors work when the antibody and antigen interact with each other
(biomolecular interaction). This causes a physical or chemical change at the
5
biointerface which is converted by the transducer to an electrical signal. Output
from the transducer is then amplified, processed and finally displayed as a
measurable digital signal [25].
2.1.1. Biosensors in Pathogen Detection
The most popular methods are, by far, those based on the polymerase
chain reaction, PCR (Bej et al., 1991). This can be explained on the grounds of
selectivity and reliability. Several types of biosensors, such as surface plasmon
resonance (SPR), quartz crystal microbalance (QCM) [26] and optical
interferometric, have been researched as alternatives to conventional detection
methods for avian influenza virus. For fluorescence detection, antibodies or DNA
may be conjugated to fluorescent compounds, the most common of which is
fluoresce in isothiocyanate (FITC) (Liet al., 2004). However, their use is very
restricted due to safety reasons. SPR biosensors (Cooper, 2003) measure changes
in refractive index caused by structural alterations in the vicinity of a thin film
metal surface. Adsorption phenomena and even antigen–antibody reaction
kinetics can be monitored using this sensitive technique. The main drawbacks are
complexity system and high cost of equipment and large size. Piezoelectric
sensors are based in the observation of resonance frequency changes on a quartz
crystal microbalance (QCM) following mass changes on the probe/transducer
surface (O’sullivan and Guilbault, 1999). It is more sensitive, reproducible and
reliable than traditional flow-through methods it is not as suitable for automation
6
[27]. Many sensing assays and detection methods are not practical, sufficiently
rapid, inexpensive, simple or robust for use in the field [28].
7
2.2. Electrochemical Impedance Spectroscopy
EIS possesses the ability to study any intrinsic material property or specific
processes that could influence the conductivity/resistivity or capacitivity of an
electrochemical system [29]. Therefore, impedance techniques are useful to
development a material of biosensor and monitor changes in electrical properties
arising from biorecognition events at the surfaces of modified electrodes. ElS is
an effective technique to probe the interfacial properties (capacitance, electron
transfer resistance) of modified electrodes [13], providing a rapid and very
sensitive label-free detection of affinity interactions of biomolecules. The change
of the capacitance and the charge transfer resistance can be measured as a result of
protein immobilization and antibody-antigen reactions on the electrode surfaces
[2]. It can also detect DNA hybridization events between probe and target
sequences [9]. In an experiment reported by Hang et al. a microarray
configuration of interdigitated electrodes was used to detecting the binding states
of DNA [30].
EIS measures impedance in status of connecting a three electrode system in
order to find out of characteristic of interface. This is done by applying a small
AC signal over a range of frequencies at a specified DC potential. Varying the
frequency changes the relative contribution of each elements in the equivalent
circuit to the overall impedance [31]. For that reason EIS method can minimize
the changes in physical properties of the measurement system by the process of
electrochemical measurement approach compared to analytical methods other
8
electrochemical applied to the measurement.
2.2.1. Impedance Data Representation (Nyquist, Bode)
After measuring an impedance of electrode chemical systems, effectively
plotting an impedace data is a very important. Because if it can be interpret the
data differently depending on whether you select a chart of any method. The
expression for impedance Z (jω)
(1)
If the real part is plotted on the X-axis and the imaginary part is plotted
on the Y-axis of a chart, we call a "Nyquist Plot" as shown in Figure 2. Nyquist
plotFigure 2.
Figure 2. Nyquist plot
9
Another popular presentation method is the Bode plot in Figure 3. The
Bode plot has some distinct advantages over the Nyquist plot. Since frequency
appears as X-axis, it is easy to understand from the plot how the impedance varies
on the frequency. The impedance is plotted as log scale of frequency on the X axis
and both the absolute values of the impedance (|Z|) and the phase-degree is plotted
on the Y-axis. It is usually a combination of a Bode magnitude plot and a Bode
phase plot.
Figure 3. Bode plot (Top: Magnitude plot, Bottom: Phase plot)
10
2.2.2. Circuit Modeling and Physical Analysis
In order to observe the change of the electrode surface as described above,
it is a normal way to interpreted and analyze the electrode in the experiments by
measuring the impedance of the material, and displaying Nyquist plot or Bode
plot. However, it is more easy to understand surface adsorption process or
physicochemical phenomena occurring at the electrode surface by equivalent
circuit modeling method compare to only displaying plot of impedance occurring
at the electrode surface. It also can be shown a quantitation value of the
parameters in detail the behavior of the surface and analyze changes.
The following Figure 4, was visualized a Randles circuit model which is
commonly used in electrochemical systems, the major symptoms of the electrode
to conform in working electrode. Z ’ and Z’’ can be calculated given an equivalent
circuit model for the system using traditional algebraic methods of circuit analysis
[32].
To illustrate an example, overvoltage at working electrode is applied in
electrochemical experiments with a three-electrode system. On the counter
electrode positive voltage is applied with redox ions in solution, the redox ion
move according to the mobility of the ions in the electric field. The mobility of
ions will be modeled as the resistance of the solution in the equivalent circuit.
When moved redox ions reach to the working electrode surface, the redox
ion cause an electrical double layer formed between the cations and electrons on
the surface of the working electrode, since the electrochemical reaction was not
11
occurred, it modeled by the capacitance components in the equivalent circuit. This
phenomenon is occurred continuously after on the surface of the movement of
ions so the circuit is connected in series with solution resistance. When redox ions
arrived at the working electrode interface, it occurred an electrochemical reaction
to get an electron by applied redox potential, capacitance component in the
working electrode will be change to a resistance formation due to electron flow.
This reaction was independent reactions from the double later capacitance. Hence
resistance of electron transfer was connected in parallel with capacitor, Because of
the transferred electrons, concentration of ions in the electrode surface and
solution concentration becomes different. It results in diffusion phenomena. It is
modeled the resistance of diffusion so called “Warburg impedance” and which, so
the electron transfer reaction and subsequently circuit connected to the serial
configuration. As explained above electrochemical phenomena can be analyzed
with electronic parameter; in equivalent circuit.
12
Figure 4. Double layer capacitance and circuit modeling
The components Rs and W in the electronic circuit represent bulk
properties of the electrolyte solution and diffusion features of the redox probe in
solution. The other two components in the circuit, Cdl and Rct, represent interfacial
properties of bioorganic materials onto the electrode surface. Thus, analysis of Cdl
and Rct could give important information about the extent of changes of the
surface properties resulting from coupling of biomaterials to the studied interface.
13
2.2.3. Electrical Double Layer
In case of applying redox potential to the electrode, negative charge and
positive charge are strongly coupled and induce the interface charge separation by
electrostatic attraction at the interface of the electrode, it is somewhat relaxed the
extent of binding by the thermal energy. Therefore, the charge electrode and the
solution will be arranged in a predetermined thickness on both sides around the
interface. 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). Figure 4(a) schematic representation of the electrical
double layer. (Figure drawn by adaptation from reference [33]). 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.
The charge transfer occurs only near the surface of the electrode in the
electrochemical reaction. Therefore, it is possible to affect the charge transfer rate
depending on the structure of the electric double layer near the surface of the
electrode.
