DISSERTATION

Jambrec Daliborka – 2016
DISSERTATION
Understanding Potential-Assisted Surface
Modification – From Self-Assembled
Monolayers to DNA Chips
Dissertation
Submitted for the degree of
Doctor of Natural Sciences (Dr. Rer. Nat.)
Faculty of Chemistry and Biochemistry
Ruhr University Bochum
Daliborka Jambrec
Bochum, October 2016

This work was carried out in the period from April 2013 to October 2016 in the Department of
Analytical Chemistry under the supervision of Prof. Dr. Wolfgang Schuhmann, Ruhr
University Bochum, Germany.
First Examiner: Prof. Dr. Wolfgang Schuhmann
Second Examiner: Prof. Dr. Axel Rosenhahn
Chair of the Examination Board: Prof. Dr. Christof Hättig
Date of the PhD Defense: 2 nd December 2016

Acknowledgements
For me, the PhD studies were a period of learning – learning about science, about
different cultures, languages and cuisines, about responsibilities, about growing up and making
important decisions.
During this period, many people passed through my life and helped me on my journey,
professionally and privately. The best example for this is my mentor and friend, Prof. Dr.
Wolfgang Schuhmann, who I owe my deepest gratitude for constantly helping me become a
better researcher and a stronger person, encouraging and supporting me in every step of my
path.
Furthermore, I am grateful to Prof. Dr. Axel Rosenhahn, for being interested in my research
and accepting to be my second examiner.
I would like to thank Anna Lauks, who helped me make the first steps in the PhD, during the
initial month of my stay in the group. It was a challenging period, starting from the fact that
she did not speak English nor did I speak German, thus I am grateful for her patience and
willingness to transfer all her experience to me. Also, I would like to thank Dr. Kirill Sliozberg
who helped me throughout the thesis with technical aspects alongside Dr. Thomas Erichsen.
I would like to express special thanks to Bettina Stetzka, for her big help with the
administration, which was not always only work related, and for having patience with my
German. Her constant smile, calmness and physical fitness are very inspirational for me.
Furthermore, I would like to thank all the people who cooperated with me during this period. I
am especially thankful to Dr. Magdalena Gebala, Prof. Dr. Fabio La Mantia and Dr. Arturo
Estrada-Vargas for their help in understanding the fundamentals of electrochemistry and DNA
sensing. I owe my thanks to Bianca Ciui for the work we started on miRNA detection, Dr.
Adrian Ruff for his help in the work with the glucose-oxidase-acridine orange intercalator, Dr.
Ĺubomír Švorc for helping me expand my knowledge about DNA to boron-doped-diamond
electrodes, Dr. Sanaz Pilehvar and Prof. Dr. Karolien De Wael for cooperation in the research
with DNA aptasensors and Vera Eβmann for the help in discovering the field of bipolar
chemistry.

I would like to thank all colleagues from my group, from older members like Lutz (whose
cheerful spirit and singing in our office always made me very happy), Piyanut and Fangyuan,
to the new ones like Danea and Nergis, for the pleasant and always dynamic intercultural
atmosphere, and for making long days in the lab not difficult. During this period I developed
an especial fondness for a few people from the group. I would like to thank Jan and his wife
Kathi for the great moments spent together, usually by cooking good food and listening to good
records. My biggest appreciation goes to the members of my “fantastic 4” – Grecia, Felipe and
Uğur, with who I shared the most beautiful and the most difficult moments of this period of
my life. We became brothers and sisters.
In the end, there are not enough words to show gratitude to my family, old and new, for their
unconditional love and support during the ups and downs throughout my stay in Germany.
Even though it was difficult to be separated from my grandma, my mom and dad, and especially
my sister, they encouraged me and cheered for me in every step of my chosen way. Hvala vam
na svemu, volim vas! My warmest thanks go to my husband – my best teammate, for giving
me peace, stability and the biggest support in pursuing my goals. VTNNSC!

“A structure this pretty just had to exist”
James Watson, “The Double Helix”, 1986
“Science and everyday life cannot and should not be separated…
it is based on fact, experience and experiment.”
Rosalind Franklin, 1940

Table of Contents
1. Introduction …..………………………………………………………………… 1
1.1 Thiol self-assembling …..………………………………………………………… 3
1.2 DNA ……………………………………………………………………………...…… 6
1.2.1 DNA in solution …….……………………………………………………………...… 9
1.2.2 DNA in front of a polarized electrode ……………………………………………..... 13
1.3 DNA immobilization …………………………………………………………….. 16
1.3.1 DNA immobilization approaches …...……………………………………………..... 16
1.3.2 Techniques for DNA microarray fabrication .……………………………………..... 18
1.4 Hybridization detection ..……………………………………………………….. 21
2. Aims of the Work ….………………………………………..…………...… 26
3. Results and Discussion ………………………………..……………...… 29
3.1 Importance of preparing the surface. Criteria for cleanliness .…...….. 30
3.2 Importance of knowing the surface ……………………………………...….. 34
3.2.1 Electrochemical impedance spectroscopy. DNA assay build-up …...…………...….. 35
3.2.2 Potential of zero charge of bare and DNA-modified electrodes ………………...….. 41
3.3 Importance of controlling the surface ...………………………………...….. 50
3.3.1 Fast and controlled formation of DNA surfaces. Optimization of ssDNA
immobilization procedure …..…………………………………………………...….. 51
3.3.2 Formation of compact thiol SAMs within minutes ……………………………...….. 67
3.3.3 Reproducible recycling of Au modified surfaces within seconds …..…………...….. 81
3.4 Potential-assisted preparation of DNA sensors …..…………………...….. 85
3.4.1 Optimization of the potential-pulse assisted immobilization method …………...….. 85
3.4.2 DNA microchip fabrication ……………………………………………………...….. 91
3.5 Intercalation as a DNA detection technique …………………………...….. 96
4. Conclusions ………………………………………………………………...… 104
5. Experimental Work ...………………………………………………....... 108

5.1 Materials and consumables …..……………………………...…….......… 109
5.2 Electrochemical setup and instrumentation ……………...……….…. 110
5.3 Preparation of gold surfaces ….………………………………………. 111
5.4 Determination of the potential of zero charge …...………………….. 113
5.5 Potential-assisted formation of self-assembled monolayers ….…….. 114
5.6 Potential-assisted desorption …………………………………………. 115
5.7 Preparation of DNA sensors ….………………………………………. 116
5.7.1 ssDNA immobilization via incubation ...…………………………………………… 116
5.7.2 Potential-assisted ssDNA immobilization ….……………………………………… 116
5.7.3 Passivation by means of incubation ...……………………………………………… 117
5.7.4 Potential-assisted passivation .……………………………………………………… 117
5.8 Preparation of DNA chips ..…………………………………………………… 117
5.9 Characterization of DNA sensors …..………………………………... 118
5.9.1 Electrochemical impedance spectroscopy ……………………………………… 118
5.9.2 Cyclic voltammetry ……………………………………………………………… 118
5.9.3 Chronocoulometry for determination of DNA coverage …………………………… 118
5.10 Hybridization and dehybridization ..……………………………………… 118
5.10.1 Detection of hybridization ………………………………………………………… 119
5.10.2 DNA coverage determination by means of FSCV ...…………………....………… 119
5.11 Intercalation .…………………………………………………………………… 120
5.12 Methods …..……………………………………………………………………… 120
5.12.1 Electrochemical impedance spectroscopy ………………………………………… 120
5.12.2 Chronocoulometry for the determination of DNA coverage ...…………………… 124
6. References ...………………………………………………………………….... 127
7. Appendix ..……………………………………………………………………… 133
7.1 List of symbols and abbreviations …...……………………………………… 133
7.2 Publications list …..……………………………………………………………… 136
7.3 Conference contributions ...…………………………………………………… 138

1. Introduction

______________________________________________________________________ Introduction
Surface modification is the process of altering of the material surface by implementing
new and desired characteristics to it. Molecular self-assembly, a process of high-level intermolecular
orientation and organization without an outer force, is a powerful tool for surface
modification. Self-assembled monolayers (SAMs) have a wide range of applications 1 , but
whether they are used for regulation of the surface wettability, as a protection layer in corrosion
inhibition, or surface functionalization for binding proteins, DNA or cells, the main requirement
in all their applications is the control of the modification process. For example, depending on
the envisaged application, the desired coverage of the layer can significantly differ.
Furthermore, the choice of numerous parameters determines both the rate and the duration of
the modification process resulting in the desired coverage.
One of the biggest applications of SAMs is in the development of biosensors. DNA
hybridization detection and DNA sensors are becoming tremendously important in diagnostics
as a result of the continuous advancement of the Human Genome Project 2,3 . A DNA sensor
usually consists of a single stranded DNA grafted on an electrode surface, essential for the
recognition of the complementary target DNA present in a sample under investigation. The
recognition process occurs by hybridization between these two DNA strands, which is
converted into a signal by the transducer part of the DNA biosensor. Transduction of the signal
can be done by different techniques – optically, piezoelectrically and electrochemically. In
recent years the development of electrochemical DNA biosensors is in the spotlight, due to their
operating simplicity, possibility of miniaturization, and thus portability, and low cost, which
makes them very attractive for mass production. Continuous investigation of fundamental
issues, such as surface characterization, tailoring of interfaces and DNA recognition, along with
the progress in the field of microfabrication will lead to powerful, yet easy-to-use DNA
diagnostic products 4 .
The sensitivity and selectivity of DNA biosensors is highly dependent on the quality of the
prepared DNA sensing surface. In order to tailor the desired DNA sensing surfaces for
envisaged sensing platforms it is of utmost importance to understand processes occurring at the
electrode during the surface modification 5 . Only in this way properties of the surface can be
controlled in a reproducible manner on a desired time scale.
The following sections will present the current state of the art in the field of self-assembly and
DNA sensing, including some historical aspects related to the “molecule of life”.
2

______________________________________________________________________ Introduction
1.1 Thiol self-assembling
The formation of self-assembled monolayers (SAMs) was first reported by Zisman et al. in the
late 1940s 6,7 . Later, in the 1980s, Nuzzo and Allara made initial studies on alkylthiols selfassembly
on gold 8 . Since then, a tremendous number of studies were devoted to understanding
and improving the formation of alkylthiol SAMs for numerous applications 1 – tailoring surface
wetting behavior (hydrophilicity/hydrophobicity), protective coatings, corrosion inhibition and
surface functionalization.
Self-assembling occurs from either a vapor or liquid phase 9 , where head groups assemble close
to the surface, while tail groups orientate away from the surface (Figure 1.1). Self-assembly of
thiols on gold surfaces is characterized by its simplicity of preparation and a high number of
available functional groups, even though it has a limited stability due to the relatively small safe
potential region of the Au-S bond.
Alkylthiols consist of an alkyl chain with a thiol head group at one end, which has a high affinity
towards gold, and a functional group of choice at the other end. Adsorbed thiols lose their
hydrogen atom from the head group and form a strong thiolate-gold bond at the interface. The
formed chains are in a trans-conformation obtaining an angle of 60-70° with respect to the
surface. The sulfur atom is in sp3 hybridization and therefore bound to three Au atoms. Au-S
bond formation occurs through the reduction of alkylthiols to alkylthiolates 10 and the following
equation presents one of the proposed mechanisms for this reaction:
R − SH + Au(0) n → R − S − Au(I) • Au(0) n + 1 2 H 2
(1.1)
Disulfides are believed to chemisorb on gold via S-S bond cleavage:
RS − SH + 2Au → 2RS − Au (1.2)
The thickness of the monolayer is controlled by the length of the alkyl chains. During selfassembly,
alkyl chains align themselves by van der Waals forces. Therefore, longer alkylthiols
result in more stable and densely packed monolayers 7 . Furthermore, longer alkylthiols adsorb
preferentially over shorter ones 11 .
1.1 Thiol self-assembling 3

______________________________________________________________________ Introduction
Figure 1.1. Self-assembly on an electrode surface.
In the beginning of the self-assembly process, while the density of molecules at the surface is
small, molecules either form a disordered mass or an ordered lying phase at the surface 1 (Figure
1.2, a). When the density increases, molecules start forming ordered three-dimensional
structures. Even though the Au-S bond is reasonably strong, the adsorbed thiols can still move
around on the electrode to heal gaps of exposed gold 12 . This so-called healing process occurs
in a second phase, where thiols slowly reorganize and order on the surface 13-15 (Figure 1.2, b).
Various parameters affect alkylthiol SAM formation such as the architecture and cleanliness of
the gold surface, temperature, thiol concentration, functional groups and solvent composition.
Commonly, organic solvents such as ethanol or DMSO are used for adsorption of alkylthiols
on gold.
Preparation of SAMs is standardly performed by immersing a gold substrate into an alkylthiol
solution. Thus, the adsorption occurs at open circuit potential (OCP) and the kinetics of the
adsorption roughly follows the Langmuir adsorption curve 1,9 . According to the Langmuir
kinetic model, the rate of adsorption is proportional to the free space on the surface:
dθ
dt = k(1 − θ) (1.3)
where θ is the partial coverage and k is the rate constant. Nevertheless, this model makes strong
assumptions such as the homogeneity of the surface (perfectly flat surface) and the absence of
interactions between adjacent molecules. In reality, surface roughness of polycrystalline gold
and thiol island formation affect the SAM formation kinetics.
1.1 Thiol self-assembling 4

______________________________________________________________________ Introduction
Externally applied potentials affect already formed SAMs, causing the change of their
structure 16 , change of wettability 17 or desorption 18,19 . The desorption of SAMs can occur by
applying rather high positive or negative potentials. Nevertheless, the potential value required
to invoke SAM desorption depends on the type of alkylthiols used (length of the alkyl chain,
head group repulsion) and the coverage, since they influence the stability of the SAM 12 . Longer
alkylthiols and more compact layers are more stable SAMs and therefore, more difficult to
desorb. Furthermore, the possibility of controlling SAM formation by applying an external
potential was observed in the 1990s 20 . Application of anodic constant potentials seems to
accelerate the kinetics of SAM formation 13,21-24 . However, this phenomenon is still poorly
understood.
Figure 1.2. Self-assembly on an electrode surface occurs through two phases: a) in the first
phase (lying phase) molecules randomly lie on the surface, and in b) the second phase
(healing phase) they slowly reorganize to form highly compact monolayers.
1.1 Thiol self-assembling 5

______________________________________________________________________ Introduction
1.2 DNA
Deoxyribonucleic acid (DNA) is the carrier of genetic instructions used for the proper
development and functioning of all living organisms. It was isolated for the first time by
Friedrich Miescher in 1869. Since then, the “molecule of life” has evoked curiosity from many
researchers around the world. The defining moment in nucleic acid research was a paper
published in The Journal of Experimental Medicine in 1944 by Oswald Avery, Colin MacLeod
and Maclyn McCarty, where they show for the first time that DNA, and not proteins, is the
material of inheritance 25 .
The controversy about the discovery of the DNA structure still remains, as it is debatable who
should get the credit for it. In 1953 five papers were published in the journal Nature, describing
and providing evidence for the double helix structure of the DNA. James Watson and Francis
Crick first suggested the correct double helix model 26,27 , based on an X-ray diffraction image
taken by Rosalind Franklin and her student Raymond Gosling. Previously, J. Watson listened
R. Franklin’s lecture about the DNA structure in 1951, and got access to her progress report
without her knowledge. In the same issue of the journal, Maurice Wilkins with his colleagues 28
and R. Franklin with R. Gosling 29,30 provided experimental evidence supporting the Watson
and Crick model. In 1962, after R. Franklin’s death, Watson, Crick and Wilkins received
together the Nobel Prize in Physiology or Medicine. Nobel Prizes are awarded only to living
recipients and it remains unknown whether Rosalind Franklin should have received the prize
as well.
DNA consists of two strands, where each strand is composed of subunits called nucleotides. A
nucleotide consists of three components – a nucleobase, a sugar (deoxyribose) and a phosphate
group (Figure 1.3). Within the same strand, nucleotides are covalently bound to each other in a
chain by connecting the sugar of one nucleotide to the phosphate group of the next one via a
phosphodiester bond, creating a sugar-phosphate backbone. Each phosphate group is bound to
the 3' carbon of the previous deoxyribose, and to the 5' carbon of the following sugar ring. The
ends of a strand are then commonly designated as 3' end or 5' end if the chain finishes with an
unbound 3' or 5' carbon, respectively. Moreover, the connection between nucleotides of two
complementary DNA strands occurs through the nucleobases via formation of hydrogen bonds.
The nucleobases are divided in two groups – pyrimidines (thymine and cytosine) and purines
1.2 DNA 6

______________________________________________________________________ Introduction
(adenine and guanine). Adenine binds only to thymine, forming two hydrogen bonds, and
cytosine binds exclusively to guanine, forming three hydrogen bonds.
Figure 1.3. Scheme of the primary DNA structure.
Two complementary DNA strands (single-stranded DNA, ssDNA) coil around the same axis
forming a double helix (double-stranded DNA, dsDNA). This way hydrophobic bases are inside
the helix and the hydrophilic sugar-phosphate backbone is orientated outside (Figure 1.4). Both
DNA strands store the same biological information, that is, the nucleobase sequence. However,
they can bind only in an antiparallel disposition, which means that the strands are orientated in
opposite directions, one in the direction 3' to 5' and the other from 5' to 3'. The 5' end has a
terminal phosphate group, while the 3' end, a terminal hydroxyl group. DNA can form three
conformations – A, B and Z. B-DNA is a right-handed conformation, where each full turn in
the helix is 3.4 nm long and consists of ten bases (thus with 0.34 nm between adjacent bases).
A-DNA is also right-handed and it occurs when the humidity is lower than 75 %. In this
conformation the base pairs are not perpendicular to the helix axis, resulting in a wider and
1.2 DNA 7

______________________________________________________________________ Introduction
shorter structure than B-DNA. On the other hand, Z-DNA is left-handed and it is stable only at
high salt concentrations in order to minimize electrostatic repulsion, since the phosphate groups
in the backbone are closer to each other.
The genetic information is stored within the double helix structure of the DNA molecule by the
sequence of nucleobases. Besides encoding the formation of cell organelles and proteins, DNA
sequences are responsible for all physical traits and disease susceptibility. Therefore, the
knowledge of DNA sequences is priceless for biological research and fields like medical
diagnosis, genetic screening, forensic biology, environmental control and biotechnology,
among others.
Figure 1.4. Scheme of the secondary DNA structure.
1.2 DNA 8

______________________________________________________________________ Introduction
1.2.1 DNA in solution
DNA is a highly charged polymer with charge density equal to two elementary charges per base
pair. Therefore, in solution DNA interacts strongly with surrounding ions that counterbalance
its charge. These ions can be grouped into different zones 31 . In the region closest to the DNA
there are so called site-bound or inner-sphere ions that share water molecules with the DNA. In
the following zone are outer-sphere ions (territorial ions) that keep their inner hydration layer
and are free to move along the DNA molecule but are kept close to it due to the electrostatic
field. And in the last zone are free ions that form an ionic cloud around the DNA 3 .
The two most known approaches employed to investigate the extent of DNA-ion interaction are
the Manning-Oosawa (MO) counterion condensation theory 32 and the Poisson-Boltzmann (PB)
equation. MO theory addresses the DNA charge compensation by focusing on the counterions
that form a so-called condensed layer around the DNA (outer-sphere ions) in order to reduce
the charge density below a certain critical value 33 . According to the theory, counterion
accumulation at the DNA surface forming a condensed layer occurs under the condition that
the charge density parameter η is > 1:
η = zl B
b
(1.4)
where z is the valence of counterions, lB is the Bjerrum length and b is the charge separation.
The Bjerrum length represents the distance between charges at which their electrostatic
interaction energy equals the thermal energy and is defined as:
l B =
e 2
4πεε 0 kT
(1.5)
where e is the elementary charge, ε is the dielectric constant of the solvent, ε0 is the vacuum
permeability and kT is thermal energy scale. This means that the counterion condensation (for
monovalent ions) occurs when the charge separation b is smaller than the Bjerrum length, which
is true for both ss- and dsDNA. Namely, lB = 0.71 nm in aqueous solutions, while the charge
separation is 0.43 nm in ssDNA and 0.34 nm in dsDNA 31 . According to the theory, condensed
counterions are still assumed to be mobile 34 .
1.2 DNA 9

______________________________________________________________________ Introduction
Accumulation of counterions occurs to a certain extent, until the charge density parameter
decreases to 1, that is, until b increases to lB. The remaining effective charge of DNA is reduced
by a factor r:
r = 1 − 1 zη
(1.6)
Thus, it is predicted that the charge of DNA in a solution with monovalent counterions is
reduced by 76 % 31 . An important condition of the MO theory is that condensation occurs only
if the relation κ -1 >> a is satisfied (where κ -1 is the Debye length and a is the polymer radius),
which means either for highly diluted solutions or an infinitely thin DNA strand. Furthermore,
the theory is valid only for vanishingly small DNA concentrations.
For finite DNA concentrations a more complicated model needs to be used that is based on
solving of the PB equation. The PB equation describes the Gouy-Chapman (GC) model, where
a charged solid comes into contact with an ionic solution creating a double layer. Due to the
thermal motion of ions the counterion layer is a diffuse layer. The remaining charge around the
DNA molecules (ions in the third zone) is described by a linearized form of the PB equation –
the Debye-Hückel equation 35 that explains the relationship between the electrostatic behavior
of DNA and the ionic strength. In the equation the screening of DNA is quantified by the Debye
length κ -1 :
κ 2 = 8πl B N A I (1.7)
κ 2 = 2e2 N A I
εε 0 kT
(1.8)
where NA is the Avogadro number (6.022 × 10 23 1/mol) and I is the ionic strength.
The first quantitative experimental studies about conformational properties of DNA as a
function of salt concentration were done by Harrington 36 , who measured the DNA radius of
gyration in dilute DNA solutions. The DNA radius of gyration (Rg) depends on the DNA
persistence length (lp):
R g = √ Ll p
3
(1.9)
1.2 DNA 10

______________________________________________________________________ Introduction
where L is the contour length of DNA. The persistence length is a basic mechanical property of
a polymer and it is a measure of its stiffness. It is defined as a characteristic length over which
the chain maintains a certain direction 37 . When the contour length is smaller than the persistence
length a polymer behaves as a rigid rod, while a flexible coil behavior is observed for contour
lengths much higher than the persistence length. The persistence length of dsDNA is considered
to be around 50 nm 38 , while the persistence length of ssDNA is only 1-2 nm 39,40 . Thus, dsDNA
is expected to behave as a rigid rod and ssDNA as a flexible chain (Figure 1.5).
dsDNA
ssDNA
L = N bp × b = 20 × 0.34 nm
L = 6.8 nm
l p ≈ 50 nm
L ≪ l p (rigid rod)
L = N b × b = 20 × 0.43 nm
L = 8.6 nm
l p ≈ 1 to 2 nm
L > l p (flexible chain)
Figure 1.5. Dependence of the DNA mechanical properties on the persistence length
shown for an example of a 20-mer DNA strand. Nbp and Nb represent the number of base
pairs and number of bases, respectively, while b represents the charge separation.
Nevertheless, the persistence length consists of two contributions:
l p = l 0 + l el (1.10)
where l0 is an intrinsic stiffness due to chain properties and lel is the electrostatic repulsion
within the chain, which depends on the ionic strength:
l el =
l B
(2bκ) 2 (1.11)
Through the Debye length, the ionic strength influences the flexibility of charged polymers,
which should be considered especially while investigating the behavior of ssDNA in solution.
In solutions of increased ionic strength, the Debye length decreases leading to a decrease in the
1.2 DNA 11

______________________________________________________________________ Introduction
electrostatic component of the persistence length. In this case, the polymer behaves more as a
flexible coil (Figure 1.6). On the other hand, decreasing the ionic strength, that is, increasing
the Debye length, leads to an increase in the rigidity of a polymer 41 . Therefore, depending on
the ionic strength, ssDNA may also manifest high rigidity.
Figure 1.6. Dependence of the polymer rigidity on the ionic strength. Figure adapted from
ref. 41 .
1.2 DNA 12

