Phd thesis ClaudioStagni
Università degli Studi di Bologna
Ciclo XIX
Tesi di Dottorato di:
Relatori :
Chiar. mo Prof. Ing.
Chiar. mo Prof. Ing.
Chiar. mo Prof. Ing.
Settore Scientifico Disciplinare: ING/INF01 Elettronica
Anno Accademico 2006/2007
label free
capacitance mesurements
low cost
”L’universo risponde
il vero
se interrogato onestamente”
C.S. Lewis
”— Don Camillo, perché ce l’hai tanto con i numeri?
— Perché , secondo me, gli uomini
non funzionano piú proprio a causa dei numeri.
Essi hanno scoperto il numero
e ne hanno fatto il supremo regolatore dell’universo. [...]
— Gesú le idee sono dunque finite?
Gli uomini hanno pensato tutto il pensabile?
— Don Camillo cosa intendi tu per idea?
— Idea, per me, povere prete di campagna, é una lampada
che si accende nella notte profonda dell’ignoranza umana
e mette in luce un nuovo aspetto della grandezza del creatore.
Il Cristo sorrise...”
G. Guareschi
1.1 Life and the Genome . . . . . . . . . . . . . . . . . . .
1.2 Genetic Analysis . . . . . . . . . . . . . . . . . . . . .
1.2.1 Tracing the genome: achieving the genomic era
1.2.2 Application of DNA-sequences detection . . . .
1.3 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . .
2.1 Role of technology and electronics . . . . . . . . . . . . .
2.2 Technology, microfabrication and micromachining of silicon and other materials . . . . . . . . . . . . . . . . . .
2.2.1 Pattern definition of high-density spots of different
probes molecules on a substrate. DNA Microarray
technology . . . . . . . . . . . . . . . . . . . . . .
2.2.2 Microfluidics on chip . . . . . . . . . . . . . . . .
2.3 Electronics and microelectronics . . . . . . . . . . . . . .
2.3.1 Electrical-addressing of conductive sites . . . . . .
2.3.2 Electronic Circuits for signal detection and processing in biosensors application . . . . . . . . . .
2.3.3 Semiconductor sensors . . . . . . . . . . . . . . .
2.4 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.1 Physics and Chemistry of the sensing principle . . .
3.2 Sensing layer formation on electrodes . . . . . . . .
3.2.1 Organosilicon derivatives . . . . . . . . . . .
3.2.2 Thiol Layers . . . . . . . . . . . . . . . . . .
3.3 Previous work on Interface capacitance sensing . . .
3.4 Impedance measurement techinique . . . . . . . . .
3.4.1 Standard Methods. Impedance Spectroscopy
3.5 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . . .
4.1 Passive Microarrays vs Active Matrices . . . . . . . . . .
4.1.1 Label-free techniques . . . . . . . . . . . . . . . .
4.2 Bio-functionalization of micro-fabricated electrodes . . .
4.2.1 Microfabricated electrodes . . . . . . . . . . . . .
4.2.2 Basic process . . . . . . . . . . . . . . . . . . . .
4.2.3 Compatibility . . . . . . . . . . . . . . . . . . . .
4.3 Experimental Results . . . . . . . . . . . . . . . . . . . .
4.3.1 Complex impedance measurements . . . . . . . .
4.3.2 Integrable ciruitry . . . . . . . . . . . . . . . . . .
4.4 Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . .
4.5 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . . .
5 Smart sensor on PCB based on µ-controller
5.1 State-of-the-art and detection principle . . . .
5.2 Hardware and Software Design . . . . . . . . .
5.3 Experimental results . . . . . . . . . . . . . .
5.3.1 Tuning . . . . . . . . . . . . . . . . . .
5.3.2 Electrical characterization . . . . . . .
5.3.3 DNA Detection Measurement . . . . .
5.4 Figures . . . . . . . . . . . . . . . . . . . . . .
for genetic
6 On Chip DNA detection based on CBCM capacitance
6.1 Related Work . . . . . . . . . . . . . . . . . . . . . . . .
6.2 Capacitance-based DNA Detection Principle . . . . . . .
6.3 Chip . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
6.3.1 Chip architecture . . . . . . . . . . . . . . . . . .
6.3.2 Sensing site circuitry . . . . . . . . . . . . . . . .
6.3.3 Physical layout . . . . . . . . . . . . . . . . . . .
6.4 Measurement set-up . . . . . . . . . . . . . . . . . . . .
6.5 Experimental results . . . . . . . . . . . . . . . . . . . .
6.5.1 Electrical Characterization of gold electrodes . . .
6.5.2 Electrode bio-modification . . . . . . . . . . . . .
6.5.3 DNA detection . . . . . . . . . . . . . . . . . . .
6.6 Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . .
6.7 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . . .
7 Capacitance measurement for DNA detection with
chip A/D conversion
7.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . .
7.2 Related Work . . . . . . . . . . . . . . . . . . . . . . .
7.3 Capacitance-based DNA Detection Principle . . . . . .
7.4 Label-free DNA chip . . . . . . . . . . . . . . . . . . .
7.4.1 Chip architecture . . . . . . . . . . . . . . . . .
7.4.2 Sensing site circuitry . . . . . . . . . . . . . . .
7.4.3 Physical layout . . . . . . . . . . . . . . . . . .
7.5 Experimental results . . . . . . . . . . . . . . . . . . .
7.5.1 Measurement set-up . . . . . . . . . . . . . . .
7.5.2 Electrical Characterization . . . . . . . . . . . .
7.5.3 DNA detection . . . . . . . . . . . . . . . . . .
7.6 Figures . . . . . . . . . . . . . . . . . . . . . . . . . . .
8 Application on tumor marker
8.1 Tumor marker analysis . . .
8.2 Immunosensors . . . . . . .
8.3 Device and methods . . . .
8.4 Experimental results . . . .
8.5 Figures . . . . . . . . . . . .
9 EEPROM memory as DNA
9.1 Introduction . . . . . . . .
9.2 Devices and method . . .
9.3 Measurement setup . . . .
9.4 Experimental results . . .
9.5 Discussion . . . . . . . . .
9.6 Figures . . . . . . . . . . .
and future
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10 Conclusions and perspectives
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DNA 105
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11 Publications
11.1 Conferences . . . . . . . . . . . . . . . . . . . . . . . . . 125
11.2 Journals . . . . . . . . . . . . . . . . . . . . . . . . . . . 126
11.3 Patents . . . . . . . . . . . . . . . . . . . . . . . . . . . 126
Chapter 1
Life and the Genome
The feature of life is its ability to reproduce itself, but this ability alone is
not enough. Crystals influence the matter around them to create structures similar to themselves but they are not alive. Life can be defined by
recognizing its fundamental interrelatedness which means that all living
things are related to each other: they all have a common ancestor in
the distant past [1]. Organisms came to differ from each other though
evolution, which can be described as a cumulative process of the following three components: inheritance, variation and selection. Evolution not
only helps us to define what life is, but also to understand how living system function. All of an organism inherited characteristics are contained
in a single messenger molecule, the deoxyribonucleic acid, or DNA. The
characteristics are defined in a simple, linear, four element code. The genetic encoding for an organism is called genotype, the resulting physical
characteristics of an organism is called phenotype. Evolutionary variation (mutation, sexual recombination and genetic rearrangements) allows
modification of a genotype, that can have large consequences in phenotype. On the other side, selection acts only on phenotypes. Diversity
generated by evolution is enormous and very evident, but, although our
understanding of the molecular level of life is less detailed, this diversity
is encoded there. In fact, protein with very similar shapes and function
can have very different chemical composition. Organism that look quite
similar be very different from their genetic code. Despite this incredible
diversity, nearly all of the same basic mechanisms are present in all or-
ganism. All living things are made of cells. The thousands of substances
that make up the basic reactions inside the cell are remarkably similar
across all living things. The genetic material that codes for all of these
substances is written more or less in the same molecular language in every
organism. The genetic material is organized in the genome which is characterized by a defined and very variable size and it is contained identical
in the nucleus of every cell of an organism. The smallest known genome is
the one of the bacterium which contains about 600,000 DNA base pairs.
Human and mouse genomes have some 3 billion base pairs. Except for
mature red blood cells, all human cells contain a complete genome. DNA
in the human genome is arranged into 24 distinct chromosomes-physically
separate molecules that range from about 50 million to 250 million base
pairs. Each chromosome contains many genes, the basic physical and
functional units of heredity. Genes are specific sequences of bases that
encode instructions on how to make proteins. Genes comprise only about
2% of the human genome; the remainder consists of non-coding regions,
whose functions may include providing chromosomal structural integrity
and regulating where, when, and in what quantity proteins are made.
The human genome, at present, is estimated to contain 30,000 genes.
Proteins, expressed by genes by means of RNA sequence, first, and subsequently amino acids, perform most life functions and even make up the
majority of cellular structures. Proteins are large, complex molecules
made up of smaller subunits called amino acids. Chemical properties
that distinguish the 20 different amino acids cause the protein chains to
fold up into specific three-dimensional structures that define their particular functions in the cell. The constellation of all proteins in a cell is
called its proteome. Unlike the relatively unchanging genome, the dynamic proteome changes from minute to minute in response to tens of
thousands of intra- and extracellular environmental signals. A protein’s
chemistry and behaviour are specified by the gene sequence and by the
number and identities of other proteins made in the same cell at the
same time and with which it associates and reacts. The DNA sequence
is the particular side-by-side double-helix arrangement of bases along
the DNA strand (see Fig. 1.1). This order defines the exact instructions
required to create a particular organism with its unique traits. Bases
are parts of nucleic acid molecules which forms the DNA. They are of
four kinds: adenine, thymine, cytosine, guanine. All of the adenines on
one side of the DNA (one single strand) recognize the thymines on the
other side, in the sense that they bind together specifically by means of
two hydrogen bonds. In fact, the guanines recognize cytosines binding
themselves by means of three hydrogen bonds. This reaction is known as
base pairing and it is the basis of two molecular mechanisms: replication
and recognition (see Fig. 1.2). Replication: related to base pairing for
the first time by Watson and Crick: April, 25. 1953 Nature”.It has not
escaped our notice that specific base pairing we have postulated immediately suggests a possible copying mechanism for the genetic material.”
Recognition: A DNA sequence is defined and described by each of the
complementary single-strands which form the double-helix. Moreover,
if they are separated, they can recognize specifically each other. This
affinity reaction is known as hybridization. Together with the number of
complementary bases which composes two single strands, many parameters influence their binding: temperature, ionic force, the number of C-G
couples (which implies a stronger bond). In the literature there exist
at least two nomenclature systems for referring to molecular elements
involved in hybridization. Both use common terms ”probes” and ”targets”. According to [2, 3] probe is the tethered nucleic acid with known
sequence, while target is the free nucleic acid sample whose sequence or
quantity is to discover. The DNA is expressed during the cell life by
transcription of selected parts copied into the RNA molecule and then
by translation of the RNA into proteins. Recognition of DNA (or RNA)
sequence by hybridization is the key molecular reaction and analysis principle for genetic research and gene-bases tests. This dissertation concerns
electronic-microfabricated solutions for automated, low-cost, easy-to-use
tools to improve speed, efficiency, reliability and diffusion of DNA-based
research and tests.
Genetic Analysis
Genetics and proteomic may lead to the development of extremely powerful tools. Benefits of genetic research in several scientific matters are
enunciated as follows, according to the HGP: In Molecular Medicine
to: improve diagnosis of disease, detect genetic predispositions to disease, create drugs based on molecular information, use gene therapy and
control systems as drugs, design ”custom drugs” based on individual
genetic profiles. In Microbial Genomics to: rapidly detect and treat
pathogens (disease-causing microbes) in clinical practice, develop new
energy sources (biofuels), monitor environments to detect pollutants, protect citizenry from biological and chemical warfare, clean up toxic waste
safely and efficiently.
Tracing the genome: achieving the genomic
Human Genome Project
In June 2000, scientists announced the completion of the first working
draft of the entire human genome. Lately, in April 2003 - the 50th
anniversary of Watson and Crick’s publication of DNA structure - the
high-quality reference sequence was completed, marking the end of the
Human Genome Project. The two-step analysis procedures which are at
the basis of genome description are i)mapping and ii)sequencing:
Mapping means to make descriptive diagrams maps of each human
chromosome Mapping involves (1) dividing the chromosomes into smaller
fragments that can be propagated and characterized and (2) ordering
(mapping) them to correspond to their respective locations on the chromosomes. A genome map describes the order of genes or other markers
and the spacing between them on each chromosome. Human genome
maps are constructed on several different scales or levels of resolution.
At the coarsest resolution are genetic linkage maps, which depict the
relative chromosomal locations of DNA markers (genes and other identifiable DNA sequences) by their patterns of inheritance. Geneticists have
already charted the approximate positions of over 2300 genes, and a start
has been made in establishing high- resolution maps of the genome.
Sequencing: After mapping is completed, the next step is to determine the base sequence of each of the ordered DNA fragments. The
completed map will provide biologists with a Rosetta stone for studying human biology and enable medical researchers to begin to unravel
the mechanisms of inherited diseases. Much effort continues to be spent
locating genes; if the full sequence were known, emphasis could shift
to determining gene function. Technological advances are leading to
the automation of standard DNA purification, separation, and detection
steps. Sequencing procedures currently involve first Constructing Clones
(Fig. 1.3). The next step is amplification, which can be performed in
vivo , through a suitable host cell or in vitro by means of the Polymerase
Chain Reaction (Fig. 1.4). PCR amplify a desired DNA sequence of any
origin (virus, bacteria, plant, or human) hundreds of millions of times in
a matter of hours The reaction is highly specific, easily automated, and
capable of amplifying minute amounts of sample. PCR has also had a
major impact on clinical medicine, genetic disease diagnostics, forensic
science, and evolutionary biology. The next step is to make the subcloned fragments into sets of nested fragments differing in length by one
nucleotide, so that the specific base at the end of each successive fragment
is detectable after the fragments have been separated by gel electrophoresis (Fig. 1.5). This detection step can be performed by means of other
techniques which may involve sequence recognition by hybridization [4].
Single Nucleotide Polymorphisms and the Hap-Map Project
When the HGP began officially in 1990 a heated debate started concerning which Genome had to be sequenced. Fortunately, geneticists were not
forced to make this choice. During the accomplishment of the project,
scientists have described not only a single human genome sequence, composed of little bits from many humans, but also some 3 million sites of
variation mapped along that reference sequence (2001). The effort in localizing and determining these variations or ’polymorphisms’, is related
to the fact that genetics simply cannot exist without understanding their
function. Genomes most often differs in terms of Single nucleotide polymorphisms (SNPs, snips) [5]. The 3 million known SNPs are found at
a density of one SNP per 1.91 kilobases. This means that more than
90% of any sequence 20 kilobases long will contain one or more SNPs.
Therefore, 93% of genes contain a SNP which implies that, at present,
nearly every human gene and genomic region is marked by a sequence
variation. In 2002 the International HapMap project started. Its goal
is to compare the genetic sequences of different individuals to identify
chromosomal regions where genetic variants are shared. The Project will
help biomedical researchers find genes involved in diseases and responses
to therapeutic drugs. Being able to ”type” individual genomes and make
comparisons will be essential to understand
• how variation shapes biochemical and cellular functions
• in illuminating past human evolution;
• dissecting the contributions of individual genes to diseases that
have a complex, multigene basis;
• know how to implement patient care in relation to genetic variation
(tissue and organ incompatibility, affecting the success of transplants);
The research of disease-related SNPs is carried on comparing the haplotypes in individuals with a disease to the haplotypes of a comparable
group of individuals without a disease. If a particular haplotype occurs
more frequently in affected individuals compared with controls, a gene
influencing the disease may be located within or near that haplotype.
Common diseases such as cancer, stroke, heart disease, diabetes, depression, and asthma usually result from the combined effects of a number of
genetic variants and environmental factors. According to an idea known
as the common disease-common variant hypothesis, the risk of contracting common diseases is influenced by genetic variants that are relatively
common in populations. More and more widely distributed genetic variants associated with common diseases are being discovered, including
variants that contribute to autoimmune diseases, schizophrenia, diabetes,
asthma, stroke, and heart attacks. Variation in genome sequences underlie differences in our susceptibility to, or protection from, all kinds of
diseases; in the age of onset and severity of illness; and in the way our
bodies respond to treatment.