14
2.2.4. Charge Transfer Reaction
The rate of heterogeneous charge transfer reaction
O + ne = R (2)
is given by the expression
(3)
where is the faradic current density, kf,b are the forward and reverse rate
constants, and CO,R are the concentrations of the reactants and product at the
interface at time [34].
A similar resistance is formed by a single kinetically controlled
electrochemical reaction. In the forward reaction in the first equation, electrons
enter the metal and metal ions diffuse into the electrolyte. Charge is being
transferred. This charge transfer reaction has a certain speed. The speed depends
on the kind of reaction, the temperature, the concentration of the reaction products
and the potential [35]. The general relation between the potential and the current
(which is directly related with the amount of electrons and so the charge transfer
via Faradays law) is:
(4)
15
Where i0 = exchange current density, CO = concentration of oxidant at the
electrode surface, CO*
= concentration of oxidant in the bulk, CR =
concentration of reductant at the electrode surface, η = over potential (Eapp –
Eoc) F = Faradays constant, T = temperature, R = gas constant, α = reaction
order, n = number of electrons involved
When the concentration in the bulk is the same as at the electrode surface,
CO=CO* and CR=CR*. This simplifies equation below
(5)
This equation is called the Butler-Volmer equation. It is applicable when
the polarization depends only on the charge-transfer kinetics. Stirring the solution
to minimize the diffusion layer thickness can help minimize concentration
polarization. When the overpotential, η, is very small and the electrochemical
system is at equilibrium, the expression for the charge transfer resistance changes
to:
(6)
From this equation the exchange current density can be calculated when Rct is
known.
16
2.3. Nanowell Electrode
2.3.1. Benefits of Nanoscale Electrodes
The past three decades have seen tremendous growth and increased
application of nanoelectrodes in fundamental electrochemistry, electrochemical
analysis, electrocatalysis, and many other research areas [10]. Micrometer size or
smaller electrode provide a lot of benifits in electrochemical studies and
applications. These benifits are more extended when electrode size is more
smaller upto nano scale [24, 36]. The major benifit of nanosale electrode is that it
can be obtained enhanced mass transfer which takes place. As shown in Figure 5,
3-dimensional diffusion becomes dominant and results in faster mass transport as
electrodes size is decreased.
Figure 5. Advantage of nanowell array electrode
Because of fast mass transfer, nanoscale electrode can be able to kinetic
measurement in the steady-state without diffusion limit. This is because the
electron transfer process is less likely to be limited by the mass transport of
reactant to the electrode surface at very high rates of mass transport [14].
17
Morf and de Rooij et al. (2006) presented theoretical calculations for the
current output of arrays of different packing densities using steady-state and
chronoamperometric responses [37]. Densely packed arrays were considered to
have an inter electrode distance of 2 times the radius. In the analysis, an array of
loosely packed electrode yielded a near ideal multiple response of a single ultra-
microelectrode and closely packed array more closely resembled the behavior
from a macro electrode of a similar surface area [24]. The experimental behavior
starts to deviate from extrapolations of behavior at larger electrodes. This point
may be viewed as the separation point between nanoscale electrodes and
microelectrodes [36].
Another benefit of nanoscale electrode is lower solution resistance [10].
Decreased charging currents and decreased deleterious effects of solution
resistance, can all be expected with nano-electrodes and enable new applications
of ultra-small electrochemical systems [38].
18
2.3.2. Fabrication Methods of Nanowell Array Electrode
Numerous methods have been employed to produce nanoelectrodes of
various shapes. Representative methods include micropipette pulling technology,
partial insulation of an electrochemically sharpenedmetal wire or carbon fiber in
photoresist, Teflon, electrophoretic paint, and glass [10]. Despite enormous
progress, the controllable fabrication of structurally well-defined nanoelectrodes
and their characterization at fast diffusion remain challenging for their successful
application [10]. One of the greatest challenges facing nanoelectrode researchers
is the preparation and fabrication of devices in order to study and realize some of
the benefits discussed above. There have been reported five main nanowell
fabricated methods which is using electron beam lithography, focused ion beam
lithography (FIB), dip-pen nanolithography (DPN), soft lithography (mCP,
molding), nanoimprint lithography (NIL).
Figure 6. Nanowell array using nanolithography
Lee et al. (2006) published nanowell arrays composed of 200 200
19
reaction zones made of gold electrodes using nanolithography technology with a
diameter of 50~200 nm, height of 150~200 nm, and interspacing of 300 nm as
shown in Figure 6 [39]. This nanowell array system was applied to detect target
DNA using electrochemical detection technologies such as square wave
voltammetry (SWV) and cyclic voltammetry (CV), allowing only one or a few
biomolecules to enter and attach to nanosized gold dots.
Kim P. et al. (2008) reported soft lithography fabrication using a
polyurethane acrylate (PUA) mold and detected a biological reactions on the level
of single lipid vesicle as shown in Figure 7 [40].
Figure 7. Molded nanowell electrode using soft lithography
Lee et al. (2009) also reported nanowell arrays using the UV
nanoimprinting method and applied this to microfluidic chips, enabling the use of
20
a localized surface plasmon resonance system as shown in Figure 8 [41].
Figure 8. Nanowell fabrication method using UV nanoimprinting
However, these nanowell fabrication process resulted in low throughput [39,
42], non-uniformity [40, 43] and high costs [41]. Therefore, further improvement
in the structure of nanowells and facilitation of the fabrication process was
required to solve these limitations from previous studies [39-43]. In this paper,
nanowell array electrode has been fabricated using a KrF laser so as to be possible
uniform construction and mass production.
21
2.4. DNA Immobilization
2.4.1. Chemisorption
DNA immobilization by covalent attachment is often used [44, 45]. Thiol-
metal interactions are frequently used to bind biomolecules covalently onto gold
surfaces. The strong affinity of the thiol groups for noble metal surfaces enables
the formation of covalent bonds between the sulfur and gold atoms.
R-SH + Au R-S-Au +e-+H
+ (7)
On the basis of this principle (chemisorption), biosensors have been
developed using thiol-modified DNA probes [46]. In the same way, DNA probes
were immobilized onto gold interdigitated electrode arrays by self-assembly of
thiol-modified ODNs [47]. DNA strands also were attached to gold micro pads
deposited on a silicon surface [46]. The 11-MUA SAM onto Au electrode is
highly passivated and stable in a laboratory environment. However, during
impedance measurements, this interface did not reach steady state until about 30
min after immersion into 50 mM PBS and 5 mM K3Fe(CN)6 during all
impedance studies, the 11-MUA covered electrode was allowed to stabilize for 30
min.[48] For this work, 11-MUA was used as SAM to passivate the gold electrode
and immobilize the streptavidin.
22
2.4.2. Covalent Attachment on Functionalized Surfaces
Covalent reactions often use carbodiimide as a reagent, with or without N-
hydroxysuccinimide (NHS), 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide
(EDC) is the most frequently used activation coupling reagent [1].
Figure 9. Schematic presentation of DNA immobilization using EDC coupling
For example, self-assembled carbon nanotube (CNT) layers were formed
on gold substrates [49]. Carboxylic acid groups were introduced to CNTs that
formed covalent bonds with amino groups at the 5’ ends of DNA probes in the
presence of EDC. Different covalent immobilization techniques also were tested
[50].