______________________________________________________________________ Introduction
1.2.2 DNA in front of a polarized electrode
In addition to the DNA-ion interaction, it is important to understand the influence of a charged
electrode on the behavior of DNA on its surface for the application in DNA sensing at electrified
surfaces. The GC model of the double layer describes how the ionic strength and the
polarization of the electrode influence the double layer structure and the potential drop in front
of the electrode. The GC equation:
Φ = 2kT
e
ln 1 + γexp ( −d
κ −1)
1 − γexp ( −d
κ −1) (1.12)
γ = tanh ( eΦ 0
4kT ) (1.13)
where Φ0 and Φ are the potentials at the electrode surface and at a distance d from the surface,
respectively, reveals that the potential distribution strongly depends on the ionic strength, where
an increase of the ionic strength leads to a steeper drop of the potential (Figure 1.7, a). Thus, a
few nm away from the surface, Brownian motion prevails over electric forces and dominates
the system response 42 . Furthermore, the model predicts a sharp potential drop for highly
charged electrodes (high Φ0), while the decline is more gradual for lower Φ0 values 43 (Figure
1.7, b).
The DNA conformation on the electrode surface can be manipulated by externally applied
potentials 44,45 . However, this is true only under certain conditions 42 . Namely, as in solutions of
high ionic strength the applied potential decays within a nm distance, the range is too short to
significantly affect grafted DNA molecules. Therefore, both negative and positive potentials do
not affect the conformation of neither ds- nor ssDNA. dsDNA exhibits a rigid conformation,
while ssDNA coils on the electrode surface (Figure 1.8, a). This remarkable difference in
conformation originates from the difference in the persistence length of dsDNA and ssDNA, as
explained in Section 1.2.1. Furthermore, in presence of filler molecules with height comparable
to the length of the DNA spacer, the dsDNA conformation is almost perpendicular with respect
to the electrode surface 42,45 . The reason for the upright conformation is the steric repulsion
between the lowest base pairs and a monolayer that backfills the gold electrode between DNA
strands. In contrast, ssDNA remains lying on the surface since, besides the very weak electrical
interactions, self-repulsion along the DNA strand is suppressed by the high ionic strength.
1.2 DNA 13

______________________________________________________________________ Introduction
Figure 1.7. a) Potential distribution in relation to distance from the electrode for different
ionic strengths. Φ0 of 100 mV was used for the calculation. Figure adapted with
permission from ref. 42 . Copyright (2010) American Chemical Society. b) Potential profile
calculated for different Φ0 values and an ionic strength of 10 mM. Figure adapted from
ref. 43 .
In solutions of intermediate ionic strength, where the Debye length spans over few DNA base
pairs the influence of the applied potential is more significant, scaling with the distance from
the electrode surface (Figure 1.8, b). In this case, it is possible to manipulate the conformation
of dsDNA by applying positive (invoking a lying conformation) or negative potentials
(invoking up-right conformation). Still, control of dsDNA is achieved with less difficulty than
of ssDNA, due to their persistence lengths, while ssDNA manipulation is only partial depending
on the Debye length. While at a positively polarized electrode ssDNA will remain in the lying
conformation, negative potentials evoke at least partially an up-right orientation. The upper part
of the ssDNA is not exposed to the electric field and it exhibits a randomized conformation.
In solutions of low ionic strength charge screening is weak resulting in high charge repulsion.
Therefore, dsDNA is not stable under these conditions. Furthermore, applied electric potentials
are very long ranged, which leads to an efficient repulsion of the ssDNA from the surface
(Figure 1.8, c). Since the persistence length depends on the ionic strength, in solutions of low
ionic strength ssDNA exhibits a rigid conformation. Thus, combining these two effects, ssDNA
1.2 DNA 14

______________________________________________________________________ Introduction
has an up-right position when negative potentials are applied and a lying conformation at
positive potentials.
Figure 1.8. Schematic representation of DNA conformation on negatively and positively
charged surfaces for a) high, b) intermediate, and c) low ionic strengths. Figure adapted
with permission from ref. 42 . Copyright (2010) American Chemical Society.
1.2 DNA 15

______________________________________________________________________ Introduction
1.3 DNA immobilization
1.3.1 DNA immobilization approaches
A DNA immobilization strategy is determined by the substrate material used for the attachment.
Over the years, various surfaces were investigated for immobilization of DNA such as among
others the hanging mercury drop electrode, carbonaceous materials, boron-doped diamond,
silver, platinum and gold. Initially, research on DNA was conducted solely on mercury drop
electrodes (in the beginning of the 1970s) and carbon electrodes (since the middle of the
1970s 46 ). Later, gold electrodes became popular, as the chemisorption of thiol-tethered DNA
showed to be a very promising method for the preparation of DNA sensors.
With respect to the type of bond formed during immobilization, DNA immobilization methods
are characterized by three main mechanisms 47 : physisorption, covalent immobilization and
chemisorption.
Physisorption is the simplest immobilization method, which is based on the adsorption of
unmodified oligonucleotides on an electrode through electrostatic forces, van der Waals
interactions, hydrogen bonds and hydrophobic interactions. The immobilization occurs either
through nucleic bases (immobilization of ssDNA, Figure 1.9, a) or the phosphate backbone
(dsDNA, Figure 1.9, b). It is characterized by a multiple site attachment, which allows for the
investigation of direct DNA oxidation and reduction. The main drawback of this method is its
sensitivity on environmental changes (pH value, temperature, ionic strength) due to the weak
attachment. Furthermore, attachment of the DNA occurs via multiple points, which prevents
further hybridization due to the restricted configurational freedom of physisorbed DNA 48 .
Carbonaceous materials and the mercury drop electrode were mostly utilized for this
immobilization technique.
Covalent immobilization results in a much stronger binding between the surface and DNA
(Figure 1.9, c). Another advantage of this method is the appropriate orientation of the probe
DNA due to the end-point attachment of ssDNA, which facilitates hybridization. In order for
the immobilization to occur, the immobilization surface (chemically or electrochemically) and
the DNA itself need to be activated, which presents a drawback of this method. Activation of
1.3 DNA immobilization 16

______________________________________________________________________ Introduction
carbonaceous materials and glass substrates was extensively investigated for the covalent
immobilization of DNA.
Figure 1.9. Schematic representation of DNA immobilization by physisorption via a)
bases, b) the phosphate backbone, c) covalent immobilization, d) chemisorption, and e)
biotin-streptavidin immobilization.
Chemisorption of thiol tethered DNA on noble metals (typically gold) occurs by the selfassembly
process that is explained in Section 1.1. This method is widely used for the
preparation of DNA sensors due to its simplicity and the formation of a relatively strong Au-S
bond (Figure 1.9, d). Biotin-streptavidin immobilization was also employed for grafting of
DNA. It occurs through two steps, namely biotinylation of either the surface or the DNA
followed by binding of the streptavidin-modified counterpart (DNA or surface, respectively)
and formation of a very strong biotin-streptavidin conjugate (Figure 1.9, e). Nevertheless, since
the surface is usually biotinylated by chemisorption of biotin-terminated SAMs, the strength of
the modification is determined by the Au-S bond strength.
1.3 DNA immobilization 17

______________________________________________________________________ Introduction
1.3.2 Techniques for DNA microarray fabrication
A practical way of detecting and diagnosing various diseases is through the detection of nucleic
acid sequences, specific for any living organism 49 . One of the possible ways to gain insight into
a DNA sequence is by using DNA sensors. The sensitivity and selectivity of DNA biosensors
is highly dependent on the quality of the prepared DNA sensing surface. Good immobilization
technique ensures high reactivity, appropriate orientation, accessibility and the stability of the
grafted probe DNA and prevent unspecific binding 47 .
DNA immobilization methods can be divided into two groups 47,50 :
- Base-by-base synthesis (light-directed synthesis), which represents a bottom-up synthesis
of DNA sequences at the surface
- Direct attachment of already synthesized DNA sequences to the surface
Both strategies are used for DNA microarray fabrication. Several base-by-base synthesis
strategies were developed, including light-directed synthesis using photolithographic masks by
Affymetrix (GeneChips), photo-mediated synthesis by Roche (NimbleGen) that used digital
masks instead, and inkjet base-by-base manufacturing (printing) of DNA probes on the surface
developed by Agilent Technologies 51 . Among these, the Affymetrix strategy is the most known
and it will be shortly described here. On the other hand, among strategies for DNA array
fabrication by direct attachment of pre-synthesized DNA sequences, spotting (printing) is by
far the most known approach. However, new technologies for DNA array production are arising
on the market, including “electronic microarrays”.
Light-directed synthesis is a complex method requiring specialized equipment, however
attractive for DNA microarray fabrication. The principle of in situ DNA synthesis by
Affymetrix consists of UV masking and light-directed chemical synthesis of DNA sequences
directly at the array, one nucleotide at the time per spot for many spots simultaneously (Figure
1.10). The surface is initially modified with a covalent linker containing a protection group that
is easily removed upon irradiation. Using a photolithographic mask, desired spots on the surface
are irradiated, removing locally the protecting groups. Subsequently, these spots are modified
with the desired protected nucleotides. This experimental sequence is repeated as many times
as necessary to obtain the desired DNA sequences on the surface and this is determined by the
1.3 DNA immobilization 18

______________________________________________________________________ Introduction
length of the DNA. On average, for sequences with 25 bases, around 100 masks are needed per
chip. Nevertheless, the obtained chips can have an extremely high density (> 10 6 spots per
array). Main drawbacks of this technique are the complex nature of chemical synthesis and a
very expensive production 51 . Furthermore, difficult customization and the possibility of
synthesizing only relatively short DNA sequences are limiting the application of the technique
for production of point-of-care devices.
Figure 1.10. Schematic representation of the photolithographic synthesis of DNA
sequences on the electrode surface. a) surface modified with a covalent linker with a
protective group, b) irradiation of desired spots on the surface, c) modification of
deprotected spots with desired nucleotides, d-h) repetition of the process of irradiation
and modification with nucleotides to fabricate spots with desired DNA sequences.
The principle of spotting (printing) of already synthesized DNA sequences consists of spotting
nano- to picoliter volume drops of DNA-containing solution onto a predefined grid by a
1.3 DNA immobilization 19

______________________________________________________________________ Introduction
specialized ink-jet printer ran by a robot (Figure 1.11). By this, spots with a diameter of 100-
150 µm are created on the surface. The number of spots is limited to prevent crosscontamination.
Therefore, the density of these microarrays is moderate with 10,000 to 30,000
spots per array. Furthermore, this technique requires strict monitoring of the production
reproducibility and quality control. Difficulties with efficiency and accuracy are another
drawback of this technique 47,51 . Nevertheless, an advantage is that the content of the
microarrays is flexible.
Figure 1.11. Schematic representation of DNA solution spotting on the electrode surface.
Electronic microarrays use an electric field to control the immobilization of DNA. The
Company Nanogen developed a Nanochip with 12 connectors controlling 400 individual sites
on a chip. The principle of immobilization consists of the transport of negatively charged DNA,
modified with biotin, through an agarose permeation layer towards specific sites, modified with
streptavidin, on the chip where a positive current is applied 52 . Even though the density of these
chips is limited to 400 spots, this is sufficient for the majority of diagnostic applications 51 .
Furthermore, the content of the microarray can be specified by the user, which decreases
microarray manufacturing costs. Biotin-streptavidin chemistry offers a very strong bonding but
comes with some limitations. The synthesis of streptavidin-modified surfaces consists of
several steps, such as the activation of the surface, immobilization of streptavidin and blocking,
which increases the production costs and time 47 . Furthermore, one of the drawbacks of using
streptavidin is the problem with unspecific interactions.
1.3 DNA immobilization 20

______________________________________________________________________ Introduction
1.4 Hybridization detection
Compared to the traditional methods for the detection of DNA sequences, such as electrophoresis
or membrane blots, DNA sensors are faster, simpler and less expensive 53 , and
therefore, present a very active research field. Presently, a wide range of DNA sensor
technologies are in development or being commercialized.
DNA sensors are based on the specific detection of a DNA sequence using a so-called probe
DNA immobilized on the surface 49 (Figure 1.12). Therefore, two main parts of a DNA sensor
are 48 :
- biorecognition interface (immobilized specific DNA sequence) that allows the detection
of a target element (complementary DNA sequence)
- transducer that transforms a detected signal into a readable output
Figure 1.12. Principle of a DNA sensor.
The selective binding, followed by conformational and structural changes in the recognition
layer, can be detected by various techniques. Depending on the transduction approach, DNA
biosensors can be optical, piezoelectric and electrochemical.
Optical DNA sensors based on fluorescence are very sensitive and DNA chips with this
detection scheme have already been commercialized 51 . One of the optical readout strategies is
the detection of an intercalating dye (e.g., ethidium bromide) upon hybridization. Furthermore,
the use of molecular beacons was explored extensively (Figure 1.13). The ends of a molecular
1.4 Hybridization detection 21

______________________________________________________________________ Introduction
beacon are complementary and form a loop through self-hybridization. Coupling of a
fluorophore and a quencher on opposite ends of a beacon results in quenching of fluorescence
when the beacon is closed and a fluorescence signal when the beacon is opened due to
hybridization with a target DNA. Another optical technique commonly used for DNA detection
is surface plasmon resonance, based on the change in the refractive index of a thin metal
substrate modified with a biorecognition interface upon hybridization. Nevertheless, the
required equipment for optical detection is sophisticated and relatively expensive and therefore
not suitable for clinical purposes and point-of-care diagnostics but rather for laboratory
applications. Furthermore, inconsistent yields of target synthesis and labelling as well as nonuniform
rates of photobleaching can result in insufficient readout accuracy required for patient
diagnosis 54 .
Figure 1.13. Optical DNA detection using molecular beacons.
Piezoelectric DNA sensors rely on the mass change upon hybridization with a target DNA that
is correlated with an increase in the fundamental resonance frequency of a crystal 55 . Quartz
crystal microbalance is the most commonly used technique for the real-time monitoring of the
hybridization process 56 . Different strategies to increase the mass change upon hybridization
were used in order to increase the sensitivity, usually by using bulky labels 57,58 .
Point-of-care diagnostics requires fast response, high sensitivity and selectivity, and operation
simplicity. These, coupled with portability, low cost and possibility of miniaturization, are
characteristics of electrochemical DNA sensors, which makes them very attractive for mass
production. Unlike bulky optical readout systems, electrochemical detection can be integrated
1.4 Hybridization detection 22

______________________________________________________________________ Introduction
on a chip 4 . Since electrochemical detection results directly into an electric signal, there is no
need for an expensive transduction equipment 56 . Electrochemical transduction of the
hybridization event can be divided into direct and indirect DNA detection.
Direct DNA detection relies either on changes in the electrical properties of the interface caused
by hybridization or a direct oxidation of nucleic bases 48 . In the 1960s Paleček and coworkers
pioneered the work on direct reduction and oxidation of DNA at a mercury electrode 46 (Figure
1.14). Later, the oxidation of purine bases of DNA was achieved using carbon, gold, indium tin
oxide and polymer-coated electrodes 56 . Since guanine is the most redox active from all DNA
bases, its oxidation was studied the most. Even though this method is quite sensitive, its main
drawback are high background currents due to the high potentials required for direct DNA
oxidation.
Figure 1.14. Reduction (red) and oxidation (blue) sites of DNA bases for direct DNA
electrochemical detection. Figure adapted with permission from ref. 59 . Copyright (2012)
American Chemical Society.
Electrochemical impedance spectroscopy (EIS) was extensively used to follow changes in the
interface properties upon hybridization. The method measures the change in the faradaic
impedance in the presence of redox species resulting from the hybridization event. Different
strategies have been used for the amplification of the signal.
Indirect DNA detection can be achieved through the use of electrochemical mediators, redox
active indicators that non-covalently interact with dsDNA, labelling of the probe or target DNA
or using sandwich type assays (Figure 1.15).
1.4 Hybridization detection 23

______________________________________________________________________ Introduction
DNA detection via redox mediators was employed using various mediators such as
[Ru(bpy)3] 2+ , [Co(phen)3] 3+ , ferrocene or [Fe(CN)6] 3-/4- (Figure 1.15, a). Catalytic oxidation of
guanine using [Ru(bpy)3] 2+ on indium tin oxide (ITO) electrodes is a well-known example 60 .
The approach involves addition of DNA into a solution containing the ruthenium complex,
while the electrode is held at a potential suitable for the oxidation of the reduced form of the
complex. The complex is regenerated by the oxidation of guanine from DNA. Consequently,
the signal is enhanced proportionally to the amount of guanine available for oxidation, since
the direct guanine oxidation is not possible at ITO electrodes.
The characteristics of non-covalently interacting indicators are a different affinity towards dsand
ssDNA. They can interact with DNA either by electrostatic binding, binding to the groove
of dsDNA or intercalate into dsDNA (Figure 1.15, b). The most widely used indicators are
daunomycin, proflavine, antraquinone, methylene blue, [Co(phen)3] 3+ and [Ru(NH3)6] 3+ . One
of the initial studies on non-covalent interaction of compounds with DNA was done by
Mikkelsen et al., using [Co(phen)3] 3+ as a dsDNA minor groove binder. The metal complex is
positively charged and attracted by negatively charged DNA, resulting in a higher current for
the more negatively charged dsDNA 61 . Furthermore, Barton and coworkers pioneered the work
on long-range charge transfer resistance through DNA, using electrochemically active
intercalators 62 (methylene blue and daunomycin). Interaction of non-electrochemically active
intercalating compounds was investigated using EIS 63 .
First covalently bound DNA markers were investigated in the beginning of the 1980s 46 .
Labelling of the DNA can be performed using the probe DNA, where the label is positioned at
the distant end of the probe (Figure 1.15, c). Due to the flexibility of the ssDNA, without the
presence of target DNA, the label is close enough to the surface and the signal is detected. Upon
hybridization the signal switches off due to the rigidity of the dsDNA. Namely, after the
hybridization the label is too far away from the surface for the electron transfer to occur.
Labelling of the target DNA is also possible, and it is usually done at the end that is close to the
surface upon hybridization (Figure 1.15, d). Therefore, after hybridization the signal is switched
on. However, this approach requires a preparation step in which the target DNA from the sample
needs to be labelled, which prolongs the assay time. In a sandwich-type assay, immobilized
probe DNA is initially hybridized with a portion of the non-labelled target DNA from the
sample, and subsequently a labelled signal DNA is hybridized with the overhang of the target
DNA (Figure 1.15, e). Enzymes are readily used as labels for this approach 64 . In this case, DNA
1.4 Hybridization detection 24

______________________________________________________________________ Introduction
detection can be performed using a redox mediator in solution. An alternative approach is to
hybridize a signal DNA with an overhang from the target DNA that is placed close to the
surface 65 . This way there is no need for the redox mediator to obtain the signal.
Figure 1.15. Indirect DNA detection using a) electrochemical mediators, b) non-covalently
interacting indicators, c-d) labelling of probe or target DNA and e) sandwich type
detection.
1.4 Hybridization detection 25

2. Aims of the Work

_________________________________________________________________ Aims of the Work
Self-assembly finds its application in many different fields of research, yet the most
used protocol is the same in all of them – simple immersion of the material of choice into a
solution containing the molecule to be assembled, followed by a spontaneous adsorption and
formation of appropriate bonds 66 . Among SAMs self-assembly of thiolated molecules on gold
is the most known.
While, for example, the aim of SAM formation of alkylthiols is generally to reach highly
compact monolayers, the immobilization of thiolated DNA is performed aiming at a wide range
of DNA coverages, depending on the envisaged sensing strategy. Nevertheless, in order to
obtain high coverages of thiolated molecules a long incubation time is required, ranging from
several hours to days 13-15 . In contrast, low coverages can be obtained in a shorter time, however,
with a significant variation of densities and substantially decreased reproducibility 67 .
To meet the demands of the market, future point-of-care devices have to link high-quality
performance with speed, simplicity and low production costs 4 . Despite its simplicity of
formation of comparably stable films, thiol chemisorption on gold is not yet the chemistry of
choice for fabrication of commercial DNA biosensors, partly due to issues with reproducibility
and the time required for modification. In order to profit from possibilities that self-assembly
offers, the dependence on the spontaneous adsorption process, that is very slow and lacks
reproducibility, needs to be eliminated. A new and simple strategy that will allow to
reproducibly control the surface modification in a desired manner, and what is equally
important, in a very short time, would significantly decrease the production costs and with this
the cost of a final point-of-care device.
Base-by-base DNA sequence synthesis made a tremendous contribution to fabrication of highly
dense DNA chips for genome sequencing. However, the complex nature of chemical synthesis
and very expensive production, coupled with a limited flexibility for customization and length
of probe sequences, makes this technique less suitable for point-of-care devices 51 . Moreover,
the very high number of individual test sites is often not necessary for specific measurements
of point-of-care applications. Furthermore, the more flexible and cheaper approach, namely
spotting of pre-synthesized DNA sequences, still suffers from several drawbacks, such as
special working conditions and problems with accuracy and efficiency. Nevertheless, new
approaches for the production of DNA chips are arising, including electrochemically driven
surface modification, demonstrating the evolution of microarray technology to more practical
platforms for diagnostic applications.
27

_________________________________________________________________ Aims of the Work
This thesis focuses on the development of a new strategy for the gold surface modification by
thiolated molecules that will enable for a controlled immobilization of molecules, regardless of
their charge in a fast manner. The main envisaged application of the proposed strategy is in the
production of DNA chips as means for overcoming limitations in the development of point-ofcare
devices.
The initial part of this work focuses on understanding the processes occurring at the interface
during surface modification, taking into consideration the interdependence of physico-chemical
properties of the investigated molecules, electrode polarization and the surrounding solution,
with the goal of finding strategies to improve the kinetics and the reproducibility of immobilization
of both intrinsically charged molecules, such as DNA, and uncharged molecules, such
as alkylthiols. The influence of potential pulses on the surface modification is explored as a
new strategy to obtain high-quality DNA sensing platforms, and the advantages of this approach
over the standard passive approach are investigated. Moreover, this work focuses on
implementing the developed strategy into the fabrication of a DNA array for multiple probe
detection. The potential pulse-assisted cleaning of Au modified surfaces is also investigated
with the aim to regenerate the Au surfaces within a very short time, while not causing any
damage to the electrode surface. The last part of the thesis focuses on the implementation of the
newly established surface modification strategy into the development of a new DNA sensing
platform based on the signal amplification via enzyme-conjugated intercalating compound as a
hybridization indicator. Challenges related to the use of intercalators are mainly reflected in the
need to prevent non-specific adsorption to obtain a desired contrast between ss- and dsDNA.
Therefore, the ability of the developed surface modification strategy to create high-quality DNA
sensing surfaces for this application is evaluated.
28

3. Results and
Discussion

_____________________________________________________________
Results and Discussion
3.1 Importance of preparing the surface. Criteria for cleanliness
Building of the automatic polishing machine was the author’s idea. The machine was
constructed by Dr. Kirill Sliozberg. Optimization of the mechanical cleaning procedure was
done by the author.
Even though the surface preparation is cardinally important for the quality and reproducibility
of the measurements especially in the field of biosensors and self-assembly, it is often
underestimated and neglected. Due to the nanoscale size of the DNA strands and its charge,
surface architecture is a parameter that can significantly increase the sensitivity of a DNA
sensor 68 . Depending on the surface pDNA loading can be higher or lower.
Even though the interest of the thesis was not to investigate the influence of the surface
roughness and the size of the furrows on the biosensor performance, special attention was given
to the surface cleaning with the aim to improve the reproducibility of measurements. The
preparation of electrodes consisted of several steps:
- Mechanical cleaning using an automatic polishing machine
- Electrochemical cleaning in H2SO4
- Characterization of the surface by electrochemical impedance spectroscopy (EIS)
- Characterization of the surface by cyclic voltammetry (CV)
By reproducibly preparing the surfaces using an automatic polishing machine (polishing setup
and procedure are explained in Section 5.3) the quality of the prepared electrodes was
significantly improved in comparison with a previously used manual polishing procedure
(Figures 3.1 and 3.2). Moreover, a reproducible roughness was achieved as one of prerequisites
for minimal signal deviation between experiments. This allowed exclusion of the surface
properties as one of the possible parameters affecting the already complex results.
3.1 Importance of preparing the surface. Criteria for cleanliness 30