Application of DNA-sequences detection
Gene tests (DNA-based tests)
Gene testing exams an organism’s DNA, taken from cells in a sample
of blood or, occasionally, from other body fluids or tissues. The most
widespread applications are the search for DNA change that flags a disease or disorder, or for DNA sequences that could describe a reaction to
drugs. A few types of major chromosomal abnormalities, including missing or extra copies or gross breaks and rejoinings (translocations), can be
detected by microscopic examination. Most changes in DNA, however,
are more subtle and require a closer analysis of the DNA molecule to
find perhaps single-base differences. Genetic tests are used for several
reasons, including: (i) carrier screening, which involves identifying unaffected individuals who carry one copy of a gene for a disease that requires
two copies for the disease to be expressed; (ii) pre-implantation genetic
diagnosis; (iii) prenatal diagnostic testing; (iv) newborn screening; (v)
pre-symptomatic testing and confirmational diagnosis of a symptomatic
individual; (vi) forensic/identity testing. For some types of gene tests,
researchers design short pieces of DNA called probes, whose sequences
are complementary to the mutated sequences (see Fig. 1.6). These probes
will seek their complement among the three billion base pairs of an individual’s genome. If the mutated sequence is present in the patient’s
genome, the probe will bind to it and flag the mutation. Another type of
DNA testing involves comparing the sequence of DNA bases in a patient’s
gene to a normal version of the gene. Cost of testing can range from hundreds to thousands of dollars, depending on the sizes of the genes and
the numbers of mutations tested. Gene testing already has dramatically
improved lives. Some tests are used to clarify a diagnosis and direct a
physician toward appropriate treatments, while others allow families to
avoid having children with devastating diseases or identify people at high
risk for conditions that may be preventable. On the horizon is a gene
test that will provide doctors with a simple diagnostic test for a common
iron-storage disease, transforming it from a usually fatal condition to a
treatable one. Limitations and ELSI (Ethical, Legal and Social Issues)
The tests give only a probability for developing the disorder. One of the
most serious limitations of these susceptibility tests is the difficulty in
interpreting a positive result because some people who carry a diseaseassociated mutation never develop the disease. Scientists believe that
these mutations may work together with other, unknown mutations or
with environmental factors to cause disease. A limitation of all medical
testing is the possibility for laboratory errors. These might be due to
sample misidentification, contamination of the chemicals used for testing, or other factors. Few treatments or preventive strategies exist for
patients testing positive for most gene tests. Unfortunately, knowledge
of a gene mutation alone is insufficient information for researchers trying to devise intervention strategies. Researchers must first understand
the normal function of the disease-associated gene(s) and determine how
the mutation disrupts that function. Therefore, many in the medical
establishment feel that uncertainties surrounding test interpretation, the
current lack of available medical options for these diseases, the tests’ potential for provoking anxiety, and risks for discrimination (by employers,
insurers, commercial institutions, schools, army) and social stigmatization could outweigh the benefits of testing. Other important ethical issues
are also very controversial, such as privacy and confidentiality, fairness
in use of information, reproductive rights.
Expression analysis
When proteins are needed, the corresponding genes are transcribed into
RNA (transcription). RNA polymerase II, together with the necessary
transcription elongation factors, travels along the DNA template and
polymerizes ribonucleotides into an RNA copy of the gene. The polymerase moves at a regular speed (approximately 30 nucleotides per second) and holds on to the DNA template efficiently, even if the gene is
very long. At the end of the gene, the RNA polymerase falls off the
DNA template and transcription terminates. The RNA is first processed
so that non-coding parts are removed (processing) and is then trans-
ported out of the nucleus (transport). Outside the nucleus, the proteins
are built based upon the code in the RNA (translation). The information
contained in the nucleotide sequence of the mRNA is read as three letter words (triplets), called codons. Each word stands for one amino acid.
During translation amino acids are linked together to form a polypeptide
chain which will later be folded into a protein. The translation is dependent on many components, of which two are extra important. First of all;
the ribosome which is the cellular factory responsible for the protein synthesis. It consists of two different subunits, one small and one large and
is built up from rRNA and proteins. Inside the ribosome the amino acids
are linked together into a chain through multiple biochemical reactions.
The second component is the tRNA, a specialised RNA molecule that
carries an amino acid at one end and has a triplet of nucleotides, an anticodon, at the other end. The anticodon of a tRNA molecule can basepair,
i.e. form chemical bonds, with the mRNA’s three letter codon. Thus the
tRNA acts as the translator between mRNA and protein by bringing the
specific amino acid coded for by the mRNA codon (see Fig. 1.7).
The study of gene expression on a genomic scale is the most obvious
opportunity made possible by complete genome sequences of the model
organisms, and experimentally the most straightforward. Four characteristics of the regulation of gene expression at the level of transcript
(RNA) abundance account for the great value and appeal of genomewide surveys of transcript levels. First, it is eminently feasible - DNA
microarrays (see Chapter 2) make it easy to measure the transcripts for
every gene at once. The second reason is the tight connection between
the function of a gene product and its expression pattern. As a rule, each
gene is expressed in the specific cells and under the specific conditions in
which its product makes a contribution to fitness. Just as natural selection has precisely tuned the biochemical properties of the gene product,
so it has tuned the regulatory properties that govern when and where the
product is made and in what quantity. The logic of natural selection, as
well as experimental evidence, partially provides that there is a sensible
link between the expression pattern and the function of its gene product. Thirty years of molecular biology have provided numerous examples
of genes that function under specific conditions and whose expression is
tightly restricted to those conditions. Historically, transcript abundance
is assayed by immobilizing mRNA or total RNA (electrophoretically separated or in bulk) on membranes and then incubating with a radioactively
labelled, gene-specific target. If multiple RNA samples are immobilized
on the same matrix, one obtains information about the quantity of a particular message present in each RNA pool. In the last ten years cDNA
arrays have altered this strategy in several ways. In an array experiment,
many gene-specific polynucleotides derived from the 3 end of RNA transcripts are individually arrayed on a single matrix. This matrix is then
simultaneously probed with fluorescently tagged cDNA representations
of total RNA pools from test and reference cells, allowing one to determine the relative amount of transcript present in the pool by the type of
fluorescent signal generated. Relative message abundance is inherently
based on a direct comparison between a ’test’ cell state and a ’reference’
cell state; an internal control is thus provided for each measurement.
Figure 1.1: ”The specific base pairing immediately suggests a possible
copying mechanism for the genetic material” (Watson and Crick, 1953).
Figure 1.2: Base Pairing is the basis of two molecular mechanisms: replication and recognition.
Figure 1.3: Cloning for Sequencing. Cloned DNA molecules must be
made progressively smaller and the fragments subcloned into new vectors to obtain fragments small enough for use with current sequencing
technology. Sequencing results are compiled to provide longer stretches
of sequence across a chromosome.
Figure 1.4: Amplification. Polymerase Chain Reaction. PCR is a process based on a specialized polymerase enzyme, which can synthesize a
complementary strand to a given DNA strand in a mixture containing
the 4 DNA bases and 2 DNA fragments (primers, each about 20 bases
long) flanking the target sequence. The mixture is heated to separate
the strands of double- stranded DNA containing the target sequence and
then cooled to allow (1) the primers to find and bind to their complementary sequences on the separated strands and (2) the polymerase to
extend the primers into new complementary strands. Repeated heating
and cooling cycles multiply the target DNA exponentially, since each new
double strand separates to become two templates for further synthesis.
In about 1 hour, 20 PCR cycles can amplify the target by a millionfold.
Figure 1.5: Standard Sequencing (Sanger Method): Dideoxy sequencing
uses an enzymatic procedure to synthesize DNA chains of varying lengths,
stopping DNA replication at one of the four bases and then determining
the resulting fragment lengths. Each sequencing reaction tube (T, C,
G, and A) in the diagram (1) contains (i) a DNA template, a primer
sequence, and a DNA polymerase to initiate synthesis of a new strand of
DNA at the point where the primer is hybridized to the template; (ii) the
four deoxynucleotide triphosphates (dATP, dTTP, dCTP, and dGTP) to
extend the DNA strand; (iii) one labeled deoxynucleotide triphosphate
(using a radioactive element or dye); (iv) and one dideoxynucleotide
triphosphate, which terminates the growing chain wherever it is incorporated. Tube A has didATP, tube C has didCTP, etc. For example,
in the A reaction tube the ratio of the dATP to didATP is adjusted so
that each tube will have a collection of DNA fragments with a didATP
incorporated for each adenine position on the template DNA fragments.
The fragments of varying length are then separated by electrophoresis
(1) and the positions of the nucleotides analyzed to determine sequence.
The fragments are separated on the basis of size, with the shorter fragments moving faster and appearing at the bottom of the gel. Sequence
is read from bottom to top (2).
Figure 1.6: Gene test with microarray technology.
Figure 1.7: Gene Expression: from DNA to protein.
Chapter 2
Role of technology and electronics
In 1 we described the promising and in some cases already achieved accomplishment of genetic research based on the detection of a specific
nucleotide sequence in solution. This can be a DNA strand extracted
from a cell of interest or a DNA copied from an RNA which have been
expressed by the cell in particular conditions (cDNA). The desire of the
author is to give a feeling of the crucial role of electronics and technology in developing efficient, innovative and mass produced devices for
nucleic-acids sequence detection. The scientific and technological areas
which are involved in designing and implementing genetic assay may be
listed as follows: (i) element 1: Technology, microfabrication and micromachining of silicon and other materials; (ii) element 2 : electronics and
micro(nano)electronics; (iii) element 3 : semiconductor sensors.
These elements would lead to
• miniaturization of reaction sites and cell (less sample and reagent)
(Element 1)
• miniaturization of measurement system (less noise and portability)
(Elements 1-2-3)
• high-parallelism (order of magnitude improvement in speed of analysis until the extreme achievements of a genome-wide screening
(elements 1-2-3)
• integration of mechanical and fluidic functions for sample handling,
delivery, mixing, purification, separation, amplification. (multifunctionality and stand-alone devices, easy-to-use devices) (element
1)- miniaturization. The small dimensions also reduce the amounts
of carriers (reagents) necessary to conduct a chemical process: for
miniaturized handling volumes are often in the nanoliter to picoliter range rather than the microliter range or larger in conventional
• batch production (low-cost mass-produced assays) (Elements 1-2-3)
• generation of electrical read-out from the sensor (a multitude of
microelectronic circuits are available for electric signal conditioning
and modification as amplification, filtering, modulation) (Elements
2-3) the electrical signal is very well suited for signal transmission
(Elements 2-3)
• high performance sensors through the employment of ultra-sensitive
electron devices (Element 3)
All of the listed characteristics, or a selection of them, may describe the so called Lab-on-a-chip technology, which aims at implementing miniaturized, stand-alone, low-sample and fast analysis tools. At
present, only part of these issues have been considered and traduced
in sensors or systems. Some general references on the issues related
are [6, 7, 8].
Technology, microfabrication and micromachining of silicon and other materials
Pattern definition of high-density spots of different probes molecules on a substrate. DNA
Microarray technology
The possibility of to attach and localize (and or address) receptors onto
a substrate in a very precise and dense way has been a fundamental innovation achieved in early 90s. This technology lead to high parallelism of
analysis with the possibility to test an eventually huge number of probes
sequence at the same time and on the same substrate. In addition, probes
sites (or features) are very little and dense. At present, 1.2 cm2 substrate
may contain more than 500000 different probe sequences [Lipshutz, 1999].
Several techniques, with very different density results, are employed. The
list that follows is in growing order of spot density/cm2
• Mechanical Micro-spotting 3000 spot/cm2 probe cDNA (500-5,000
bases long) are immobilized to a solid surface such as glass using
high-speed robot spotting and exposed to a set of targets either
separately or in a mixture. This method, traditionally called DNA
microarray, is widely considered as developed at Stanford University [9].
• Electric-Filed assisted immobilization (Nanogen, San Diego, California) (5000 sites/cm2 ) (see chapter 1) [10]
• Electro-immobilization by copolymerization (10000 sites/cm2 ) (Leti,
Grenoble, France) [11]
• Inkjet technology (IBM) (10000 sites/cm2 ) [Bietsch, 2004]In this
approach, a DNA sample is loaded into a miniature nozzle equipped
with a piezoelectric fitting (or other form of propulsion) which is
used to expel a precise amount of liquid from the jet onto the substrate. After the first jetting step, the jet is washed and a second
sample is loaded and deposited to an adjacent address. A repeat
series of cycles with multiple jets enables rapid microarray production.Bubble Jet Technology (Canon) (20000/cm2) [Okamoto,
2000]Photolithography (390000 features/cm2) (Affymetrix, Santa
Clara, California) [Lipshutz, 1999]. The probe is synthesized either
in situ : an array of oligonucleotide (20 80-mer oligonucleotides) or
peptide nucleic acid (PNA) (labelled sample DNA, hybridized, and
on-chip) or by conventional synthesis followed by on-chip immobilization. The array is exposed to the identity/abundance of complementary sequences are determined. The method, ”historically”
called DNA chips, was developed by Affymetrix, Inc, which sells
its photolithographically manufactured wafers under the GeneChip
trademark. Probe synthesis occurs in parallel, resulting in the addition of an A, C, T, or G nucleotide to multiple growing chains
simultaneously. To define which oligonucleotide chains will receive
a nucleotide in each step, photolithographic masks, carrying 18 to
20 square micron windows that correspond to the dimensions of
individual features, are placed over the coated wafer. The windows
are distributed over the mask based on the desired sequence of each
probe (see Fig. 2.1).
The techniques listed above are employed in the so-called microarrays devices. Microarrays are defined here as monolithic, flat surfaces
that bear multiple probe sites and each bear a reagent whose molecular recognition of a complementary molecule can lead to a signal that
is detected by an imaging technology, most often fluorescence. Literature references to microarrays before 1995 concerned arrays of electrodes
rather than arrays of different molecules. The first molecular microarray,
reported in 1991, was composed of peptides, not DNA, and was not even
identified as a microarray [12].
Microfluidics on chip
A number of basic fluidic components have been assembled in different ways to perform various chemical measurements. Many of these are
based on electrokinetic transport principles, and include valves, mixing
structures, chemical reactors, and chemical separation channels. In addition, chemical separation mechanisms have been miniaturized, including
free-solution and gel electrophoresis, solvent programmed chromatography, isoelectric focusing, isotachophoresis, and two-dimensional separations based on liquid chromatography and free-solution electrophoresis.
Surface interactions have been exploited for solid-phase extraction to
process samples and for hybridization of target DNAs, and nanoliterscale reactors have been demonstrated for continuous flow, stopped flow,
and thermal cycling reactions [13] Chips integration microfluidics may be
made from plastic, glass, quartz, or silicon. Bulk and surface micromachining performed with sophisticated etching, patterning and deposition
techniques are at the basis of channel implementation. One of the most
interesting microfabricated implementation in chip is the PCR molecular
amplification. This approach has been widely investigated exploiting the
good properties of thermal conductivity of silicon and the possibility to
easily integrate thermal resistances [14, 15] (see Figure 2.2).
• Cantilever-array production for mass and stress molecular detection techniques. Recent works have reported the observation that
when specific biomolecular interactions occur on one surface of a
microcantilever beam, the cantilever bends. The recent discovery of the origin of nanomechanical motion generated by DNA hybridization and protein-ligand binding provided some insight into
the specificity of the technique. In addition, its use for DNA-DNA
hybridization detection, including accurate positive/negative detection of one-base pair mismatches, was also reported [16]. This tech-
nology readily lends itself to formation of microarrays using wellknown microfabrication techniques, thereby offering the promising
prospect of high-throughput protein analysis (see Fig. 2.3).
• Porous material for sensing based on high surface/volume ratio.
High surface/area materials technology is directed primarily towards the creation of inexpensive low bulk volume/area media for
applications that involves chemical reaction on surfaces. High surface areas provide a mechanism to achieve detection sensitivities
that are in the range of parts per billion on a short time scale [17].
Electronics and microelectronics
Electrical-addressing of conductive sites
Some examples of use of electrical addressing of probe-sites are here described:
• Functional probes immobilization: CEA-LETI and Cisbio International developed a CMOS platform able to address and polarize
128 test sites to perform electrochemical immobilization of probes
by means of a pyrrole bearing an oligonucleotide. A local electrocopolymerization allows oligonucleotide probes to attach permanently on the electrode. Biocompatibility issue requires specific
insulation layer and surface metal materials to form the electrodes.
The design of the cells has been modified to be compatible with
the pyrrole electro copolymerization step. It means that the voltage applied during this step has to be fully withstood by the CMOS
• Enhanced immobilization and hybridization by electrical polarization: Nanogen’s technology utilizes the natural positive or negative charge of most biological molecules. Applying an electric current to individual test sites on the NanoChip microarray enables
rapid movement and concentration of the molecules. Through electronics, molecular binding onto the NanoChip-microarray is accelerated up to 1000 times faster than traditional passive methods.
Nanogen’s technology involves electronically addressing biotinylated probe samples for immobilization, hybridizing complementary DNA and applying stringency to remove unbound and nonspecifically bound DNA after hybridization [18, 10] (see Fig. 2.4)
Electronic Circuits for signal detection and
processing in biosensors application
Infineon developed a fully electronic 16×8 sensor array. The chip is based
on a standard 0.5 µm CMOS process extended with additional process
steps to form sensor electrodes made of gold. A single sensing site within
these array consists of interdigitated gold electrodes (width = spacing
= 1µm) arranged within a circular compartment (diameter = 250 µm).
Using a microspotter, single-stranded DNA probe molecules are spotted
and immobilized at the Au surface. After immobilization, an analyte containing target molecules to be detected is applied to the whole chip and
hybridization occurs in case of matching DNA strands. For read-out,
a redoxcycling based electrochemical sensor principle is used: After a
washing step, a suitable chemical substance (p-aminophenyl-phosphate)
is applied and electrochemically redox-active compounds are created by
an enzyme label (alkaline phosphatase) bound to the target DNA strands.
Applying simultaneously an oxidation and a reduction potential to the
interdigitated gold electrodes, a redox current between these electrodes
occurs whose magnitude depends on the amount of double-stranded DNA
at this sensor position. The circuits within each position allow operation
over five decades, sensor currents between 10-12 A and 10-7 A are detectable [19] (see Fig. 2.5).