For this study, the above-mentioned covalent bonding reaction was used.
SAM binds with streptavidin and coated on the nanowell gold electrode using the
EDC and NHS as an activation coupling reagent.
23
2.4.3. Streptavidin-Biotin Interactions
The formation of streptavidin-biotin complexes is useful in a variety of
applications [51]. This specific binding is largely used to immobilize enzymes,
antibodies, or DNA. Biotin is a small molecule that binds with a very high affinity
to the streptavidin binding sites (Ka = 1015
M-1
). Moreover, streptavidin is
tetrameric proteins that have four identical binding sites for biotin. Streptavidin
with an isoelectric point (pI) equal to 5 is thus preferably used over avidin, which
has a pI of 10.5, to avoid nonspecific interactions. The avidin (or streptavidin)-
biotin interaction is often used to develop DNA biosensors. For this study, biotin-
labled DNA bound to streptavidin through streptavidin-biotin interactions [52].
24
3. Materials and Methods
3.1. Materials
A biotinylated 18-mer DNA oligonucleotide, 5-biotin-ATG GAG AAA
ATA GTG CTT-3, was used as the probe DNA. The sequences 5-AAG CAC
TAT TTT CTC CAT-3 and 5-ATG GAG AAA ATA GTG CTT-3 were used as
the target DNA and non-complementary DNA, respectively. All DNA synthesis
reagents were obtained from Bioneer Co. (Daejeon, Korea). Streptavidin,
ethanolamine hydrochloride, 11-mercapto undecanoic acid (11-MUA), 1-ethyl-3-
(3-dimethylaminopropyl) carbodiimide (EDC), N-hydroxysuccinimide (NHS)
and all other chemicals of analytical grade were purchased from Sigma-Aldrich
(St. Louis, MO, USA).
Table 1. Sequences of 18-mer oligonucleotide
Name Sequence
Probe 5-biotin-ATG GAG AAA ATA GTG CTT-3
Target 5-AAG CAC TAT TTT CTC CAT-3
Non-complementary 5-ATG GAG AAA ATA GTG CTT-3
25
3.2. Methods
3.2.1. Nanowell Array Electrode Fabrication
A 6-inch Si wafer was pre-cleaned using piranha solution (H2SO4:H2O2 =
1:4) and hydrogen fluoride (HF, 1%, wt/vol) to remove the organic contaminants
and native oxide layer on the Si substrate. A silicon dioxide (SiO2) layer of 300
nm thickness was deposited on the Si substrate by plasma-enhanced chemical
vapor deposition (PECVD, Plasmalab 80 Plus, Oxford, UK). Additionally, a
titanium (Ti) layer of 30 nm thickness as an adhesion layer and a gold (Au) layer
of 300 nm thickness as a bottom electrode were deposited by sputtering (SRN-120,
Sorona, Korea). After micro-patterning, the Ti/Au layers were etched by an
inductively coupled plasma (ICP) etcher (MultiplexICP, STS, UK), and the
photoresist (PR) was removed by a microwave asher (Enviro II, ULVAC, USA).
Figure 10. Fabrication of nanowell-based electrochemical sensor
26
A SiO2 layer of 200 nm thickness was deposited again onto an Au electrode
by PECVD to define the nanowell array patterns. SiO2 was used for 2 main
purposes: to define the array pattern and to act as an insulating layer. A KrF
stepper (PAS 5500/300C, ASML, Veldhoven, The Netherlands) was used for
nanoscale patterning after PR coating. The exposed SiO2 areas were finally etched
using an ICP etcher (ICP380, Oxford, UK), and the PR was removed.
Figure 11. Photograph of fabricated samples (Top) and SEM image of
nanowell electrode (Bottom)
Fifty-seven very uniform and well-fabricated chips were obtained in a 6-
inch wafer. The size of a single chip was 21 × 10 mm2, and each chip consisted of
27
a couple of nanowell areas of size 4 × 2 mm2
in Figure 11.
The morphology of the fabricated nanowell array electrode were observed
using field emission scanning electron microscopy (FE-SEM, S-4800, Hitachi,
Japan) with an acceleration voltage of 15 kV (Figure 11).
28
3.2.2. Immobilization of Probe DNA on the Nanowell Coated with
Streptavidin
For immobilization of probe DNA, all electrodes were treated with acetone
and cleaned with ethanol and deionized (DI) water. After N2 blowing for
cleanning a nanwell array electrode, SAM was prepared on a nanowell array
electrode by incubating 10 mM of 11-MUA for 1 hour at room temperature. Then,
50 mM EDC and 50 mM NHS in 1 mL sodium acetate buffer (pH 5.5) were added
to form active ester functional groups. Streptavidin in phosphate-buffered saline
(PBS; 10 μg/mL) was immobilized on SAM for 30 min at room temperature. The
unreacted functional ester groups were blocked by treating with 1 M ethanolamine
for 30 min. Figure 12 shows a schematic procedure for the hybridization of the
probe with the target DNA on the nanowell array electrode.
Figure 12. Schematic procedure of probe/target DNA hybridization assay on
nanowell array
Step1: Streptavidin attachment to carboxylate group of 11-MUA after
EDC/NHS treatment.
Step2: Immobilization of probe DNA on streptavidin/SAM structure
Step3: Hybridization of probe and complementary target DNA
29
For immobilization, 10 nM of biotinylated probe DNA was added onto the
streptavidin/SAM and incubated at room temperature for 30 min to allow the
streptavidin-biotin reaction to occur. After the immobilization procedure of probe
DNA, the electrodes were rinsed 3 times with PBS buffer and N2 blowing for
further treatment. The target DNA at concentrations ranging from 1 pM to 1 µM
was hybridized with the probe on the nanowell array for 15 min by applying 50
μL DNA.
3.2.3. Physical Characterization of DNA Hybridization
To understand and characterize hybridized DNA structure, AFM was used.
In biological applications, the most appealing advantage of the AFM as a high-
resolution microscope in comparison with other techniques such as SEM and
TEM, is that it allows measurements of native biological samples in
physiological-like conditions, avoiding complex sample preparation procedures
and artifacts connected to them [53].
The AFM is a member of the family of scanning probe microscopes (SPM).
A probe is scanned in close proximity to a surface in order to gather information
on various types of properties. The AFM uses a sharp tip, which is mounted on a
cantilever, that will deflect due to forces acting on it and this deflection can be
measured using a laser reflected from the cantilever on a deflection detector [54].
The prepared hybridized DNA was analyzed by AFM using Autoprobe (XE-150,
Park Scientific Instruments, Suwon, Korea), equipped with Proscan software and
30
a tactile profilometer with a curvature radius of 350 μm and a resolution of 10 nm.
An aluminum-coated silicon AFM cantilever with an oscillation frequency of 150
kHz and a spring constant of 4.5 N/m was used for AFM imaging.