_____________________________________________________________
Results and Discussion
Figure 3.1. Optical microscope images of the electrode surface for a) a new electrode as
received from a company, b) an electrode after polishing by hand with 3 µm, and c) 3 µm
and 1 µm diamond polishing pastes.
Figure 3.2. Optical microscope images of an electrode surface after cleaning with an
automatic polishing machine following each cleaning step using: a) 3 µm, b) 1 µm, c) 0.5
µm and finally d) 0.1 µm diamond polishing pastes.
3.1 Importance of preparing the surface. Criteria for cleanliness 31

_____________________________________________________________
Results and Discussion
In addition to the roughness factor (for its determination see Section 5.3), the criterion for
efficiency of electrochemical cleaning in H2SO4 was the shape of the CV. The electrode surface
was considered clean only if the oxidation peaks of individual gold facets 69 were observed upon
electrochemical cleaning as shown in Figure 3.3.
Figure 3.3. Representative CV of a clean gold electrode. Electrochemical cleaning was
performed in 0.5 M H2SO4 at a scan rate of 100 mV/s.
Following the mechanical and electrochemical cleaning of the electrode surface, EIS and CV
using the ferro/ferricyanide couple as redox probe were used to verify the electrode cleanliness
and to allow for a comparison of separate electrodes. Afterwards, criteria were selected for the
desired charge transfer resistance (Rct) derived from EIS and the peak current in the CV as a
prerequisite for the use of prepared electrodes for subsequent surface modification.
Representative EIS and CV of clean gold electrodes are shown in Figure 3.4.
Modification of electrodes was performed immediately after mechanical and electrochemical
cleaning and characterization of the bare electrodes since it was observed by EIS that upon
storage of bare gold electrodes in air, H2SO4 or HClO4 unknown contaminants adsorb on the
electrode surface (data not shown).
3.1 Importance of preparing the surface. Criteria for cleanliness 32

_____________________________________________________________
Results and Discussion
Figure 3.4. a) EIS and b) CV of a clean 2 mm diameter gold electrode performed in
K3[Fe(CN)6]/K3[Fe(CN)6] (5 mM each) in 10 mM phosphate buffer (PB), 20 mM K2SO4.
EIS measurements were performed by applying a DC potential of 220 mV superimposed
by a 5 mV AC perturbation. The frequency range was scanned from 30 kHz to 10 mHz.
CV was performed at 100 mV/s scan rate.
3.1 Importance of preparing the surface. Criteria for cleanliness 33

_____________________________________________________________
Results and Discussion
3.2 Importance of knowing the surface
In order to tailor optimal DNA-modified surfaces for the envisaged sensing platforms it is of
utmost importance to understand processes occurring at the electrode surface during the DNA
assay build-up. Only in this way, the properties of the surface can be controlled in a reproducible
manner. Electrochemical impedance spectroscopy is an excellent technique for this due to
which it gained a wide popularity in the field of bioelectroanalysis 70,71 . It is a non-destructive
technique allowing one to monitor surface modification without altering the system’s response.
Moreover, due to its sensitivity it is a great tool for following very subtle changes in the surface
architecture. EIS is a very informative technique that allows in depth investigation of processes
occurring at the electrified interface by sequentially following each step of the build-up of DNA
assays 5,72 .
Furthermore, to understand the behavior of DNA strands in front of an electrode surface we
need to investigate not only physico-chemical properties of the DNA itself but it is essential to
also observe the electrode and the surrounding solution as important components of the system.
Since the DNA is essentially a negatively charged polyelectrolyte (depending on the ionic
strength of the surrounding solution), the charge of the surface has a significant impact on the
behavior of DNA at the interface. Therefore, it is of great importance to know how the surface
is polarized (positive or negative) upon application of different potentials. Consequently, the
potential of zero charge (pzc) of the bare polycrystalline electrode was determined.
Additionally, the influence of the surface modification with DNA on the pzc was also
investigated.
3.2 Importance of knowing the surface 34

_____________________________________________________________
Results and Discussion
3.2.1 Electrochemical impedance spectroscopy. DNA assay build-up
Electrochemical impedance spectroscopy is based on applying a DC potential that is commonly
an open circuit potential superimposed with an AC potential of small amplitude. Measuring of
the resulting AC current signal allows sampling of modulus and phase of the response. The
impedance is usually presented by plotting the real and imaginary components in a Nyquist
plot. The principle of the method is explained in detail in Section 5.13.1.
In this study, a modified Randles equivalent electric circuit was used for modelling the behavior
of the electrode surface during each step of the DNA assay build-up (Figure 3.5). In the circuit,
a solution resistance (Rs) is connected in series with a constant phase element (CPE), which
represents the double layer by taking into account the roughness of a polycrystalline gold
electrode, and an impedance of a faradaic reaction (consisting of a charge transfer resistance,
Rct, and a Warburg element representing the semi-infinite linear diffusion of electroactive
species to a flat electrode, W). Since the alteration of Rct is most pronounced as compared to
other electric circuit elements, the change of Rct was followed during surface preparation.
Figure 3.5. Randles equivalent electric circuit used for fitting of Nyquist plots obtained
during the DNA assay preparation. Rs represents the solution resistance, Rct is the charge
transfer resistance, CPE is a constant phase element representing the double layer and W
is the Warburg element representing diffusion.
The interpretation of the properties of DNA-modified gold electrode surfaces by means of EIS
was based on the repulsion of a negatively charged free-diffusing redox mediator, namely
[Fe(CN)6] 3-/4- , from the DNA-modified electrode in a solution of moderate ionic strength 3 . The
extent of the DNA charge screening depends on the ionic strength of the solution above a certain
3.2 Importance of knowing the surface 35

_____________________________________________________________
Results and Discussion
threshold 34 . Thus, DNA is screened more in solutions of higher ionic strength, decreasing by
this the ability of the DNA to repel molecules of the redox mediator. Therefore, in order to
achieve the desired sensitivity, the ionic strength of the working solution should be low enough
to allow for DNA to manifest high enough effective negative charge to significantly block the
redox mediator. On the other hand, the ionic strength needs to be high enough to not affect the
stability of the double helix during measurements with dsDNA-modified electrodes.
Figure 3.6 represents a schematic view of the electrode surface during the preparation of a DNA
sensor via a two-step immobilization method. Initially, a thiol-tethered ssDNA is immobilized
on the electrode creating a negatively charged interface (Figure 3.6, a and b). Consequently, the
approach of the redox mediator is hindered and the electron transfer rate decreases. In EIS this
is observed as an increase in Rct (Figure 3.7). The increase in Rct depends on the amount of
immobilized DNA, where a higher increase is observed for a higher ssDNA coverage. It should
be noted that the obtained EIS response is a result of the repulsion of the redox mediator by
both negatively charged immobilized ssDNA and unspecifically adsorbed DNA strands 73 , as
well as steric hindrance caused by lying ssDNA that physically blocks the access of the redox
mediator. Due to the fact that ssDNA behaves as a flexible coil and that it orientates randomly
on the electrode surface, especially at lower ssDNA coverage this response lacks in
reproducibility.
Figure 3.6. Scheme of the electrode surface during the build-up of the DNA sensor: a)
bare electrode, b) ssDNA-modified electrode, c) ssDNA/thiol-modified electrode.
Therefore, upon immobilization the electrode surface is covered with a mixture of DNA strands
that are chemisorbed via a Au-S bond and DNA strands that are bound to the surface through
the DNA backbone or bases. Additionally, grafted DNA strands also adsorb to the electrode by
3.2 Importance of knowing the surface 36

_____________________________________________________________
Results and Discussion
coiling on the surface due to their persistence length and flexible behavior. This leads to a steric
hindrance towards the hybridization with target DNA. Thus, in the next step of the DNA sensor
the build-up the ssDNA-modified surface is passivated using self-assembling properties of thiol
molecules (Figure 3.6, c), as originally proposed by Tarlov et al. 74,75 . The formed thiol
monolayer removes unspecifically adsorbed ssDNA from the electrode surface and forces
grafted DNA to obtain an orientation that is more favorable for hybridization. It should be noted,
that the DNA still does not obtain a perpendicular position with respect to the surface but rather
lies down on the thiol layer due to its flexibility. Consequently, since the flux of the redox
mediator to the electrode surface is facilitated, a decrease in Rct is commonly observed in EIS
measurements upon the passivation step (Figure 3.8). On the other hand, tDNA can also
unspecifically bind on the bare electrode surface (data not shown) and therefore this step is very
important for preventing the unspecific adsorption of the target DNA and with this a false
signal.
Figure 3.7. Typical Nyquist plots representing the Rct increase upon ssDNA immobilization.
ssDNA immobilization was performed in 10 mM PB with 450 mM K2SO4
containing 1 µM ssDNA by incubation at 37 °C. EIS was measured in 10 mM PB with 20
mM K2SO4 containing equimolar concentrations (5 mM) of K3[Fe(CN)6] and K4[Fe(CN)6].
A DC potential of 220 mV was applied superimposed with a 5 mV AC perturbation. The
frequency range was from 30 kHz to 10 mHz.
3.2 Importance of knowing the surface 37

_____________________________________________________________
Results and Discussion
Taking into account that the reproducibility of the immobilization step as detected by EIS is
quite low, it is more informative to observe the ssDNA/thiol-modified surface. However, since
this response contains the contribution of both passivation and DNA immobilization, it should
be compared with the Rct obtained for an only thiol-modified electrode. Namely, Rct of the
ssDNA/thiol-modified electrode needs to be higher than the Rct of the electrode modified only
with the passivating thiol implying that a detectable amount of DNA was immobilized.
Figure 3.8. Change in the Rct value upon surface passivation with MCH. ssDNA
immobilization was performed for 15 min as stated in Figure 3.7. MCH passivation was
done by incubation for 19 h in 10 mM PB containing 20 mM K2SO4 and 10 mM MCH.
EIS measurements were performed as stated in Figure 3.7.
Therefore, after ssDNA immobilization by incubation for 15 min and subsequent passivation,
the obtained Rct is negligibly higher as compared to the Rct of the electrode modified only with
the mercaptohexanol (MCH) for a time equal to the passivation duration (Figure 3.9). This
implies that the amount of immobilized DNA is very low and not detectable by means of EIS.
On the other hand, for longer immobilization times (2 h and 8 h) a significant increase in the
Rct value is observed as compared to the thiol-modified electrode reflecting higher DNA
coverages.
3.2 Importance of knowing the surface 38

_____________________________________________________________
Results and Discussion
One of advantages of the electrochemical impedance spectroscopy is that it enables label-free
detection of DNA hybridization, circumventing the need for dye, enzyme or redox labelling of
the target DNA 71 . However, a careful design of the electrode architecture needs to be achieved
to improve the sensitivity of the overall system. Upon DNA hybridization an increase in Rct is
generally observed. However, there are different interfacial phenomena caused by the
hybridization process that influence the access of the redox mediator to the surface.
Figure 3.9. Comparison of the Rct obtained for a MCH-modified electrode with the Rct
obtained for ssDNA/MCH-modified electrodes using different incubation times. The
MCH-modified electrode was obtained by incubation of the bare gold electrode in 10 mM
PB containing 20 mM K2SO4 and 10 mM MCH for 19 h. The ssDNA/MCH-modified
electrode was prepared by initial ssDNA immobilization followed by subsequent MCH
passivation (see Figures 3.7 and 3.8, respectively). EIS measurements were performed as
stated in Figure 3.7.
EIS measurements are performed in an intermediate ionic strength to assure that the DNA is
not fully screened by counterions and that it still manifests some effective negative charge.
Therefore, as a result of the hybridization, the amount of negative charge in front of the
electrode surface increases leading to a stronger repulsion of the redox mediator and an increase
3.2 Importance of knowing the surface 39

_____________________________________________________________
Results and Discussion
of the Rct. Moreover, the presence of additional DNA strands leads to an additional steric
hindrance decreasing the access of the redox mediator and contributes to an increase in Rct with
a strong dependence on ssDNA coverage. When the coverage is very low, additional DNA
strands do not notably alter the surface architecture and the redox mediator access is negligibly
hampered. At high ssDNA coverage which is however low enough to allow for maximum
hybridization efficiency, the approach of the redox mediator to the electrode surface is more
likely affected.
Nevertheless, due to the low persistence length of ssDNA it behaves as random coil lying on
top of the thiol passivation layer (Figure 3.10) still preventing the access of the redox mediator
to the electrode surface. The structure of the formed dsDNA is more rigid as compared to the
one of the ssDNA, leading to a more upright orientation of the dsDNA and facilitating the
approach of the redox species. The degree of this influence depends on the length of DNA
strands. Short DNA is expected to exhibit only a little impact while longer DNA causes a more
significant change of the electrode architecture.
Figure 3.10. Interfacial changes upon DNA hybridization. DNA obtains a more upright
orientation due to the rigidity of the dsDNA. Furthermore, the amount of the negative
charge increases and an additional steric hindrance arises due to an increased number of
strands. Thiol passivation layer was intentionally not shown for better clarity.
3.2 Importance of knowing the surface 40

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Results and Discussion
3.2.2 Potential of zero charge of bare and DNA-modified electrodes
Theoretical discussions were done with Dr. Fabio La Mantia. Parts of this section were
published in ref. 5 : “D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, Angew. Chem. Int.
Ed. 2015, 54, 15064-15068; Angew. Chem. 2015, 127, 15278-15283.”.
The potential of zero charge is a fundamental property of an electrode-electrolyte interface.
Therefore, for the in-depth understanding of the surface behavior it is of high importance to
know the value of the pzc. At the pzc the excess charge density on the electrode surface equals
zero 76,77 . The term “potential of zero free charge” (pzfc) is also used for the same value in order
to differentiate it from the so-called potential of zero total charge (pztc) that refers to the case
when a specific adsorption occurs on the interface involving charge transfer. The pztc is defined
as a potential at which the sum of free electronic excess charge density and charge density
transferred in the adsorption processes equals zero 76 . In order to better understand these terms,
three different situations at the electrode-electrolyte interface are shown in Figure 3.11, when
the applied potential equals pz(f)c, pztc or any other value.
In the literature there is a substantial variation of reported values for the pzc of gold electrodes 78 .
Due to the polycrystalline nature of the Au surface, the use of different electrolyte solutions and
the fact that the pzc is very sensitive to surface impurities, many different values were proposed.
Therefore, we determined the pzc for our system with thoroughly cleaned Au electrodes in the
same solution that was later used for the modification of the surface.
The most common way to determine the pzc is to find the minimum of the differential
capacitance in the Cd-E curve, predicted by the GC theory 77 . It should be noted that for systems
in which specific adsorption occurs the capacitance minimum provides the value of the pztc.
Since we used a solution that contains sulfates, the obtained value evidently represents the pztc.
However, for convenience of the readers, the general term pzc is used in the following text.
Based on the definition, the pzc can be described by equation 3.1:
σ m = dC d
dE = 0 (3.1)
3.2 Importance of knowing the surface 41

_____________________________________________________________
Results and Discussion
where σm represents the excess charge of the metal and Cd the differential capacitance. This
equation explains that at the pzc the Cd-E curve shows a minimum. In order to find this potential,
we need to evaluate the nature of the differential capacitance and on what parameters it depends.
surface at the potential of zero (free)
charge
σ m = 0
σ i = 0
σ ddl = 0
surface at the potential of zero total
charge
σ m + σ i = 0
σ ddl = 0
surface at any other potential
σ m + σ i + σ ddl = 0
Figure 3.11. Schematic representation of the electrode-electrolyte interface upon
polarization at different potentials. σm, σi and σddl stand for the excess charge in the metal,
compact layer and diffuse double layer, respectively.
3.2 Importance of knowing the surface 42

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Results and Discussion
According to the GC model of the double layer 43 , Cd consists of two contributions (Figure 3.12,
a), namely the capacitance of the compact layer (Ci) and the capacitance of the diffuse double
layer (Cddl):
1
= 1 + 1
(3.2)
C d C i C ddl
Figure 3.12. Schematic representation of a) potential profile at the electrode surface, and
b) expected behavior of Cd in dependence on the applied potential and the ionic strength
according to the GCS theory. Φ represents the potential distribution and d distance from
the electrode surface. Figures adapted from 43 .
This means that Cd is determined by the lower value of these two components. In theory, Ci
does not depend on the ionic strength nor on the potential. On the other hand, Cddl depends on
both of these values under certain conditions. When the applied potential is far away from the
pzc, Cd is determined by Ci regardless on the ionic strength. However, when the potential
approaches the pzc (Figure 3.12, b), Cddl becomes significant at high ionic strength and the term
1/Cddl in Equation (3.2) becomes negligible. In that case, Cd is again determined by Ci and only
a constant capacitance of Ci is observed. However, at very low ionic strength Cddl is small and
the term 1/Cddl becomes significant as compared with 1/Ci. Cd is then determined by Cddl, which
is represented as a minimum in the Cd-E curve. It should be noted, that this model is based on
certain assumptions like the independence of Ci on the ionic strength and the potential, that
result in some discrepancies in real systems. In reality, a Cd-E curve does not necessarily follow
the predicted ideal U-shape.
3.2 Importance of knowing the surface 43

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Results and Discussion
In order to find the capacitance minimum in the Cd-E curve, potentiodynamic electrochemical
impedance spectroscopy (PDEIS) was used. This technique consists of a set of EIS
measurements performed for a desired range of potentials. Initially, EIS was performed for all
solutions to identify the suitable frequency range for the determination of the pzc. This was
done to verify whether the system can be fitted to the RC equivalent circuit and that there are
no artifacts in the frequency range of interest. Very high frequencies were excluded since a high
frequency region of the impedance spectra is very sensitive to the cell geometry 79 . To be
specific, the solution resistance is smaller for the electrode edge than for the center of the
electrode, leading to a mixed behavior of the system which cannot be fitted with a simple RC
circuit. In order to minimize this behavior, a setup that assures homogeneous distribution of
current lines was constructed, where the counter electrode completely surrounds the working
electrode and the reference electrode is placed below them (Chapter 5.4, Figure 5.4). Moreover,
very low frequencies were also avoided and with this any possible faradaic side-reactions.
Furthermore, a capacitive bridge (shunt capacitor 80 ) was implemented and a Pb/PbF2 reference
electrode was used that shows high impedance 81 in the measuring setup to eliminate possible
instrumental artifacts due to very low ionic strengths and high frequencies. After these
optimization steps Nyquist plots with the characteristic behavior for a system with the solution
resistance and the double layer capacitance connected in series (Figure 3.13) were obtained. In
an ideal case, a Nyquist plot of this system should consist of a vertical line with the intercepting
the real axis at the solution resistance. However, as shown in Figure 3.13 a frequency dispersion
which is common for polycrystalline electrodes is observed. There are various theories
discussing the reasons for this phenomena, among which are the roughness of polycrystalline
electrodes, that creates inhomogeneities in the current density along the surface, or adsorption
effects 77,82 .
After performing PDEIS for all solutions in the chosen frequency range and the potential range
from -0.2 to 0.7 V (vs. Ag/AgCl/3 M KCl), the determination of Cd was done by extracting the
imaginary impedance component from the Nyquist plots for each sampled potential value and
calculating Cd values using the following equations for the RC equivalent electric circuit:
Z = R −
j
ωC d
(3.3)
3.2 Importance of knowing the surface 44

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Results and Discussion
C d = 1
(3.4)
2πfZ
′′
where ω is the angular frequency.
Figure 3.13. Example of a Nyquist plot used for the system characterization and
calculation of Cd. The presented curve was recorded for a potential of 0.8 V vs. Pb/PbF2/5
M KF (0.2 V vs. Ag/AgCl/3 M KCl) in the frequency range from 4 kHz to 100 Hz. The
measurement was performed in 10 mM PB with 20 mM K2SO4. A capacitive bridge (2 µF)
was used. The beige line represents the fit to the RC equivalent circuit.
Figure 3.14 shows the obtained Cd values for three electrolyte solutions of different ionic
strengths calculated at three different frequencies. Comparing different ionic strengths, no clear
minimum is observed (Figure 3.14, a-c). On the contrary, for the highest ionic strength a
maximum was observed in one part of the curve. This behavior was previously observed for
high ionic strength electrolytes 83 . Furthermore, looking at the individual ionic strength, the
same trend can be observed for all three chosen frequencies. However, when different ionic
strengths for the same frequency are compared (Figure 3.14, d-f), the biggest difference is
observed between different ionic strengths for a frequency of 100 Hz (Figure 3.14, d).
3.2 Importance of knowing the surface 45

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Results and Discussion
Figure 3.14. Determination of Cd in at varying potentials of the bare gold electrode. Cd
was measured from the imaginary impedance component for different ionic strengths: a)
10 mM PB with 20 mM K2SO4, b) 1 mM PB with 2 mM K2SO4 and c) 0.1 mM PB with 0.2
mM K2SO4 and for different frequencies: d) 100 Hz, e) 200 Hz and f) 500 Hz.
3.2 Importance of knowing the surface 46

_____________________________________________________________
Results and Discussion
Taking into account Equation (3.2), it is assumed that the investigated ionic strengths are not
low enough to result in the desired capacitance minimum from the Cd-E dependence alone.
Namely, for the highest ionic strength (10 mM PB, 20 mM K2SO4) Ci dominates the overall
capacitance response, however, at the lowest investigated ionic strength (0.1 mM PB, 0.2 mM
K2SO4) obviously contributions from both Ci and Cddl are significant. Therefore, by subtraction
of these two values Cddl can be obtained:
1
C ddl =
1
C
− 1 (3.5)
d C i
Finally, by plotting the Cddl-E dependence (Figure 3.15), the pzc can be determined from the
capacitance minimum at a potential of 0.5 V (vs. Ag/AgCl/3 M KCl), where the most
pronounced minimum is observed for a frequency of 100 Hz. The non-ideal U-shape of the
obtained curves is probably a consequence of the polycrystalline nature of the gold surfaces and
the contribution of all crystallographic facets.
Figure 3.15. Determination of the pzc of the bare gold electrode. Cddl values were
calculated using the data from Figure 3.14 for different frequencies.
To investigate the effect of the modification of the surface with DNA on the pzc value,
determination of the pzc (pztc) of the DNA-modified electrode (pzc (DNA)) was performed.
This was done in the same manner as for the pzc determination of bare gold, except that Ci was
3.2 Importance of knowing the surface 47

_____________________________________________________________
Results and Discussion
replaced by the capacitance of the DNA monolayer (CDNA). Cd is represented by the following
equation:
1
= 1 + 1
(3.6)
C d C DNA C ddl
Upon DNA immobilization, Cd (Figure 3.16) and subsequently Cddl values (Figure 3.17) were
obtained in relation to the applied potential using PDEIS. In this case, the investigated potential
range was shifted to -0.4 V to 0.5 V (vs. Ag/AgCl/3 M KCl) to avoid the cleavage of the Au-S
bond at higher potentials. The pzc (DNA) was determined to be around 0.1 V (vs. Ag/AgCl/3
M KCl) showing that the pzc shifts to more negative values due to the surface modification
with DNA. Unlike in the case of the bare polycrystalline gold electrode, for the DNA-modified
electrode a well-defined minimum was observed. A possible explanation is a more
homogeneous surface after the modification with the DNA layer. Furthermore, the broad
minimum is in agreement with the GC double layer model that assumes peak broadening with
decreasing ionic strength 43 .
Figure 3.16. Determination of Cd at varying applied potentials at the DNA-modified
electrode. Impedance was measured for two ionic strengths: 10 mM PB with 20 mM
K2SO4 and 0.1 mM PB with 0.2 mM K2SO4. Cd was calculated at a frequency of 10 Hz.
DNA immobilization: 0.5/-0.2 V pulse profile (10 ms pulse duration, vs. Ag/AgCl/3 M
KCl), 5 min in 1 µM DNA solution prepared in 10 mM PB with 450 mM K2SO4. Procedure
explained in the Section 3.3.1.
3.2 Importance of knowing the surface 48

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Results and Discussion
Figure 3.17. Determination of the pzc of the DNA-modified electrode. Cddl values were
calculated using the data from Figure 3.16 obtained at a frequency of 10 Hz.
3.2 Importance of knowing the surface 49