Motorola (Clinical Microsensors Division) developed a novel electronic detection format for nucleic acids that utilizes electronic instrumentation including a disposable DNA chip (see Fig. 2.6). Printed circuit board technology is used to manufacture chips with 14 exposed gold
electrodes, each of which is wired to a connector at the chip edge. A
solder mask defines the electrode diameter (250 or 500 µm) and covers the lead to the connector. DNA capture probes are deposited as a
mixed solution with the other components of the self-assembled monolayer (SAM). After deposition, chips are rinsed, dried, and sealed in a
housing for hybridization. Signalling probes containing ferrocene moieties,5 redox active metal centres that facilitate the detection of nucleic
acid targets in homogeneous assays that eliminate the need for separate
labelling and washing steps, form the basis of this electronic detection
platform. Hybridization is measured by alternating current voltammetry
(ACV) [20].
Semiconductor sensors
The most relevant applications of semiconductor sensors in biomolecular
detection can be summarized as follows:
Electrolyte/Insulator/Semiconductor structures
These include (i) one-dimensional vertical capacitor similar to MOS structures where metal has been substituted with an electrolyte conductive solution; (ii) Filed effect transistors where the gate has been substituted as
well. Such devices have been widely employed in the last twenty years in
electrochemistry and analytical chemistry for ion and pH sensing. More
recently great efforts have been done to detect DNA hybridization occurring at the interface between the electrolyte and the insulator. The
detection is based on field-effect of DNA negative charge on the gate
oxide [21, 22, 16, 23].
Other sensing devices
Several semiconductor sensing techniques have been tested for molecular
detection. Some of them are here listed:
• Surface and Bulk Acoustic Wave Sensors [24]
• Fiber Optics [25]
• Nanowires [26]
Figure 2.1: Photolithographic technique for in situ probe synthesis
(Affymetrix, Santa Clara, California).
Figure 2.2: (A) Mask design for the portable PCR-CE microchip. The
glass microchannels are indicated in black, the RTD and microfabricated
electrodes are in green, and the heater (located on the backside of the
device) is shown in red. The PCR chamber is loaded through reservoirs
a and b. Reservoir c is the co-inject reservoir, d is the cathode, e is the
waste, and f is the anode. (B) Schematic side view of a PDMS microvalve.
(C) Exploded view of the assembly of the PCR-CE microchip, showing
PDMS microvalve construction and PDMS gaskets [15].
Figure 2.3: Bending of a cantilever due to the change in surface stress
after target specific binding.
Figure 2.4: Motorola e-sensor.
Figure 2.5: Integrated gold electrodes on silicon chips fro DNA detection.
Infineon Technologies.
Figure 2.6: Nanogen Nanochip.
Chapter 3
Physics and Chemistry of the sensing
A capacitor is formed when two conducting plates are separated by a nonconducting medium. The capacitance depends on the size of the plates,
the distance between the plates and the properties of the dielectrical
layer, according to following equation:
S0 ǫǫ0
With S0 = surface of the electrode, d = thickness of the dielectrical
layer, ǫ0 = dielectric constant and ǫ = relative dielectric constant. The
dielectric constant is a physical parameter, while the relative dielectric
constant depends on the material. It should be noticed that there is
a large difference between the dielectric constant of water (about 78)
and that of an organic coating (4-8). The formation of ultrathin organic films, e.g. alkanethiols, on bare metal electrodes leads to changes
of the electrode capacitance. To preserve electrical neutrality, an opposite charge in the bulk solution has to compensate for the electrode
charge and an electric double-layer is formed. The capacitance of the
double layer depends on many variables including electrode potential,
temperature, ionic concentrations, types of ions, oxide layers, electrode
roughness, impurity adsorption. Further, this double layer can ideally be
considered as a plate-capacitor. For description of this effect the GouyChapman or Stern model is often used [27]. When the self-assembly of
the thiol molecules on the metal surface is complete, the double layer is
shifted because of the new dielectrical layer. A further capacitor cthiol
is added to the electrical circuit. The total capacitance can be described
by two capacitors cthiol and cdl in a serial arrangement, where cthiol
represents the capacitance of the alkanethiol and cdl the capacitance of
the ionic double layer (Gouy-Chapman layer) of the bulk phase. While
the capacitance of the ionic double layer depends on the electrolyte concentration the capacitance of the thiol layer is nearly independent on
the electrolyte concentration. By measuring the electrode capacitance
of a bare metal electrode, the capacitance of the electric double-layer is
obtained which according to the Gouy-Chapman theory is proportional
to c1/2 . A simplified circuit diagram is shown in Fig. 3.1. Values for
the electrode capacitance of bare gold electrodes are in the range of 1
to 100 µF/cm2 . This is related to the influence by Faradayic currents,
which simulate a pseudo-capacitance [28] and in some cases can be much
higher than the electrode capacitance. Self-assembled alkanethiols monolayers form an insulating coating. That additional capacitor requires a
new circuit diagram. The major part of the potential drop will be in
this case inside of the dielectric layer. Therefore, the capacitance of the
Gouy-Chapman layer can be neglected. Also in the experiment the influence of the electrolyte concentration on the capacitance of alkanethiol
layers was found to be very small. Because of the high ohmic resistance
of the alkanethiol layer (about 500kΩ [29], the electric current after application of an AC voltage is almost purely capacitive, therefore the circuit
diagram can be simplified. The advantage of the alkanethiol covered
electrodes for capacitive biosensors is that electron transfer is strongly
limited or even totally blocked, which allows an undisturbed capacitance
measurement. Also, ion adsorption, electrolyte activity of redox active
substances is minimized. After formation of the dielectric self-assembled
alkanethiol monolayer, further molecules can be immobilized or adsorbed
on this basis. Adsorption of molecules leads to an increase of the dielectrical thickness of the organic layer, and therefore decreases the electrical
capacitance. If the specific capacitance of the uncovered receptor (no
analyte is present) layer is C0 , follows for the capacitance of this layer
covered partly with analyte (specific capacitance Cads )
S0 c0 + Cads
With Sads = surface covered with analyte.
Sensing layer formation on electrodes
The modification of a transducer surface with an organic interface is
one of the critical steps in biosensor preparation. It can be achieved by
polymer-coating, by electrochemical polymerisation (e. g. polypyrrole),
or of self-assembled monolayers. SAMs are molecular assemblies that
are formed spontaneously by the immersion of an appropriate substrate
into a solution. A series of SAMs are known in literature, including
organosilicone on hydroxylated surfaces as SiO2 on Si, Al2 O3 on Al or
glass [30], sulphur containing compounds (R-SH, RS-SR, R-S-R) on metal
surfaces, amines on platinum [31].
Organosilicon derivatives
The reaction of alkanesilanes of the general structure RSiX3, R2SiX2
or R3SiX, where R is a carbon chain that can be functionalised with,
e. g., amino, carboxy or pyridyl groups, and X is chloride or hydroxy
with hydroxylated surfaces as SiO2 , SnO2 or T iO2 , is well established.
A schematic view of this structure is shown in Fig. 3.2. The formation
of the self-assembled monolayer is an in situ formation of polysiloxane,
which is connected to surface silanol groups via Si-O-Si bonds. The
typical preparation is very easy: immersion of a hydroxylated surface
into an organic solution of the organosilicon derivates for a few hours.
Even multilayers, i.e. three-dimensional polymer networks, can be built
by introduction of further hydroxy groups (e. g. tetrahydroxysilane)
into the structure. In consideration of these qualities, a large variety of
sensors and biosensors are nowadays described in the literature [32].
Thiol Layers
The first paper on the formation of self-assembled monolayers of dialkanesulfides (RS-SR) on gold surfaces was published in 1983 by [33]. In
the following years, numerous publications on this topic displayed its interesting properties for a variety of applications. In this chapter a short
overview is presented. Chemisorption of sulfur-containing compounds,
especially thiols (R-SH), on gold [34], silver [35], have been investigated.
Also, alkanethiols on copper and silver are described to form well ordered
and closely packed systems, the main scientific interest was focused on
gold surfaces. The fact that gold does not have a stable oxide, and thus
can be handled in ambient conditions, is a major reason. Further advantages of this noble metal are high oxidation stability, easy cleaning
procedures, well-defined surfaces, i.e. Au (111) or Au (100), and nearly
defect-free SAMs. Bain et al. in 1989 reported on the effect of the
formation of ultrathin layers on gold by comparison of different organic
compounds, for example disulfides, sulfides, isonitriles or trialkanephosphines. They found the thiol group to undergo the strongest interaction
with the gold surface, which are even 150 kJ/mol or higher [36]. This
is in the range of covalent bonds as known in organic chemistry (e. g.
S −S 226 kJ/mol). The structure of long-chain chemisorbed alkanethiols
with more than nine methylene groups is significally different from the
structure of short-chain alkanethiols. Long-chain compound are postulated to exist at room temperature in crystal-like periodicity, while for
short-chain compounds a liquid state is claimed. The symmetry of sulfur atoms in a monolayer of alkanethiols on mono-crystallised (111) gold
is hexagonal, with a S-S spacing of about 5 Å, and a calculated area
of 21.4 Å2 Fig. 3.3) [37]. The chains are in all-trans conformation and
have an angle of inclination in the range of 25 to 30◦ Fig. 3.3 [38]. Also,
the preparation of thiol monolayers on a metal surface is simple. Clean
surfaces (wafers or wires) are dipped into dilute thiol solutions (typically
10−3 mol/l). The choice of the organic solvent, ethanol or chloroform are
mostly used, and the incubation time depend on the solubility and the
chain length of the thiol. Alkanethiols show no further changes after 6 to 8
hours. Some working groups connect a formation of good layers already
after one hour. In case of disulfides the adsorption time is about two
times longer. This close-packed layer have excellent insulator properties,
which can be proved, i.e. by cyclovoltametry [34], ellipsometry, contact
angle or impedance measurements. SAMs are nearly defect-free and produce an ideally polarizable ultrathin basis. Alkanethiols with carboxy-,
amino- or other head groups can be used as coupling positions for the
immobilization of receptor molecules by a variety of methods. Therefore,
self-assembled monolayers of alkanethiols seem to be an optimal system
for the preparation of capacitive (affinity) sensors.
Previous work on Interface capacitance
Capacitive interface detection of DNA hybridization is a label-free, fully
electrical technique, thus it represents in principle the simplest, most
direct, hence also best solution. In this field, bio-functionalized gold
electrodes have been used in connection with electrochemical impedance
spectroscopy performed with a standard three-electrode system. A significant example of this approach [29] exploits single crystal gold surface modified with a self-assembled mono-layer of alkanethiols, used as
linker molecules for DNA probes to form a compact, stable and dense
layer, that featured an area of 8.56 mm2 and a capacitance of about
2 µF/cm2 . However, three-terminal techniques are too complex for costeffective fully integrated solutions. Therefore, two- electrodes alternative,
such as that of interest for this work, have also attracted significant interest. For example a two-electrode set-up using 2.4 mm2 gold- and an
Ag/AgCl counter- electrode has been studied. In this case, gold was
covered with a 24-base oligonucleotides layer, realized by means of alkanethiols, and the interface exhibited ideal capacitive behaviour in the
10 − 100 Hz frequency range [39]. As for the effects of DNA hybridization, a capacitance decrease before and after insertion of 179-base long
target molecule has been observed [Berggren, 1999] in a flow cell containing a gold surface of 7.065 mm2 covered with 26-base complementary
oligonucleotides. Furthermore, interface capacitance has been measured
in a three-electrode set-up by applying a 50 mV potentiostatic step and
measuring the current in the circuit. The interface capacitance, as well
as the series resistance of the system, have been extracted observing the
transient response. Comparisons between surface capacitance values reported in the literature can be misleading as they are strongly affected
by the electrolyte parameters, by the alkanethiol chain length and most
of all by the effective surface which depends on the electrode surface
roughness. Nevertheless, capacitance of bio-modified gold-electrodes is
expected to vary between 1 and 5 µF/cm2 . As far as integrable capacitance methods are concerned, recently an advance in the state of the
art has been made [40] demonstrating that DNA hybridization can be
detected by measuring interface capacitance with a two-gold-electrode
system (without the use of a reference electrode). In any case, all the
works mentioned above have been based on measurements performed
on gold electrodes larger than 2 mm2 . In contrast, several attempts to
perform DNA detection on micro-fabricated devices has been based on
Electrolyte-Insulator-Semiconductor structures [21, 22, 16, 41, 23]. In
these cases a reference electrode acting as control gate has been placed
in the solution to allow electrical polarization. Probe molecules have been
attached covalently or by absorption on the insulator surface (or on a thin
metallic layer deposed on it) to form a sensing molecular layer affecting
the interface potential. The binding event can be detected by impedance
measurements of the this structure with a two- or three-electrode setup.
Ideal capacitive behaviour has been observed [21] between 102 and 104 Hz.
Alternatively, the field-effect has been used on transistor-like devices by
polarizing the semiconductor and by measuring the current change flowing between drain and source. The EIS approach is very attractive,
because the transduction device integrates the sensing element and can
be easily reduced in dimensions. However, it presents problems related
to the counterions screening while detecting intrinsic molecular charge at
the insulator/solution interface.
Impedance measurement techinique
Standard Methods. Impedance Spectroscopy
Part of the experiments shown in this work were performed by measuring the changes of electrode capacitance or even the whole impedance
spectra. It offered the possibility to monitor the immobilization of DNA
probes as well as the DNA target hybridization on the electrodes. Electrochemical impedance is a general parameter for the measurement of
circuit elements with a complex behaviour, which do not follow the basic concept of the ideal resistor (R = V /I). It is usually measured by
applying an AC potential to an electrochemical cell and measuring the
current through the cell. When a sinusoidal potential excitation is applied, the response to this potential is an AC current signal, containing
the excitation frequency and it is harmonics. This current signal can be
analyzed as a sum of sinusoidal functions (Fourier series). Usually, electrochemical impedance is measured by using a small excitation signal.
Therefore lock-in amplifiers or frequency response analyzers can be used.
In a linear or pseudo-linear system, the current response to a sinusoidal
potential will be a sinusoid at the same frequency but shifted in phase
(see Fig. 3.4). The excitation signal, expressed as a function of time, has
the form:
E(t) = E0 cos(ωt)
where E(t) is the potential at time t, E0 is the amplitude of the
signal and ω is the radial frequency (ω = 2πf ). The response signal I(t)
is shifted in phase and has a different amplitude I0 .
I(t) = I0 cos(ωt + φ)
According to previous equations a generalized Ohms law for the impedance
of the system was formulated.
E0 cos(ωt)
= Z0
I0 cos(ωt + φ)
cos(ωt + φ)
where the impedance Z is expressed in terms of a magnitude Z0 and
a phase shift φ. Using Eulers relationship:
exp(iφ) = cos(φ) + isin(φ)
the impedance can be expressed as a complex function. Therefore,
the potential (E), current response (I) and impedance (Z) are described
E(t) = E0 exp(iφt)
I(t) = I0 exp(iφt)
= Z0 exp(iφt) = Z0 (cos(φ) + isin(φ))
Three major plotting graphs are used in impedance analysis: the
Nyquist plot (real part of impedance vs. imaginary part), and the Bode
plot (impedance with log frequency vs. both absolute value of impedance
and phase-shift). Electrical circuit theory differentiates between linear
and non-linear systems. Electrochemical cells in general have a non-linear
behaviour. For such systems a doubling the voltage (input) will not be
followed by double the current (output). If the system is non-linear, the
current response will contain harmonics of the excitation frequency. However, electrochemical systems can be considered in first approximation as
pseudo-linear, if only a small AC signal is applied to the electrochemical cell; a pseudo-linear segment of the cells current vs. voltage curve.
Electrochemical impedance spectroscopy can be used for capacitance and
conductivity measurements. While changes in solution resistance of the
electrochemical cell causes changes in conductance, capacitance changes
only occur because of modification of the electrodes surface. An exhaustive compendium of Impedance Spectroscopy can be found in [42].
Figure 3.1: Adsorption and desorption of alkanethiols on an electrode.
The self-assembly results in a dielectrical layer with insulating properties,
which leads to a decrease of capacitance. cthiol represents the capacitance of the alkanethiol and cdl the capacitance of the ionic double layer
of the bulk phase. If the thiols desorb from the electrode opposite direction molecules desorb from the structure, the capacitance increases. This
could be achieved for example by applying a very negative potential, see
electrically addressable immobilization.
Figure 3.2: Self-assembled monolayer (1) and multilayer (2) of alkanetrihydroxysilane on a hydroxylated surface. The molecules in the layers are
connected both to each other and to the surface by chemical bonds.
Figure 3.3: Proposed coverage scheme and orientation of alkanethiols on
a gold (111) surface.
Figure 3.4: (left) Sinusoidal current response in a linear system. (top)
Vector diagram of the phase-sensitive current analysis. The AC current
can be presented as a sum of two vectors: conductive current (a) and capacitance current (b). A phase angle θ can be calculated as arctan (b/a).
If the conductive component is zero, θ = 90◦ , it corresponds to ideal
capacitance behaviour.