31
3.2.4. Electrochemical Characterization of Target DNA
Hybridization
After nanowell array electrode fabrication for H5N1 DNA detection, EIS
was used to monitor probe and target DNA hybridization using an Ivium Stat
(Ivium Technology, Eindhoven, The Netherlands) at room temperature. After
probe and target DNA hybridization as described in Figure 12, electrochemical
measurement was performed using a conventional 3-electrode system.
Figure 13. Counter, reference, working electrode and resorvoir
A glass Ag/AgCl eletrode with a diameter of 6 mm and a length of 5 cm
was used as the reference electrode and a Pt coil electrode from BASi analytical
instruments (West Lafayette, IN, USA) was used as the counter electrode (Figure
13). All electrode was integrated into the electrode holder. The position of the
three-electrode system should be fixed accurately for electrochemical
measurements. Otherwise, the voltage applied to the electrodes of nanowell array
32
varies depending on the distance from reference electrode, and the generated
current in the nanowell electrode also varies depending on the electrode location.
Thus a three electrode holder fixed position between the three electrodes was
designed. The distance between the working electrode and the reference electrode
should be designed closer in order to reduce the interference of the solution
resistance. Detailed schematic design is shown in Figure 14(a). Designed
electrode holder with a distance of 13 mm between the working electrode and the
reference electrode and the distance between counter electrode and reference
electrode is 26 mm. Top view of the combined electrode holder was shown in
Figure 14(b).
(a) (b)
Figure 14. Schematic design of electrode holder
EIS test was performed in ferri/ferrocynide solution which is contained in
the reservoir. The reservoir volume made of polycarbonate is 1200 µL and height
is 40 mm.
33
Figure 15. EIS measurement in 3-electrode system
The impedance of the DNA on immobilized electrode surfaces was
determined for redox ions using ferri/ferrocyanide (5 mM each) in PBS buffer
(Figure 15). EIS was recorded from 1 MHz to 0.1 Hz, using a modulation voltage
of 50 mV AC signal amplitudes. At the end of each test, the nanowell array
electrode was rinsed 3 times with PBS buffer and N2 blowing for further treatment.
The Z-view modeling program (Scriber and Associates, Charlottesville, VA, USA)
was used to fit the curve and the parameters of the measured impedance spectra
was extracted in order to investigate sensitivity of nanowell electrode.
34
4. Results and Discussion
4.1. Determination of Probe DNA Concentration
A semicircle diameter is different and dependent on a electrode interface
condition such as probe DNA immobilization, target DNA hybridization. Katz et
al, and Suni et al reported that proteins, antibody and DNA can be most
sensitively detected through the increase in the charge transfer resistance (Rct)
with increasing DNA concentration, since this causes an increase in the layer
thickness on the electrode, reducing the rate of electron transfer [17, 48].
We performed EIS for determination of the probe DNA optimization. It is
nessasary to know the proper probe DNA concentration for sensitive H5N1
detection sensor. If probe DNA concentration is too low then, it is very sensitive ,
However it might be hard to detect hybridization at higher tareget DNA
concentration. Lim at el. (2009) reported that the oligonucleotide density on the
surface might influence the hybridization efficiency; on one hand a very low
number of oligonucleotide reduces the number of potential binding sites, and on
the other a very high density might lead to sterical hindrance [55]. As shown in
Figure 16, we performed EIS in order to optimize assay for determinaton of probe
DNA concentration. First, probe DNA is immobilized at different concentrations
respectively, EIS is perfomed in a different probe DNA concentration levels. A
total 4 of probe is analyzed for determining the probe concentration. All results
have a same solution resistance at high frequency region shown in Figure 16.
35
Diffusion line from 1 pM to 10 nM at the low frequecny is observed. However,
diffusion reaction line for 1 µM of probe DNA was disappeared. As probe DNA
concentration is increased, semicircle diameter is increased because the probe
DNA inhibits charge transfer between ferricyanide and the nanowell electrode.
For 10 pM probe DNA, the low diffusion region in the Nyquist plot of impedance
sprectrum was shown. 45 º line represent diffusion in the low frequency range. In
contrast, 1 µM of probe DNA didn’t show the diffusion properties at low
frequency. Diffusion line is not displayed when 1 µM of probe DNA was
immobilized and it probably can be explained; Probe DNA are immobilized too
densely on the electrode surface and surface is fulled with negative charged DNA
backbone. Therefore, there is no movement of the redox ion electrode surface into
negative charged surface and diffusion transport toward nanoelectrode can not be
happened due to non-increased concentration difference of ferrocyanide.
Therefore, the 1 µM of probe DNA is not used to detect target DNA. The amount
of probe DNA was determined by using diffusion characteristic at the low
frequency. A main strength of nanowell electrode is fast mass transfer.
36
Figure 16. Nyquist plot of probe DNA immobilization on nanowell electrode
A flat structure electrode at 1 µM of probe DNA immobilization was
characterized by AFM. Figure 17 shows a surface morphology of 1 µM of probe
DNA. The analyzed line profile data on the right reveals that the 200 nm of
nanowell depth size was not obseved. It was caused by immobilization of
streptavidin/SAM structure and the formation of a high concentration probe DNA
in the nanowell structure. We confirmed that the different charge transfer kinetics
depends on the concentration of probe DNA-immobilized in nanowell electrodes.
Charge transfer is fairly rapid on the lower concentration of probe DNA (Figure
16). 10 nM of probe DNA was decided and used to maximize the hybridization of
target DNA concentration.
37
Figure 17. Surface morphology characterization of 1 µM probe DNA
immobilization on nanowell array electrode using AFM
Additionlal experiment is performed in the macro electorde as the same
way in order to understand the characteristics of the electrode. The Figure 18
shows change in impedance due to the probe DNA on the macro electrodes.
Significantly large semicircles are observed from 10 nM of probe DNA. The
charge transfer rate decreases by (Rct) charge transfer resistance increases when it
was analyzed by Randles circuit model. Ferrocyanide ions is not easy to move
into the macro electrode structure compared to nanowell electrode. This large
background semicircle is disadvantage for developing sensitive DNA biosensors.
DNA sensitivity in nanowell improved significantly compared to that in macro
38
electrode. Probe and target DNA hybridization on the nanowell array was studied
by EIS using ferri/ferrocyanide as a probe anions in the bulk solution.
Figure 18. Nyquist plot for DNA concentration on a macro electrode
39
4.2. Surface Morphologies and Characterization of DNA
Hybridization by AFM
Figure 19 shows the formation of a 40 40 nanowell grid with 500 nm
diameter and 200 nm depth on a nanowell array electrode. In the circular dark area
is the a gold electrode inside nanowell structure, the bright part is a SiO2 resist
layer which is formed outside of nanowell structure.
Figure 19. Surface morphology characterization of nanowell array electrode
using AFM for bare nanowell array electrode
Figure 20 shows an AFM image for the hybridization of target DNA (1
pM) on the streptavidin structure immobilized on SAM. In contrast to Figure 19, a
bright spot is seen in the nanowell due to the strepavidin/SAM-coated layer and
the hybridization of target DNA. There are also bright spots on the outside of the
nanowell in Figure 20; these might be caused by adsorption of non-specific biding.
However, these do not contribute to the electrochemical signals due to the fact that
they are adsorbed on the SiO2 resist layer not a gold electrode area.