_____________________________________________________________ Results and Discussion
3.3 Importance of controlling the surface
The most popular strategy for the immobilization of thiolated molecules on gold surfaces is a
spontaneous formation of Au-S bonds via self-assembly achieved by a simple immersion of the
gold electrode into a solution containing a desired thiol 66 . This strategy is widely used for both
SAM formation of alkylthiols and immobilization of thiolated DNA. The aim of SAM
formation of alkylthiols is generally to achieve the highest possible coverage and to obtain
compact layers with a high blocking ability. On the other hand, desired DNA coverage depends
on the envisaged DNA detection strategy and it can range from low to high coverages 66,84,85 .
Nevertheless, in order to obtain high coverage of thiolated molecules long incubation times are
required, ranging from several hours to days 13-15 . In contrast, low DNA coverages can be
obtained in a short time, but with the drawback of significant variation of densities 67 .
Therefore, a new immobilization strategy needs to be introduced that allows to reproducibly
control the surface modification in a desired manner and what is equally important, in a very
short time. The new approach needs to eliminate the dependence on the spontaneous selfassembly
that is very long and lacks reproducibility. Thus, the possibility of surface control by
potential-assisted surface modification was investigated using both thiolated DNA and
alkylthiols as examples of intrinsically charged and uncharged molecules, respectively.
Furthermore, with the aim of using the envisaged potential pulse-assisted immobilization
method for the preparation of DNA arrays, this approach has to allow array modification with
multiple DNA probes. Due to the need for an electrochemical system (reference and counter
electrodes in addition to the chip working electrode) to perform potential pulsing and the size
of the individual electrodes on an array (usually μm dimensions) it is obvious that more than
one electrode of the array needs to be exposed to the solution used for modification. Therefore,
to prevent crosstalk between electrodes, each electrode needs to be cleaned prior to the
modification. Thus, potential pulse-assisted cleaning of Au modified surfaces was investigated
with the aim to regenerate Au surfaces within a very short time, while not causing any damage
to the surface.
3.3 Importance of controlling the surface 50

_____________________________________________________________ Results and Discussion
3.3.1 Fast and controlled formation of DNA surfaces. Optimization of ssDNA
immobilization procedure
Theoretical discussions were done with Dr. Magdalena Gebala and Prof. Dr. Fabio La Mantia.
Parts of this section were published in ref. 5 : “D. Jambrec, M. Gebala, F. La Mantia, W.
Schuhmann, Angew. Chem. Int. Ed. 2015, 54, 15064-15068; Angew. Chem. 2015, 127, 15278-
15283.” written by the author. Figure adapted from ref. 5 .
The amount and accessibility of immobilized ssDNA on the electrode surface greatly influences
the later hybridization process. Therefore, well-defined reproducible and controlled DNAmodified
surfaces are a prerequisite for the development of optimized DNA sensors 3 . Even
though ssDNA-modified surfaces are the foundation for all envisaged applications, the
mechanism of ssDNA immobilization has not been fully elucidated yet. DNA immobilization
is a complex process depending on many parameters such as ionic strength, strand length, and
the valence of the ions screening the charge of the DNA strands 86 . By introducing varying
applied potentials, this process becomes even more complex. While it is known that certain
potentials (positive or negative with respect of the pzc) induce the bending or lifting of DNA at
the electrified interfaces, the reasoning that the DNA itself is attracted by a positively charged
3.3 Importance of controlling the surface 51

_____________________________________________________________ Results and Discussion
electrode or repelled by a negatively charged electrode is not justified for a set of parameters
usually employed for surface modification with DNA.
Figure 3.18. Nyquist plots with typical Rct values obtained for different immobilization
times where the immobilization was performed at OCP (incubation method). The MCHmodified
electrode was prepared by incubation of the bare gold electrode in a solution of
10 mM MCH with 10 mM PB and 20 mM K2SO4 for 19 h at 37 °C. ssDNA immobilization,
MCH passivation and EIS measurements were performed as stated in Figures 3.7 and 3.8.
Experiments were performed using different electrodes.
DNA immobilization performed at OCP by simple immersion of the gold electrode into a DNA
containing solution is diffusion controlled. Initial immobilization of DNA strands is relatively
fast, while with an increasing amount of immobilized ssDNA grafting of additional strands
becomes energetically unfavorable and the diffusion of new strands is hindered. Since DNA
strands can unspecifically adsorb onto the electrode, lying DNA strands sterically hinder the
approach of new strands and with this the formation of Au-S bonds. Figure 3.18 presents the
immobilization kinetics obtained by performing ssDNA immobilization at OCP followed by
EIS. The immobilization efficiency was investigated by studying the surface after the
passivation step (for reasons explained earlier, Section 3.2.1) and evaluated by comparing the
3.3 Importance of controlling the surface 52

_____________________________________________________________ Results and Discussion
Rct of ssDNA/MCH-modified electrodes with the Rct of an electrode modified only with MCH
under the same conditions that were used for the passivation of the surface. When the incubation
is done only for 15 min a minor increase in Rct is observed as compared with the MCH-modified
electrode implying that a negligible amount of DNA is immobilized. Rct increases notably after
2 h of incubation, however, by prolonging the incubation time to 4 and 8 h the immobilization
kinetics drastically slows down, which is manifested in the EIS plots as a small additional
increase in Rct.
In order to design a potential-assisted immobilization method that will lead to a significantly
improved immobilization kinetics and the formation of well-defined and reproducible DNAmodified
surfaces, the influence of several parameters on the behavior of DNA in the vicinity
of the electrode surface needs to be taken into consideration (Figure 3.19). As it was shown
earlier (Section 3.2.2), the pzc shifts to more negative potential values due to the modification
of the electrode with DNA. This shift requires a careful selection of the applied pulse potentials
to obtain control of the immobilization process.
Figure 3.19. Scheme representing dominating parameters influencing the behavior of
DNA at the electrode surface.
Furthermore, in order to understand the behavior of DNA at the surface of a polarized electrode
we need to observe the potential profile at the electrode surface developed upon applying a
certain potential. By applying a potential more positive or negative with respect to the pzc, an
excess charge is created on the metal side of the interface. As a response a double layer is
3.3 Importance of controlling the surface 53

_____________________________________________________________ Results and Discussion
formed in the electrolyte in order to compensate for this excess charge. A potential profile is
formed and the potential decays with the distance from the electrode as it is explained in Section
1.2.2. Even though the Gouy-Chapman double layer theory does not take into account the linear
potential drop within the Helmholtz plane, it shows the difference in the system response in
relation to parameters such as the ionic strength of the solution or the applied potential. In order
to increase the immobilization kinetics and DNA coverage, DNA immobilization is often
performed in solutions of high ionic strength. However, according to the GC model, the
potential drop is steeper with increasing ionic strength. Therefore, under conditions of high
ionic strength a significant potential drop is observed in the immediate proximity of the
electrode surface. Thus, only a small fraction of a DNA strand in close vicinity of the electrode
can be affected by the applied potential.
Moreover, DNA is a highly negatively charged polyelectrolyte that strongly interacts with
surrounding ions resulting in charge compensation, i.e., DNA screening (described in Section
1.2.1). In the case of monovalent cations, the charge at a DNA strand is screened by counterions
accumulating around the DNA strand in two layers, namely a condensed layer and additional
ions in a second sphere 34 . Therefore, due to the absence of an effective net charge, a DNA strand
cannot be directly affected by the applied potential as it is generally suggested.
These observations imply that the polarized electrode neither attracts nor repels DNA strands
directly, but rather affects the ions in the vicinity of the electrified interface. Namely, during
the charging of the electrochemical double layer, ions have to rearrange in both Helmholtz
planes and the diffuse layer. Thus, when the electrode is polarized to negative values with
respect to the pzc, cations move towards the electrified interface while anions move towards
the bulk of the solution and vice versa. This suggests that switching fast enough between these
two situations creates a “stirring effect” that effectively exceeds the Debye length in front of
the electrified interface. Furthermore, efficient stirring should also move DNA strands present
in close proximity to the electrode surface including their condensed ion cloud. This way the
immobilization will not be diffusion controlled but driven by the migration of ions in front of
the electrode. Based on this hypothesis, we created a potential pulse-assisted immobilization
method that consists of fast switching between potentials more positive and more negative with
respect to the pzc. Figure 3.20 demonstrates the principle of the measurements conducted
during this study.
3.3 Importance of controlling the surface 54

_____________________________________________________________ Results and Discussion
Figure 3.20. a) Potential-time dependence representing potential pulses during potentialassisted
immobilization and b) corresponding current-time response.
It becomes evident that in order to find an efficient pulse profile the following conditions need
to be satisfied:
- to find appropriate pulse intensities
- to find an appropriate pulse duration
Applied potential intensities need to be on the one hand within the stable potential window of
the Au-S bond and on the other hand high enough to evoke an efficient stirring to bring the
DNA towards the surface. Namely, even though according to the G-C theory the potential drop
in front of the electrode is steeper for higher applied potentials, the absolute value of the
potential Φ at a fixed distance from the electrode is higher for a higher applied potential (i.e.,
potential at the OHP, Φ0). This difference is more pronounced in close proximity to the
electrode surface. To demonstrate the influence of potential intensities in a pulse profile on the
efficiency of the immobilization process, different pulse profiles were compared (Figures 3.21
and 3.22). For the purpose of this study the total duration of the immobilization was kept
constant to 15 min and the pulse duration was kept at 10 ms. Using the same upper potential
(0.5 V vs. Ag/AgCl/3 M KCl) and varying the lower potential, three potential differences were
defined: 300 mV (0.5/0.2 V pulse profile), 500 mV (0.5/0 V pulse profile) and 700 mV (0.5/-
0.2 V pulse profile). Figure 3.21 shows how the chosen potential differences relate to both pzc
3.3 Importance of controlling the surface 55

_____________________________________________________________ Results and Discussion
(bare) and pzc (DNA). Namely, all pulse profiles have an upper potential equal to the pzc (bare).
Furthermore, they all relate differently to the pzc (DNA). For the 300 mV pulse profile both
potentials are always positive with respect to the pzc (DNA). For the 500 mV pulse profile
pulsing occurs between positive and slightly negative potentials with respect to the pzc (DNA)
and in the 700 mV pulse profile both positive and negative potentials with similar amplitude
with respect to the pzc (DNA) are applied.
Figure 3.21. Selected potential pulse profiles with respect to the pzc (bare) and pzc (DNA):
ΔE = 300 mV (pulse profile 0.5/0.2 V), ΔE = 500 mV (pulse profile 0.5/0 V) and ΔE = 700
mV (pulse profile 0.5/-0.2 V). Figure adapted from ref. 5 .
By applying a 300 mV pulse profile a small increase in Rct as compared to the Rct of the MCHmodified
electrode is observed showing that obviously only a very small amount of DNA was
immobilized (Figure 3.22, blue curve). However, the amount of immobilized DNA still exceeds
the amount of DNA immobilized by means of the incubation method after the same time (Figure
3.22, black curve). This shows that the investigated pulse profile accelerates the immobilization
of DNA. After the initial immobilization of a minor amount of DNA, the pzc shifts from pzc
(bare) to a more negative value (pzc (DNA)). Therefore, since the pulsing occurs only between
positive potentials relative to pzc (DNA) and the potential difference is small, only a minor ion
stirring effect is achieved and upon initial DNA immobilization, the DNA strands remain lying
on the electrode surface. The electrode is physically blocked by the initially immobilized DNA
strands and the access of additional DNA molecules is impeded.
3.3 Importance of controlling the surface 56

_____________________________________________________________ Results and Discussion
A similar result is observed for the pulse profile with 500 mV potential difference (Figure 3.22,
yellow curve). Although the potential difference is higher than for the 0.5/0.2 V pulse profile,
the applied negative potential is apparently not sufficiently low to induce effective enough ion
stirring and push up already immobilized DNA strands that are lying on the surface as a random
coil due to their low persistence length. Therefore, the access of new strands is hindered and a
poor immobilization yield is observed. On the other hand, further increase of the potential
difference to 700 mV (pulse profile 0.5/-0.2 V) leads to a significantly improved immobilization
yield (Figure 3.22, green curve). Evidently, this pulse profile is vigorous enough to increase the
ion flux and to induce improved ion stirring. Thus, already immobilized DNA strands can be
lifted from the electrode surface and can create space for new strands to approach the surface.
As explained earlier, even though for higher applied potentials the potential drop is steeper, at
a certain distance from the electrode, the DNA is affected by an absolute higher potential.
Figure 3.22. Comparison of different pulse profiles used for potential-assisted DNA
immobilization. MCH-modified electrode was prepared as stated in Figure 3.8. ssDNA
immobilization was performed for 15 min by incubation at OCP or using different pulse
profiles (0.5/0.2 V, 0.5/0 V and 0.5/-0.2 V) with 10 ms pulse duration. MCH passivation
and EIS measurements were performed as stated in Figures 3.7 and 3.8. Figure adapted
from ref. 5 .
3.3 Importance of controlling the surface 57

_____________________________________________________________ Results and Discussion
However, the selected potential values are not the only parameter that determine the efficiency
of the potential-assisted immobilization. Namely, after increasing the potential difference to
900 mV (pulse profile 0.5/-0.4 V) while keeping the pulse time at 10 ms no significant increase
in Rct is observed (data not shown) suggesting that the pulse duration needs to be adjusted with
respect to the selected pulse profile.
In order to demonstrate the effect of the pulse duration on the efficiency of immobilization,
different pulse times were investigated (1, 10, and 100 ms) keeping the total immobilization
time and the pulse profile constant (Figure 3.23). The total duration was 5 min and the 0.5/-0.2
V pulse profile was used. Decreasing the pulse time to 1 ms significantly increases the number
of cycles to 150,000 for the total immobilization time of 5 min. However, at 1ms pulse time the
lowest Rct was obtained among all tested pulse times. When the pulse duration is prolonged to
10 ms and the number of cycles is 15,000, a significant increase in the immobilization efficiency
is observed as shown in a substantial increase in Rct. By prolonging the pulse duration even
more to 100 ms, only 1,500 cycles are completed and Rct again drops to lower values.
Figure 3.23. Comparison of different pulse durations used for potential-assisted DNA
immobilization. ssDNA immobilization was performed for 5 min using the 0.5/-0.2 V pulse
profile with 1, 10 and 100 ms pulse duration. MCH passivation and EIS measurements
were performed as stated in Figures 3.7 and 3.8. Figure adapted from ref. 5 .
3.3 Importance of controlling the surface 58

_____________________________________________________________ Results and Discussion
The obtained results suggest that when the pulse duration is too short to allow for a whole DNA
strand to be pulled down via ion stirring during a single pulse (e.g. 1 ms), only a small fraction
of the DNA in the vicinity of the electrode that is already oriented with the anchor group towards
the electrode surface is grafted (Figure 3.24). On the other hand, if the pulse duration is too long
for the investigated molecule length (e.g. 100 ms), a DNA strand will be brought completely to
the surface during a fraction of a single pulse and it will remain lying on it until a negative pulse
is applied. This creates less efficient ion stirring and, in addition, reduces the number of pulse
cycles. For the investigated system a pulse time of 10 ms is apparently long enough for complete
DNA strands to be pulled down, hence allowing for the formation of the Au-S bond. Moreover,
the 10 ms pulse time is short enough to allow for a high number of potential pulse cycles.
Figure 3.24. Schematic representation of the influence of the pulse duration on the DNA
immobilization efficiency.
After presenting the effect of applied potential intensities and the duration of a single pulse on
the efficiency of potential-assisted DNA immobilization, a model that explains the processes
occurring during grafting of DNA strands supported by applied potential pulses can be proposed
(Figure 3.25). Due to the short distance to which applied potentials have an effect at a given
ionic strength during the application of a single positive potential pulse, cations that are in the
vicinity of the electrode are repelled from the electrode surface including counterions
surrounding DNA strands that are located within the layer of influence. By this, an increased
effective charge at the DNA strand is achieved. The extent of the layer of influence depends on
the intensity of applied potentials and pulse duration. If the pulse is long enough, a wider layer
3.3 Importance of controlling the surface 59

_____________________________________________________________ Results and Discussion
of influence is attained with a higher applied potential. Conversely, anions in the same layer are
attracted towards the electrode and with this a small fraction of a DNA strand that is in
immediate proximity to the electrode is pulled towards its surface. As a result, the remaining
part of a DNA strand that was outside of the area of influence is brought closer to the electrode
surface. If the same potential is still applied the next part of the DNA strand undergoes the same
process; DNA counterions are repelled together with cations from the solution and the rest of
the DNA strand is brought closer to the surface. This process continues as long as the positive
potential pulse is applied. If the duration of a pulse is long enough, the complete DNA strand
is sequentially confined on the electrode surface. Therefore, regardless of the orientation of the
DNA strand, the anchor group is brought close enough to the electrode surface for the formation
of the Au-S bond.
Figure 3.25. Scheme presenting the zipper-like pulling of a DNA strand towards the
electrode surface during potential-assisted DNA immobilization. Figure adapted from
ref. 5 .
The positive potential pulse is followed by a potential jump to a negative potential with respect
to the pzc (DNA). If an appropriate negative potential is applied, anions in the vicinity of the
electrode are repelled from the surface together with the remaining negatively charged DNA
backbone that was left unscreened. By this, the grafted DNA strand are lifted towards the bulk
of the solution. Thus, additional space is created for new DNA strands to approach to the
3.3 Importance of controlling the surface 60

_____________________________________________________________ Results and Discussion
surface. Repetitive switching between positive and negative potentials leads to an increase in
the amount of immobilized DNA due to the created ion stirring that facilitates the approach of
DNA to the surface and the formation of the Au-S bond.
As it was shown previously, that DNA immobilization at OCP exhibits a very slow kinetics.
The proposed potential-assisted method seems to overcome diffusion limitations of the
incubation method by controlling the immobilization kinetics via migration of ions causing ion
stirring. In order to compare the developed method with the immobilization of DNA at OCP,
potential-assisted DNA immobilization was performed using the optimal pulse profile (0.5/-0.2
V with 10 ms pulse duration) while varying the total immobilization time and subsequently
passivating the surface by incubation in MCH solution for the same duration as for DNA
immobilization at OCP (19 h).
Comparing the kinetics of these two methods it is clear that the efficiency of potential-assisted
DNA immobilization is tremendously higher as compared to the incubation method (Figure
3.26). A significant increase of the Rct value is obtained after immobilization for only 15 min
as compared to the MCH-modified electrode. The DNA coverage determined for this
immobilization time using the chronocoulometric method proposed by Steel et al. 75 was
(6.85±0.47) × 10 12 DNA molecules/cm 2 , which is considered to be within an optimal DNA
coverage range for application in DNA sensors 67,75,87,88 . On the other hand, using the incubation
method only a negligible Rct increase is observed after 15 min of immobilization as compared
to the MCH-modified electrode. For 2 h of immobilization at OCP, a coverage of (4.65±0.26)
× 10 12 DNA molecules/cm 2 was determined, which is 47 % lower than the amount of
immobilized DNA achieved within 15 min by the potential-assisted method. Increasing the
incubation time results in a significant decrease in the immobilization kinetics. In order to
obtain the same Rct value that is achieved in only 15 min using the potential-assisted method,
immobilization for about 8 h is needed when the electrode is modified using the commonly
used incubation method.
Figures 3.26b and 3.27 show that using the potential-assisted DNA immobilization method
much higher Rct values, that is, higher DNA coverages, can be achieved within the investigated
immobilization duration. This may be valuable for applications such as the investigation of
DNA repair proteins 66,84 .
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Figure 3.26. Comparison of DNA immobilization methods. a) Nyquist plots with Rct values
obtained for different immobilization times with the immobilization being performed by
incubation and the potential-assisted method. b) Comparison of the immobilization
kinetics using these two methods. ssDNA immobilization, MCH passivation and EIS
measurements were performed as stated in Figures 3.7 and 3.8. Rct values are obtained
from EIS measurements for ssDNA/MCH-modified electrodes by fitting Nyquist plots to
the [R(Q[RW])] equivalent circuit. Figure adapted from ref. 5 .
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_____________________________________________________________ Results and Discussion
Figure 3.27. Change of DNA coverage with the duration of potential-assisted DNA
immobilization. DNA coverage was determined as discussed in Section 5.9.3.
In order to prove that the accelerated immobilization kinetics is indeed due to a generated ion
stirring, for which the application of alternating between positive and negative potentials is
important, DNA immobilization was also investigated applying constant potentials. During the
course of the immobilization, a constant potential of 0.5 V or -0.2 V was applied and the total
immobilization time was kept constant (15 min).
As expected, applying a constant negative potential of -0.2 V (vs. Ag/AgCl/3 M KCl) during
DNA immobilization leads to a very small increase in Rct upon ssDNA immobilization (Figure
3.28, a). Under these conditions, charging of the double layer at the electrode surface occurs
relatively fast and the equilibrium of the system is reached. Therefore, no significant effect of
the applied potential is expected if the DNA is considered as a screened molecule. Similar
results are expected as in the case of immobilization at OCP. If a simple electrostatic
attraction/repulsion model is considered, no change of Rct with respect to the bare electrode
should be observed since the negatively charged electrode would repel the DNA strands and
fully prevent DNA immobilization. However, applying a constant positive potential would be
expected to make a high impact on the immobilization efficiency, since a positively charged
electrode is assumed to attract negatively charged DNA and facilitate the grafting process. This
is not observed and applying a constant potential of 0.5 V (vs. Ag/AgCl/3 M KCl) during DNA
immobilization leads to a similar result as in the case of the negatively charged electrode, where
3.3 Importance of controlling the surface 63

_____________________________________________________________ Results and Discussion
only a slightly higher Rct value is observed. Furthermore, the resulting blocking of the electrode
surface is much lower than in the case of potential pulse-assisted DNA immobilization (Figure
3.28).
Figure 3.28. Comparison of the potential pulse-assisted immobilization method with
immobilization at constant potentials. ssDNA-modified electrodes were characterized
using a) EIS and b) CV. ssDNA-modified electrodes were obtained by potential-assisted
DNA immobilization using the 0.5/-0.2 V pulse profile with 10 ms pulse duration and by
applying a constant potential of -0.2 V or 0.5 V vs. Ag/AgCl/3 M KCl for 15 min. ssDNA
immobilization was performed in 10 mM PB with 450 mM K2SO4 and 1 µM ssDNA. EIS
measurements were performed as stated in Figure 3.7. CVs were performed in 10 mM
PB, 20 mM K2SO4 containing 5 mM of K3[Fe(CN)6] and K4[Fe(CN)6] at 100 mV/s scan
rate.
Finally, ssDNA/MCH-modified electrodes obtained by DNA immobilization using the
potential pulse-assisted method (0.5/-0.2 V vs. Ag/AgCl/3 M KCl, 10 ms pulse duration) or
immobilization supported by a constant positive potential (0.5 V vs. Ag/AgCl/3 M KCl) were
compared (Figure 3.29). The total immobilization and passivation times were kept constant. As
expected, the Rct value obtained by applying 0.5 V during immobilization is comparable to the
Rct value obtained for immobilization performed by a simple incubation. Furthermore, this
amount is significantly lower as compared to the potential-pulse assisted immobilization, which
additionally supports the suggested mechanism behind the developed approach.
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_____________________________________________________________ Results and Discussion
Figure 3.29. Nyquist plots with typical Rct values obtained by comparing the potentialassisted
immobilization method with immobilization performed by simple incubation or
supported by applying a constant potential. MCH passivation was done by incubation for
19 h at 37 °C.
Besides the ability of the potential-assisted immobilization method to tremendously accelerate
the immobilization kinetics and to achieve much higher DNA coverages in a shorter time as
compared to the incubation method or applications of constant potentials, a high reproducibility
of the surface modification needs to be attained. Figure 3.30 presents average Rct values
obtained with the potential-assisted immobilization method using different pulse profiles. In all
cases Rct values are obtained with a standard deviation below 5 %, showing that the developed
immobilization protocol leads to a highly reproducible surface modification. Thus, the
envisaged ssDNA surface coverage can be pre-selected by choosing the number of applied
potential pulse cycles.
Looking at the results presented in this chapter it can be concluded that understanding the
behavior of ssDNA in front of the electrode surface affected by the surrounding electrolyte and
the polarization of the electrode is essential for the development of highly reproducible DNAmodified
surfaces. The developed potential pulse-assisted immobilization strategy leads to
high-quality DNA-modified surfaces helping by this to overcome difficulties in DNA sensor
3.3 Importance of controlling the surface 65