Chapter 4
This Chapter presents the experimental characterization of two-terminal
micro-fabricated capacitors for microarrays with electrical sensing of labelfree DNA. So far [40], such a concept has been demonstrated only in experimental set-ups featuring dimensions much larger than those typical of
micro-fabrication. Therefore, this work investigates: a) the compatibility
of the silicon microelectronic processes with biological functionalization
procedures; b) the effects of parasitics when electrodes have realistic dimensions; c) measurement stability and reproducibility; d) the possibility of a fully integrated, stand-alone device. The obtained results clearly
indicate that two-terminal capacitive sensing with fully integrated electronics represents a realistic prospective for DNA label-free detection/
Passive Microarrays vs Active Matrices
Passive DNA chips based on optical detection of the hybridization of
labelled DNA [43] and capable of testing hundreds of thousands dif-
ferent probes in parallel are technologically mature devices already on
the market. The targets are labelled with chromophores molecules and
spread on the array surface. Successively, all this material is removed
and only the (labelled) targets that have hybridized with complementary probes remain immobilized on their specific sensing sites Finally,
the device is observed optically to localize the chromophores, revealing
where hybridization has taken place, i.e. the ”optical” signature of the
analyzed DNA. Often, the resulting optical image is not immediately
recognizable, because hybridization can involve many adjacent sites resulting in a complex optical pattern to be resolved by means of adequate
image processing. This technology presents two main drawbacks: cost
of the instrument to detect and resolve the optical signal; necessity of a
labelling step, with the need of additional reagents and the possibility of
sample pollution. To eliminate the former problem, methods based on
the generation of electrical signals upon hybridization have been developed. However, some of these techniques are not label-free, as they use
mediator elements to generate the electrical signals [44, 45].
Label-free techniques
Our interest is for label-free devices with fully integrated electrical reading, a subject largely still at the research and development level. Within
this category, another distinction can be made between devices exploiting active elements as sensor and transducer (producing electrical signals)
and those featuring passive sensing elements, in practice capacitors made
sensitive to DNA hybridization. The former case is scientifically very interesting, but has significant disadvantages in terms of noise, sensitivity
and repeatability. Therefore, we are currently pursuing a solution based
on electrical measurements of DNA sensitive capacitances, already shown
to be viable at the level of discrete lab prototypes. However, current
research on label-free, fully electrical devices has not yet satisfactorily
investigated the main problems posed by micro-fabrication, namely: a)
compatibility of bio-functionalization procedures with silicon microelectronic processes; b) effects of the large parasitics, inevitably present in
the structures; c) stability and repeatability of the measurements; d)
possibility to integrate reading circuits together with the sensors (so as
to achieve fully integrated, stand-alone devices).
Bio-functionalization of micro-fabricated
This Section investigates the compatibility of bio-functionalization procedures with silicon micro-electronic processes. To this aim, first the
basic principles of the functionalization techniques are briefly illustrated,
then the key problems are analyzed.
Microfabricated electrodes
The experiments here presented have been performed on micro-fabricated
test-chips featuring 1800 µm2 gold electrodes deposited by sputtering and
contacted with buried aluminum wires. Two different chip types have
been tested. The first-one (STM CH268A) hosted an array of twenty
identical gold electrodes of 1800 µm2 and a unique counter electrode
surrounding the others. The second one (STM CH402A) instead presented forty-eight electrodes and the counter. The surface structure of
electrodes have been observed with Atomic Force Microscopy (Fig. 4.1)
Basic process
The basic processes involved in gold bio- functionalization for DNA detection are illustrated in Fig. 4.2. First, a careful pre-cleaning must be
performed (step 1) to allow efficient attachment of tethered probes by
means of Au-S bonds. Gold is extremely reactive and has to be freed
from organic and inorganic molecules. Cleaning can be attained by oxygen plasma [Berggren, 1999], chemical etching with hot piranha solution
bathes (1 : 3 H2 O2 /H2 SO4 ) [39, 46], ultrasonic bathes [Peterlinz, 1996],
electrochemical stripping techniques [29]. Mechanical polishing with alumina [47] or silicon carbide powders [29] has sometimes been employed to
improve surface cleaning and to provide better electrode flatness. In fact,
Losicet al. have found that surface roughness can cause an increase of defects in the molecular layer. Immediately after cleaning, the surface must
be bio-functionalized with the sensing layer of probe-molecules (step 2).
Several biochemical procedures have been proposed and tested to bind
covalently short oligonucleotide chains on gold surfaces and create dense
and homogeneous bio-molecular layers. The most efficient techniques
employ an anchoring alkanethiol chain attached to an extreme of the sequence of nucleic acid. The thiol group provides a covalent bond with
gold atoms and the small alkanic chain acts (depending on its length)
to form a compact layer. Molecules can be spotted by micro-spotting,
inkjet printing [Bietsch, 2004] or deposited on the surface by immersion (dipping) in a micro-molar concentration solution [46]. The solvent
where molecules are dissolved should contain salts to avoid electrostatic
repulsion between charged molecules during layer formation. Then, the
surface must be rinsed with ultra pure water or buffers to remove noncovalently attached probe molecules and to obtain a layer of ordered and
well-oriented receptors (step 3). Then, gold can be dried and briefly
stored. DNA bio-affinity recognition reaction (step 4) is performed by
immersing the electrodes modified with different probes sequence in a
buffer solution with suitable physical and chemical parameters (temperature, ionic strength, pH). The solution conditions are extremely important to ensure efficient and specific sensing. The formation of the double
helix is a reversible reaction, thus the target can be eventually desorbed
from the surface by a heat treatment which consists of rinsing the electrode with hot solution of low ionic strength to observe an increase in
capacitance (opposite signal) (step 5). At this stage, hybridization specific binding can be repeated. Beside bio-chemical procedures, electrical
characterization also involves immersion of gold electrodes in a conductive (ionic) solution and polarization with low-voltage levels. Electrical
measurements are usually performed before and after complementary sequences recognition.
Some of the procedures illustrated in Fig. 4.2 may permanently damage
the silicon chip hosting the electrodes as well as the devices and the interconnects. For example, concerning step 1, polishing and strong acid
surface cleaning should be avoided. In fact, electric tests of wires conductivity performed after 30 sec of piranha surface cleaning indicated
extensive damages (see Fig. 4.3). Thus low-impact bio-functionalization
procedures must be developed to limit damages. To this purpose, the
cleaning procedures have been restricted to ethanol bathes, ultrasonic
bathes and oxygen plasma cleaning, which are standard processes for silicon chips. Immobilization of single-stranded probe sequences has been
achieved by exposing the electrode surfaces for 72 h to a 3 µM molecule
concentration in a phosphate saline buffer solution (10 mM phosphate
buffer and 0.3 M NaCl). However, a much shorter immersion time
can be sufficient, provided a proper optimization of the chemical conditions; in particular a 90% layer formation [Gergiadis, 2000] can be
obtained with a 5 h functionalization process in a 1 M NaCl buffer.
Nevertheless, an ideal immobilization process would be driven by microspotting or inkjet printing techniques to limit the wet chip surface.
Probe molecules have been synthesized as 30-base thiol-modified oligonucleotides (5’gatcatctacgccggacccgggcatcgtgg-3’) with (CH2 )6 − SH at
the 5’ position. After exposure the electrodes have been rinsed thoroughly with Milli-Q deionized water and immersed in the annealing buffer
(0.01 M EDT A, phosphate 0.01 M and 0.5 M NaCl) for electrical measurement. Hybridization has been performed at room temperature by
spreading overnight on the chips a 3 µM sample molecule concentration
in the annealing buffer. Molecules have been expected to bind to probes
on the surface thanks to their strand of complementary sequences (target
sequence: 5’-ccacgatgcccgggtccggcgtaga-3’). The chips have been rinsed
with the same buffer and measured again. The processing steps outlined
above can be used as the basis for buildin a front-end biofunctionalization process, which is fully compatible with standard CMOS-type silicon
Experimental Results
The results have been described by means of an equivalent circuit to
investigate the role of parasitic capacitances, interface charge-transfer
resistances and series resistances; the stability and reproducibility of the
capacitive elements involved in detection. In addition, we have also investigated another measurement approach, comparing the result obtained
with the impedance analyzer to those obtained by means of the chargebased capacitive technique described in [48, 40], that is suitable for integration on the same chip as the capacitive sensors (although in the
present work no attempt has been made to integrate sensors and reading
Complex impedance measurements
The bio- functionalization process introduces insulating layers between
electrode and solution and the value of the capacitive component of the
surface impedance is strongly affected by their chemical-physical characteristics. When DNA hybridization occurs, target molecules are blocked
near the surface and change the structure of the interface. In particular
they increase the distance between electrode and solution, thus modifying surface impedance. The aim of the proposed DNA sensing technique
is to measure two-electrode cells and to analyze the capacitive behaviour
of the interfaces before and after hybridization to observe the change that
are lead by hybridization.
A high precision impedance analyzer (Agilent 4294A) has been employed to characterize electrochemical cells formed by two micro-gold
electrodes integrated on the same chip and immersed in a conductive
solution. The cells have been electrically characterized. In the frequency
range from 100 Hz to 1 MHz (400 measurement points distributed in
logarithmic scale) a 20 mV sinusoidal signal amplitude has been applied
between the two electrode. For each frequency, data describing complex
admittance Y (ω) have been obtained.
Y (ω) =
= G(ω) + jCP (ω)
Where ω is 2πf . The Nyquist plot (−Im(Z(ω)) vs Re(Z(ω))) corresponding to a system of two identical gold electrodes is traced in Fig. 4.3
starting from 1 MHz down to 100 Hz. The experimental points in the
frequency range from 1 MHz to 10 kHz have been traced in Fig. 4.4
a describing an almost complete semicircle. For lower frequencies a second capacitive element dominates and the resistance in parallel to such
capacitance is very high, since the second semicircle is left incomplete
for frequency higher then 100 Hz. The behaviour of the two-electrode
system can be interpreted with a simple four-element equivalent circuit
(see Fig. 4.5) [42], that has been used to fit the experimental points as
shown in in the same figure (Bode plot). A small capacitive element
(CG) is originated by the geometrical capacitor made of the two electrodes and by other parasitics as wire capacitances. This value depends
on the configuration of the two electrodes and by the permittivity of the
solution. Moreover, the solution is conductive and it gives a resistive parameter dependent on salts concentration. This contribute can be added
to the other parasitic series resistance of the system (wires) and counted
in a general parameter RS . When the interface is not exchanging charge
with the solution, the two gold/solution interfaces are characterized by
a high resistance parameter (RIN T ) and a capacitive parameter (CIN T ).
RIN T is expected to be very high for microelectrodes and CIN T , which
depends strongly on the bio-layer characteristics, is expected to have a
surface density between 0.1 and 1 pF/µm2 . The capacitive elements of
the interface CIN T is frequency dependent. Nevertheless, they can be
modelled and extracted as Constant-Phase Elements (CPEs), which can
be described by a couple of parameters Q and α as [MacDonald, 1987]:
CP E =
The Z(ω) is described by two main RC contributes: a higher one
related to CIN T and RINT and a smaller one that comes from CG and
RS. According to the complex impedance theory [MacDonald, 1987], the
RIN T -CIN T couple will be measured at low frequencies, where RC of
higher value dominates. These parameters are defined by the contributes
of both interfaces in solution. In particular, if the electrodes were identical, CIN T would be the half of each single interface capacitance, while
RIN T would be the double of each interface resistance. Results have been
obtained on chip STM CH268A with an electrode configuration employing one of the electrodes of the array (see Fig. 4.1) and the surrounding
counter one. The CPE originated by the gold/solution interfaces can
be extracted by fitting the Bode plot of −Im(Z(ω)) at low frequency.
Fig. 4.6 shows the Bode plot of the imaginary part related to a chip with
non-modified gold electrodes and a chip modified with single-stranded
receptors chains. A decreasing value of (Q) after bio-modifications has
been found, as already reported [40]. This difference can be observed at
low frequency where the interface parameters plays a dominant role. Experimental curves related to bio-modified gold electrodes before and after
complementary DNA recognition have been plotted (Fig. 4.7). From the
extracted values reported in Tab 4.1 (extraction frequency range: 40 Hz2000 Hz) a further decrease of (Q) can be deduced after target molecules
surface adsorption. The stability of bio-modified gold surfaces has been
tested over time. The imaginary part of the impedance is plotted in
Fig. 4.8 for three measurements taken every fifteen minutes. For chip
STM CH268A, impedance plots of gold electrodes before immersion in
solution have been traced to define the role of resistance and capacitance
parasitics. Geometric capacitance can reach 80 pF when measuring connecting an electrode and his counter electrode. Moreover, 10 pF have
been measured for geometric capacitance of two near gold array electrodes. Parasitic capacitances are not negligible compared to interface
constant phase elements. However, Impedance Spectroscopy can easily
separate in frequency these capacitance contributes thanks to the large
difference between the two resistances. On the contrary, a total capacitance measurement technique could not. In this case, issues related to
the parasitics variability between different chips should be taken into
account as well as their stability during time.
Integrable ciruitry
This section investigate the possibility to measure the variation of capacitance induced by DNA hybridization by means of a simple circuits, suitable for easy integration in standard CMOS technology, in the prospective of a fully integrated, stand- alone DNA micro-arrays.
The results of the previous section suggest that interface capacitance
components in microfabricated devices should be significantly affected by
DNA hybridization. Therefore, we have applied the Charge-Based Capacitance Measurement (CBCM) technique [40] to the microfabricated
devices of this work to verify that it is not invalidated by the presence
of large parasitics and unexpected phenomena at chip level. The capacitance value obtained with this technique accounts for the total capacitance of the two-electrode system [40], which is a composition of all
capacitive contribution of the chip. Moreover, the interface capacitances
CIN T has a non-ideal capacitive behavior and they are correctly described
by a Constant Phase Element (CPE) which takes into account the frequency dependent behaviour. Consequently the measured result of capacitance values represents an average effect. Consequently, no attempt
has been made to compare absolute capacitance value obtained with
the two employed measurements set-up. Nevertheless, consistent variation after DNA detection of the Q parameter extracted from impedance
measurements and the total capacitance obtained with the CBCM setup
have been found. The CBCM technique has been employed to detect the
decrease of the interface capacitive component after hybridization. Two
couples of identical array-electrodes of chip STM CH402A have been contacted and immersed in solution to form the electrochemical cell. The
results obtained are reported in Tab. 4.2 showing a capacitance decrease
between 18% and 37% when complementary sequences hybridize with
probes on the surfaces after an overnight complementary-sequence solution exposure. When non-complementary sequences are used or when no
DNA at all in present in the sample, no decrease of interface capacitance
can be detected but only a slight increase.
Table 4.1: Extracted values of interface CPE before and after hybridization. A decrease of the Q parameter is observed, while an increase of the
parameter is observed.
Before Hyb.
After Hyb.
Q (pF)
Q (pF)
Table 4.2: Percentage decrease of two-electrode interface capacitances
after specific binding between probes on electrodes and complementary
target sequences. In case of no complementary sequence or of DNA-free
sample no descrease can be detected. Instead a slight increase can be
observed. The last row shows the negligible decrease of capacitance due
to non specific adsorption of DNA target on bare gold electrodes (no
probes where previously attached on the surface).
Tested sample
Decrease of Capacitance (%)
Complementary sequence
Complementary sequence
Complementary sequence
Non-Complementary Sequence
DNA-free sample
DNA-free sample
DNA target on bare electrodes
Figure 4.1: Atomic Force Microscopy Measurement of the surface of the
chip in correspondence of gold exposed electrodes.
Figure 4.2: Basic Process of bio-modification of gold microfabricated
electrodes fro probe deposition, target detection and successive desorption.
Figure 4.3: Evidence of extensive damages caused by piranha cleaning of
the chip surface.
Figure 4.4: Nyquist plots (−Im(Z(ω)) vs Re(Z(ω))) of a system composed by two identical gold electrodes for frequency range 1 MHz100 Hz. The experimental points in the frequency range from 1 MHz
to 10 kHz have been traced in part (a) describing an almost complete
semicircle. For lower frequencies a second capacitive element dominates
and (b)the resistance in parallel to such capacitance is very high, since
the second semicircle is left incomplete for frequency higher then 100 Hz
(see c).
Figure 4.5: 5 Fit of complex impedance measurements (b) performed
with the simple four-element circuit of (a). Extracted electrical parameters to fit the experimental points in the whole frequency range are: RS =
309020Ω; RP = 2.3751011Ω; QIN T = 2.88710−10 F ; αIN T = 0.91873;
QG = 6.62610−11F ; αG = 0.89826.
Figure 4.6: Bode plot of the imaginary part related to a chip with nonmodified gold electrodes and a chip modified with single-stranded receptors chains. A decreasing value of (QIN T ) after bio-modifications has
been found.
Figure 4.7: Two couples of bio-modified gold electrodes have been tested
before and after complementary DNA recognition. A decrease of electrode interface CPE due to the change in chemical and physical characteristics of the sensing biolayer.
Figure 4.8: Imaginary part of the impedance plotted for three measurements taken every fifteen minutes.