40
Figure 20. Surface morphology characterization of nanowell array electrode
using AFM for hybridization of probe/target DNA (1 pM)
The analyzed line profile data in Figure 20 reveals that the nanowell
depth size was reduced from 200 nm to 15 nm compared to the depth size of bare
nanowell array electrode. It caused by immobilization of streptavidin on SAM
structure and due to the formation of DNA hybridization complex.
These two AFM images and line profile show that the nanowell array
pattern was well oriented on the SiO2 surface and that the formation of the
hybridization complex of probe and target DNA occurred satisfactorily.
41
4.3. Characterization of Nanowell Electrode Sensitivity
The electron transfer of the redox couple [Fe(CN)6]3−/4−
through nanowell
electrodes can be used as an effective indicator for electrode surface
characteristics. Therefore, each surface modification step was monitored. EIS was
performed in order to compare the characteristics of the macro electrode and
nanowell array electrode.
Figure 21. Characterization of nanowell electrode sensitivity by EIS
Nyquist plot shows: (a) macro electrode, (b) probe DNA on nanowell (c)
hybridization of 100 nM of target DNA
Figure 21 shows Nyquist plots of the three impedance spectrum at
different stages. (a) is status of probe DNA immobilization on a macro electrode
without nanowell structure. Immobilization of probe DNA exhibited a large
42
semicircle in macro electrode structure not having any nanowell structure (Figure.
21(a)). The Nyquist plot for the macro gold electrode does not show a linear part
at low frequency despite the same area as the nanowell. It is confirmed that after
the immobilization of streptavidin and probe DNA on the macro electrode, macro
electrode system was changed from diffusion-controlled system to very slow
kinetic-controlled system due to the effective blocking ability of probe DNA and
streptavidin/SAM structure, preventing the redox reaction of Fe(CN)63−/4−
[33].
Figure 21(b) shows the impedance spectroscopy of immobilization of probe DNA
on nanowell electrode. The semicircular Nyquist impedance spectra are observed
and the size is small compared to the macro electrode (Figure. 21(a)). It indicates
the charge-transfer resistance at the electrode/electrolyte interface decrease due to
nanowell electrode area, showing fast diffusion characteristic. This means
nanowell electrode is faster kinetic-controlled system and more sensitive system
compared to macro electrode. 100 nM of target DNA is hybridized on the
nanowell array electrode and the diameter of semicircle is increased owing to the
hybridization of target DNA and probe DNA (Figure. 21(c)). This phenomenon
indicates that the charge-transfer resistance at the electrode/electrolyte interface
increases as more DNA added to the electrode surface. The characteristics of
nanowell electrode and macro electrode were further analyzed by Bode plot. The
Bode plot shows the transfer function of frequency was plotted as a log-frequency
axis
43
Figure 22. Characterization of nanowell electrode sensitivity by EIS
Bode plot shows: (a) macro electrode, (b) probe DNA on nanowell, (c)
hybridization of 100 nM of target DNA
X-axis indicated the frequency, the y-axis is the impedance and |Z| is shown
as a log scale (Figure 22). A higher impedance value in macro electrode are
observed at the low-frequency region (<1.59Hz). This means that the macro
electrode current flow in the low frequency is not sensitive as nanowell electrode.
Generally, the current signal from kinetic reaction is proportional to the larger
working electrode surface in the Cottrell behavior system. Even though area of the
nanowell structure is much larger than the macro electrode owing to SiO2 resist
layer. Nanowell electrode system shows much higher sensitive system than the
macro electrode owing to 3-dimensional diffusion structures. Effect of target DNA
was also analyzed using 100 nM of target DNA. After the hybridization process of
44
target DNA in the nanowell electrode, impedance is higher than after
hybridization process as shown in the Figure 22(b), (c). The diameter of Nyquist
semicircles increases as probe DNA hybridizes with target. This phenomenon
indicates that the charge transfer resistance at the electrode/electrolyte interface
increases as more layers of hybridization complex are added to the electrode
surface. Therefore, ferrocyanide ions are more inhibited to move toward the
electrode surface after the hybridization of target DNA. We performed a
simulation using the equivalent circuit by Z-view program; mainly changing the
parameters at the electrode surface was seen. It is easily understandable the
physical aspects of nanowell electrode using EIS data.
45
4.4. Development of Equivalent Circuit Model
Equivalent circuit fitting is a quite useful to understand an immobilization
status of H5N1 DNA sensor, particularly aspects of physical meaning. Through
the possible relationship between equivalent circuit element values and physical
characteristics of the EIS system, 4 kinds of models were examined, modified,
and compared to develop a model with good fitting properties and logical physical
significance.
Figure 23. Randles circuit model
Figure 24. Nyquist plot indicating impedance of the Randles circuit[56]
The Randles circuit is one of the simplest models, indicating charge
46
transfer for a redox species attached to a monolayer (Figure 23). Similar circuits
have been successfully used in several reports of biosensors [57, 58]. Fitting
constraints were imposed such that further iterations were stopped when the chi-
square (χ2) change was less than 0.001% compared to the previous iteration. The
goodness of fit was assessed from minimum χ2, correlation matrix and relative
error distribution plots, less than 5% fluctuations between the experimental and
fitted data were assumed to be satisfactory in confirming the validity of the
equivalent circuit [59].
For more accurate analysis, the recorded impedance spectra were
numerically modelled using the equivalent circuit shown in Figure 25, which
includes 3 parameters: the solution resistance (Rs), Rct, and the double-layer
capacitance (Cdl) for the electrode/solution interface [21].
47
CASE 1
Figure 25. Nyquist and Bode plot of fitting results with equivalent circuit (Rs
= 543 ohm, Rct = 49,875 ohm Cdl = 4.56 × 10 −7
F)
A simplified Randles circuit is always showing a semicircle (Figure 25).
The solution resistance was found by reading the real axis value at the high
frequency intercept. The curve fitting result is not suitable well at low frequencies.
Impedance data |Z| was fitted well in Bode magnitude plot. However, plot of
phase angle was not showing good fitting results. All simulated parameters are
listed numerically in case 1 (Table 2). The chi-sqr value in the last column is the
statistical analytic result for case 1. The value of 0.74 for chi-sqr means that test
result was not fitted well compared to other cases. Equivalent circuit parameter
fitting was not proper because of use of ideal capacitor and absence of Warburg
element for diffusion.
First, capacitors in EIS experiments often do not behave ideally. Instead, they act
0 25000 50000 75000
-75000
-50000
-25000
0
Z' (ohm)
Z''
(oh
m)
base (probe DNA)FitResult
10-1 100 101 102 103 104 105102
103
104
105
Frequency (Hz)
|Z|
base (probe DNA)FitResult
10-1 100 101 102 103 104 105
-75
-50
-25
0
Frequency (Hz)
the
ta
48
like a CPE as defined below [60].
Due to surface heterogeneity, the impedance Z(ω) of such a nonideal
layer can be expressed as Z(ω) = CPE−1
(jω)−n
, where ω is a circular frequency
and n parameter varies from 0 to 1. When n is close to 0, CPE is essentially a
resistance. If n=1, the CPE is a pure capacitance and electrode is considered as
ideal [17, 61, 62]. For a constant phase element, the exponent n is less than one.