_____________________________________________________________ Results and Discussion
preparation. Therefore, the proposed method may become a new standard DNA immobilization
procedure to obtain desired surface coverages in a very short time as a prerequisite for the
development of highly sensitive and reproducible DNA hybridization assays with
electrochemical read-out.
Figure 3.30. Reproducibility of the potential-assisted immobilization method. Data
obtained from EIS measurements for ssDNA/MCH-modified electrodes by fitting Nyquist
plots to a [R(Q[RW])] equivalent circuit. The pulse profile for DNA immobilization was
performed for 15 min using a 10 ms pulse duration. The black dashed line represents the
average fitted Rct value for the MCH-modified electrode. Error bars represent the
standard deviation between measurements (n > 3). Figure adapted from ref. 5 .
3.3 Importance of controlling the surface 66

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3.3.2 Formation of compact thiol SAMs within minutes
Data analysis was done based on discussions with Dr. Felipe Conzuelo and Dr. Arturo Estrada-
Vargas. Parts of this section were published in ref. 89 : “D. Jambrec, F. Conzuelo, A. Estrada-
Vargas, W. Schuhmann, ChemElectroChem 2016, 3, 1484-1489.” written by the author. Figure
adapted from ref. 89 .
A typical approach for a DNA sensor buildup is a two-step immobilization strategy, in which
initially a probe DNA is grafted on the surface and subsequently the electrode is covered by a
thiol SAM. Thiol passivation forces the grafted DNA to lift up and to obtain a more favorable
orientation for the hybridization process concomitantly removing unspecifically bound DNA
strands. Additionally, the blocking ability of the passivation step in the DNA sensor preparation
plays an important role in the efficiency of an envisaged DNA sensing scheme. It prevents
unspecific adsorption of undesired molecules employed in the detection scheme and determines
the signal of the negative control. Therefore, controlling of the blocking ability of the modified
surface is crucial for construction of very sensitive and reproducible DNA sensors.
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_____________________________________________________________ Results and Discussion
Even though SAM formation has been an extensively investigated topic for decades until now,
the implementation of procedures for a fast and controlled SAM deposition is of great interest,
particularly when looking at future of point-of-care diagnostics devices, for which fast and
cheap techniques for sensor fabrication are becoming more and more important in order to
decrease production costs. The process of self-assembly supposedly consists of two
phases 22,90,91 , during which in the initial stage thiol molecules randomly cover the surface, and
in a second phase they slowly organize on the surface to form a monolayer. Although the twophase
process is widely accepted, there are conflicting reports on the kinetics of SAM
formation 92 .
Most approaches for SAM formation involve a self-assembly process at open circuit potential.
This is a simple method with, however, quite low reproducibility 24 and slow kinetics. It has
been reported in the literature that the application of constant positive potentials accelerates the
immobilization kinetics of long thiol chains providing compact SAMs in a shorter time 13,21,92 .
Nevertheless, further improvement of the immobilization process is still needed, especially
focusing on short-chain thiols commonly used for sensor fabrication 22 . Therefore, the
application of the potential pulse-assisted immobilization method for alkylthiol derivatives as
examples for uncharged molecules was investigated.
One of the most reliable methods for the investigation of the self-assembly process is the
measurement of the double-layer capacitance since it precisely describes the SAM adsorption
properties. The interfacial capacitance of a gold-SAM-electrolyte interface consists of the
capacitance of the SAM and the capacitance of the diffuse layer connected in series. The overall
capacitance is determined by the smaller one, which is the capacitance of the SAM. EIS is a
commonly used technique for the determination of the capacitance of a system. Real time
measurement of electrochemical impedance allows the continuous determination of the
capacitance with time and hence a detailed investigation of the SAM formation kinetics.
Subramanian and Lakshminarayanan 92 used real time impedance monitoring to study the selfassembly
mechanism of thiols by applying a constant DC potential superimposed with an AC
signal at a single frequency. This measuring principle was modified in that way, that a pulsetype
potential modulation employed for the above described potential-pulse assisted
immobilization method was applied as DC potential.
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Figure 3.31. Schematic representation of the setup used for real-time impedance
measurements during potential-pulse assisted acceleration of SAM formation: A – signal
in, R – AC current magnitude, θ – AC current phase, AD/DA – analog-digital/digitalanalog
conversion for potential application and data acquisition, RE – reference electrode,
CE – counter electrode, WE – working electrode. The potentiostat receives the DC
potential signal as potential pulses from the function generator and superimposes an AC
signal with a single high frequency generated by the oscillator in the lock-in amplifier.
The recorded current is fed back to the lock-in amplifier and the AC current at the
excitation frequency is amplified providing the magnitude and the phase of the resulting
AC response. Figure adapted from ref. 89 .
The real-time impedance measuring setup consists of several components: a potentiostat, a
function generator, a lock-in amplifier, an AD/DA card and an electrochemical cell (Figure
3.31). The function generator is used to create a pulse-type DC potential modulation and to
apply it to the external potential input of the potentiostat. In order to be able to apply potential
pulses for accelerated SAM formation while simultaneously applying an AC frequency for
impedance measurements, a summing amplifier is used to superimpose the square wave DC
signal with a high-frequency AC signal. It should be noted that the AC signal needs to be of a
significantly higher frequency with respect to the pulse time of the DC signal to prevent any
influence of the small AC perturbation on the behavior of the investigated system. The current
3.3 Importance of controlling the surface 69

_____________________________________________________________ Results and Discussion
measured by the potentiostat is sent to the lock-in amplifier input after current-to-voltage
conversion and the magnitude and the phase of the AC current response at the frequency of the
AC excitation signal are obtained. Using these values, the imaginary impedance component and
subsequently the interfacial capacitance can be calculated using the following expression,
derived from an RC series equivalent electrical circuit (for detailed calculation see Section 5.5):
C = −1
ωZ ′′ (3.7)
where ω is angular frequency and -Z" is the imaginary component of impedance.
Figure 3.32. Mercaptoundecanol (MCU) SAM formation kinetics performed at constant
potential equal to the previously measured OCP of the measuring solution (0 V vs.
Ag/AgCl/3 M KCl). Measurement was performed in 10 mM PB with 20 mM K2SO4
containing 1 mM MCU (30 % ethanol).
As mentioned earlier, the kinetics of SAM formation at OCP is quite slow after the initial fast
phase (Figure 3.32). During the first hour a relatively fast decrease in capacitance is manifested
after which a significant deceleration of the immobilization kinetics is observed. Even after
more than 12 h a compact thiol monolayer did not form, which is evident in the figure from the
absence of a capacitance plateau. The measurement was conducted by applying a constant
potential equal to the previously measured OCP. After determining the OCP, the measurement
was performed applying a constant potential equal to the OCP superimposing an AC signal of
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_____________________________________________________________ Results and Discussion
high frequency. The fact that the OCP does not significantly change over the course of the SAM
formation by simple incubation (data not shown) justifies this approach.
As in the previous section, the influence of different parameters on the potential-pulse assisted
SAM formation is investigated: potential intensities in the potential-pulse cycle and duration of
the potential pulse (Figure 3.33). Emphasis in this chapter is on the dependence of these
parameters on the length of a alkyl chain of the thiol derivative.
Figure 3.33. Scheme of investigated potential pulse profiles (ΔE = 400 mV, 0.3/-0.1 V; ΔE
= 700 mV, 0.5/-0.2 V; ΔE = 900 mV, 0.5/-0.4 V, all vs. Ag/AgCl/3 M KCl) and pulse
durations (1 ms, 10 ms, 100 ms and 10 s). Figure adapted from ref. 89 .
It was previously shown that the pzc shifts from 0.5 V vs. Ag/AgCl/3 M KCl (pzc of the bare
electrode) towards more negative values due to surface modification with DNA strands. The
same behavior is expected in the case of thiol SAM formation, since it was reported that the
pzc of a thiol-modified electrode is around 0.1 V 93 (vs. Ag/AgCl/3 M KCl). Therefore, this shift
in the pzc value was taken into account during the selection of the potential pulse intensities.
Using the ΔE = 400 mV pulse profile (0.3/-0.1 V vs. Ag/AgCl/3 M KCl) potentials of +200 mV
and -200 mV relative to the pzc value of the thiol-modified surface are applied, while for ΔE =
700 mV (0.5/-0.2 V vs. Ag/AgCl/3 M KCl) +400 mV and -300 mV are applied and for ΔE =
900 mV (0.5/-0.4 V vs. Ag/AgCl/3 M KCl) +400 mV and -500 mV are applied.
Figure 3.34 shows capacitance curves obtained during the formation of a MCU SAM using
different potential-pulse profiles. It presents the influence of potential intensities on the kinetics
3.3 Importance of controlling the surface 71

_____________________________________________________________ Results and Discussion
of SAM formation. All experiments were performed with a pulse duration of 10 ms and only
the intensities of the applied potentials within a pulse profile were varied. The capacitance
values are normalized by the real electrode surface area derived from CVs in sulfuric acid.
Using the ΔE = 400 mV pulse profile (0.3/-0.1 V vs. Ag/AgCl/3 M KCl) a tremendous
acceleration of the immobilization kinetics is observed as compared to SAM formation at OCP.
The monitored capacitance reaches a plateau after about 90 min implying that a fully SAM
covered electrode surface was obtained (Figure 3.34, blue curve). Increasing the potential
difference to 700 mV (0.5/-0.2 V pulse profile) an even faster immobilization kinetics is
achieved, and the capacitance plateau is reached within minutes (Figure 3.34, orange curve).
With a higher potential intensity within the applied pulse profile, the potential drop in the
vicinity of the electrode is faster generating a higher concentration gradient and thus a more
efficient ion stirring. This obviously results in a much faster immobilization kinetics of the
alkylthiol molecules. However, the efficiency of an applied pulse profile does not only depend
on the selected potential values relative to the pzc, but also on the pulse duration and the
migration of ions in solution as it was above described for the potential-assisted immobilization
of DNA. If the potential pulse is too short to allow the formation of a sufficient concentration
gradient, increasing of potential intensities while keeping the pulse time constant results in a
slower SAM immobilization kinetics. This is shown in Figure 3.34 (yellow curve) for the case
when increasing the potential difference to 900 mV (0.5/-0.4 V pulse profile) while keeping the
same pulse time of 10 ms.
In contrast to DNA immobilization, the evaluated alkanethiols are uncharged molecules, the
principle of potential pulse-assisted SAM formation is similar to the potential pulse-assisted
DNA immobilization. Namely, during a single potential pulse, a certain potential drop is
generated in the vicinity of the electrode surface. By switching to a potential with opposite sign
with respect to the pzc, the excess of ions in front of the electrode moves away from the surface
and exchanges with ions of opposite charge. Fast switching between positive and negative
potentials creates an ion stirring in the vicinity of the surface, moving along thiol molecules
that are within this layer. If a potential pulse is sufficiently long, a whole thiol molecule can be
pulled to the surface during a single pulse regardless of its orientation, facilitating the formation
of the Au-S bond. Thus, potential pulse-assisted thiol immobilization, like in the case of charged
molecules, is driven by the migration of ions in the vicinity of the electrode rather than the
diffusion of thiols, what is the case during of self-assembly at OCP 94 .
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Figure 3.34. Capacitance kinetics curves obtained for potential-pulse assisted SAM
formation of MCU using different potential pulse intensities. Pulse time was kept constant
at 10 ms for all experiments. All measurements were performed in 10 mM PB, 20 mM
K2SO4 (30 % ethanol) containing 1 mM MCU. Figure adapted from ref. 89 .
Figure 3.35. Capacitance kinetics curves obtained for potential-pulse assisted SAM
formation of MCU using the 0.3/-0.1 V pulse profile while varying the pulse time. All
measurements were performed as explained in Figure 3.34. Figure adapted from ref. 89 .
3.3 Importance of controlling the surface 73

_____________________________________________________________ Results and Discussion
The duration of a pulse has to allow the system to respond to the perturbation invoked by
applying a certain pulse profile. In order to achieve the highest SAM formation efficiency, the
duration of a given potential pulse should be long enough to allow for an appropriate
concentration gradient to form and a whole thiol molecule to be brought to the electrode surface.
By this the formation of the Au-S bond is achieved regardless of the orientation of a thiol
molecule. However, the pulse time should be also short enough to allow for a high number of
pulsing cycles. Figure 3.35 demonstrates the influence of the pulse duration on the SAM
formation kinetics. For the investigated system (MCU, ΔE = 400 mV) 10 ms pulse duration
manifests as the optimal pulse time. Prolonging the pulse time leads to a less efficient ion
stirring and with this a slower immobilization kinetics (Figure 3.35, yellow and green curves).
On the other hand, decreasing the pulse time prolongs the time necessary to obtain a compact
monolayer (Figure 3.35, blue curve) even though the number of stirring cycles during the
overall immobilization time is increased.
In order to demonstrate the influence of the length of the investigated thiol derivative on the
optimization of the potential-pulse profile for potential-assisted immobilization, Figure 3.36
compares alkylthiols of three different lengths: MCH (6 carbon atoms), MCU (11 carbon atoms)
and MCHD (mercaptohexadecanol, 16 carbon atoms). In all three cases, a more efficient ion
stirring and faster SAM formation kinetics are achieved using a higher pulse potential
difference, suggesting that the optimal potential difference does not depend on the molecule
length as long as the chosen pulse duration is appropriate. In contrast, the optimal pulse duration
depends on the molecule length, more precisely, on its diffusion coefficient. For the
immobilization of short and intermediately long molecules (Figure 3.36, a and b) the 10 ms
pulse duration is sufficiently long. Prolonging the pulse time leads to a lower number of pulse
cycles per time and slower adsorption kinetics. The optimal potential-pulse assisted
immobilization procedure for short and intermediate thiols employs the 0.5/-0.2 V pulse profile
with 10 ms pulse duration. In case of longer thiols such as MCHD (Figure 3.36, c) 10 ms is too
short to allow for the whole alkyl chain to be brought to the interface, which results in a slower
adsorption kinetics. By prolonging the pulse duration to 10 s a much faster SAM formation rate
is observed. Therefore, the optimal immobilization kinetics of longer thiols is achieved using
the 0.5/-0.2 V pulse profile with 10 s pulse duration. At these conditions the capacitance reaches
a plateau within only 5 min.
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Figure 3.36. Optimization of the potential-pulse assisted SAM formation kinetics for three
alkylthiol derivatives of different length: a) MCH, b) MCU, and c) MCHD. All
measurements were performed as explained in Figure 3.34. Figure adapted from ref. 89 .
According to the Helmholtz model of the double layer the interfacial capacitance of a SAMmodified
electrode is inversely proportional to the thickness of the SAM, that is, the length of
the alkyl chain of the thiol derivative 43 :
C = εε 0
δ
(3.8)
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_____________________________________________________________ Results and Discussion
where ε has the value 2 for compact chains and δ is the thickness increment (0.11 nm per carbon
atom for a 30° tilted alkyl chain 92 ). This supports the results from Figure 3.37, in which it is
shown that the capacitance of the MCH SAM is the highest and the one of a MCHD SAM the
lowest. Furthermore, it was demonstrated previously 95 that for a low number of carbon atoms
(less than 8) the relation between thiols of different length may not be linear, which is seen also
in Figure 3.37. Considering the relationship derived from the Helmholtz model, capacitance
values of 2.7, 1.5 and 1.0 µF/cm 2 are expected for SAMs of MCH, MCU and MCHD,
respectively. The experimentally determined capacitance values are smaller, however, they are
in the same order of magnitude.
Figure 3.37. Capacitance values of fully covered thiol monolayers in dependence on the
inverse number of carbon atoms of the investigated thiol derivatives. Data obtained from
Figure 3.36. Figure adapted from ref. 89 .
From the capacitance curves it is possible to calculate the coverage kinetics of the investigated
alkylthiols using the following equation:
θ = (C − C 0)
(C f − C 0 )
(3.9)
where C0 is the initial capacitance of the bare electrode, Cf is the capacitance of a fully covered
monolayer and C is the capacitance at any time. The initial capacitance was calculated to be
6.77 µF/cm 2 as determined in background electrolyte in absence of any thiol. Calculated
coverage curves are presented in Figure 3.38, where it is shown that using the optimal potential-
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pulse assisted immobilization procedures, tailored for the specific molecule length, compact
SAMs can be obtained within minutes.
Figure 3.38. Coverage curves calculated from optimal capacitance curves for all three
thiol derivatives: MCH 0.5/-0.2 V (10 ms pulse duration), MCU 0.5/-0.2 V (10 ms pulse
duration), MCHD 0.5/-0.2 V (10 s pulse duration) according to Equation (3.9). Figure
adapted from ref. 89 .
It was previously reported that by applying constant potentials an acceleration of the SAM
formation kinetics of aliphatic thiols (C12-C14) can be achieved 21,92 . For example, Ma and
Lennox reported to obtain a complete chemisorption in only 15 min upon applying a constant
potential 13 . They assessed the SAM integrity by evaluating the blocking ability towards electron
transfer of a redox mediator and observed a full blocking of electron transfer for SAM formation
supported by applying constant potentials from 0.2 V up to even 1 V. The obtained result is
surprising since it is known that relatively high anodic potentials of above 600 mV induce Au-
S bond cleavage 19 . Nevertheless, for the purpose of comparison, our potential-assisted SAM
formation method was tested using this approach. A typical CV obtained by using the optimized
procedure for potential-assisted SAM formation of MCU is presented in Figure 3.39, where it
can be seen that a complete blocking of electron transfer is achieved already within seconds.
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Figure 3.39. CVs for the evaluation of blocking of the electrode surface. a) Before and
after MCU SAM formation using the potential-pulse assisted method. b) Rescaled CVs
obtained after different times of surface modification. 0.5/-0.2 V pulse profile with 10 ms
pulse duration was used. Immobilization was performed in 10 mM PB, 20 mM K2SO4
containing 1 mM MCU (30 % ethanol). CVs were performed in 10 mM PB, 20 mM K2SO4
containing 5 mM of K3[Fe(CN)6] and K4[Fe(CN)6] at 100 mV/s scan rate.
To evaluate how constant potentials influence the self-assembly of thiols with an –OH
functional group in aqueous solutions, a control experiment was performed by applying a
relatively high constant potential of 0.5 V (vs. Ag/AgCl/3 M KCl) during thiol chemisorption,
that is still below the potential range of Au-S bond cleavage. As can be seen in Figure 3.40, a
significant acceleration of the immobilization kinetics was observed as compared with the SAM
formation at OCP. However, the SAM formation kinetics is still substantially slower than the
observed kinetics obtained by the potential pulse-assisted immobilization method. A constant
capacitance value is reached after 3 h (Figure 3.40, grey curve), while with our approach
compact monolayers are obtained within only 5 min (Figure 3.40, orange curve). It should be
noted that, in contrast to the vast majority of reported procedures for electrochemical-assisted
thiol immobilization 13,21,22,92 , the proposed technique for SAM deposition is performed in
aqueous solutions ensuring high compatibility with biomolecules used for further surface
modification.
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Figure 3.40. Comparison of the formation kinetics of MCU SAMs performed at a constant
potential equal to the previously measured OCP (0 V vs. Ag/AgCl/3 M KCl, black curve),
a constant potential of 0.5 V (grey curve) and by using the potential-assisted method with
a 0.5/-0.2 V potential-pulse profile with 10 ms pulse duration (orange curve).
Measurements were performed in 10 mM PB with 20 mM K2SO4 containing 1 mM MCU
(30 % ethanol). Inset: Immobilization kinetics of MCU at a potential equal to the OCP
shown for 12 h. Figure adapted from ref. 89 .
The obtained results for the kinetics of potential-assisted SAM formation suggest that the
stirring process generated by pulsing enhances both, the approach of thiol molecules towards
the electrified interface to allow formation of the Au-S bond and the packing and reorganization
of adsorbed molecules. It is important to understand that this process is not likely driven by
diffusion of thiols, but rather by migration of ions from the solution that carry along thiol
molecules towards the surface. Furthermore, a Langmuir isotherm cannot be used to fit our
experimental data, which supports the assumption that the process is not diffusion but rather
migration driven. The developed potential pulse-assisted SAM formation method is promising
for a fast and reproducible surface modification that may be implemented in diverse
applications especially related to the production of point-of-care devices.
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Moreover, the fact that the potential-assisted immobilization method tremendously accelerates
the immobilization kinetics of uncharged molecules as well, and not only of intrinsically
charged molecules such as DNA, proves that simple DNA attraction/repulsion by the polarized
surface is a far too simple model to explain the mechanism of DNA immobilization supported
by potential application.
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3.3.3 Reproducible recycling of Au modified surfaces within seconds
Parts of this section are in the manuscript that is in preparation: “Fully potential-controlled
preparation of DNA chips” with coauthors: D. Jambrec, Y. U. Kayran and W. Schuhmann.
Electrochemical SAM desorption has a wide range of applications including cleaning of gold
electrodes to allow reusability of the surface (i.e. sensor recycling) 18 or modifying surface
wettability and fabrication of mixed SAM layers 96,97 . The interest for electrochemical
desorption of SAMs in this thesis is the development of a very fast and efficient desorption
approach that allows cleaning of individual electrodes on a chip prior to their modification with
appropriate DNA strands. Since DNA and alkylthiols that are used for surface modification
have different anchoring molecules, where alkylthiols form only a single Au-S bond and the
DNA forms six bonds Au-S from three disulfides within the anchoring site (see Section 3.4.1),
we tested the efficiency of the developed potential pulse-assisted desorption approach on both
types of molecules.
Figure 3.41. Removal of DNA from the electrode surface (2 mm gold electrode). a) EIS
and b) CV were used for the characterization of the surface at each step. DNA
immobilization: incubation for 2 min in 1 µM DNA solution (10 mM PB, 450 mM K2SO4).
Potential pulse-desorption: 0.9/-0.9 V vs. Ag/AgCl/3 M KCl, 10 ms pulse duration, 30 s.
Measurements were performed in 10 mM PB, 20 mM K2SO4 containing 5 mM of
K3[Fe(CN)6] and K4[Fe(CN)6]; CV at 100 mV/s scan rate.
3.3 Importance of controlling the surface 81