Chapter 5
Smart sensor on PCB based
on µ-controller for genetic
Smart sensor, i.e. embedded systems integrating transducers, signal conditioning and processing, represent the end point of an evolutionary path
that starts from the discovery and understanding of the sensing principle, followed by the development of laboratory set-ups based on high-cost
measurement equipment. This paper describes the design and implementation of a smart sensor for DNA hybridization detection. In particular,
as far as the sensing principle is concerned, this work is based on twoelectrode measurements of interface capacitance variations induced by
hybridization of target DNA with suitable probes immobilized on the
capacitors’ electrode [40]. So far, this sensing and transduction principle
has been studied using laboratory instrumentation for impedance measurement and the generation of input waveforms. All the required signal
generation, conversion and processing circuitry has been designed and
integrated on a single board (PCB) by means of standard components
(COTS). The result is a low-cost, low-power and compact system, with
significant advantages compared to the complex and expensive optical
scanning devices used today to read state-of-the-art DNA micro- arrays.
Thus, the smart sensor presented here represents a significant step forward toward the long-awaited new generation of low-cost point-of-care
systems for genetic screening and diagnosis [49]. As for performance, the
system descrided in this chapter has been compared against a measurement setup based on laboratory instrumentation both for the cases of
test capacitances, as well as for the detection of DNA in real operating
conditions. Results have been satisfactory in both cases. In particular,
the smart sensor is as accurate as the laboratory setup for measuring test
capacitances, and it has been shown to reliably detect DNA hybridization
in a number of experiments with real biosamples.
State-of-the-art and detection principle
Biosensors, i.e., sensors incorporating biological elements, are finding increasing use in clinical diagnosis and biomedicine, food production and
analysis, bacterial and viral analysis, and many others. Within this
broad field, DNA sensors aim at the detection of the highly-specific
binding reaction between two molecules of DNA (hybridization). A detailed treatment of the biochemical basis of hybridization is beyond the
scope of this paper (the interested reader is referred [50] for an excellent introductory treatment). For our purpose, it is sufficient to understand that DNA molecules can be described as linear sequences of
four different basic components (called nucleotides). As known, DNA
can be found in two states: single-stranded (i.e., a single linear sequence) and double-stranded. During hybridization, two single-stranded
molecules bind to become a double-stranded molecule. This chemical
reaction takes place (with high probability) only if the nucleotide sequences of the two single-stranded molecules match on a nucleotide-bynucleotide basis. DNA biosensors use custom-designed single-stranded
DNA molecules as recognizing elements (probes) for sub-sequences in
the DNA of test samples. If hybridization between the probe and the
sample DNA can be reliably detected, then the genetic signature of the
organism can be determined by virtue of the one-to-one affinity relationship. State-of-the-art of DNA (gene) sensors for high throughput analysis
are microfabricated matrices (microarrays) requiring labeling of the DNA
targets in order to generate optical or electrochemical signals [4]. However, the fluorescent molecules commonly used for this purpose require
expensive external instruments. Furthermore, the need of sample pretreatment as well as increased background noise represent further drawbacks [6]. Several label-free approaches have been intensively studied
in recent years, such as, for instance, micro-gravimetric and the electrochemical approaches, measuring the molecular-layer mass difference of
the molecules before and after hybridization [51, 52], and the changes
of electrical properties (impedance) of the substrate/electrolyte interface, respectively. The latter approach has been implemented starting
from standard methodologies for electrochemical interface analysis, such
as chronoamperometry [53], impedance spectroscopy [29], or exploiting
field effect devices [23]. The main disadvantage of these techniques is the
use of three electrodes (two for the measurements and one as potential
reference) that increases the complexity of the system. Two-electrodes
alternatives, such as that of interest for this work, have also attracted
significant interest. For example, a two-electrode setup using 2.4mm2
gold- and a Ag/AgCl counter-electrode has been studied and found to
exhibit ideal capacitive behavior in the 0 − 100Hz frequency range [39].
From this point of view, an important advance in the state-of-the-art has
recently been made [40] demonstrating that DNA hybridization can be
detected by measuring interface capacitance with a two-electrode system
(avoiding the use of a reference electrode). Our work builds upon these
As anticipated, our system is based on an electrochemical approach
featuring the immobilization of single-stranded DNA molecules (probes)
of known sequence on metal electrodes. This process is normally called
surface bio-modification or electrode functionalization. The sensing process is performed by dipping the electrodes in a physiologic solution
containing target DNA molecules to be captured by the probes in case
of sequence matching (hybridization). The binding event changes the
impedance of the electrode-solution interface [40], and in particular its
capacitive component, which is strongly affected by the type of molecules
attached to the surface (single or double-stranded). The interface can
be modeled with capacitance and resistance as shown in Fig. 5.1. The
presence of DNA molecules at the electrode surface strongly affects the
chemical-physical characteristics of the electrode-solution interface. We
are interested in the capacitance variation that takes place when target
DNA molecules in a liquid sample are captured by known DNA probes
on the electrode surface. Measuring the capacitance changes induced by
hybridization implies recognizing the occurrence of hybridization, hence
the nucleotide sequence of the target DNA.
Capacitance variations can be measured in many ways [54, 29]. In
this work we have used a Charge-Based Capacitance Measuring (CBCM)
technique, which has been shown to provide high accuracy even the case
of small on-chip parasitic capacitance [55]. The principle of this technique
is to charge and discharge the impedance under test, using a voltage pulse
of known amplitude ∆V at an appropriate frequency f (period T = 1/f ),
and to measure its equivalent capacitance from the average current in a
half-period (Fig. 5.2).
In fact, if interfaces can be modeled as simple capacitors, the average
current can be expressed as:
= 2C∆V f,
T /2
T /2
where IAV G is the the measured average current, and C is the only
unknown quantity. In our case, in order to increase precision and reduce
noise capacitance extraction is performed by measuring IAV G for different frequencies and calculating the slope of the linear interpolated curve
IAV G vs f . To this purpose, the current signal is converted into a digital
voltage signal and processed by a µ-controller to implement averaging
operations and determine the capacitance value. The main advantage
of CBCM technique compared to standard complex impedance measurements based on analog sinusoidal signals is that it lends itself to a simple
implementation using digital input signals. In fact, the analog part can be
reduced to an I/V converter followed by an A/D converter. A laboratory
setup implementing the CBCM technique requires a signal generator, to
control the switches, and a current meter.
Hardware and Software Design
The purpose of our design is to develop a low cost system to replace all
the laboratory instruments that are connected to the cell, as well as the
data-processing software running on the PC. The output of the system
is in digital form, and can be either directly visualized on a LCD or forwarded to other subsystems through standard digital links (e.g. RS232,
USB, etc.). Our implementation is based on µ-controller directly processing the signals produced by the measurement cell. The µ-controller
also controls pulse generation (frequency and amplitude) by mean of a
DAC converter, processes data sampled by the ADC converter and calculates IAV G in charge and discharge cycles. In the next sub-section we
describe in details both the hardware and software design of our smart
sensor. The system is based the Atmega128 µ-controller [56], controlling a DAC device, Max536 [57], and an ADC device, MAX186 [57],
(Fig. 5.3). The I/V conversion is performed with a precision resistor and
before digital conversion the output signal is amplified by means of an
OpAmp (TL081) [58].
At the end of the measurement, the capacitance value is available
in digital form on a serial interface. More in details, the Atmega128 is
an 8-bit µ-controller, with low power consumption and very low cost.
It is based on RISC architecture with 32 8-bit general-purpose registers, 6 I/O configurable ports, one SPI interface and two serial ports
(USART protocol). It also provides useful devices for embedded usage,
like counters and a timer, associated with interrupt generation. The
µ-controller is the core of the system, since it controls time and synchronization among all components. The possibility to program this device
is essential to guarantee the flexibility of our smart sensor. The DAC
features four out-channels and 16 bits program word. The output signal
is a 12 bits data which represents a voltage within the range 2.048V and
−1.2V as inputs of the cell. The upper value (2.048V ) is generated by
an ADR290 [59] and the lower value (−1.2V ) by means of a zener diode
LT1034 [60]. The control word is sent every time it is necessary to change
the output signal, exploiting a standard SPI protocol. The TL081 component, used to amplify the output signal, is a low noise OPAMP and it
is connected with a variable feed-back resistance to regulate gain. The
ADC is a 12 bits and 8 channels converter. In case of both positive
and negative values (as in our case), a 2’s complement integer between
−VREF ADC /2 to +VREF ADC /2 is generated. The resolution of the ADC
is 1mV with VREF ADC = 4.096V . At every sampling instant, the µcontroller sends 8 control bits and receives 16 data bits of which only 12
are significant. Flexibility and portability of the sensor are guaranteed
by the programmability of the system (programmed in C and Assembler). The µ-controller determines frequency and amplitude of the input
signal, generated by the DAC, calculates the average current for each
frequency, hence the capacitance according with Eq. 5.1. As anticipated,
the CBCM technique assumes that charge and discharge transients are
completed in each half-period of the input signal: hence, the maximum
frequency that can be used depends on the capacitance value of the cell.
For this reason, a preliminary step is necessary to determine the maximum frequency (fM AX ) of the input waveform (the typical range is from
5Hz to 100Hz and the system does not perform measurements out of
this range). In standard lab set-up, this step is usually performed manually or semi-automatically. In our system, instead, a fully-automated
self-tuning procedure has been implemented that, starting from a given
upper frequency, verifies if the transient are completed by comparing
partial areas of the output signal in the proximity of the end of the half
period. If the difference between these areas exceeds pre-defined thresholds, the frequency is decreased and the procedure is repeated (Fig. 5.4).
When the system finds the maximum frequency, the test is stopped and
a variable is set. This self-tuning procedure, lasting about 1s, ends with
the generation of a vector of suitable input frequencies.
After self-tuning, capacitances can be correctly measured. To do that,
it is necessary to calculate average currents for different frequencies and
then calculate capacitance from equation 5.1. In order to measure the
IAV G parameter, the system measures an average voltage by mean of
trapezoidal integration,
N −1
1 X
V dt =
ci+1 + ci ·
T i=1
where V is the voltage output of the cell, ci is the i−sample, N is the
number of samples for every half-period, and h is the distance between
two samples (the voltage output signal is sampled at a maximum rate
of 40kHz and converted to a digital value). The VAV G parameter is
proportional to average current thanks to the resistor (RL ) between the
output cell and ground (Fig. 5.3):
Moreover, the IAV G can be written taking into account the gain due
to the OP-AMP (Av ) and the same resistor RL as:
= 2C∆V f RL Av = mf,
T /2
T /2
where m can be evaluated by the slope of the IAV G vs f plot by mean
of regression linear method. Finally, the capacitance can be calculated
2∆V Av RL
where RL is the current-to-voltage resistance as previously indicated.
Capacitance is calculated sampling 10 half-period voltage transients.
Each transient is sampled up to 250µs before the end of the half-period
to enable the µ-controller interrupts, disabled during ADC sampling. In
fact, an internal counter generates an interrupt twice in a period to warn
the µ-controller to change the DAC output voltage. Nevertheless, if the
frequency has been defined according to the previous tuning process, the
transient is expected to end at least 625µs before the end of the halfperiod.
Experimental results
Since cheap and off-the-shelf components can be satisfactorily used, the
cost of the system is approximately 100 times smaller than that of a standard laboratory set-up. The power consumption is 520mW in both idle
state and operating state without important variations during measurements. The system has been tested on a wide range of discrete capacitances (from 10pF to 4.7µF ) and it has also been employed to determine
the equivalent capacitance of an electrochemical two-electrode cell to detect DNA hybridization. Fig. 5.5 shows a photo of the system set-up.
The tuning procedure has been tested with 2.2µF and 4.7µF discrete
capacitances to verify the correct estimate of maximum frequency.
The tuning procedure determines the maximum frequency allowing the
RC circuit (load resistor-cell capacitance) to be completely charged and
discharged. This is the highest frequency for which ( 5.4) is verified.
Hence, the linear relationship between f and ∆V predicted by Equation 5.4 is observed for all frequencies lower than fM AX . The value of
fM AX can be observed on the IAV G −f plot (Fig. 5.6) clearly showing the
deviation from the linear behavior at high frequencies due to incomplete
charging and discharging of the capacitance. In order to confirm the robustness and accuracy of the self-tuning procedure, we have evaluated
the deviation from the linear behavior at the frequency determined by
the system which results always below 1%.
Electrical characterization
Our system has been tested and compared to National Instrument PXI [61]
measuring discrete capacitances of 10pF - 4, 7µF , that corresponds to the
expected electrode area ranging from 2∗103 µm2 to 1cm2 that are characteristic of macro and micro electrodes, respectively. The macro-electrode
range was chosen from 350nF to 4.7µF while the micro-electrodes range
was chosen from 10pF to 700pF , according to available data [53] which
associates 20µF capacitance for 1cm2 to an electrode/solution interface.
The current to voltage conversion resistor was 750Ω for larger capacitances and 2.3MΩ for smaller ones, respectively. For large capacitances
the accuracy of our smart sensor and of PXI are comparable, corresponding to a relative percentage error below 3, 5%. For lower capacitances
the smart sensors achieves better accuracy than PXI. In particular, for
a 150pF capacitance value, relative error is 0.5% and 1.5%, respectively.
For low capacitances a calibration curve is necessary for both set-ups
to compensate parasitic which can be estimated 50pF for the PXI and
2 − 3pF for the smart sensor. Therefore, the calibration is necessary in
particular for the implementation based on the PXI (50 pF compared to
2 − 3pF ). The reduction of parasitic is a consequence of the more compact on-board implementation. Nevertheless, exploiting the calibration
curves it is possible to eliminate the systematic error.
DNA Detection Measurement
The system of this work has also been employed to detect when the hybridization reaction occurs. The system has been connected to a cell
of two 0.8cm2 gold electrodes immersed in an electrolyte solution. The
equivalent capacitance measured with our setup of in case of bare gold
electrodes is 19.04µF/cm2 (as expected from literature [53, 29]). Biomodification of the gold surface leads to an interface capacitance decrease. The entity of the variation is related to the biochemical nature
of the molecular layer formed on the surface and to its conformation. In
fact, a lower capacitance value has indeed been measured for the cells
that hosted bio-modified electrodes.
The cell used for this experiments was composed by electrodes modified overnight with single-stranded DNA probe molecules (3µM in saline
solution). Successively, experiments were performed through the following steps. 1) The capacitance of the cell has been measured. 2) The cell
has been immersed in a 70◦ C solution containing target molecules and
cooling down the cell to room temperature. 3) Capacitance has been measured at the end of the binding reaction after extensive rinsing with saline
solution to remove unbound molecules. 4) The sample molecules can be
removed by heat-treatment (electrode regeneration). 5) The equivalent
capacitance of the regenerated electrodes has also been measured and
compared with measurements at steps 1 and 3.
As shown in Fig. 5.7, four repeated measurements have been performed in 6 minutes on the same electrodes couple for each step and the
average values are compared. As expected, measurements at steps 1 and
5 both correspond to electrode interfaces modified with a layer of probe
molecules: in fact their capacitances differ by less than 1.31% (3.10 and
3.06µF respectively). In contrast, the presence of target molecules captured by probes on the interface (measurements at Step 3) leads to a
capacitance decrease of 16, 5% (2, 59µF ).
Therefore, the described smart sensor has several advantages with re-
spect to commercial systems based on optical scanners. First, it avoids
the use of high-cost detectors and the pre-treatment of the sample to
introduce marker molecules. Moreover, the programmability of the system µ-controller guarantees the flexibility and self-tuning capability of
the smart sensor. In particular, the system could host and manage additional components, such as a wireless interface or other sensors. The
system has been tested and compared in terms of accuracy with a highcost laboratory implementation of the same charge-based capacitance
measurement technique, showing competitive performance. Finally, capacitance measurements performed on typical set-up, show the capability
of our system to detect DNA hybridization reaction occurring in on the
surface of gold electrodes.
Figure 5.1: Electrical model of Electrode/Solution interface.
Figure 5.2: Schematic description of Capacitance-Based Capacitance
Measuring technique.
Figure 5.3: Sensor block diagram where are indicated connection between
the core of the system, µ-controller, and the devices controlled by it, ADC
and DAC.
Figure 5.4: Tuning principle.
Fluidic cell
Gold electrode
Figure 5.5: The photo shows the PCB implementation of the system and
the fluidic cell used for the experiments.
4,7 µF
2,2 µF
f, Hz
Figure 5.6: Graphical evaluation of fM AX . A deviation from the linear behavior is observed around 40Hz and 70Hz for 4.7µF and 2.2µF
respectively, which indicate that for highest frequencies capacitance measurements cannot be correctly performed.
Capacitance, µF
Single-Stranded DNA
DNA De-Hybridization
DNA Hybridization
Time, min
Figure 5.7: Experimental results of hybridization and de-hybridization
capacitance measurements.
Chapter 6
On Chip DNA detection
based on CBCM capacitance
In the last decade, miniaturized arrays for gene-based tests (known as
DNA microarrays) have been introduced in genetic research. These are
fabricated on glass or quartz slides where different DNA probe molecules
are immobilized in a two-dimensional array of small sites. Some of these
devices, implemented with photolithographic techniques, are able to test
a whole genome, with densities of a million sites per square centimeter [4]. The DNA strands of the organism under test are preliminary
marked with fluorescent molecules. Their presence at specific sites of the
array is measured by an optical scanner or a fluorescence microscope,
after binding to matching probe molecules. These devices have also been
employed for population genotyping [62] and research on cancer predisposition [63]. Nevertheless, the high-cost of the scanner and the processing
steps required to tag the samples with markers (also called labels), pose
critical limits to the use of these tools in point-of-care analysis. For
these reasons, a large effort is devoted to the development of devices
suitable for low-cost mass production and ease to be used not only in
highly specialized laboratories. A single-chip solution, implementing a
detection technique with direct electrical read-out and avoiding labeling
of the DNA target molecules, would significantly enhance portability, as
well as high-parallelism, on-site sensing and data processing.