The "double layer capacitor" on real cells often behaves like a CPE, not a
capacitor. While several theories (surface roughness, “leaky” capacitor, non-
uniform current distribution, etc.) have been proposed to account the non-ideal
behavior of the double layer, it is probably best to treat n as an empirical constant
with no real physical basis.
49
Instead of an ideal capacitor, the CPE is used to compromise errors due to
microscopic roughness and atomic scale inhomogeneity in surfaces [63].
CASE 2
Figure 26. Nyquist and Bode plot of equivalent circuit-based fitting results
(Rs = 550 ohm, Rct = 56,836 ohm, CPE = 6.37 × 10 −7
F, CPE-P = 0.92)
After changing a capacitor to CPE for fitting impedance data (Figure 26),
the value of the chi-sqr was decreased significantly from 0.74 to 0.02 (Table 2).
The size and height of the semicircular became similar (Figure 26). However, the
error at a low frequency range has not been improved yet in the Bode plot. The
Warburg impedance characteristic was appeared as a diagonal line with an slope
of 45° at a low frequency (Figure 26). |Z| is given by
(8)
0 25000 50000 75000
-75000
-50000
-25000
0
Z' (ohm)
Z''
(oh
m)
base (probe DNA)FitResult
10-1 100 101 102 103 104 105102
103
104
105
Frequency (Hz)
|Z|
base (probe DNA)FitResult
10-1 100 101 102 103 104 105
-75
-50
-25
0
Frequency (Hz)
the
ta
50
where YO is a parameter characteristic of the medium [64]. For a reversible redox
couple, in which both reduced and oxidized forms have similar D and K and are at
the same concentration cro/2 in solution, YO is given by
(9)
Where cro is the bulk concentration of the redox species. D is the
diffusivity of ion in the membrane, K is the partitioning coefficient. More general
expressions for YO may be found in [34, 65].
The Warburg open and Warburg short element was added to get an
improved fitting data (Figure 27 and Figure 28). Warburg Open (WO). The
formula for Zw is similar to ZO but the hyperbolic cotangent function replaces
hyperbolic tangent.
(10)
The expressions for YO and B are identical to Eq. (9). They may be important for
analysis of diffusion through defects in a thin film, e.g. cracks or pinholes[66].
The finite thickness of the electrolyte solution has been well known. It
requires that the regular Warburg (valid for an infinitely thick layer) be replaced
with the so-called porous Warburg short element[34, 67]. Warburg Short (Ws)
element is described by a formula involving the complex hyperbolic tangent
function[67] with YO is given by the same expression as for the regular Warburg
51
Eq. (8) and B = . The parameter B2 has the meaning of the characteristic
time of diffusion through the electrolyte. It is easily seen that at high frequencies
(f >> B−2
) ZO becomes identical with the regular Warburg impedance ZW (Eq. (7)).
At low frequencies (f << B−2
) ZO will be proportional to the steady-state diffusion
resistance of the film to the redox species: Importantly, at high frequencies ZO ≈
ZW is determined, while at low frequencies ZO ≈ RO is determined by the
absolute permeability
(11)
As shown in Table 2, when using the Warburg short, the results of the
fitting came out as the best. For that reason the target DNA and the non-
complementary DNA hybridization was measured and analyzed by EIS with the
equivalent circuit model (CASE 4) [68].
52
CASE 3
Figure 27. Nyquist and Bode plot of Randles circuit-based fitting results (Rs =
557 ohm, Rct = 55,694 ohm, CPE = 6.32 × 10 −7
F, CPE-P = 0.92, W1-R =
13586 W1-T = 5.363, W1-P = 0.83)
0 25000 50000 75000
-75000
-50000
-25000
0
Z' (ohm)
Z''
(oh
m)
base (probe DNA)FitResult
10-1 100 101 102 103 104 105102
103
104
105
Frequency (Hz)
|Z|
base (probe DNA)FitResult
10-1 100 101 102 103 104 105
-75
-50
-25
0
Frequency (Hz)
the
ta
53
CASE 4
Figure 28. Nyquist and Bode plot showing Randles circuit-based fitting
results (Rs = 557 ohm, Rct = 56,400 ohm CPE = 6.19 × 10 −7
F, CPE-P = 0.92,
W1-R = 41010, W1-T = 13.97, W1-P = 0.92)
Table 2. Elemental parameters values obtained from fitting of various
equivalent circuit models
Circuit Rs
(Ohm)
Rct
(Ohm)
CPE
(F)
CPE-P W1-R W1-T W1-P Chi-Sqr
Case 1 543 49,857 4.26E-07 N/A N/A N/A N/A 0.743
Case 2 550 56,836 6.37E-07 0.92041 N/A N/A N/A 0.022
Case 3 557 55,694 6.23E-07 0.92323 13586 5.363 0.8312 0.022
Case 4 557 56,400 6.19E-7 0.92481 41010 13.97 0.92458 0.014
0 25000 50000 75000
-75000
-50000
-25000
0
Z' (ohm)
Z''
(oh
m)
FitResult
10-1 100 101 102 103 104 105102
103
104
105
Frequency (Hz)
|Z|
base (probe DNA)FitResult
10-1 100 101 102 103 104 105
-100
-75
-50
-25
0
Frequency (Hz)
the
ta
54
4.5. Hybridization of Probe DNA and Non-
Complementary DNA
EIS spectra of non-conpmentary DNA samples after incubation process at
various concentrations from 1 pM to 1 μM were analyzed. The Nyquist plots (Z
vs. Z; Z = real impedance and Z = imaginary impedance) corresponding to the
charge transfer process and the diffusion process are shown in Figure 29.
(semicircle in the high-frequency region; straight line in the low frequency region).
It can be easily observed that solution resistance was not moved.
0 10 20 30 40 50
-0
-10
-20
-30
-40
-50
Z'' (
ko
hm
)
Z' (kohm)
base probe DNA
1 pM
10 pM
100 pM
1nM
10 nM
100 nM
1 uM
Figure 29. Nyquist plot of probe DNA and non-complementary DNA
hybridization
Figure 29 shows the Nyquist plot at variant non-complementary DNA
concentrations. An increased semicircle in Rct was not observed when the non-
complementary DNA concentration was increased. A semicircle diameter of probe
DNA without hybridization was smallest value for all cases. It is probably caused
55
by non-specific binding of non-complementary to the probe DNA and results in
reduce a charge transfer ability on the nanowell electrode surface. Warburg line is
same for all cases. No additional increase or decrease did not observed in the
Nyquist plot. So this system is probably controlled by diffusion because redox
couple moved toward nanowell electrode owing to concentration gradient of
redox ion. Concequently, a Total of 8 concentrations is preformed including probe
DNA (see the legend of base probe DNA, black) all plot line is similar pattern in
the entire frequency range. Any significant increase of parameters in Nyquist plot
was not observed in non-complementary H5N1 DNA in the any parameter of
Nyquist plot. Figure 30 shows a Bode plot of non-complementary DNA
hybridization with target DNA. No significant increase or decrease at the various
DNA concentrations did not observed over the whole frequency. (0.1 Hz to 1
MHz) At 1 MHz of high frequency, theta value is over 90 º, normally frequency is
higher and conductance value (impedance of capacitance value) is close to zero in
the following equation.