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The typical approach for SAM desorption is the polarization of SAM-modified Au surfaces to
potentials that are either too positive or too negative, leading to the cleavage of the Au-S bond,
that is, oxidative or reductive desorption, respectively. However, the exact potential values
needed in order to provoke SAM desorption are under discussion for decades and a substantial
variety of values is reported. According to our knowledge, applying both positive and negative
potentials within the same desorption process by pulsing between these potentials was not yet
reported. Even though fairly fast desorption times are reported in the literature 18 (down to 30 s)
the aim of the study was to further improve the efficiency of desorption.
Applying very high potentials seems a reasonably simple approach to achieve SAM desorption,
however, another important condition for the development of a desorption process is to leave
the electrode surface undamaged after polarization, especially in the case of multi-electrode
chips where the Au layer thickness is in the order of nm. Therefore, the potential pulse-assisted
desorption approach was developed by exposing the electrode to very short potential pulses
within the range of ms. The modified electrode was exposed to potential pulsing between
potential values of 0.9 V and -0.9 V (vs. Ag/AgCl/3 M KCl) with a pulse time of 10 ms. The
influence of the applied pulse profile on a DNA-modified electrode is presented in Figure 3.41.
Prior to desorption, a bare electrode was modified with ssDNA by incubation for 2 min to mimic
the conditions that would be observed on the chip. Namely, during potential-assisted DNA
immobilization on a selected chip electrode, the rest of the electrodes would be exposed to the
same solution and the DNA molecules would immobilize at OCP. In order to prevent later
cross-talk between electrodes, these surfaces need to be cleaned prior to further modification
with a desired probe DNA. Even though short DNA immobilization at OCP leads to a minor
Rct increase (Figure 3.41, a) and consequently to a small DNA coverage, it can behave as an
impurity significantly affecting the surface modification. Subsequent modification with a DNA
probe of interest is then leading to a less efficient hybridization since the probe DNA coverage
is higher than the optimal one. Even though the optimized immobilization procedure is
employed the hybridization will be hindered due to the contribution from the leftover DNA.
After the treatment of the DNA-modified electrode with the developed desorption method for
30 s Rct decreases to the value of the bare electrode and the voltammogram obtains the same
shape as for the bare electrode, showing that all DNA was efficiently removed from the surface.
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Figure 3.42. Recycling treatment of an Au electrode. Three cycles of surface modification
with a MCU SAM and subsequent desorption are shown. a) bare electrode; b) after 1 min
of potential-pulse assisted SAM formation (0.5/-0.2 V vs. Ag/AgCl/3 M KCl, 10 ms); c)
after 30 s potential-pulse desorption (0.9/-0.9 V vs. Ag/AgCl/3 M KCl, 10 ms); d) after
SAM formation; e) after desorption (30 s); f) after SAM formation; g) after desorption (5
s). Measurements were performed in 10 mM PB, 20 mM K2SO4 containing 5 mM of
K3[Fe(CN)6] and K4[Fe(CN)6] at 100 mV/s scan rate.
Furthermore, to investigate whether the developed method can desorb also highly compact
SAM layers, desorption of MCU SAMs was investigated (Figure 3.42 and Figure 3.43). A MCU
SAM was formed by the potential-pulse assisted method (0.5/-0.2 V, 10 ms pulse duration, total
time 1 min) leading to a full blocking of electron transfer from the free-diffusing redox mediator
(Figure 3.42, a and b). Then, potential-pulse assisted desorption was performed for 30 s
resulting in a complete regeneration of the gold surface (Figure 3.42, c), showing that the
employed desorption procedure efficiently removes also compact thiol layers from the electrode
surface. To verify whether the electrode surface remains undamaged after the used desorption
procedure, immobilization/desorption cycles were repeated 3 times on the same electrode. The
passivation of the surface by the MCU SAM is very reproducible and the surface becomes fully
blocked (Figure 3.43, a). By using the developed desorption method the surface gets fully
regenerated after each cycle (Figure 3.43, b), which proves that the employed desorption
approach is harmless for the Au surface. Finally, the first two desorption cycles were performed
3.3 Importance of controlling the surface 83

_____________________________________________________________ Results and Discussion
for 30 s, while the last desorption cycle was performed for only 5 s, resulting in the same effect,
namely the full regeneration of the surface. Shorter desorption times were not investigated.
The potential pulse-assisted desorption allows regeneration of Au surfaces within only 5 s while
preserving the surface quality for further modification, which makes this method suitable for
the application in DNA microarrays.
Figure 3.43. CV characterization during the removal of a MCU SAM from the electrode
surface. Data was taken from Figure 3.42 and plotted to compare each cycle of a) SAM
formation by the potential-pulse assisted immobilization method; b) surface
characterization before and after regeneration by potential-pulse desorption.
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3.4 Potential-assisted preparation of DNA sensors
Experiments for the optimization of the DNA sensing surface were done by Bianca Ciui under
the supervision of the author. Experiments with multi-electrode chips were performed together
with Yasin Uğur Kayran. Parts of the chapter are included in the manuscript that is in
preparation: “Fully potential-controlled preparation of DNA chips” with coauthors: D.
Jambrec, Y. U. Kayran and W. Schuhmann.
3.4.1 Optimization of the potential-pulse assisted immobilization method
Optimization of the potential-pulse assisted immobilization method to achieve highly
reproducible and very fast immobilization of intrinsically charged (DNA) or uncharged
molecules (thiols) provides a very efficient technique for preparation of DNA sensing
platforms, for which the desired ssDNA coverage and blocking ability can be reached within
minutes. Moreover, due to a very short modification time while maintaining a high level of
reproducibility and the ability to electrochemically remove undesired layers from the electrode
surface within seconds, the developed protocol has a potential for application in DNA array
production.
Figure 3.44. Scheme of the anchoring site used for immobilization of ssDNA.
Nevertheless, implementation of the developed technique into a DNA sensor or a DNA array
production requires further investigation. It is important to know whether the stability of a
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formed DNA film is affected by the subsequent potential-pulse assisted passivation steps,
whether the passivating thiol replaces the grafted DNA from the surface, and whether the
application of potential pulses within a subsequent passivation step causes a desorption of
previously grafted DNA molecules.
Figure 3.45. Nyquist plots of ssDNA-modified electrode before and after pulsing in 10 mM
PB (450 mM K2SO4) using the 0.5/-0.2 V (vs. Ag/AgCl/3 M KCl) pulse profile with 10 ms
pulse duration. EIS measurements were performed as explained in Figure 3.7.
To assure that the passivating thiol does not remove the immobilized DNA, a special anchoring
molecule with six binding sites is used for the immobilization of ssDNA, where six Au-S bonds
are formed per individual DNA strand (Figure 3.44). Since the thiol molecules form only one
Au-S bond they cannot replace ssDNA that is much more stable 98 .
The effect of the pulsing itself on the stability of the DNA film was investigated by exposing a
ssDNA-modified electrode to potential pulsing performed in buffer solution alone (10 mM PB,
450 mM K2SO4). The same potential profile (0.5/-0.2 V, 10 ms pulse duration) that is employed
in the passivation step of short or intermediate length thiol was used. Figure 3.45 shows that
upon pulsing for 1 min the DNA layer remains unaltered, i.e., no visible desorption occurs.
After 2 min, a small decrease in Rct is observed suggesting that the DNA film starts slightly to
desorb even though it cannot be excluded that the change in Rct may originate from
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reorganization of the DNA at the surface. Pulsing for 5 min creates a much stronger effect on
the DNA film inducing a significant decrease of Rct, which is most likely connected to DNA
desorption.
Figure 3.46. CV measurements made before and after surface modification with a) MCH
and b) MCU and subsequent FSCV measurements made with c) MCH- and d) MCUmodified
electrodes before and after incubation in Fc-tDNA solution. SAM formation was
done using the potential-pulse assisted method (0.5/-0.2 V vs. Ag/AgCl/3 M KCl, 10 ms
pulse duration) for 1 min in 1 mM thiol solution (10 mM PB, 20 mM K2SO4, 30 % ethanol).
CV measurements were performed in 10 mM PB, 20 mM K2SO4 containing 5 mM of
K3[Fe(CN)6] and K4[Fe(CN)6] at 100 mV/s scan rate. FSCV measurements were
performed in 10 mM PB with 450 mM K2SO4 at 1 V/s scan rate.
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Even though the employed pulse profile is within a safe potential window of the Au-S bond,
fast pulsing still provokes a partial desorption of DNA molecules, albeit at a much slower rate
than with the pulse profile employed for potential-pulse assisted desorption (0.9/-0.9 V, 10 ms).
This raises the question about why the same pulse profile results in a significantly accelerated
DNA and thiol immobilization. The reason is that the immobilization rate is apparently much
higher than the desorption rate and it prevails, leading to a very fast immobilization kinetics.
After DNA immobilization, the following passivation step should be performed within a time
window that provokes a negligible desorption of the previously formed DNA layer, which is in
this case 1-2 min. In order to find the optimal passivation procedure, potential-assisted
immobilization of MCH and MCU was performed for a duration of 1 min and the obtained
layers were compared with respect to their integrity using CV. Furthermore, the ability of
formed layers to block the unspecific adsorption of tDNA was investigated by measuring the
ferrocene signal from ferrocene labelled tDNA (Fc-tDNA) using fast scan cyclic voltammetry
(FSCV) upon incubation of modified electrodes into the Fc-tDNA solution. Since the SAM
formation of MCU is more efficient as compared with MCH in the investigated time, which is
observed as a better blocking of the redox mediator (Figure 3.46, a) and absence of unspecific
adsorption of the Fc-tDNA (Figure 3.46, b), MCU was chosen for the potential-pulse assisted
passivation step in the build-up of the DNA assay. This finding is expected since longer
alkylthiols produce a more ordered SAM with less defects 12,99 .
A crucial step in designing a DNA sensor is the optimization of the ssDNA coverage, since it
determines both the hybridization efficiency and its kinetics. To demonstrate the use of the
potential-pulse assisted immobilization method for the preparation of a DNA sensor, the ssDNA
coverage was optimized for hybridization detection based on the measurement of the ferrocene
(Fc) signal from a Fc-tDNA. Both ssDNA immobilization and subsequent MCU passivation
were performed using the potential-assisted method (0.5/-0.2 V, 10 ms pulse duration). The
ssDNA immobilization duration was varied and passivation was kept at 1 min. Sensor
preparation was characterized by CV, following each step of the build-up (an example is shown
in Figure 3.47, a). The prepared sensors were subjected to hybridization with a fully
complementary Fc-tDNA. FSCV was used for the detection of the ferrocene signal (Figure
3.47, b) and the calculation of dsDNA coverage according to equations from Section 5.11.1.
The obtained results are shown in Figure 3.48. When the ssDNA coverage on the surface is low
that adjacent DNA strands do not interfere with each other neither sterically nor
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_____________________________________________________________ Results and Discussion
electrostatically, a maximal hybridization efficiency of ideally 100 % is obtained providing a
sufficient concentration of tDNA. Therefore, an increase of the ssDNA coverage leads to an
increase in the hybridization yield and in the current signal obtained from labelled tDNA
(observed for DNA immobilization up to 2 min). However, further increase of the ssDNA
concentration causes a steric and electrostatic hindrance depending on the ionic strength
towards the hybridization process and the efficiency of hybridization decreases leading to a
lower signal. Therefore, for the employed detection scheme (hybridization with Fc-tDNA, 20-
mer probe DNA) the optimal ssDNA coverage is around 2-3 × 10 12 molecules/cm 2 , achieved
by immobilization of ssDNA for 2 min using the developed potential-assisted immobilization
method.
Figure 3.47. a) CV characterization of the surface during sensor preparation, and b)
FSCV measurements before and after hybridization of the ssDNA/MCU-modified electrode
with Fc-tDNA. ssDNA immobilization: 10 mM PB, 450 mM K2SO4, 1 µM DNA; p.a.
immobilization (0.5/-0.2 V vs. Ag/AgCl/3 M KCl, 10 ms pulse duration), 2 min. MCU
passivation: 10 mM PB, 20 mM K2SO4, 1 mM MCU (30 % ethanol), 0.5/-0.2 V vs.
Ag/AgCl/3 M KCl (10 ms), 1 min. CV and FSCV measurements were performed as
explained in Figure 3.46.
This chapter demonstrates that using the developed potential-pulse assisted immobilization
technique DNA sensor preparation can be done in as little as 3 min depending on the desired
DNA coverage. Compared to the standard sensor preparation praxis where surface modification
3.4 Potential-assisted preparation of DNA sensors 89

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usually occurs over a period of two days, the ability to reproducibly prepare DNA sensors
within minutes presents a significant improvement.
Figure 3.48. Dependence of dsDNA coverage on the total duration of ssDNA immobilization.
After performing ssDNA immobilization (0.5/-0.2 V vs. Ag/AgCl/3 M KCl pulse
profile, 10 ms pulse duration, 10 mM PB with 450 mM K2SO4 and 1 µM DNA) for a
predefined time, all electrodes were passivated with MCU (1 min, 0.5/-0.2 V vs. Ag/AgCl/3
M KCl pulse profile, 10 ms pulse duration, 10 mM PB with 20 mM K2SO4 and 1 mM
MCU) and subjected to hybridization with Fc-tDNA (10 min incubation, 37 °C, 10 mM
PB, 450 mM K2SO4, 1 µM DNA). FSCV:10 mM PB, 450 mM K2SO4, 1 V/s. Coverage
determination as described in Section 5.9.3.
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3.4.2 DNA microchip fabrication
DNA chips are becoming tremendously important in diagnostics 2,3 . They allow for a
simultaneous analysis of a large number of target molecules and have become a crucial tool in
genomic analysis. The development of an electrochemically produced DNA microarrays may
provide highly controlled surfaces complemented with low production costs and time.
Therefore, the applicability of the developed potential-assisted immobilization technique for
the production of DNA microarrays was investigated using a 32-electrode chip comprising of
(70 × 70) µm 2 gold electrodes (Figure 3.49) by modifying the chip with two different DNA
sequences.
Figure 3.49. a-e) Schemes of the chip modification sequence: a) bare chip; b) after
potential-assisted immobilization of columns (C) 1 and 2 with the FRIZ sequence; c)
potential-assisted MCU passivation of C1 and C2; d) potential-assisted immobilization of
C5 and C6 with an E. coli sequence performed after the potential-assisted cleaning; e)
potential-assisted cleaning of C3 and C4 and subsequent potential-assisted MCU
passivation of C3-C8; during each step the whole chip was exposed to each solution; f)
scheme of the final chip modification with the region (a) representing ssFRIZ/MCUmodified
electrodes, region (b) representing MCU-modified electrodes, region (c)
representing E. coli/MCU-modified electrodes and region (d) representing ssFRIZ/E.
coli/MCU-modified electrodes; picture of the chip with the zoomed picture of the 32
electrodes (taken from ref. 100 ).
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DNA chips were prepared employing potential-assisted immobilization and desorption methods
by initially immobilizing a DNA sequence 1 (20-mer, “FRIZ sequence” in the further text) on
the first two columns of the chip for 30 s with a subsequent passivation with MCU for 1 min
(Figure 3.49, a-c). Since the whole chip was immersed in the same solution the other electrodes
were subjected to immobilization at OCP. In order to prevent any undesired signal on these
columns, prior to the immobilization with the sequence 2 (42-mer, “E. coli sequence”) columns
5 and 6 were cleaned by the potential-assisted desorption for 5 s. Afterwards, they were
modified with an E. coli sequence for 30 s (Figure 3.49, d). Subsequently, columns 3 and 4
were cleaned with the potential-assisted desorption (5 s) and finally columns 3-8 were
passivated with MCU (Figure 3.49, e). It should be noted that the aim of the experiment was
not the optimization of the assay parameters but rather a proof of concept. Therefore, the
employed immobilization and passivation times are not optimized. Independently it was
observed that an immobilization time of only 5 s and a passivation time of 3 s are sufficient to
achieve optimal DNA coverage for the employed system and the desired passivation of the
electrode surface, respectively (data not shown). Compared to macro electrodes, the optimal
modification time is much shorter, which can be explained by the improved diffusion profile of
microelectrodes.
By designing the experiment in the explained manner the obtained DNA chip is supposed to
consist of 4 different regions: region a modified with FRIZ/MCU, region b modified only with
MCU, region c modified with E. coli/MCU and region d where both FRIZ and E. coli
sequences are immobilized by incubation at OCP and a subsequent passivation is done by the
potential-assisted method (Figure 3.49, f). The whole chip was exposed to each solution leaving
the possibility of contamination between different regions. Therefore, to verify the quality of
preparation, the chip was subjected to hybridization using both target DNA sequences, FRIZ
and E. coli tDNA, labelled with ferrocene. The results obtained by FSCV are shown in Figure
3.50. In general, it can be observed that each region that was exposed to the same conditions
shows reproducible results among individual electrodes (Figure 3.50, a).
Figure 3.50, b and c compare representative electrodes from each region upon hybridization
with FRIZ and E. coli tDNA sequences, respectively. In order to understand what these figures
show, each region is discussed for both cases. Region a is the only region where FRIZ ssDNA
immobilization was performed via the potential-assisted method and therefore should show the
hybridization signal from Fc-labelled FRIZ-tDNA. Indeed, this is observed on each electrode
3.4 Potential-assisted preparation of DNA sensors 92

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Figure 3.50. a) Response after hybridization of a complete chip with FRIZ (orange curve)
or E. coli sequences (green curve) and baseline corrected FSCVs presenting cathodic
peaks after hybridization with b) FRIZ and c) E. coli sequence.
from region a (Figure 3.50, b, orange curve). The obtained signal is not very high since the
amount of hybridized DNA is in the order of 10 11 molecules/cm 2 . The reason for this is that the
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_____________________________________________________________ Results and Discussion
immobilization time employed in the experiment (30 s) is not optimal and already leads to a too
high ssDNA coverage for the employed detection scheme resulting in a decreased hybridization
efficiency. By doing this the rest of the electrodes were longer exposed to immobilization by
incubation allowing to investigate, how a possible contamination influences the quality of the
chip preparation. After potential-assisted immobilization of FRIZ ssDNA and passivation with
MCU, region a was also exposed to E. coli ssDNA. There was no signal upon hybridization
with E. coli tDNA showing the high integrity of the formed FRIZ/MCU films.
Region b was exposed to FRIZ ssDNA solution, and subsequently to the MCU passivating
solution and finally to the E. coli ssDNA solution, with immobilization occurring by incubation
in all cases. Prior to the final potential-assisted passivation with MCU, the electrodes were
cleaned using the potential-assisted desorption method. Thus, this region should be modified
only with MCU. Figure 3.50, b and c confirm this by showing the absence of any signal from
FRIZ and E. coli tDNA (grey curve). In order to prove that the absence of signals is due to a
successful desorption and not due to an undetectable amount of immobilized DNA, region d
was exposed to same conditions, without performing any desorption step. That region should
be modified with both FRIZ and E. coli ssDNA sequences and passivated completely with
MCU. In both cases (Figure 3.50, b and c, beige curve) a very small amount of tDNA is detected
upon hybridization with both FRIZ and E. coli tDNA, however, this signal is negligible as
compared to the signal obtained in the regions with potential-assisted immobilization.
Nevertheless, the amount of contaminating DNA is measurable, proving that desorption in
region b was done successfully.
Finally, region c was initially exposed to FRIZ ssDNA solution at OCP, after which the
electrodes were cleaned by potential-assisted desorption and subsequently immobilized with E.
coli ssDNA and passivated with MCU by the potential-assisted method. Therefore, it is the only
region that should show a signal upon hybridization with the E. coli tDNA sequence (Figure
3.50, c, green curve). The absence of any peak upon hybridization with FRIZ tDNA (Figure
3.50, b, green curve) also indicates that desorption was performed successfully in this region as
well.
Potential-assisted methods for immobilization and desorption allow for a fully
electrochemically controlled preparation of DNA chips within a very short time. The big
advantage of the potential-assisted immobilization method is that it provides the desired DNA
coverage and blocking of the surface tremendously faster than the process at OCP and by this
3.4 Potential-assisted preparation of DNA sensors 94

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it does not allow undesired contamination, which guarantees a high quality of the prepared
DNA chips. Additionally, the construction of DNA chips in this way allows for the use of even
smaller electrodes. The size of the electrode does not need to be limited by the size of the droplet
used for modification or its evaporation rate since there is no need for covering individual
electrodes for their modification with the solution of interest.
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3.5 Intercalation as a DNA detection technique
Synthesis of the intercalator was done by Dr. Adrian Ruff. Results from this chapter are
included in the manuscript to be submitted: “Amperometric detection of dsDNA via acrydinium
orange modified glucose oxidase” with coauthors: D. Jambrec, A. Ruff, W. Schuhmann, written
by A. Ruff (synthesis part) and the author.
In order to implement electrochemical sensing platforms into point-of-care diagnostic devices
plenty of improvement still needs to be done. Besides the design of the sensing platform, the
applied strategy for the sensing process itself is of tremendous importance defining the
sensitivity of the chosen approach. Even though covalent labelling of target DNA using dyes
or enzymes significantly enhances the sensitivity of DNA detection schemes, it demands
sample preparation and lacks of simplicity 101 . In contrast, label-free electrochemical detection
methods provide several advantages such as miniaturization, fabrication of low cost devices,
operation simplicity and rapid detection time. Therefore, they are of particular interest in
personal diagnostics using point-of-care devices 102 .
Among label-free approaches, sandwich-type assays, in which a third labelled signaling
sequence hybridizes the overhang of the target DNA sequence, have been extensively
investigated. Nevertheless, this strategy requires the use of different sequences for each
investigated target DNA, which increases the complexity and cost of the assay. Furthermore,
3.5 Intercalation as a DNA detection technique 96

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EIS was employed for label-free detection of DNA hybridization. Even though this technique
is very sensitive to surface modification, it is not yet sensitive enough to achieve the desired
sensitivity. On the other hand, the use of compounds that are able to intercalate into the DNA
helix is very promising. One of the main advantages of intercalators is that these molecules can
be used universally with any given DNA sequence. This significantly decreases the complexity
of the system and allows for the application in multiple probe DNA chips. Therefore, this
chapter focuses on the investigation of a new intercalation compound, an acridine orange based
intercalator (AO) conjugated with glucose oxidase (GOx) from Aspergillus niger.
Synthesis of a ferrocenyl labelled intercalation compound for the detection of double stranded
DNA (dsDNA) was reported recently 102 . However, this strategy is limited by low currents in
case of low dsDNA coverages due to the oxidation of the 1e - redox-species, i.e., the ferrocene
derivative. On the other hand, redox-enzyme conjugated intercalation compounds offer much
higher efficiency at low dsDNA coverages due to the intrinsic catalytic amplification by
catalytic redox-conversion of a specific substrate and thus the continuous recycling of a redox
mediator/redox probe by the enzyme. DNA hybridization detection schemes based on redoxenzymes
modified with an intercalation compound were previously reported 103,104 in which the
authors used the well-known biotin-streptavidin chemistry. By constructing an intercalator that
is covalently bound to an enzyme, one assay preparation step is avoided as compared to
previously described intercalator-enzyme conjugates 103,104 ) in the DNA detection scheme.
Figure 3.51. Scheme of the synthesized acridine orange-glucose oxidase intercalating
compound.
Since acridine orange is a well-established intercalator selective towards dsDNA 105,106 it was
chosen as a basis to synthesize an intercalator-enzyme conjugate. Glucose oxidase is a widely
3.5 Intercalation as a DNA detection technique 97

_____________________________________________________________ Results and Discussion
used enzyme in amperometric biosensors. In the presence of glucose and a suitable redox
mediator it provides enhanced catalytic currents due to the continuous recycling of the redox
probe 107 . The modification of GOx is readily achieved by reductive amination of the
deglycosylated enzyme. Therefore, the acridine orange-glucose oxidase conjugate (AO-GOx)
was synthesized by binding the acridine orange moiety via a flexible 14-atom-long tether to
deglycosylated GOx via reductive amination using NaBH4 (Figure 3.51).
To assure that the enzyme remains active towards the oxidation of glucose even after the harsh
conditions employed during synthesis, electrochemical characterization of AO-GOx was
performed. CVs were recorded in PB containing ferrocene methanol as redox mediator and
glucose as substrate for GOx before and after addition of AO-GOx. The observed catalytic
current upon addition of the intercalating compound confirms that the enzyme remained intact
and that it can still transfer electrons to the electrode via the redox mediator (Figure 3.52).
Figure 3.52. Electrochemical characterization of the AO-GOx intercalating compound.
CVs recorded at a bare gold electrode in a solution of 10 mM PB, 450 mM K2SO4
containing 1 mM ferrocene methanol and 100 mM glucose without AO-GOx (grey line)
and with increasing concentrations of AO-GOx (black lines). Measurements were made
with a scan rate of 10 mV/s in a potential window of -0.1 to +0.45 V vs. Ag/AgCl/3 M KCl.
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Figure 3.53. Scheme of the characterization sequence: a) ssDNA immobilization (1 min,
0.5/-0.2 V pulse profile with 10 ms pulse duration); b) passivation with MCU (1 min, 0.5/-
0.2 V pulse profile with 10 ms pulse duration); c) hybridization with Fc-tDNA (10 min
incubation); d) dehybridization (H2O, 10 min); e) hybridization with a non-labelled tDNA
(10 min incubation); f) intercalation with AO-GOx (15 min, incubation) and detection of
the glucose oxidation current; g) dehybridization and removal of AO-GOx (H2O, ethanol,
~10 min); h) evaluation of the interaction of AO-GOx with ssDNA (15 min, incubation)
by amperometric detection. Chronoamperometric measurements were conducted at an
applied potential of +400 mV. FAD: flavin adenine dinucleotide. All potentials vs.
Ag/AgCl/3 M KCl; drawing not to scale.
In order to optimize DNA detection via the synthesized AO-GOx a thorough characterization
sequence was employed, which is presented in Figure 3.53. The DNA assay build-up consisted
of an initial ssDNA immobilization (Figure 3.53, a) followed by passivation of the surface with
MCU (Figure 3.53, b). The created DNA sensing platform was then subjected to hybridization
3.5 Intercalation as a DNA detection technique 99