The chip features a number of sensing sites, each containing two gold
electrodes [64] that can be independently selected by means of on-chip addressing circuitry, while the read-out electronics has been realized essentially by means of external standard components in order to exploit the
advantages of large experimental flexibility. The surface of the electrodes
is bio-modified (functionalized) by covalent binding of single-stranded
DNA probes. During in-field operation, the capture of complementary
DNA strands from the liquid samples is signaled by a variation of the
capacitance between electrode pairs, which can be measured by connecting them (through on-chip addressing logic) to an external capacitancemeasurement circuit.
The chip is bonded on a PCB (Printed Circuit Board) and is connected to external processing and storage devices, resulting in a complete
analysis system that has been characterized and tested for DNA recognition. To this end, the capacitor electrodes have been functionalized and
their capacitance variation has been measured after exposition to both
complementary and non-matching target molecules (to test specificity).
Reversibility has also been tested by measuring sensor capacitance before and after sensing molecule regeneration, obtained by temperature
treatment. The results demonstrate that a DNA sensor based on the
electronic detection technique described in this paper can be fabricated
as a single chip with standard CMOS technology augmented with gold
deposition for the capacitor electrodes.
Related Work
DNA microarrays allow highly parallel and low-cost analysis. In fact,
they exploit the capability to fabricate a large number of miniaturized
detection sites on a substrate and to extract information after exposition
to the target DNA to be detected/recognized. In general, such devices
work as described in Fig. 6.1. Each site is specifically bio-functionalized
by means of DNA probe molecules of known sequence immobilized on
its surface. Target molecules in the sample solution bind only to probes
with complementary sequences (hybridization): thus, their presence at
specific sites reveals their composition.
Most of current microarrays require optically active labels attached to
the target DNA and fluorescence sensing. Some innovative microarrays
employ electrochemical labels (redox-active molecules or enzymes) that
produce an electrical current through conductive electrodes in case of
hybridization ([65, 66, 19]).
Label-free techniques are intensively investigated, since they avoid
expensive reagents and pre-treatment steps. To this end, recently, several
approaches have been proposed, based on piezoelectric materials ([24,
67]), microfabricated cantilevers ([68, 69]). Other techniques detect the
variations of electrical properties of electrode/solution interfaces induced
by DNA recognition. Within this category, a distinction can be made
between devices where sites exploit passive components or semiconductor
sensors ([21, 22, 16, 41, 23, 70]).
The approach of interest in this work is based on interface capacitance
measurements, whose conventional implementation exploits a three-electrodes
setup allowing bio-functionalized gold electrodes to be investigated with
impedance spectroscopy ([29] [47]).
Much simpler two-electrodes configurations, such as that used in
this work, have also attracted significant interest. For example, a twoelectrode set-up using 2.4 mm2 gold- and a Ag/AgCl counter- electrode
has been found to work properly in capacitive detection at 20 Hz [39].
As far as integrable capacitance methods are concerned, it has already
been shown that DNA hybridization can be detected by measuring interface capacitance with a system making use of two (gold) electrodes only
(i.e. without the use of a reference terminal) [40]. Furthermore, more
recently it has been shown that reliable measurements can be done exploiting passive microfabricated electrodes [71]. Arrays made of passive
electrodes are pad-limited, as each electrode pair requires two pads for
connecting to external measurement circuits.
In such a context, this work represents an advance in the state of
the art in that it presents a silicon chip featuring micrometric capacitors
and on-chip logic for selecting an multiplexing the on a limited number of I/O pins. The chip has been implemented with standard CMOS
technology (with the addition of gold deposition for capacitor electrodes).
Capacitance measurements are performed by means of the Charge-Based
Capacitance Measurement (CBCM) technique [72], which is suitable for
low-complexity, small-area on-chip implementation.
The chip includes addressing circuitry, while the read-out electronics
has been realized externally by means of standard components in order
to exploit the advantages of large experimental flexibility (although it
could eventually be easily integrated with the rest of the system).
Capacitance-based DNA Detection Principle
Under suitable electrochemical conditions, bio-modified metal interfaces
in saline solution exhibit a capacitive behavior. This is the case if gold
electrodes are modified with short DNA strands immobilized with alka-
nethiol linkers on the surface ([39, 29]). In this case, capacitance values of electrode/solution interface have been estimated between 1 and
20µF/cm2 (although this value strongly depends on electrode surface
treatment). A first-approximation model of the electrode/solution interface is the equivalent circuit depicted in Fig. 6.2 (left) ([42, 73]), where:
RS depends on the interface and solution characteristics; RP is related to
the insulating properties of the interface (for well-formed layers it can be
considered infinite); CP is mainly affected by the physical and chemical
characteristics of the insulating bio-layer immobilized on the surface. In
this model, the geometrical capacitance formed by the electrodes would
be in parallel to CP but, since it is several orders of magnitude smaller
than the interface-capacitance, its contribution is negligible: thus it is
not considered here.
This simple model provides an intuitive explanation of the basic sensing principle exploited in our work. Considering a capacitor formed by
two neighboring electrodes of our chip, when a complementary DNA
strand binds with the surface probes and the DNA duplex is formed,
the solution counterions present in the liquid solution are pushed further away from the polarized metal surface: this increases the distance
between the charge inside the electrodes and the ions in the electrolyte:
thus CP decreases ([47, 39, 40]) (Fig. 6.2, right).
When measuring a capacitance with CBCM technique, an input square
wave is applied to one electrode, while the current required to completely
charge (or discharge) the capacitance is measured at the second electrode.
The measured capacitance value equals the product of the average value
of: the output current, the applied voltage step, the frequency of applied
clock signal. With this technique, the measurement of the capacitive
component of an impedance is independent from the resistive components provided that: (a) the charging and discharging current transients
have the time to settle during the period of the square wave (i.e. all the
charge stored in the reactive component is transferred); (b) the static current flowing through the RP and RS resistive path is negligible compared
with the average output current (i.e. most of the charge transfer is due to
the reactive component). Previous work on CBCM has demonstrated
that both these requirements can be met in a practical experimental
setting [40].
It is important to stress however that the model of Fig. 6.2 is a highly
idealized view of more complex physical phenomena, namely charge transfer at the electrode-solution interface in presence of biological molecules
immobilized on the electrode surface. Thus, we cannot assume, for instance, that capacitances and resistances in Fig. 6.2 are constant over a
wide frequency range. Sensing of impedance variations requires therefore
careful selection of the frequencies used for CBCM excitation waveforms.
This section describes the architecture and the operation of the test chip
designed, fabricated and characterized in this work. The chip contains
a number of sensing sites and the addressing circuitry, while read-out is
realized by means of by external components in order to systematically
explore different measurement configurations and parameters. The chip
is packaged on a PCB to form a system suitable for test measurements
and characterization.
Chip architecture
Essentials of chip architecture and signal flow are shown in Fig. 6.3.
Both analog and digital input signals are required. The former ones
determine the voltage range used for the measurements, while the latter
ones include a seven bit address to select the sensing sites. An external
clock is used to provide the measurement frequency, which is the same as
the external signal, as well as to synchronize the different blocks. Each
sensing site features a block that, controlled by external clock, generates
a four square-wave signal with high timing precision.
The on-chip address signals are generated by means of two decoders.
If one sensing site is selected, the input signals coming from the pulse
generator are enabled and measurement is performed. Only the output
signal of the selected site is connected to the shared output pad (Fig. 6.3).
At the output pad, the transient current is converted by means of
an external circuit into a voltage signal, that is sampled and processed
by a PC to calculate the capacitance value. With respect to an on-chip
implementation of the I/V conversion, this external solution allows larger
flexibility in experimenting different measurement parameters.
Sensing site circuitry
The circuit implemented for each sensing sites, illustrated in the dotted
box in Fig. 6.4, includes a pulse generator and four switches which selectively connect the electrodes of the sensing capacitances to one of three
different reference voltages: V +, V − and VREF . Between the pulse generator and the switches a multiplexer is inserted which allows the clock
pulses to drive the electrodes of the sensing element only if it is selected.
The switches are implemented with both n- and p-channel transistors in
order to exploit full voltage range.
Exploiting its proximity with the switches, the (local) pulse generators provide very precise signals controlled by the external clock. A
critical timing issue has to be considered at this regard in that Ck1,
Ck1− and Ck2, Ck2− must not close the switches at the same time in
order to avoid shorting V + and V −. For this reason, a suitable delay is
inserted and non-overlapping clock signals are used (as shown in the inset
of Fig. 6.5) in order to avoid that two transistors at the same time see a
gate voltage larger than their threshold. The delay used for this purpose
is 1 ns and is obtained by means of the pulse generator block reported
in Fig. 6.6, also showing the output buffer following the multiplexer.
The switches allow to connect the electrodes of the sensing capacitors
either to V − and to VREF , or to V + and the voltage about VREF at the
negative input of the opamp shown in Fig. 6.6 (see also Table 6.1). When
switching between these two bias conditions, a voltage step is applied at
the sensing electrodes and a transient current is generated. With the
circuits shown in Fig. 6.4, the charging and discharging currents are
driven to two different output pads, OUT 1 and OUT 2, thus allowing
measurement of the average value using only one of them.
Physical layout
The chip is fabricated in 6” n-well 0.5 µm CMOS process with two
metal layers. The gold electrodes are deposited after standard CMOS
processing on silicon nitride (Si3 N4 ) as described in [64]. Annealing
steps are introduced to guarantee good CMOS performance after gold
deposition [64]. The area of the three sensing capacitors used in this
work is 1 mm × 1 mm and the distance between them is 500 µm,
while that of total chip, featuring 44 electrode couples of different size,
is 6.5 mm × 4.6 mm (Fig. 6.7). The parasitic capacitance between gold
electrodes and substrate has been estimated to be 6 aF/µm2 therefore
it is several order of magnitude smaller than electrode/solution interface
capacitance (about 35 nF as shown in section 6.5).
The pin-out includes analog as well as digital signals, power supply
and ground. Table 6.1 provides all I/O signals, with indications of the
label and the value of the voltage applied during the measurements.
The main analog signals are V +, V −, determining the voltage step
(V +)−(V −) = ∆V ) applied to the electrodes. Digital inputs include the
seven bit address signals and one external clock to set the measurement
frequency. The power supply voltage VDD is the same for the whole chip.
In order to reduce interferences and noise, analog ground is VREF =
VDD /2. Our choice of VDD and VREF allows to generate only positive
signals (e.g. V +, V −), while the digital ground can be made negative
respect to VREF . In fact, the signal voltage levels VS1 = V + = VREF +
0.1V , and VS2 = V − = VREF − 0.1V are generated by external voltage
sources related to digital ground.
Measurement set-up
The chip described in the previous section is glued to a PCB featuring
large contacts on one side to be connected to standard instrument and
small ones on the other side connected to the chip by means of gold
bonding wires (Fig. 6.8). A passivation process protects both wires and
chip from the saline solutions used in the experiments. To this purpose,
epoxy glue is placed on the critical electrical parts. Since to bind the
DNA probes to the gold electrodes, these must be accurately cleaned
with plasma oxygen and epoxy glue is damaged by this process a further
SU-8 layer is added to protect the glue from treatment.
In our electrical measurements, the transient discharging current is
driven to an external I/V converter (Fig. 6.9) implemented off-chip by
means of the operational amplifier OP A103AM [61] and precision resistors. The voltage signal at the op-amp output is sampled by means of a
National Instruments DAQ board (PCI-6032E [61]). Finally, the average
value of the output voltage (Vavg ) is calculated by means of a trapezoidal
integration and the average discharge current (Iavg ) is then obtained as:
Iavg = Vavg /RC .
Iavg = 2C∆V f + Iof f set .
Iavg can be expressed as:
Here, ∆V , ((V +) − (V −)), and f are controlled parameters, while C
is the capacitance to be extracted from experimental data.
For higher precision, the value of Iavg is measured at 9 different frequencies (from 60 Hz to 100 Hz and step of 5 Hz for each measurement)
as the average of 10 charging transient.
Finally, C is computed by least squares fitting on equation 6.2 providing the slope m:
Experimental results
Electrical Characterization of gold electrodes
Our chips have been tested in wet conditions at room temperature. After
cleaning, a saline solution, TE (Tris 10 mM, EDTA 1 mM) 0, 3 M NaCl pH 7,
is spread on the electrode surface and CBCM measurements are performed to characterize differences between electrode couples. In a first
phase, test measurements are performed to set two important parameters, namely RC , which determines the I/V conversion factor, and the
frequency range to be used in the measurements.
From section 6.2, we recall that CBCM frequency should be low
enough to reach steady state in cell current, and that the DC current
flowing through the RP and RS resistive path should be negligible compared with the average output current. By analyzing current transients
in response to voltage steps, we determined that steady state was reached
in a few tenths of ms. We therefore set the period of the excitation waveform 1 order of magnitude larger than transient time. Moreover, RP and
RS have been estimated with a standard impedance spectroscopy measurement and found to be about 10 MΩ and 10 Ω, respectively. RP is
large enough to ensure that the DC current flowing between the electrodes falls below the noise level of the experimental set-up. Thus, both
requirements for applicability of CBCM are met1 .
However, the experimental measurements of CP , in contrast with the
first-approximation model of Fig 6.2, (left), have shown a large frequency
dependence (Fig. 6.10). The behavior can be attributed to electrochemical effects taking place at the electrode-solution interface of the
sensing device. These effects can, in fact, be accounted by modeling CP
with a Constant Phase Element [74]. Indeed, Fig. 6.10 reports a good fit
of our experimental data obtained with the described model.
It is worth pointing out that detection does not require determining
a single value of the sensing capacitance, but to distinguish between
the device behavior in the presence rather than in the absence of DNA
hybridization and, as shown later, this is indeed the case.
The I/V converter is implemented with the operational amplifier
OP A103AM [61] featuring a gain-bandwidth product of 106 , a value
To further strengthen this point, PSPICE simulations have been done considering the circuit represented in Fig. 6.2, (left), with a constant value of CP . With
this model, the capacitance is then extracted from the (average) transient discharge
current as done in the experiments (described in Section 6.4) and the results indicate
a dependence on f smaller than 7% in the frequency range from 10 Hz to 1 kHz).
of interest in the perspective of fully integrated electronics. This characteristic plays a crucial role in the choice of the frequency range. The
optimal value of f for the CBCM voltage excitation frequency to be used
in the measurements is chosen taking into account the frequency dependence shown in Fig. 6.10 and also the characteristics of the I/V converter.
Lower frequencies (hence longer periods) lead to long measurements, with
lots of noise. On the other hand, when frequency is too high, small gains
of the I/V converter lead to small signals, again critically affected by
In practice, if f is 1 kHz, the gain-bandwidth product of our amplifier
would require RC = 1 kΩ. Unfortunately, in this case, gain would be
rather low, and the Signal to Noise ratio (S/N) would be too low (about
5 dB). On the other hand, if f < 10 Hz, then RC > 105 kΩ and gain
would be high enough, but the measurements would be too noisy, as
shown in by the error bars of Fig. 6.10.
Therefore, the optimal condition of the I/V converter circuit is RC =
10 kΩ (leading to S/N = 34 dB and total measurement time around 20 s
for each site), and f in the range 10−100 Hz. The frequency dependence
of CP suggests to work at a measurement frequency as high as possible.
Therefore, we have used frequencies in the range 60 − 100 Hz.
Typical experimental results, obtained with equation 6.3, are shown in
Fig. 6.11 and in Table 6.2. ”Couple 1” refers to the electrode plates on the
left of the chip while ”Couple 2” refers to those on the right. The standard
deviation reported in Fig. 6.11 come from several measurements for each
sensing site. Relative differences between capacitance measurements for
two different couples on the same chip are evaluated to be less than
4.2%. Moreover, statistics on data shows that the capacitance values
measured on the same chip but on two different couples are almost the
same. Differences among several chip are due mainly to differences in
the level of cleaness of the electrodes after the cleaning step.
Electrode bio-modification
First, the gold electrodes have been cleaned by exposure for 20 minutes to
plasma oxygen at 200 W . After this step, single stranded DNA molecules
modified with alkanethiol groups are immobilized on the gold electrodes
by covalent S-Au bonds (a 3 µM DNA 1 M Na2 HP O4 solution is spread
on the electrode surface for 18 hours). Two different probe molecules of
the same length (25-mer) and thiol modified with a chain of 6 carbon
atoms as a spacer are bound to different electrodes on the same chip
(electrode couple on the left and on the right). To this end, two separate
droplets are placed by mean of microliter pipettes to obtain sites which
will and will not experience hybridization reaction. Before measurements,
the gold surfaces are extensively rinsed with ultra pure water to remove
molecules that are not covalently bound to the gold electrodes or to the
passivation layer of the chip.