( )
The higher frequency, |Z| value is smaller but theta value is bigger.
Capacitance (pseudo-capacitance component) is shown. It is not a value resulted
from kinetic or diffusion reactions. It is caused from electrode surface
characteristic generated by too high frequency. So we exclude this signal to
interpret interface/electrolyte reaction for accurate analysis.
56
10-1
100
101
102
103
104
105
106
0
10
20
30
40
50|Z
| (k
oh
m)
Frequency (Hz)
base probe DNA
1 pM
10 pM
100 pM
1 nM
10 nM
100 nM
1 M
10-1
100
101
102
103
104
105
106
-0
-20
-40
-60
-80
-100
-120
Ph
as
e a
ng
le ()
Frequency (Hz)
base probe DNA
1 pM
10 pM
100 pM
1 nM
10 nM
100 nM
1 M
Figure 30. Bode Plot of probe DNA and non-complementary DNA
hybridization
In this study, we focused on the double layer capacitance and charge
transfer resistance between the nanowell surface and the electrolyte solution. In
57
fact, modification at this interface by immobilization of probe DNA to the
conducting surface will lead to a change in capacitance and charge resistance
whose magnitude will depend on the nature and coverage of the recognition
element[13]. For the more acute analysis equivalent circuit fitting was performed.
The best-fit impedance parameters are given in Table 3. The average solution
resistance value is 542 ohm, a large difference has not occurred with increasing
concentration of non-complementary DNA.
In Rct parameter which shows the characteristics of the charge transfer
reaction in the interface/electrolyte, the Rct value varied from 1.4% to 8.0%. A W
parameter shows the characteristics of diffusion reaction at the electrode surface
and WR value decreased to 76% at 1 pM non-complementary DNA. A CPE
parameter shows the characteristics of the electrode surface and the CPE value
increased to maximum 11% at 1 µM of non-complementary DNA.
58
Table 3. Elemental parameters values obtained from equivalent circuit fitting
DNA
concentration*
Rs
(Ohm)
Rct
(Ohm)
W1-
R(+)
W1-
T(+)
W1-
P(+)
CPE
(F)
CPE-P
Chi-sqr
Base probe
DNA 484 35,506 38,637 18.5 0.61 6.5E-07 0.92 0.02
1 pM
511 36,013 9,384 2.5 0.59 6.4E-07 0.92 0.02
10 pM
488 36,121 9,069 2.3 0.61 6.2E-07 0.92 0.02
100 pM
445 38,343 4,601 1.8 0.72 6.2E-07 0.90 0.03
1 nM
485 37,183 7,686 2.4 0.64 6.1E-07 0.92 0.03
10 nM
511 37,661 10,862 2.3 0.62 6.6E-07 0.91 0.02
100 nM
484 38,243 9,339 2.2 0.62 6.3E-07 0.92 0.02
1 µM
489 36,881 15,792 2.9 0.55 7.2E-07 0.91 0.03
*EIS was measured at various concentration of non-complementary DNA (Figure
29)
59
4.6. Hybridization of Probe DNA and Target DNA
Figure 31 shows the results of Nyquist plot for the hybridization event
between probe and target DNA with various concentrations. It was observed that
solution resistance was not moved, and the diameter of semicircle increased at the
higher concentration of the target DNA. The semicircle is represented by Rct
connected with diffusion impedance series [34, 65]. It is inversely proportional to
the concentrations of redox species and is relatively small for fast reversible
reactions, such as Fe(CN)6 3− ↔ Fe(CN)6
4−.
Similar to other reprorted modified electrodes[69], covering of nanowell
electrode with target DNA brought steric hindrance effect on the electron transfer
of [Fe(CN)6]3−/4−
through electrode, resulting in reduced electron transfer speed,
and an obvious increase of the Rct value. Since the solution pH (7.0) is above the
theoretical pI of streptavidin (5), the streptavidin carries negative charge, whose
effect on the Rct value would be relatively small compared to the huge steric
hindrance effect. Moreover, Rct increase can be explained by the accumulation of
negative charge from the DNA backbone after hybridization. This causes a higher
barrier for the negatively charged ferri/ferrocyanide anions ([Fe(CN)6 3−/4−
]) and
results in reduced charge transfer ability on the nanowell electrode surface,
leading to an increase in the Rct value. Tersch et al. (2011) reported that the
increase of Rct in EIS depends on the increase in length of the oligonucleotides
[22]. A negligible change in RS was observed during the target concentration
60
variation as shown in Table 4, results demonstrate that the solution resistance was
not affected by the target concentration. The average solution resistance value is
562 ohm, which has a 40 ohm values difference in non-complementary DNA
(Table 3). A large difference has not occurred with increasing concentration of
target DNA can be distinguished. The solution resistance value was not changed
in the three-electrode system if electrochemical system is a stable.
The CPE value was only reduced to 11% in maximum at 100 µM
concentration of target DNA (from 0.62 µF to 0.55 µF), In high target DNA
concentrations, the diffusion-controlled part of impedance spectroscopy did not
appear at lower frequency. On the contrary, the impedance plot of the low
concentration of target DNA exhibited an obvious linear correlation of Z’ and Z’’
at lower frequency region, which is the diffusion-controlled process
corresponding to Warburg impedance. Disappearance of the straight line at low
frequency from 10 nM to 1 µM is observed. The difference in the impedance low
concentration and high concentration target DNA hybridization on nanowell
electrode surface manifested the effective blocking ability of target DNA to the
redox reaction of Fe(CN)63−/4−.
As suggested by Gautier et al. [7], probe DNA on
the nanoelectrode surface generated a negatively charged backbone that reduced
the electron ability to penetrate into electrode by passing electrolyte. electrolyte to
penetrate the electrode, eventually eliminated the response of the Fe(CN)63−/4−
anion effectively. The depth of the nanowell is reduced as probe DNA hybridize
with variant target DNA and reduces the effect of 3d-diffusion. Therefore,
61
Warburg impedance can be neglected when the higher target DNA concentration
increases. However, it was not used for quantitation of target DNA in this study.
At the same time, the changes in Rct were much larger than those in other
impedance components (Table 4).
On the other hand Rct value shows a tendency to increase as the target DNA
concentration increases. Rct value of the probe DNA is different between the non-
complementary and target DNA. Indicating that the absolute value of the Rct is not
a parameter for quantitation, However, Rct-based difference (ΔRct) on probe DNA
is possible to use a sensor parameter as quantitative detection of target DNA.