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with Fc-tDNA to indirectly determine the ssDNA coverage which is possible due to a low
ssDNA coverage, Figure 3.53, c. After subsequent dehybridization (Figure 3.53, d), free ssDNA
was again hybridized with a fully complementary non-labelled tDNA sequence (Figure 3.53, e)
and subsequent intercalation of AO-GOx was performed (Figure 3.53, f) followed by chronoamperometric
measurement. Afterwards, tDNA and AO-GOx were removed from the electrode
by H2O and ethanol (Figure 3.53, g), the remaining ssDNA/MCU-modified surface was again
immersed into the AO-GOx solution (Figure 3.53, h) and the signal of the negative control was
measured by subsequent amperometric experiments.
Control of the ssDNA coverage and minimization of unspecific adsorption was achieved by
careful surface preparation based on the developed potential-assisted surface modification
method. The envisaged pDNA coverage is lower than in the case of a direct detection of
hybridization (e.g., via detection of Fc-tDNA). Due to the large dimension of the enzyme
(diameter of the native enzyme is around 7 nm 108,109 ) and with this the entire intercalation
compound, ssDNA coverage must be low to prevent steric hindrance of neighboring DNA
strands towards intercalation. Shortening the immobilization time to obtain ssDNA coverage in
the order of 10 11 molecules/cm 2 provides enough space for the intercalating compound to
interact with dsDNA.
After the immobilization of ssDNA the electrode was passivated with MCU by applying the
potential-assisted method. One of the main challenges in DNA detection schemes involving
intercalators is to reduce the signal of the negative control, i.e. the interaction of the intercalator
with the ssDNA/thiol-modified surface. This interaction can occur either with the ssDNA or
from unspecific adsorption of the intercalator on the electrode surface. Since it is reported that
acridine orange solely intercalates into dsDNA 105,106 interaction with the ssDNA is expected to
be minimal. Thus, it is important to provide well-passivated surfaces that prevent significant
unspecific adsorption of the intercalator to allow for a high signal-to-noise ratio. DNA detection
via GOx amplification is based on the re-oxidation of the redox mediator at the electrode
surface. Therefore, the electron transfer between the mediator and the electrode needs to be
allowed. Previously it was reported that MCU provides the best compromise between electrode
passivation and permeability 100 . Applying the developed approach for potential-assisted
passivation of the surface with MCU for 1 min it was possible to clearly differentiate between
dsDNA and ssDNA-modified electrodes and obtain a high signal-to-noise ratio.
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In order to perform the experimental sequence as explained in Figure 3.53, the possibility of
removal of the intercalating compound was tested using the ssDNA/MCU-modified electrode.
Upon incubation with AO-GOx, a chronoamperometric measurement was performed in
phosphate buffer containing the redox mediator (ferricyanide, 1 mM) by applying a potential
of +400 mV (vs. Ag/AgCl/3 M KCl). A small interaction of the intercalator with the
ssDNA/MCU modified electrode was detected, shown in Figure 3.54 as a catalytic current upon
addition of glucose. Afterwards the electrode was treated with ethanol and water to remove
unspecifically bound intercalator and the chronoamperometric measurement was repeated. The
absence of any catalytic current upon addition of glucose in this case indicates that the
intercalator was completely removed from the surface.
Figure 3.54. Chronoamperometry with a ssDNA/MCU-modified electrode a) after
incubation with AO-GOx for 15 min, and b) after removal of AO-GOx. A potential of
+400 mV (vs. Ag/AgCl/3 M KCl) was applied; 10 mM PB containing 450 mM K2SO4 and
1 mM ferricyanide. Glucose was injected as shown on the figure.
Furthermore, to investigate how the intercalation process influences the stability of the formed
double helix, a control experiment was performed by following each step of the assay by means
of FSCV. Namely, initial hybridization of the ssDNA/MCU electrode was performed using FctDNA
and subsequent intercalation was done by leaving the labelled tDNA on the surface.
FSCV does not show any significant change in the response after intercalation, except of a small
shift of the oxidation peak. The DNA coverage remains unchanged suggesting that the
intercalation process does not impede the stability of the double helix (Figure 3.55).
3.5 Intercalation as a DNA detection technique 101

_____________________________________________________________ Results and Discussion
Figure 3.55. FSCV recorded for a ssDNA/MCU (grey line) and dsDNA/MCU-modified
electrode before (black line) and after intercalation with AO-GOx into the dsDNA (green
line). FSCV measurements were conducted in 10 mM PB solution containing 450 mM
K2SO4 with a 1 V/s scan rate. DNA immobilization and subsequent passivation were
performed by potential-assisted immobilization (0.5/-0.2 V vs. Ag/AgCl/3 M KCl pulse
profile with 10 ms pulse duration), each for 1 min. Immobilization was performed in 1
µM DNA solution in 10 mM PB with 450 mM K2SO4. Passivation was performed with 1
mM MCU in 10 mM PB with 20 mM K2SO4 (30 % ethanol). Hybridization was done by
incubation for 10 min in 1 µM Fc-tDNA solution in 10 mM PB with 450 mM K2SO4 at 37
°C. Intercalation was performed by drop coating of an aliquot of an AO-GOx solution (in
100 mM PB) on the dsDNA/MCU-modified electrode (incubation time: 15 min).
Figure 3.56 shows the possibility to clearly differentiate between dsDNA and ssDNA-modified
electrodes via AO-GOx intercalation using the developed DNA detection scheme based on
potential-assisted surface modification to control the ssDNA coverage and the passivation of
the surface. Normalized steady state currents with respect to dsDNA-modified electrodes are
extracted from I-t curves after the addition of glucose (40 mM) measured with ssDNA and
dsDNA-modified electrodes. After removal of tDNA and AO-GOx the electrodes were again
incubated in the AO-GOx solution and the amperometric measurement was repeated in order
to investigate the interaction of the AO-GOx with ssDNA (negative control). Higher currents
3.5 Intercalation as a DNA detection technique 102

_____________________________________________________________ Results and Discussion
can be detected for dsDNA-modified electrodes (right column) than for the negative control
(left column). This shows that the 14-atom-long and flexible tether between the AO-moiety and
the enzyme allows for an at least partial intercalation of AO into dsDNA. AO-GOx is able to
intercalate into dsDNA and upon addition of glucose, the enzyme catalyzes the oxidation of
glucose to gluconolactone. The reduced enzyme transfers the electrons via a ferricyanide to the
electrode surface. The obtained catalytic current proves the presence of AO-GOx and with this
DNA hybridization.
Figure 3.56. Background corrected and normalized currents measured with ssDNA/MCU
(left) and dsDNA/MCU (right) electrodes exposed to AO-GOx solution upon addition of
glucose (40 mM) into the electrolyte solution. Chronoamperometric detection was
performed in 10 mM PB containing 450 mM K2SO4 and 1 mM ferricyanide at an applied
potential of +400 mV (vs. Ag/AgCl/3 M KCl). Preparation of electrodes was performed as
explained in Figure 4.55. Error bars represent standard deviation between measurements
(n = 3).
It should be noted that the absolute currents vary for different electrodes, however, the ratio
between the currents obtained for dsDNA and ssDNA-modified electrodes is rather constant
(see error bar for the left column). Therefore, following the developed procedure, well defined
DNA-modified surfaces are obtained that ensure a high signal-to-noise ratio. The conversion of
glucose by the enzyme and the continuous regeneration of the redox probe ensure a high signal
amplification. Thus, the novel sensing platform allows for the clear differentiation between dsand
ssDNA even at low surface coverages.
3.5 Intercalation as a DNA detection technique 103

4. Conclusions

______________________________________________________________________ Conclusions
Self-assembly is a very powerful tool for surface modification, providing solutions for a large
number of applications from water repellent car windshields to the inhibition of corrosion in
industry. This thesis presents the development of a new strategy to achieve controlled selfassembly
of charged and uncharged thiolated molecules on gold surfaces in a very fast manner.
The main envisioned application is the development of a low-cost, yet powerful electrochemical
technique for the production of DNA chips as point-of-care devices. The explored strategy
consists of the potential-pulse assisted acceleration of the surface modification process by
applying a pulse-type potential modulation, using carefully selected potential pulse intensities
and an optimized pulse duration. Once optimized, this strategy increases significantly the
immobilization kinetics in a reproducible way as compared to passive self-assembly or
immobilization supported by the application of constant potentials.
The quality and cleanliness of the material used for surface modification is highly important for
the reproducibility and sensitivity of the envisaged sensing devices. Therefore, special attention
was given to the preparation of the surface used for modification. This resulted in a surface
preparation protocol defining certain criteria as a prerequisite to achieve a reproducible surface
architecture, reflected by the reproducible roughness factor and signals obtained from the
characterization via electrochemical impedance spectroscopy and cyclic voltammetry.
In order to tailor optimal DNA sensing surfaces, it is important to understand the processes
occurring at the electrode surface during the DNA assay build-up, taking into consideration the
physico-chemical properties of the investigated molecules, electrode polarization and the
surrounding solution. EIS was used to sequentially follow each step of the build-up of DNA
assays and understand how the variation of different parameters alters the quality of the surface
modification. The influence of surface modification by DNA on the value of the potential of
zero charge of polycrystalline gold was investigated by determining the pzc of bare gold before
and upon its modification with DNA. It was observed that the pzc shifts towards more negative
values due to the surface modification with DNA. The importance of this shift was evident in
the study for the selection of appropriate potential-pulse profiles to achieve high immobilization
rates of both DNA and alkylthiol derivatives as representatives of intrinsically charged and
uncharged molecules, respectively.
Based on previous findings showing that in solutions of high ionic strength charge screening of
DNA is significant and the potential profile in front of an electrode surface upon electrode
polarization is fairly steep, it is evident that the widely accepted model explaining the influence
105

______________________________________________________________________ Conclusions
of the polarized electrode on the behavior of DNA in its proximity is far too simple.
Attraction/repulsion of the DNA by the electric field in front of the polarized electrode can very
unlikely be responsible for an improved immobilization rate due to the shortness of the Debye
length. Thus, a new model is proposed, suggesting that the polarized electrode rather affects the
ions in the vicinity of the electrified interface. When the electrode is polarized to negative values
with respect to the pzc, cations move towards the electrified interface while anions move
towards the bulk of the solution. In contrast, the opposite behavior is obtained when the
electrode is polarized to more positive potentials with respect to the pzc. Switching fast enough
between these two situations creates a “stirring effect” that effectively exceeds the Debye length
in front of the electrified interface and pulls along DNA strands present in close proximity to
the electrode surface including their condensed ion cloud. Hence, immobilization is not
diffusion controlled but driven by the migration of ions in front of the electrode. Parameters
that are playing a crucial role in achieving a significant improvement in the immobilization
kinetics are the applied potential intensities and the duration of an individual pulse. The applied
potential intensities need to be on the one hand within the stable potential window of the Au-S
bond and on the other hand high enough to evoke an efficient stirring to bring the DNA towards
the surface. The potential-pulse duration needs to be long enough to allow for an appropriate
concentration gradient to form and for a whole molecule to be pulled down to the electrode
surface, hence allowing for the formation of the Au-S bond regardless of the orientation of the
molecule. On the other hand, it needs to be simultaneously short enough to allow a high number
of potential pulse cycles.
Using the developed strategy, alkylthiol SAM formation can also be significantly accelerated,
leading to the formation of compact SAMs within minutes. It was shown that, besides the
potential-pulse intensities and the pulse duration, the length of a molecule influences the
efficiency of the immobilization. Optimization of the potential difference does not depend on
the molecule length, as long as the appropriate pulse duration is chosen, since longer potential
pulses are needed to bring down longer molecules to the electrode. The fact that this strategy
tremendously accelerates the immobilization of uncharged molecules as well supports the
conclusion that only DNA attraction/repulsion by the polarized electrode is a far too simple
model to explain the mechanism of DNA immobilization supported by applied potentials.
The developed strategy was subsequently implemented into the development of a DNA sensor,
using the developed potential-pulse assisted method for both DNA immobilization and thiol
106

______________________________________________________________________ Conclusions
passivation steps to create a DNA sensing platform. The optimization of the protocol was
performed by investigating the influence of subsequent passivation step and the potential
pulsing itself on the stability of DNA-modified surfaces and by choosing the right thiol for
passivation.
Furthermore, the applicability of the developed method for the production of DNA chips was
investigated using a 32-electrode array chip. Using this approach, the whole chip is exposed to
each DNA solution used for the modification of a certain number of electrodes, avoiding the
need for a localized positioning of DNA solutions on individual electrodes as in light-directed
synthesis or spotting procedures that require expensive and sophisticated equipment. Due to the
goal of making multiple probe DNA chips, it is clear that electrodes need to be cleaned prior to
the modification with a selected DNA sequence. Thus, a method for potential-pulse assisted
cleaning of Au-modified surfaces was developed, which allows a very fast and efficient
regeneration of individual electrodes without causing any damage to the surface. The fact that
by using our approach the Au surface can be easily regenerated without jeopardizing the quality
of the following surface modification, points also to chip recycling as another advantage of the
proposed approach. Until now research in this field did not address this topic, but relatively
soon, in the world of private medicine, where everyone will own point-of-care devices, the
question of reusability will become an important issue. Furthermore, construction of DNA chips
in this manner opens the door towards the use of even smaller electrodes, since their size does
not need to be limited by the size of the droplet used for modification or its evaporation rate.
Finally, the developed technique was implemented into the development of a new DNA sensing
platform based on signal amplification via an enzyme-conjugated intercalating compound
(glucose oxidase-acridine orange) as hybridization indicator. Using the potential-pulse assisted
surface modification strategy, DNA/thiol surfaces were fabricated to obtain lower probe DNA
coverage and very efficient blocking of unspecific adsorption, which resulted in a significant
contrast between ss- and dsDNA. The developed DNA sensing strategy finds its application in
multiple probe DNA chips, since the synthesized intercalating compound can universally
interact with all DNA sequences present on the chip and, by this, simultaneously increase the
sensitivity on all electrodes.
Even though this work is primarily orientated towards DNA chip development, the developed
potential-pulse assisted surface modification method is an equally promising concept for
applications in many other research fields, such as protein binding or investigation of cells.
107

5. Experimental
Work

________________________________________________________________ Experimental Work
5.1 Materials and consumables
The following materials were bought from Sigma-Aldrich Chemie (Germany): 6-mercapto-1-
hexanol, 11-mercapto-1-undecanol, [Ru(NH3)6]Cl3, K3[Fe(CN)6] and K4[Fe(CN)6]. 16-
mercapto-1-hexadecanol was from Frontiers Scientific (United States). KH2PO4, K2HPO4 and
K2SO4 were from VWR International (Germany). Glucose was purchased from AppliChem
(Germany) and was used one day after the preparation of the solution to allow for mutarotation.
All reagents were of analytical grade and used as received. Polishing cloths were bought from
Leco (USA) including polishing pastes with 3 µm, 1 µm and 0.5 µm particle size. 0.1 µm
particle size polishing paste was from Struers (Germany). Non-modified (bare gold) 32-
electrode chips were bought from FRIZ Biochem (Germany).
Glucose oxidase from Aspergillus niger (type X-S, lyophilized powder, 100000–250000 U/g)
was purchased from Sigma-Aldrich and stored at -22 °C. All DNA probes were purchased from
FRIZ Biochem (Germany) and are listed in Table 5.1.
Table 5.1. Summary of DNA sequences used in all experiments. Sequences marked with –
SS were used as probe DNA, the rest of sequences were employed as target DNA. DTPA
stands for dithiophosphoramidite. Fc stands for ferrocene. A – adenine, T – thymine, C –
cytosine, G – guanine.
Internal name
Sequence
FRIZ 12-SS
5’ TGC GGA TAA CAC AGT CAC CT TTTTTTT (DTPA)3
FRIZ 5 tDNA
5’ AGG TGA CTG TGT TAT CCG CA
FRIZ 5-Fc
5’ (Fc)4 AGG TGA CTG TGT TAT CCG CA
E. coli-SS
E. coli-Fc
5’ GTC AAT GAG CAA AGG TAT TAA CTT TAC TCC
CTT CCT CCA TTA TTTTTTT (DTPA)3
5’ (Fc)4 GGA GGA AGG GAG TAA AGT TAA TAC CTT
TGC TCA TTG CG
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________________________________________________________________ Experimental Work
Synthesis of the acridine orange-glucose oxidase intercalator was done by Dr. Adrian Ruff
following procedures from Biver et al. 105 (for the modification of acridine orange) and
Schuhmann et al. 110 (for deglycosylation of glucose oxidase).
5.2 Electrochemical setup and instrumentation
All measurements were done in an electrochemical cell consisting of a polycrystalline gold
working electrode (2 mm diameter, CH Instruments, USA), a platinum auxiliary electrode
(Goodfellow, Germany) and a homemade Ag/AgCl (3 M KCl) or Pb/PbF2 (5 M KF) reference
electrode. Unless stated otherwise all electrochemical measurements were performed with an
Autolab PGSTAT302N containing a frequency response analyzer (Metrohm-Autolab,
Netherlands).
Figure 5.1. a) CV representing the deposition of PbF2 (2 nd cycle) performed at 20 mV/s in
0.5 M NaF solution. b) Picture of a home-made Pb/PbF2 reference electrode.
Preparation of the lead-lead fluoride reference electrode was done according to 81 . A ceramic
frit was introduced at the tip of a Pasteur pipette and the surrounding glass was melted to seal
it and prevent later electrolyte leakage. Subsequently, the tip was polished to expose the frit and
allow later ionic contact with the electrolyte solution. The pipette was filled with 5 M KF
solution and a piece of cotton was placed in the pipette construction to prevent clogging of the
electrode by detached PbF2. A lead wire was connected to a copper wire and the connection
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________________________________________________________________ Experimental Work
was isolated with a shrink-tube. PbF2 was electrochemically deposited on the lead wire in a 0.5
NaF solution by cyclic voltammetry (Figure 5.1, a). During the first scan the potential was
initially decreased to negative values in order to remove the native oxide from the surface of
the lead wire and by this improve the deposition of lead fluoride. The deposition was done for
20 cycles at a scan rate of 20 mV/s. After deposition the electrode was kept in the same solution
and chronoamperometry was performed for 2 h at 0.2 V vs. Ag/AgCl (3 M KCl) in order to
improve the stability of the deposited PbF2 film. Finally, the lead wire with deposited PbF2 was
inserted in the prepared pipette and fixed (Figure 5.1, b). The prepared electrode was usually
left overnight to assure stabilization of the electrode potential. The electrode potential was
controlled before use and it was around -600 mV vs. Ag/AgCl (3 M KCl).
5.3 Preparation of gold surfaces
Figure 5.2. Home-made polishing machine.
Polycrystalline gold electrodes were mechanically and electrochemically cleaned before each
experiment in the following manner. Electrodes were mechanically polished using a homemade
automatic polishing machine built by Dr. Kirill Sliozberg (Figure 5.2). The polishing
machine consisted of a rotating disk covered with an appropriate polishing cloth and a holder
for two electrodes that was simultaneously moved in x direction keeping the electrodes
perpendicular to the surface of the rotating disk. By this polishing was performed in all
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________________________________________________________________ Experimental Work
directions equally, allowing the formation of homogenous surfaces. The procedure was
controlled by a home-made software. Polishing clothes were soaked in water prior to polishing
and covered homogeneously with an appropriate polishing paste. Electrodes were polished with
diamond pastes of 3, 1, 0.5 and 0.1 µm particle size successively, using separate polishing cloths
for each paste. Polishing time depended on the state of the electrode.
Afterwards, the gold electrodes were electrochemically cleaned by cyclic voltammetry in 0.5
M H2SO4 until reproducible voltammograms were obtained. Cycling was performed in the
potential range of 0 V to 1.6 V vs. Ag/AgCl (3 M KCl) at 100 mV/s scan rate. The roughness
factor was kept around 1.4 and its determination was done according to 111 . The method is based
on the determination of oxygen adsorption on a gold surface. During a positive potential scan
chemisorption of oxygen occurs, where different gold facets have different oxidation potentials
(Figure 5.3). During the negative potential scan oxides are reduced, which can be seen as a
single reduction peak. Assuming that oxygen atoms form a monolayer on the gold surface, the
roughness factor of an electrode can be determined using the following equations:
R = A real
A g
(5.1)
A g = r 2 π (5.2)
A real = Q exp
Q theor
(5.3)
Q exp =
peak area
ν
(5.4)
where Ag is the geometrical surface area of the electrode (0.0314 cm 2 for a 2 mm diameter gold
electrode), Areal is the real surface area, Qexp is the charge involved in the reduction of gold
oxide, Qtheor is 482 μC/cm 2 and it is a calculated charge value for chemisorption of an oxygen
monolayer on the surface of polycrystalline gold, based on the density and atomic weight of
gold 112 , υ is the scan rate of the measurement and peak area is the area under the reduction peak
from the CV measurement obtained by integration of the peak (Figure 5.3, area marked in
green).
After mechanical polishing and electrochemical cleaning, the electrodes were immediately
characterized by EIS and CV to verify their cleanliness prior to modification (procedures
explained in Sections 5.9.1 and 5.9.2, respectively).
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________________________________________________________________ Experimental Work
Figure 5.3. Cyclic voltammogram (20 th cycle) of a bare polycrystalline gold electrode in
0.5 M H2SO4 at 100 mV/s scan rate. The area under the reduction peak is marked in green.
5.4 Determination of the potential of zero charge
The determination of the potential of zero charge was done via potentiodynamic
electrochemical impedance spectroscopy (PDEIS) in PB/K2SO4 solution of different ionic
strength – 10 mM PB with 20 mM K2SO4, 1 mM PB with 2 mM K2SO4 and 0.1 mM PB with
0.2 mM K2SO4. Prior to experiments, the solutions were purged with argon for at least half an
hour and the electrodes were prepared as explained in Chapter 5.3. A capacitive bridge
(capacitance of 2 µF) was used during measurements to avoid artifacts from the potentiostat
(Figure 5.4). The electrochemical setup consisted of a big cylindrical Pt counter electrode
surrounding the working electrode and a Pb/PbF2 reference electrode that was placed below the
working electrode.
Determination of the pzc of the bare electrode was done immediately after electrode preparation
to avoid any contamination of the surface. PDEIS was done for a potential range of -0.2 V to
0.7 V vs. Ag/AgCl (3 M KCl) with 30 mV potential steps. EIS was performed at each potential
step for a different range of frequencies, depending on the ionic strength and each potential was
superimposed with an AC perturbation of 10 mVpp amplitude.
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________________________________________________________________ Experimental Work
Figure 5.4. Scheme of the electrochemical setup used for determination of the pzc. The
implemented capacitive bridge (Cb) is shown in the figure. CE – counter electrode, RE –
reference electrode, WE – working electrode.
Determination of the pzc of the DNA modified surface was done by initially immobilizing
ssDNA (from 1 µM solution in 10 mM PB with 450 mM K2SO4) on a clean electrode surface
for 5 min via potential-assisted immobilization (0.5/-0.2 V (vs. Ag/AgCl/3 M KCl) following
a pulse profile with 10 ms pulse duration (procedure explained in detail in Section 5.7.2).
Afterwards the electrode was thoroughly rinsed with the immobilization buffer and water,
respectively to remove any loosely bound DNA strands, and placed in an appropriate solution
for pzc determination. PDEIS was performed for a potential range of -0.4 V to 0.5 V vs.
Ag/AgCl (3 M KCl) with 30 mV potential steps. EIS measurements for discrete potential values
were done as in the case of the bare electrode.
5.5 Potential-assisted formation of self-assembled monolayers
Preparation of gold electrodes for thiol self-assembling was done as explained in Chapter 5.3.
After characterization with EIS and CV (see Sections 5.9.1 and 5.9.2) gold electrodes were
immediately used for potential-assisted SAM formation of thiols of different length – MCH,
MCU and MCHD (1 mM in phosphate buffer, 20 mM K2SO4 with 30 % ethanol), using several
114