Target DNA solution (3 µM DNA 30-mer and TE 0, 3 M NaCl pH 7)
is heated up to 80◦ C, spread on the electrodes and cooled down to room
temperature (for about 30 minutes). Finally the sample is rinsed in the
same saline solution (TE 0, 3 M NaCl pH 7) in order to remove the
unbound DNA target (Fig. 6.12).
In order to verify the biological steps previously described, we have
performed an independent standard optical detection test based on fluorescence molecules bound to DNA molecules. Probe molecules are labeled with fluorescein and target molecules with tetrametil-rhodamine.
In order to perform quantitative analysis after electrical measurements
the DNA molecules, both probes and targets have been removed from
the surface by mean of β-mercaptoethanol 12 mM solution for 10 hours.
Then, to increase the optical signal, the pH of the sample solution has
been brought to 10 adding a solution NaOH 10 M (the pH has been
checked with the HI-9023 pH meter). Before DNA experiment a calibration curve has been done with the LSB-50 fluorescence spectrometer
(Perkin Elmer, Boston, USA) using DNA sample solutions, (12 mM βmercaptoethanol and pH 10 adding NaOH10 M solution as previously
described) of different known concentrations. Results of optical detection of our unknown samples compared with the calibration curve show
that the density of our probes layer is about 1013 molecules/cm2 , a value
close to that reported in the literature [75]. Moreover, we have tested the
efficiency of the hybridization reaction in the case of complementary and
a-specific target molecules: the first one indicates that the 80% of the
probes react with target forming the double helix while the second one
indicates that less than 10% of probe react with target, demonstrating
the good quality of our process.
DNA detection
DNA detection is demonstrated comparing measurements on electrode
couples subjected to the same reaction but with different DNA strands
bound on the surface, complementary and non complementary to target
molecules, respectively (the latter for negative control). All measurements are performed in the same saline solution of the hybridization step
(TE 1X 0, 3 M NaCl pH 7). Since capacitances exhibit significant mis-
matches, a measurement after functionalization is performed and these
values are used as a reference to be compared with the results obtained
after (tentative) hybridization. The capacitance absolute values after
functionalization show a decrease around 25% respect to clean gold for
each couple. Results of 3 experiments over all the chip previously measured after cleaning are summarized in Fig. 6.13, where positive variations correspond to capacitance decreases.
Fig. 6.13 clearly indicates that significant capacitance decreases are
found as a result of hybridization, while smaller or even opposite variations occur in the other case, thus demonstrating the capability of
our approach to distinguish between hybridization and non-hybridization
After measurements, the sensing layers have been regenerated removing all target molecules. For this purpose, a rinsing step with hot ultrapure water is performed to break hydrogen bonds between the two DNA
strands, leaving only the probe oligonucleotide chains covalently bound
on the surface. After this step, new capacitive measurements are performed and Fig. 6.14 illustrates the measured capacitance variations.
As expected, an increase of interface capacitance is observed where
hybridization had previously occurred, while electrodes with non-complementary
probes show only negligible variations. Moreover, it is worthwhile to note
that the absolute variations of capacitance in de-hybridized chips are of
the same amount occurring in hybridized ones.
Since the applications envisaged in this work, normally make use of
DNA amplification by means of PCR-Polymerase Chain Reaction, tests
of the device sensitivity to DNA concentration , thus no attempt has
been made in this direction.
In view of a simple and effective DNA analysis, a significant effort
is currently dedicated to the development of single-chip, fully-electronic
devices and this paper illustrates a step forward in this direction since
it demonstrates the viability of a simple approach suitable for stand
alone and portable diagnosis equipments based on two-electrode capacitive measurements implemented with micro-fabricated capacitors.
In particular, this work presents a test chip, fabricated with standard
CMOS technology with the addition of gold deposition for the sensing
capacitors. The chip includes the addressing circuitry, while standard
external components are used to realize the read-out electronics, in order to improve measurement flexibility and help the search for optimum
The chip has been fully characterized and measurements have been
performed exposing the device to target DNA solutions. The results
clearly indicate that DNA hybridization produces a significant decrease
in the capacitance of the sensing capacitors, that can be safely recognized
by the on-chip read-out circuitry.
The paper then suggests that a chip, potentially containing thousands
of sensing sites, with fully integrated circuitry for DNA recognition can
be fabricated exploiting the approach described in this work.
Table 6.1: Description of the chip pinout.
N. Pin
0, VDD
0, VDD
Power Supply
V + = 2.6V
OUT 1 = VDD /2
V − = 2.4V
OUT 2 = VDD /2
Table 6.2: Capacitance measurement on 6 different chip
Chip 1 ⇒ CAP (nF )
Chip 2 ⇒ CAP (nF )
Chip 3 ⇒ CAP (nF )
Chip 4 ⇒ CAP (nF )
Chip 5 ⇒ CAP (nF )
Chip 6 ⇒ CAP (nF )
Standard Deviation (nF )
Figure 6.1: Schematic representation of microarray principle.
Figure 6.2: Left: Schematic representation of the lumped-element electrical model of the metal/solution interface of the capacitors used in this
work. Right: Schematic illustration of the DNA hybridization process
and the induced (further) displacement of counterions within the liquid
Figure 6.3: Schematic representation of the system and signal flow. A
and D indicate analog and digital signals, respectively.
Figure 6.4: Schematic of the circuit associated to each sensing element.
Figure 6.5: Schematic representation of the signals flow used in the experiments.
Figure 6.6: Schematic plot of the block used to generate not overlapping
clock signals.
Figure 6.7: Chip photograph.
Figure 6.8: Photo of the PCB used for measurements. The chip is highlighted in the box on the right.
Figure 6.9: Schematic representation of the measurement set-up.
Figure 6.10: Measured capacitance vs charge/discharge frequency on
clean gold electrodes. The continuous line shows the fitting.
Capacitance (nF)
Chip Number
Figure 6.11: Capacitance measurements of electrode couples on different
Figure 6.12: Schematic representation of the functionalization and hybridization steps.
Figure 6.13: Capacitance variations due to specific and a-specific bindings (upper and lower bands of measured capacitances, respectively).
Positive values indicate capacitance decrease.
Figure 6.14: Capacitance variations due to de-hybridization process
in the case of complementary (lower) and non-complementary (higher)
probes previously subjected to hybridization reaction. Negative values
indicate capacitance increase.
Chapter 7
Capacitance measurement for
DNA detection with on chip
A/D conversion
Miniaturized arrays for gene-based tests, known as DNA microarrays,
have drastically changed the way genetic analysis and research are performed, by enabling the user to perform a huge number of analysis in
parallel. These devices are typically based on glass slides where different
DNA probe molecules with known sequences are located on the surface
within a two-dimensional array. The DNA strands within a sample to
be analyzed are marked with fluorescent label molecules during preparation of the sample liquid. A label is an extra molecule added to a DNA
segment that reveals the presence of target DNA bound on a site of the
array. After binding to matching probe molecules during incubation of
the sensor array, the label’s presence within a specific site of the array
is detected by an optical scanner or a fluorescence microscope. Stateof-the-art optical DNA microarrays can test a whole genome, as they
achieve densities of a million sites per square centimeter [4]. Moreover,
they have also been employed for population genotyping [62] and research
on cancer predisposition [63]. Nevertheless, the high cost of the scanner,
the sensibility of optical systems and the processing steps needed to label
the samples pose critical limits to widespread diffusion and point-of-care
usage. For these reasons, significant effort is being devoted to develop
devices suitable for low-cost mass production and use outside highly specialized laboratories. A solution implementing direct electrical read-out
and avoiding labeling of the DNA target molecules would significantly
enhance portability while maintaining high-parallelism, as well as on-site
sensing and data processing. In fact, the miniaturization of sensor chip
and readout unit opens the way for completely new applications and
markets in the field of bio-molecular diagnostics. Due to the ease of use
of these highly integrated microsystems, point of care diagnostics, e.g. in
the doctor’s office, comes into reach.
The main technical advantage of monolithically integrated, fully electronic DNA sensor devices is the capability of signal processing in the
direct proximity of the sensor. This results in the highest sensitivity with
respect to the transducer signal. Furthermore, CMOS allows to integrate
large number of sensors on a single die requiring only few electrical connections to the outside world, which significantly eases the packaging of
the devices.
The focus of this paper is to present a CMOS DNA-chip featuring 128
sensing sites which implement a label-free fully-electronic capacitance
measurement technique. Each sensor site in the 8 × 16 array consists of
two interdigitated gold electrodes [64] and an integrated measurement
circuit. The single sensor can be independently selected by means of
on-chip addressing circuitry. The output of the measurement result is a
digital signal which can directly be read by a computer.
The surface of the gold electrodes is bio-modified (functionalized) by
covalent binding of single-stranded DNA probes. During in-field operation, the hybridization events of complementary DNA strands, probe and
target molecules, is detected by a variation of the capacitance between
the sensor site’s electrode pair, which is measured by a circuit below the
sensor electrode pair using a capacitance to frequency conversion.
For measurement and characterization purposes, the chip has been
bonded on a circuit board which is then directly connected to a computer.
For verification of the biochemical sensor operation, the electrodes have
been functionalized and their capacitance variation has been measured
both in case of complementary and non-matching target molecules in
order to test specificity. The results clearly demonstrate that the DNA
sensor array based on label free, capacitive detection technique can successfully detect specific DNA hybridization.
Related Work
DNA microarrays, allowing highly parallel and low-cost analysis, exploit
the capability to fabricate a large number of miniaturized detection sites
on a substrate and to extract information from each of them after exposure to the solution containing the target DNA. Each site is specifically
bio-functionalized, i.e. the sensor surface is equipped with single stranded
DNA oligonucleotides with a known sequence, which covalently bind to
the sensor’s surface. Target molecules in the sample liquid selectively
bind to probes with complementary sequence, resulting in hybridization
of the two oligonucleotide strands. Their presence at specific sites reveals
the composition of the sample solution.
Most of today’s detection techniques used within microarrays require
molecular signaling labels attached to the target DNA. The associated
sensing system detects the presence of light emitting labels on sensor
sites where hybridization has taken place by means of an optical scanner.
Some innovative microarrays employ electrochemical labels resulting in
an electrical current through sensor electrodes during readout in case of
hybridization events ([65, 66, 19]).
Label-free techniques offer significant advantages in terms of costs,
since they avoid the expensive reagents and pretreatment steps required
to attach labels. Recently, a number of approaches have been proposed
based on mass changes [24, 67, 68, 69]) or electrical properties of electrode/solution interfaces induced by DNA hybridization ([22, 16, 23, 70]).
Our approach is based on interface capacitance measurements. The
sensing principle can be summarized as follows: DNA hybridization leads
to a capacitance decrease at the electrode-to-solution interface due to
the replacement of electrolytic solution with high dielectric constant by
organic DNA oligonucleotides with low dielectric constant. This phenomenon has been studied extensively in the past, using a conventional
electrochemical three electrodes system, by applying a 50 mV potentiostatic voltage step and measuring the electrode current [47]. In [29], biofunctionalized gold electrodes were characterized with impedance spectroscopy.
Since the capacitive measurement principle is not based on reduction
or oxidation processes, two-electrodes configurations, as that used in this
work, are sufficient. This simplifies the measurement set-up and on chip
implementation. As far as integrable capacitance methods are concerned,
it has already been proven that DNA hybridization can be detected by
measuring interface capacitance with a system making use of two electrodes only, i.e. without the use of a reference terminal [40, 39]. Furthermore, more recently it has been shown that reliable measurements can be
done exploiting microfabricated electrodes [71]. In this case, however,
capacitance parameters were measured by means of external instruments
and off-chip circuits, since electrodes were deposited on a passive silicon
In comparison to passive solutions, where each transducer element
has to be connected to the readout device individually, the use of active
CMOS sensor chips is required in applications where a large number of
analysis has to be performed in parallel. Then, not only the sensitivity
of the single sensor sites is increased, but also the electrical interconnect
to the outside world is reduced, e.g. to a five pin serial interface [76].
In general, detection principles based on electrical or electrochemical signals are very sensitive to variations of the electrolyte or interface
properties and compared to optical methods as DNA detection. These
measurements required differential measurements and this also has implications on the circuit design of CMOS sensor arrays, such that different
kinds of reference sensors and related signal processing should be used
for improving the signal to background ratio of the active sensors.
Along this line of development, this work represents a further advance
in the state of the art in that it presents a silicon chip featuring an array
of sensor electrode pairs for capacitive measurement and integrated measurement circuits monolithically integrated in a standard CMOS chip.
Capacitance-based DNA Detection Principle
Under proper electrochemical conditions, bio-modified metal interfaces in
a saline solution exhibit an almost ideal capacitive behavior. This is the
case if gold electrodes are used modified with short DNA strands immobilized with alkanethiol ([39, 29]), in which the capacitance value of the
electrode-solution interface has been estimated between 1 and 20µF/cm2 ,
even if this value strongly depends on electrode surface treatment and
roughness. The electrode-solution interface can be modeled by the equivalent circuit depicted in Fig. 7.1 (left) ([42, 73]), where RS depends on
the interface and on the solution characteristics and RP is related to the
insulating properties of the interface. For dense layers RP is very high
and can be considered negligible; CP is mainly affected by the physical
and chemical characteristics of the insulating bio-layer immobilized on
the surface.
In our set-up, the dipole formed by two bio-modified electrodes in
solution exhibits an equivalent capacitance whose value is given by the
series of the two interface-capacitances, one for each electrode, in parallel
with the geometrical capacitance formed by the two electrodes. The lat-
ter is several orders of magnitude smaller than the interface-capacitance,
hence its contribution is negligible. As mentioned above, it has been
observed that when a complementary DNA strands bind with the surface probes, CP decreases [47, 39, 40]. When the DNA duplex is formed,
the solution counterions attracted to the polarized metal surface are displaced [53]. This increases the distance between the charge inside the
electrode and the ions in the electrolyte, thus decreasing the interface
capacitance (Fig. 7.1, right).
Among the many techniques available to measure capacitances, we
have opted for a technique based on the conversion of the capacitance
value to a frequency value.
The measurement technique used in the mixed-signal circuit is based
on a very simple principle, as shown in Fig. 7.2. A periodic current excitation, IREF , is provided to the electrodes. The electrodes respond to the
current pulses by changing their voltage difference in a transient waveform whose time constant is dominated by the capacitive component of
the electrode-solution impedance. The inter-electrode potential is monitored for crossings of two fixed reference values, +VREF and −VREF .
The crossings of +VREF and −VREF produces a square waveform at the
output of the comparator whose frequency is proportional to the rate of
change of the voltage. Considering RS negligible, since saline solution is
0, 3 M NaCl, the crossing frequency follows the following relation
= 2RC ln
where R and C represent the contribution of the capacitance,CP , and
parallel resistance, RP , of both electrode/solution interface. Moreover,
if the frequency is not too low (i.e. IREF is not too small) the first order
Taylor approximation of the logarithm returns the following equation:
The frequency is measured by means of a counter which is enabled to
count the edges of the square waveform for a fixed time. The data stored
in the counter can be read and the interface capacitance value of each
electrode can be computed.
One major advantage of this technique is that the digital read-out of
the frequency is almost trivial. In fact, an accurate estimate of the frequency value is obtained by counting the number of reference crossings in
a given time interval. If the interval is long enough, very good frequency
resolution can be obtained, assuming that the frequency is constant over
the measurement interval. The circuit does not directly measure capacitance, but converts transient time into frequency of reference crossings.
Label-free DNA chip
In this section, we describe the architecture and the detailed operation of
the chip, based upon the capacitance measurement technique described
in the previous section.
Chip architecture
The chip architecture and signal flow are summarized in Fig. 7.3. The
chip interface requires both analog and digital I/O signals. Voltage references are analog DC signals, and they determine the voltage ranges (VREF
eq. 7.1 and eq. 7.2) used for measurements. Electrodes are selected using digital addressing lines A0 − A6. Control logic includes reset and
clock signal. The output is fully digital, as the chip performs internally
capacitance measurement and analog-to-digital conversion. Each sensing site features an analog part which converts the impedance value into
frequency and outputs the data digitally. Thus, 128 capacitance measurements and analog-to-digital conversions can be performed in parallel, and
the results are then multiplexed on the shared output using the address
signals which are generated on-chip by means of two decoders (Fig. 7.3).
At the output pad, the data are processed by a PC to calculate the
capacitance value of the interface for each pixel.
Sensing site circuitry
The mixed-signal circuit implemented for each sensing site is illustrated
in Fig. 7.4. It includes an analog part (current source, comparator and
switches) and digital part (22 bit counter). The current pulses required
for the circuit operation are provided by a current source based on a
cascode stage followed by a pair of switches that alternatively connect
the mirror to the electrodes in a current push or pull mode. The pMOS source is part of an in-sensor site p-MOS current mirror biased by
an n-MOS current source. The n-MOS current sources bias voltages are
generated by an external circuit realized on a PCB (see section 7.5.1). By
adjusting the signal VM IRROR it is possible to change the reference current
used for the measurements, thereby increasing the range of capacitance
values that can be measured. In our case VCASC is set to 2 V to guarantee
that n-MOS transistors work in the saturation region. If we neglect
transistor mismatches and parasitic, IIN is equal to IOU T and to IREF .
The switches are implemented with parallel n- and p-channel transistors
in order to have full voltage range.
Comparison of the electrode difference voltage with the reference difference voltage level is performed by a high-gain differential CMOS stage.