0 5 10 15 20 25 30 35 40 45 50 55 60 65 70 75 80
0
-5
-10
-15
-20
-25
-30
-35
-40
Z'' (
ko
hm
)
Z' (kohm)
Base probe DNA
Target DNA 1pM
Target DNA 10 pM
Target DNA 100 pM
Target DNA 1 nM
Target DNA 10 nM
Target DNA 100 nM
Target DNA 1 M
Figure 31. Nyquist plot of probe DNA and target DNA hybridization
62
10-1
100
101
102
103
104
105
106
102
103
104
105
/Z/
(oh
m)
Frequency (Hz)
Base probe DNA
Target DNA 1 pM
Target DNA 10 pM
Target DNA 100 pM
Target DNA 1 nM
Target DNA 10 nM
Target DNA 100 nM
Target DNA 1 M
10-1
100
101
102
103
104
105
106
-0
-20
-40
-60
-80
-100
-120
Ph
as
e a
ng
le ()
Frequency (Hz)
Base probe DNA
Target DNA 1 pM
Target DNA 10 pM
Target DNA 100 pM
Target DNA 1 nM
Target DNA 10 nM
Target DNA 100 nM
Target DNA 1 uM
Figure 32. Bode plot of probe DNA and target DNA hybridization
Figure 32 shows the Bode plots of |Z| impedance and theta degree. A
linearly relationship is observed in the medium-frequency range (10 Hz ~ 1 kHz)
63
with log |Z| vs. log f the function is a straight line with a slope approximately -1. It
indicates that sensor have a mainly capacitive properties (the phase angle becomes
close to 90°) at the medium range. The log |Z| remains nearly unchanged at low
and high frequencies (where the resistive components dominate) and decreases at
medium frequencies from 10 Hz to 1 kHz.
In the Bode phase plot, the approach to pure capacitive behavior is
usually identified with theta approaching to the negative 90 degree [34].
Accordingly, the value of the can be used to evaluate the effectiveness of redox
ion diffusion in nanoelectrode at the medium-frequency region. That is, the
smaller the phase angle, the better the capacitive performance and, hence, the
faster the ions diffuse [70]. When the frequency is from 1 Hz ~ 10 kHz, where the
impedance behavior of target DNA would be kinetic-controlled, the phase angle
( ) of probe DNA is always smaller than that of target DNA (10 nM, 100 nM, 1
µM). This result indicates more rapid diffusion of redox ions in the
immobilization of probe DNA than that of target DNA.
64
Table 4. Elemental parameters values obtained from equivalent circuit fitting
Target DNA
concentration
Rs
(Ohm)
Rct
(Ohm)
W1-
R(+)
W1-
T(+)
W1-
P(+)
CPE
(F)
CPE-P
Chi-sqr
Base probe
DNA 607 53,897 5143 2.2 0.75 6.2E-07 0.92 0.02
1 pM
552 55,946 1949 2.5 0.89 5.9E-07 0.94 0.02
10 pM
630 60,824 2093 2.3 0.87 6.0E-07 0.93 0.02
100 pM
558 64,059 342.6 9.9 1.00 5.5E-07 0.95 0.03
1 nM
621 66,100 6798 3.1 1.01 5.7E-07 0.94 0.03
10 nM
490 77,836 1755 3.1 1.00 5.9E-07 0.97 0.02
100 nM
574 82,828 77787 13.9 0.94 6.5E-07 0.97 0.02
1 µM
461 81,631 218.2 9.9 1.00 5.5E-07 0.97 0.03
*EIS was measured using various concentrations of target DNA (Figure 31).
Differences in probe DNA and target DNA charge transfer resistance
(ΔRct) was plotted as a function of target DNA concentration. When probe DNA is
hybridized with target DNA, not only is the thickness of the molecular layer
increased, but the phosphate backbone of DNA also pushes away ferricyanide
anions from the electrode surface [71, 72]. In contrast to the behaviour of Cdl,
only a small (11%) increase in the capacitance was observed after DNA
hybridization.
65
Figure 33. ΔRct based quantitative analysis of probe/target DNA hybridization.
ΔRct (Rct of target DNA – Rct of probe DNA) was plotted as functions of target
DNA concentration
As shown in Figure 33, a quantitative curve for the detection of
complementary target DNA ranging from 1 pM to 1 µM was observed. In contrast,
ΔRct showed a negligible increase in the case of non-complementary DNA. The
detection limit of the H5N1 target DNA detection system on the nanowell array
electrode was 1 pM, which is much more sensitive than that of other detection
techniques such as surface plasmon resonance and quartz crystal microbalance [1].
ΔRct was chosen as the sensor parameter for DNA detection. and was used as the
H5N1 DNA detection parameter in EIS. The change in impedance at a fixed
frequency could be used as a DNA quantitating system in a biosensor.
66
5. Conclusions
A nanowell array electrode-based electrochemical quantitative system for the
detection of H5N1 DNA without amplification was developed using the EIS
method. Hybridization complex formation of probe and target DNA and the
hybridization event was demonstrated and characterized by AFM. The
hybridization of probe DNA with target DNA in the nanowell resulted in an
increase in ΔRct when analyzed by EIS, and non-complementary DNA yielded
negligible changes in ΔRct compared to complementary target DNA. The limit of
detection of H5N1 target DNA is 1 pM and it can be applied to H5N1 target DNA
detection in the biological field, providing an effective method with high
sensitivity and selectivity.
67
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국문 초록
나노웰 전극을 이용한 H5N1 DNA 의 전기화학적 정량 분석
연세대학교 대학원
의공학과
차민석
본 연구에서는 나노웰 전극을 이용하여 H5N1 조류 인플루엔자
DNA 를 정량적으로 검출하는 전기화학 분석 시스템을 개발하였다. 나노웰
전극은 KrF 레이저를 이용하여 전극 지름이 500 nm 인 전극을 사용하였고,
DNA 센서는 금전극 위에 자가조립단층과 스트렙타아비딘을 고정한 뒤,
최종적으로 biotin 이 표지된 18-mer 탐침 H5N1 DNA 를 스트렙타아비딘에
고정화 되도록 하여 DNA 검출 센서를 제작 하였다. 나노웰 표면에서의 DNA
교잡 특성을 파악하기 위해서 원자간력현미경과 임피던스분광법을 사용하였다.
원자간력현미경 이미지를 통해서는 H5N1 DNA 교잡 상태에 대한 물리적 표면
특성을 관찰하여 DNA 교잡 후에 나노웰의 깊이가 200 nm 에서 15 nm 로
줄어드는 것을 확인하였고, 임피던스 분광법을 통해서는 H5N1 표적 DNA
농도의 증가에 따른 표면 특성 변화를 측정하였다. 또한 등가회로 모델링
분석을 통하여 DNA 농도 증가에 민감하게 반응하는 인자가 전하전달저항
(Rct) 임을 확인하였다. 특히, 탐침 DNA 의 전자전달저항 값 (Rct_probe)과 표적
DNA 가 교잡된 상태의 전자전달저항 (Rct_target) 값의 차이 (ΔRct)가
정량적인 H5N1 DNA 검출을 위한 주요한 분석인자로 사용될 수 있음을
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확인하였다. 즉, 표적 DNA 의 교잡되는 농도가 높아질수록 ΔRct 값이 증가하는
형태로 나타남을 임피던스 분광법을 통해 확인하였다. H5N1 DNA 검출
시스템의 검출 한계 농도는 1 pM 이고, 본 검출 시스템을 통해 H5N1
DNA 농도가 1 pM 에서 1 µM 까지 정량적인 검출이 가능함을 확인하였다.
핵심단어: 나노웰 전극, H5N1 DNA, 임피던스 분광법, 전하전달저항