________________________________________________________________ Experimental Work
pulse profiles: 0.3/-0.1 V, 0.5/-0.2 V and 0.5/-0.4 V (vs. Ag/AgCl/3 M KCl) with various pulse
durations of 1 ms, 10 ms, 100 ms and 10 s.
Real-time impedance measurements were performed during potential-assisted SAM formation
to investigate the kinetics of self-assembly. This was done in a way that the applied pulse
profiles were considered as DC potentials which were superimposed with an AC signal of 1
kHz frequency and 5 mVrms amplitude. Experiments were performed using a setup that
consisted of a potentiostat with a summing amplifier (IPS AJ, Germany), a function generator
(Agilent 33120A, USA), a lock-in amplifier (EG&G Instruments 7265 DSP, USA) and a 16 bit
CIO-DAS 1602/16 AD/DA board (Plug-In electronic, Germany).
Determination of the capacitance was done by calculating initially the imaginary impedance
component (-Z'') using the recorded magnitude (|Z|) and phase (θ) values:
|Z| =
A AC
k|Z| meas
(5.5)
−Z ′′ = |Z|sinθ
where AAC is AC amplitude, k is the correction coefficient taking into account the current range
of the potentiostat (10 mA/10 V) and |Z|meas is the measured value of the magnitude. Using the
value of the imaginary impedance component the capacitance was calculated using the
following expression, derived from an RC series equivalent electrical circuit:
C = −1
ωZ ′′ (5.6)
Data sampling was done at a rate of 100 ms per data point and recorded with a home-made
software. Oscillations of the recorded signals coming from the applied pulse potential were
suppressed by applying a digital filter to the experimental plots. Data analysis and curve fitting
was performed with Origin.
5.6 Potential-assisted desorption
Potential-assisted desorption was performed using ssDNA-, MCU- or ssDNA/MCU-modified
electrodes. The experiment was done in 10 mM PB with 450 mM K2SO4 by applying a 0.9/-0.9
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________________________________________________________________ Experimental Work
V (vs. Ag/AgCl/3 M KCl) potential pulse profile with 10 ms pulse duration, for 5 or 30 s. Upon
desorption, the electrode was rinsed with the same buffer and water.
5.7 Preparation of DNA sensors
5.7.1 ssDNA immobilization via incubation
Immobilization of ssDNA via incubation was done immediately after cleaning and
characterization of a gold electrode as explained in Chapter 5.3. The electrode was placed in an
Eppendorf tube containing 200 μL of ssDNA solution (1 μM in 10 mM PB, 450 mM K2SO4)
and kept in a thermo-mixer (HTC BioTech, Germany) at 37 ºC. The electrode was sealed within
the tube using parafilm to prevent solution evaporation. The duration of the immobilization
procedure depended on the desired DNA coverage. Afterwards, the electrode was thoroughly
rinsed with the immobilization buffer (10 mM PB with 450 mM K2SO4) and water to remove
any loosely bound ssDNA.
5.7.2 Potential-assisted ssDNA immobilization
Figure 5.5. Electrochemical setup used for potential-assisted sensor preparation.
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________________________________________________________________ Experimental Work
Potential-assisted ssDNA immobilization was performed right after the electrode preparation,
in an electrochemical setup using small volumes (100-200 μL) as shown in Figure 5.5.
Measurements were performed in the same solution as the immobilization via incubation at
room temperature using different potential pulse profiles with various pulse durations. The total
duration of the immobilization process depended on the desired DNA coverage. In order to
remove unspecifically bound DNA from the electrode surface, electrodes were rinsed after
modification with the immobilization buffer.
5.7.3 Passivation by means of incubation
Passivation via incubation was done by placing the electrode in an Eppendorf tube containing
500 μL of a thiol derivative solution (in 10 mM PB, 20 mM K2SO4). The electrode was sealed
within the tube to prevent solution evaporation and kept in a thermo-mixer (HTC BioTech,
Germany) at 37 ºC for 19 h unless stated otherwise. Afterwards, the electrode was thoroughly
rinsed initially with absolute ethanol and then water to remove any loosely bound MCH.
5.7.4 Potential-assisted passivation
Potential-assisted passivation was performed using a 0.5/-0.2 V pulse profile with 10 ms pulse
duration. It was done in the same electrochemical setup as used for potential-assisted
immobilization (Figure 5.5). Passivation by MCH was done in a 10 mM solution with 10 mM
PB and 20 mM K2SO4, while passivation by MCU was performed in a 1 mM solution with 10
mM PB and 20 mM K2SO4. Measurements were performed at room temperature for 1 min
unless specified otherwise. In order to remove loosely bound thiols from the electrode surface,
the electrodes were rinsed with absolute ethanol and then water after modification.
5.8 Preparation of DNA chips
Prior to use, DNA chips were cleaned with piranha solution (cc. H2SO4 with 30 % H2O2, 3:1)
for 10 min. After thoroughly rinsing with water, the chips were electrochemically cleaned in
H2SO4 as explained in Chapter 5.3. Modification of chips was done using potential-assisted
procedures for immobilization, passivation and desorption. The detailed experiment sequence
is explained in Section 3.4.2.
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________________________________________________________________ Experimental Work
5.9 Characterization of DNA sensors
5.9.1 Electrochemical impedance spectroscopy
EIS measurements were performed after each assay preparation step in 10 mM PB with 20 mM
K2SO4 (pH 7.4) containing equimolar concentrations of K3[Fe(CN)6] and K4[Fe(CN)6] (5 mM
each). Experiments were conducted at the equilibrium potential of the redox couple (DC
potential, +220 mV vs. Ag/AgCl/3 M KCl) superimposed by an AC perturbation of 10 mVpp
amplitude. The frequency range from 30 kHz to 10 mHz was scanned unless stated otherwise.
The charge transfer resistance was determined by fitting the Nyquist plots to an [R(Q[RW])]
Randles equivalent circuit.
5.9.2 Cyclic voltammetry
Characterization of the surface before and during the assay preparation was performed also via
cyclic voltammetry in 10 mM PB with 20 mM K2SO4 (pH 7.4) containing equimolar
concentrations of K3[Fe(CN)6] and K4[Fe(CN)6] (5 mM). Measurements were performed at 100
mV/s scan rate in the potential window from -0.1 V to 0.5 V (vs. Ag/AgCl/3 M KCl).
5.9.3 Chronocoulometry for determination of DNA coverage
Determination of ssDNA coverage was done using the chronocoulometric method developed
by Steel and coworkers 75 . A potential step from 0 V to -0.4 V (vs. Ag/AgCl/3 M KCl) was
applied for 500 ms at ssDNA/MCH modified electrodes immersed in two different solutions,
initially in 10 mM PB solution containing 20 mM K2SO4 and subsequently in the same buffer
containing additionally 100 µM [Ru(NH3)6] 3+ . The resulting charge was measured in both cases.
Prior to measurements solutions were purged with argon for at least 30 min.
5.10 Hybridization and dehybridization
DNA hybridization was performed via incubation with the target strand after characterization
of ssDNA/thiol modified electrodes. The electrode was initially rinsed with the hybridization
buffer (10 mM PB, 450 mM K2SO4) and then placed in an Eppendorf tube containing 200 μL
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________________________________________________________________ Experimental Work
of the labeled or non-labeled target DNA solution (1 μM in 10 mM PB, 450 mM K2SO4) and
kept in a thermo-mixer (HTC BioTech, Germany) at 37 ºC for 10 min or 1 h. The electrode was
sealed within the tube to prevent solution evaporation. After hybridization the electrode was
rinsed with the hybridization buffer to remove any unspecifically bound tDNA.
Dehybridization of dsDNA was performed by incubation of the dsDNA/thiol electrode in water
for 5-10 min. The efficiency of dehybridization was confirmed using Fc-tDNA by measuring
the Fc redox peak before and after dehybridization using fast-scan cyclic voltammetry.
5.10.1 Detection of hybridization
Direct hybridization detection was done using different detection methods, depending whether
the tDNA was labeled or not. Hybridization with non-labeled tDNA was detected by means of
EIS following the change of Rct upon hybridization. EIS parameters were the same as in Section
5.9.1.
Hybridization with labeled tDNA was detected using FSCV in 10 mM PB with 450 mM K2SO4
at a scan rate of 1 V/s in the potential window from -0.05 V to 0.55 V (vs. Ag/AgCl/3 M KCl)
under argon atmosphere.
5.10.2 DNA coverage determination by means of FSCV
Direct determination of tDNA coverage or indirect determination of pDNA coverage (in case
of low coverages) was done by integration of the cathodic peak of the Fc label from FSCV
measurements. The cathodic peak area can be used to calculate the transferred charge:
Q =
peak area
ν
(5.7)
where peak area is the background corrected charge under the cathodic wave and v is the scan
rate used in the FSCV measurement. Then the surface coverage can be calculated:
Γ =
Q
nFA
(5.8)
where n is the number of electrons transferred in the redox process (in this case 4 since four
ferrocene moieties were used per DNA strand), F is the Faraday constant (96485 C/mol) and A
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is the real surface area of the electrode. In all experiments with the exception of multi-electrode
chip measurements 2 mm diameter gold electrodes were used having a geometrical surface area
of 0.0314 cm 2 and a real surface area calculated as explained in Section 5.3. The surface
coverage is obtained in mol/cm 2 .
Finally, to determine the amount of DNA molecules the following equation should be used:
where NDNA is DNA coverage obtained in molecules/cm 2 .
N DNA = Γ × N A (5.9)
5.11 Intercalation
Intercalation of AO-GOx was carried out at room temperature for 15 min by drop casting 10
µL of the intercalator solution (in 100 mM PB) on ssDNA/thiol or dsDNA/thiol-modified
electrodes. The electrode was subsequently rinsed with 10 mM PB containing 450 mM K2SO4
to remove unspecifically bound intercalator molecules.
Chronoamperometric measurements were performed in a 10 mM PB with 450 mM K2SO4
solution containing 1 mM ferricyanide that was purged with Ar for at least 30 min prior to
experiments. During measurements a constant potential of +400 mV (vs. Ag/AgCl/3 M KCl)
was applied and after stabilization of the background current 40 mM glucose was injected to
assure enzyme saturation. Measurements were done under argon atmosphere.
5.12 Methods
5.12.1 Electrochemical impedance spectroscopy
Impedance is a measure of the resistance of the system towards an alternating current and it
extends the concept of resistance to AC circuits. It is a complex function that can be represented
in two ways (Figure 5.6), namely by the real and the imaginary impedance components or by
the modulus and the phase shift:
Z = Z ′ + jZ ′′ (5.10)
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Z = |Z|exp(jθ) (5.11)
Z ′ = |Z|sinθ Z ′′ = |Z|cosθ (5.12)
where j = √−1, Z ′ is the real impedance component, Z′ ′ is the imaginary impedance
component, |Z| is the magnitude and θ is the phase shift between potential and current.
Electrochemical impedance spectroscopy is based on the application of a DC potential that is
superimposed with a small amplitude AC potential and measurement of the resulting AC
current signal:
E = E AC sin(ωt) (5.13)
i = i AC sin(ωt + θ) (5.14)
where E AC and i AC are potential and current amplitude, respectively. Except in the case of a pure
resistor, the measured current has a phase shift as compared with the applied potential (Figure
5.6). The impedance of the system is:
Z = E i = E ACsin(ωt)
i AC sin(ωt + θ) = |Z| sin(ωt)
sin(ωt + θ)
(5.15)
Figure 5.6. a) Phase diagram of the applied potential and measured current; b) impedance
vector diagram.
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As a result of the measurement, modulus and phase are sampled and the impedance is presented
usually in a Nyquist plot, since it provides fast information about system parameters, or
alternatively in a Bode plot when the frequency dependence needs to be investigated (Figure
5.7).
Figure 5.7. Representation of EIS data: a) Nyquist plot, b) Bode plot.
Figure 5.8. Current vs. potential dependence in electrochemical systems. A region
exhibiting quasi-linear behavior is shown.
For reliable data acquisition, two conditions need to be satisfied. Ideally, during EIS
measurements, the system should remain stationary, that is, the system parameters should not
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change during the experiment. Furthermore, since the interpretation becomes much more
complicated for non-linear systems, it is important to perform EIS experiments in a quasi-linear
regime for electrochemical systems (Figure 5.8). This condition determines the amplitude of
the applied AC potential since a smaller amplitude results in a more linear system. Depending
on the system, E AC is typically between 1 to 10 mV.
Table 5.2. Common electrical elements and equivalent electrical circuits, their impedance
and corresponding Nyquist plots.
Equivalent electrical circuit Impedance Nyquist plot
Z = R
Z =
1
jωC
Z = R −
j
ωC
1
Z = 1 R + jωC
EIS data are generally analyzed by fitting to an equivalent electrical circuit. Electrical elements
commonly used for the construction of equivalent electrical circuits are resistor, capacitor,
constant phase element and Warburg impedance. When designing an equivalent electric circuit
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it is important to understand the system under investigation, meaning that the electrical elements
need to have a physico-chemical meaning. Table 5.2 shows basic circuit models and their
impedance.
The so-called Randles equivalent circuit is commonly used to model interfacial electrochemical
reactions under semi-infinite linear diffusion control. A constant phase element (CPE) is
commonly used instead of a real capacitor due to frequency capacitance dispersion (Figure 5.9).
The CPE consists of two elements, Q0 that models the capacitance and n that represents the
degree of frequency dispersion with values from 0 to 1, where 1 represents pure capacitive
behavior and 0 pure ohmic resistance.
Figure 5.9. Randles equivalent circuit and corresponding Nyquist plot.
5.12.2 Chronocoulometry for the determination of DNA coverage
Determination of DNA coverage by means of chronocoulometry is based on the determination
of the charge corresponding to the amount of a cationic redox marker ([Ru(NH3)6] 3+ ) noncovalently
bound to DNA strands 75 that is later used to calculate the amount of immobilized
DNA on the electrode surface. Since measurements are performed at low ionic strength (10 mM
PB with 20 mM K2SO4), trivalent [Ru(NH3)6] 3+ can exchange native counterions surrounding
the DNA and non-covalently bind to phosphate residues on DNA strands. To ensure accurate
determination of the amount of DNA strands on the electrode surface, the measurements need
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to be conducted under saturation of the redox marker. That is, the method works under the
assumption that [Ru(NH3)6] 3+ exchanges all counterions screening the DNA.
Figure 5.10. Cyclic voltammogram of a bare gold electrode immersed in 10 mM PB
containing 20 mM K2SO4 and 100 μM [Ru(NH3)6]Cl3 in the potential window -0.4 to 0.1
V (vs. Ag/AgCl/3 M KCl) at 100 mV/s scan rate.
In order to measure the charge, a potential step is applied from a potential at which a negligible
reduction of the redox mediator is observed (0 V vs. Ag/AgCl/3 M KCl, Figure 5.10) to a
potential that corresponds to the diffusion limited current for the reduction of all surface
confined redox species (-0.4 V vs. Ag/AgCl/3 M KCl, Figure 5.10). Initially, the charge is
measured in a background solution without the redox marker, to obtain the double layer charge.
Subsequently, the measurement is repeated in the same solution containing additionally the
redox molecule [Ru(NH3)6] 3+ . The charge measured in this solution is given by the integrated
Cottrell equation:
Q = 2nFAD 1
2 C ∗
π 1 2
t 1 2 + Q dl + nFAΓ 0
(5.16)
where D is the diffusion coefficient of the redox molecule (cm 2 /s), C* is the bulk concentration
of [Ru(NH3)6] 3+ (mol/mL), Qdl is the double layer charge and the term nFAΓ0 is the charge
related to the reduction of surface confined species. Making an assumption that the double layer
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charge remains the same regardless of the presence of [Ru(NH3)6] 3+ the term nFAΓ0 can be
determined as the difference in the intercept at t = 0 from a Q-t 1/2 graph (Figure 5.11).
Figure 5.11. DNA coverage determination via chronocoulometry. Curves are shown for a
ssDNA/MCH-modified electrode in the absence (grey curve) or presence (green curve) of
[Ru(NH3)6] 3+ in solution.
Since it is assumed that each molecule of [Ru(NH3)6] 3+ binds to three phosphate groups from
the DNA backbone, the amount of DNA strands present on the electrode surface is then
calculated using the following equation:
Γ DNA = Γ 0 ( z m ) N A (5.17)
where z is the charge of the redox molecule and m is the number of phosphate groups at the
DNA strand.
126

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132

________________________________________________________________________ Appendix
7. Appendix
7.1 List of symbols and abbreviations
A
real surface area
a
polymer radius
AC alternating current
AD/DA analog-digital/digital-analog conversion
AO acridine orange
AO-GOx conjugate of acridine orange and glucose-oxidase
b
charge separation
C
interfacial capacitance
C* bulk concentration
C0
Cd
Cddl
CE
Cf
Ci
CPE
CV
d
D
DC
dsDNA
e
EAC
EIS
F
Fc
Fc-tDNA
initial capacitance
differential capacitance
capacitance of the diffuse double layer
counter electrode
capacitance of a fully covered monolayer
capacitance of the compact layer
constant phase element
cyclic voltammetry
distance from the surface
diffusion coeficient
direct current
double stranded DNA
elementary charge
potential AC amplitude
electrochemical impedance spectroscopy
Faraday constant
ferrocene
ferrocene-labelled tDNA
133

________________________________________________________________________ Appendix
FSCV
GC
GOx
I
iAC
k
L
lB
lel
lo
lp
m
MCH
MCHD
MCU
n
NA
Nbp
Nbp
NDNA
OCP
PB
pzc
pzc (DNA)
pzfc
pztc
Q
Qdl
Rct
RE
Rg
Rs
fast-scan cyclic voltammetry
Gouy-Chapman (theory)
glucose-oxidase
ionic strength
current AC amplitude
rate constant
contour length of DNA
Bjerrum length
electrostatic repulsion within a polymer
intrinsic stiffness of a polymer
persistence length
number of phosphate groups
mercapto-1-hexanol
mercapto-1-hexadecanol
mercapto-1-undecanol
number of electrons
Avogadro number
number of base pairs
number of bases
DNA coverage
open circuit potential
Poisson-Boltzmann (equation)
potential of zero charge
potential of zero charge of DNA-modified
electrode
potential of zero free charge
potential of zero total charge
charge
double layer charge
charge trasfter resistance
reference electrode
radius of gyration
solution resistance
134

________________________________________________________________________ Appendix
SAM
ssDNA
tDNA
W
WE
z
|Z|
Z'
-Z"
Γ
δ
ε
ε0
η
θ
κ -1
ν
σm
φ
φ0
ω
self-assembled monolayer
single stranded DNA
target DNA
Warburg constant
working electrode
valence
modulus
real impedance component
imaginary impedance component
surface coverage
thickness increment
dielectric constant
permeability of vacuum
charge density
partial coverage; phase shift
Debye length
scan rate
excess charge of a metal
potential at a distance d from the surface
potential at the surface
angular frequency
135

________________________________________________________________________ Appendix
7.2 Publications list
Published papers:
D. Jambrec, F. Conzuelo, A. Estrada-Vargas, W. Schuhmann, “Potential Pulse-Assisted
Formation of Thiol Monolayers within Minutes as a Promising Technique for Fast and
Controlled Electrode Surface Modification”. ChemElectroChem, 2016, 3, 1484-1489.
D. Jambrec, R. Haddad, A. Lauks, M. Gebala, W. Schuhmann, M. Kokoschka, “DNA
Intercalators for Detection of DNA Hybridization. SCS(MI)-MP2 Calculations and
Electrochemical Impedance Spectroscopy”. ChemPlusChem, 2016, 81, 604-612.
A. Estrada-Vargas, D. Jambrec, Y. U. Kayran, V. Kuznetsov, W. Schuhmann, “Differentiation
Between Single and Double-Stranded DNA by Local Capacitance Measurements”.
ChemElectroChem, 2016, 3, 855-857.
D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, “Potential-assisted DNA
Immobilization as a Prerequisite for Fast and Controlled Formation of DNA Monolayers”.
Angew. Chem., 2015, 127, 15278–15283; Angew. Chem. Int. Ed., 2015, 54, 15064–15068.
L. Švorc, D. Jambrec, M. Vojs, S. Barwe, J. Clausmeyer, P. Michniak, M. Marton, W.
Schuhmann, “Doping Level of Boron-Doped Diamond Electrodes Controls the Grafting
Density of Functional Groups for DNA Assays”. ACS Appl. Mater. Interfaces, 2015, 7, 18949-
18956.
S. Pilehvar*, D. Jambrec*, M. Gebala, W. Shuhmann, K. De Wael, “Intercalation of Proflavine
in ssDNA Aptamers: Effect on Binding of the Specific Target Chloramphenicol”. Electroanal.,
2015, 27, 1836-1841.
V. Eßmann, D. Jambrec, A. Kuhn, W. Schuhmann, “Linking Glucose Oxidation to Luminol-
Based Electrochemiluminescence Using Bipolar Electrochemistry”. Electrochem. Comm.,
2015, 50, 77-80.
* these authors contributed equally to this work
136

________________________________________________________________________ Appendix
In preparation:
D. Jambrec, A. Ruff, W. Schuhmann, “Amperometric Detection of dsDNA via Acrydinium
Orange Modified Glucose Oxidase”.
D. Jambrec, Y. Uğur Kayran, W. Schuhmann, “Fully potential-controlled preparation of DNA
chips”.
D. Jambrec, W. Schuhmann: “Potential-assisted DNA/thiol co-immobilization as a tool for
DNA chip production”.
D. Jambrec, B. Ciui, C. Cristea, W. Schuhmann: “Potential-assisted dehybridization as a DNA
detection tool employing a double-stranded ligation strategy”.
137

________________________________________________________________________ Appendix
7.3 Conference contributions
Oral presentations:
D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, “Diving into the Mechanism of
Potential-Assisted ssDNA Immobilization”, 66th Annual Meeting of the International Society
of Electrochemistry, 4 th -9 th October 2015, Taipei (Taiwan), best presentation prize.
D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, “Highly controlled and fast potentialassisted
ssDNA immobilization”, Summer meeting on bio-electrochemistry (SMOBE 2015),
17 th -20 th August 2015, Antwerp (Belgium), best presentation prize.
D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, “Highly Controlled and Fast Formation
of ssDNA-Modified Surfaces Using a Potential-Assisted Immobilization Method”, XXIII
International Symposium on Bioelectrochemistry and Bioenergetics, 14 th -18 th June 2015,
Malmö (Sweden).
D. Jambrec, M. Gebala, F. La Mantia, W. Schuhmann, “Potential assisted immobilization of
DNA probe on Au electrode surfaces”, Bioenergy Summer School, 28 th September-4 th October
2014, Ile d`Oleron (France).
Poster presentations:
D. Jambrec, F. Conzuelo, A. Ruff, A. Estrada-Vargas, W. Schuhmann, “Optimizing DNA
Assays. DNA Sensor Preparation in Minutes”, 67th Annual Meeting of the International
Society of Electrochemistry, 21 st -26 th August 2016, The Hague (Netherlands).
D. Jambrec, F. Conzuelo, A. Estrada-Vargas, W. Schuhmann, “Potential pulse-assisted
formation of thiol monolayers within minutes as a promising technique for fast and controlled
electrode surface modification”, 16 th International Conference on Electroanalysis, 12 th -16 th
June 2016, Bath (UK).
138

________________________________________________________________________ Appendix
D. Jambrec, M. Gebala, W. Schuhmann, “Fast Potential-Assisted Immobilization of
ssDNA on Au Electrode Surfaces”, 65th Annual Meeting of the International Society of
Electrochemistry, 31 st August-5 th September 2014, Lausanne (Switzerland), best poster award.
D. Jambrec, M. Gebala, W. Schuhmann, “Potential assisted immobilization of
DNA probe on Au electrode surfaces”, Surface Modification for Chemical and Biochemical
Sensing 2013, 8 th November-12 th November 2013, Łochów (Polland).
139