The output of the gain stage is then buffered by a simple CMOS inverter,
thereby becoming a digital signal. Then number of oscillations within a
given time is then counted by a digital 22-bit counter clocked. After
a user-settable time period, the counter is stopped, and the final count
value can be transferred to the output via a shift-register. The shift
register is clocked from an external source resulting in a serial output of
the counter’s data bits at the chip’s output. The counter/shift register
circuit is similar to that described in [19]. As illustrated in Fig. 7.5, a
”reset” signal is activated before each measurement to set all the counters to zero. Then, a ”count” signal is set at high value for a fixed time,
called integration time, enabling the counter. Finally, a clock signal is
used to read the data stored in the shift register of the counter.
The described circuit guarantees that each pixel oscillates continuously at a frequency determined by the interface capacitance. It is important to note, that with this free-running oscillator concept, the data
from all sites are sampled simultaneously.
Physical layout
The chip is fabricated in 6” n-well 0.5 µm CMOS process with three metal
layers, the oxide thickness is 15 nm and supply voltage is 5 V . The gold
electrodes are deposited after standard CMOS processing. After the gold
deposition, an annealing step is introduced applying N2 /H2 at 350◦C for
30 minutes in order to guarantee sufficiently low values of the interface
state density at the silicon/silicon dioxide interface [64]. Figure 7.6 shows
a tilted SEM image of the interdigitated gold electrodes demonstrating
the good quality of the deposition process. The sensor sites consist of
interdigitated electrodes with 1.2 µm line width and spacing, the diameter of the circular arrangement is 200 µm. The chip provides an
8 × 16 array of these sensors and the pitch is 250 µm. Total chip is
6.4 mm × 4.5 mm (7.8, bottom-left). The electrical interface includes
analog as well as digital signals. Two different power supply and ground
pads are also implemented on chip. A shield connected to ground is
introduced to reduce noise between analog and digital circuits (Fig. 7.7).
Experimental results
Measurement set-up
The chip has been tested in a laboratory setup as shown in Fig 7.8 and
Fig. 7.9. The chip is attached onto a PCB, and electrically contacted
using gold bonding wires. Before performing a DNA experiment, the
gold electrodes are accurately cleaned with oxygen plasma in order to
facilitate uniform covalent binding of the thiolated DNA probes to the
gold surface. A two-chambers fluidic cell protects both bonded wires and
the chip’s I/O pads from contact with saline solutions used during the
experiments (Fig. 7.8). The different agents used, for instance, to rinse
the electrodes after hybridization or to feed the biological sample to the
sensor surface, are injected in the two chambers by means of single use
In our electrical measurements, the DC voltages necessary to perform
measurements are generated by an external circuit implemented on the
PCB. The output and digital input data from the counters are handled by
means of a National Instruments DAQ board (PCI-6534E [61]). A LabView software records the frequency data of the sensor sites, calculates
the capacitance values and writes the data to a file (Fig 7.9).
Electrical Characterization
The electrical behavior of our chip is tested using discrete precision capacitors (error less than 1%) within a range from 330 pF to 10 nF . In
Fig. 7.10, measured results are shown for three different reference currents. As expected a linear behavior is obtained, the slope is 0.9837 and
the error bars are negligible.
The influence of parallel resistive components is shown in Fig. 7.11.
There, measured frequencies of four test circuits are plotted as a function of IREF . Solid lines correspond to purely capacitive load, 1 nF and
4.7 nF , respectively, dotted lines to capacitances in parallel with resistances, 1 nF with 200 kΩ and 4.7 nF with 200 kΩ, respectively. The
plot clearly shows that a parasitic parallel resistor RP does not affect the
function of the integrated sensors. Moreover, the parasitic capacitance
can be determined precisely, by evaluating the oscillation frequencies for
different reference currents. If the application allows the use of large reference currents and hence high oscillation frequencies, the effect of the
parasitic resistor can be neglected. The curve converges asymptotically
to the ideal capacitive line for large reference currents (7.11).
Moreover, we have performed measurements for a fixed capacitance,
1 nF and three different RP : 100 kΩ, 200 kΩ and 680 kΩ in order to
characterize the circuit under realistic non ideal conditions. Fig. 7.12
shows measurement results in a frequency vs reference current plot. We
can observe that only at low frequencies and low parallel resistance values,
lower than 680 kΩ, the parallel resistance influences the measurement
leaving ideal linear behavior.
The resistance per unit surface value of RP using gold electrode modified with self assembled monolayer (SAM) in solution is approximated
with 15 kΩ cm2 [29]. Since the area of our electrodes is approximately
0, 03 mm2 , we expect a parallel resistance of around 50 MΩ. This verifies
the condition of approximately ideal capacitive behavior for our sensors.
All the DNA hybridization measurements have been performed with IREF
is 1 µA and VREF is 200 mV .
DNA detection
DNA detection is demonstrated by comparing measurements on electrodes in the two isolated chambers exposed to the same solution. In
one chamber only bare electrodes are present (reference pixels) while in
the second one DNA probes are immobilized on the gold surface (functionalized electrodes) (step 1). The use of reference pixels is important
for correct measurements. The gold surface of the reference pixels is not
functionalized with probe molecules, hence, DNA strands on the sample
cannot bind to the surface. These electrodes provide the background signal, which includes all phenomena occurring at the electrodes/solution
interface, with the exception of DNA hybridization. Then, solutions
containing DNA molecules, non complementary (step 2) and complementary (step 3) to the immobilized probe molecules, are injected in the
two chambers and measurements are performed after each step.
Frequency measurements, which takes typically around 1 s, have been
performed for 8 minutes for each step in order to obtain a stable value.
The capacitance value is calculated by means of equation 7.2 from the
average value of the last 2 minutes of measurements. Since capacitance
and frequency are inversely proportional, standard deviation, σ, can be
calculated as follows:
Each step is analyzed by taking the difference signal between the
gold electrodes (reference pixels) and the functionalized ones which, after
the first step, are subjected to the same treatment. Typical trend of
measurement is shown in Fig. 7.13 where the transients and the final
stable values are clearly shown. Drifts in interface impedance have been
studied in depth in the literature. Cyclic voltammetry has shown that
the transient behavior is due to complex ions double-layer structures
occurring also in bare electrodes at the electrode/solution interfaces and
evidenced by the polarization [29]. Moreover, a decreasing trend in time
is also present along the different steps measurements. Decreasing trends
like that shown in Fig. 7.13 are typical in DNA detection, and they are
clearly observed also with other detection technologies [77].
In Fig. 7.14 typical results for different pixels are shown. The three
columns indicate the difference between bare gold and the functionalized
electrodes for each step. Important is to note that the behavior of the
single electrodes are coherent. In particular, the measured differential
capacitance in the case of non-specific binding is only slightly reduced as
compared to the value observed at functionalized electrodes before exposure to the sample liquid. This means that the small amount of nonspecific binding of target molecules to the DNA probes or non-specific
adhesion of organic molecules to the sensor’s surface is negligible. In fact,
the direct comparison of the average differential capacitances and their
standard deviations of the step 1 and the step 2 shows that the detection
provide the same value, in terms of statistical significance. On the other
hand, a direct comparison of step 1 and step 3 shows that the complementary DNA binding is detected with statistical significance. Moreover,
little discrepancies of the single detection by different pixels but referred
of the same step are, again, indistinguishable within a statistical range.
The average behavior of the pixels is shown in Fig. 7.15. It confirms
the ability of the chip to reliably detect the hybridization process on the
pixel’s sensor electrode. The standard deviation of these values reflects
statistical variations in the electrical behavior of the different pixels on
one chip but also statistical variations of the electrode’s functionalization
SAM quality. However, the measurement results clearly show, that specific detection of label free DNA target molecules can be performed with
statistical significance. Moreover, similar average values and error bars
overlapping of the non-specific binding (step 2) compared to only probe
state (step 1) demonstrates that the system is capable of rejecting false
positives due to non-specific deposition of sample DNA on the electrode
Figure 7.1: Left: Electrical metal/solution interface model. Right: DNA
hybridization process and displacement of counterions.
Figure 7.2: Measurement principle: interface capacitance determines the
frequency of the electrodes charging and discharging transients. A comparator compares the inter-electrode potential with a reference voltage
VREF producing a digital signal at its output whose frequency is inversely
proportional to capacitance.
Figure 7.3: Schematic representation of the system and signal flow. A
and D indicate analog and digital signals, respectively.
Cascode stage
Switching point: |VA-VB|= Vref
22 bit
shift out
- Count
- Reset
- Clock
Figure 7.4: Schematic of the circuit associated to each sensing element.
The current source is implemented by a current mirror circuit. The
comparator features two differential input stages and an high-gain output
stage. Finally a 22-bit counter and shift register samples and stores
measurement data.
Figure 7.5: Signal flow of counter input/output.
Figure 7.6: SEM image of intedigitated gold electrodes. Inset: zoom of
single finger electrode.
Figure 7.7: Pinout of the chip.
Output PCB pads
Bonding wires
Chip is glued on a PCB
Figure 7.8: Photo of the PCB used to contact pads with the glued chip,
bonding wires (top) and the applied fluidic cell (bottom-right). The cell
determines two separated areas on the chip (bottom-left) which can be
functionalized with different probes.
PC, National DAQ,
• Analog reference signals
• Power supply
• Ground
• Addressing
• Digital Control Signals
Graphic plot:
128 pixel
(f D 1/C)
Figure 7.9: Schematic representation of the measurement set-up. Voltage
reference signals and power supply are generated by circuitry on the PCB.
Digital control signals are provided by a PC. The LabView interface
manages all the parameters involved in the measurements and shows
directly on the screen the measurement results of the whole array.
Meas_CAP (nF)
Nom_CAP (nF)
Figure 7.10: Measured capacitance vs nominal capacitance.
frequency (kHz)
4.7 nF
4.7 nF & 200 kȍ
1 nF
1 nF & 200 kȍ
current (µA)
Figure 7.11: Frequency vs reference current showing the influence of
the parallel resistor for two different capacitance values. A non linear
behavior is evidenced in low frequencies regime.
frequency (kHz)
1 nF
1 nF & 200 kȍ
1 nF & 680 kȍ
1 nF & 100 kȍ
current (µA)
Figure 7.12: Frequency vs reference current showing that a significant
influence of the parallel resistance on the measurement result occurs only
at low current values and at RP values lower than 680 kΩ.
Average of functionalized gold electrodes
8 minutes of
Average of gold electrodes
frequency (kHz)
(probes vs gold)
(a-specific reaction)
specific reaction
30 minutes without measurements
to wait biological reaction
Time (min)
Figure 7.13: Frequency changes of the average of reference electrodes
(continuous line), and the average of functionalized electrodes (dashed
line) show a larger gap after DNA hybridization step considering the
stable value reached at the end of the transient.
gold electrode-functionalized electrode
gold electrode-functionalized electrode after aspecific reaction
gold electrode-functionalized electrode after specific reaction
ǻC (nF)
Pixel 1
Pixel 2
Pixel 3
Figure 7.14: Typical variations for several pixels among functionalized
electrodes and the average value of reference gold electrodes. Capability
to distinguish between specific and a-specific binding is shown for each
Gold electrode - Functionalized electrode
Gold electrode - Functionalized electrode after
aspecific binding
ǻC (nF)
Gold electrode - Functionalized electrode after
specific binding
Figure 7.15: The average behavior of all the pixels confirms that a-specific
and specific binding are distinguishable.
Chapter 8
Application on tumor marker
and future steps on DNA
Tumor marker analysis
Early diagnosis of cancer is crucial for the successful treatment of the
disease. Highly sensitive methods are urgently needed for measuring
cancer diagnosis markers present at ultra-low levels during early stages
of the disease. Such methods should facilitate early detection and an adequate selection of the treatment of diseases and should lead to increased
patient survival rates. Existing diagnostic tests (e.g., ELISA) are not
sensitive enough and detect proteins at levels corresponding to advanced
stages of the disease. Smaller, faster, and cheaper (one-step) devices are
highly desired for replacing time-consuming laboratory-analysis. Making
analytical results available at patient bedside within few minutes will
greatly improve the monitoring of cancer progress and patient therapy.
Advances in molecular biology have led to a much understanding of potential biomarkers that can be used for cancer diagnosis. The realization
of point-of-care cancer diagnostics thus requires proper attention to the
major challenge of multi-target detection. Arrays of biosensors, detecting protein signature patterns or multiple DNA mutations, can be used
to help screening and guide treatment. Innovative biosensor strategies
would allow cancer testing to be performed more rapidly, inexpensively,
and reliably in a decentralized setting. In this review article I will discuss
the use of electrochemical biosensors for decentralized clinical testing and
the prospects and challenges of using such devices for point-of-care cancer
Abnormal concentrations of certain proteins can indicate the presence of
various cancers. For the past two decades Heineman’s group has developed highly sensitive enzyme electrochemical immunoassays [78]. Such
protocols rely on labeling of the antibody (or antigen) with an enzyme
which acts on a substrate and generate an electroactive product that
can be detected amperometrically. Enzyme immunosensors can employ
competitive or sandwich modes of operation. In addition to enzyme labels, it is possible to use metal markers and redox tags for electronic
transduction of antigenantibody interactions.
The development of electrochemical immunosensors present the main
aim of reduce costs and increase the preventive analysis in tumor diseases. Our approach is similar to what described in previous chapters of
this thesis and is based on capacitance measurement. In particular, we
started from CBCM technique based on commercial devices, as already
described, and gold electrode.
Device and methods
The measurement system is exactly the same described in chapter 5
changing the I/V conversion resistance to 4.7 kΩ since gold electrode
are changed.In fact, the gold electrode used for the following experiment
were made by Olivetti InkJet in Arnad (Italy). Fig. 8.1 shows the layout
of the chip where it is ease to note 5 different sensing sites addressable
by means of 4 pads. From chemical point of view binding of antibody
on gold surface has been solved previously binding a layer of mercaptoundecanoic acid on gold. This layer expose a carboxylic group which
allow after activation with EDC (1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride) and NHS (N-Hydrohysuccinimide)(Fig. 8.2), to
bind the amino group of the antibody (Fig. 8.3). After that incubation
for 4 hours with antigen (SCCA-specific for liver cancer) has been performed. Buffer solution used in all experiment is PBS (Phosfate Saline
Buffer). A second way to functionalized the surface exploiting 3-glycol
chain on the top of molecules before carboxylic group was tested in order
to avoid aspecific binding on surface as demonstrated in [79]. Moreover
this last surface modification has been used also to improve stability of
capacitance measurement for DNA detection. In this case we have used
the same DNA oligonucleotide already described in this thesis but probe
molecules where modify with amino group instead that thiol group in
order to exploit the same protocol used for binding antibody.
Experimental results
Typical measurements in time after the first way of funzitionalization are
shown in Fig. 8.4. A standard deviation of 9.2% is observed. This behavior reduce the possibility to increase the sensitivity of our system in
particular for protein where the aspected capacitance variation could be
smaller because of a more distance respect to the gold electrode. Comparing this measurements whith the ones sown in Fig. 8.5 that are obtained with the second way of gold modification described in the previuos
paragraph, it is ease to observed that stability of measurement in time
increase of one order of magnitude. In fact, standard deviation in this
secon case is 1%. Therefore, all the following results are obtained using
the more stable functionalization process. In order to test the capability of our system to detect Ab/Ag reaction without aspecific signal,
we have tested the system on which where already bound Ab molecules
with BSA (Bovine Serum Albumin) and afterward we add the specific
molecules (SCCA). Results are shown in Fig. 8.6 where you can observed
a small decrease after BSA incubation while a largest increase after incubation with SCCA. Small values of standard deviation allow to clearly
and statistically distinguish capacitance variation. Moreover, also on
DNA detection the previous treatment with 3-glycol molecules increase
the capability to distinguish where hybridization reaction occurs compared to aspecific binding. Following the same protocol of Ab binding on
3-glycol molecules and the hybridization protocol described in the previous chapters, Fig. 8.7 shows measurements results on 4 different sensing
sites of the chip. An increase in capacitance value is observed over all
the chip after probe binding (probably due to reorganization of the SAM
layer). Finally, where hybridization occurs a decrease is observed (site 1
and 2) while where aspecific binding (site 3) and only buffer solution
(site 4), a small increase and a negligible variation respectively, is observed. This preliminary experiment encourage to proceed on this way
of stabilization of capacitance measurement in order to increase reproducibility and sensitivity of our detection device also for genetic analysis.
Observing the graphs differences among sites still have to be explained
and improve working on surface chemistry but the trend and low standard
deviation of measurement demonstrate the possibility of this technique
to improve detection of biological reaction.
Figure 8.1: Photograph of the sensing chip.
Figure 8.2: Activation of carboxylic group before antibody binding.
Figure 8.3: Antibody binding on gold surface modified with mercaptoundecanoic acid.
capacitance [nF]
time [min]
Figure 8.4: Typical measurement in time of mercaptoundecanoic acid.
capacitance [nF]
time [min]
Figure 8.5: Typical measurements in time of 3-glycol modified gold surface.
capacitance [nf]
Figure 8.6: Aspecific binding signal compared with specific SCCA reaction.
Capacitance [F]
Figure 8.7: Results on DNA detection exploiting 3-glycol previous layer
on 4 different sites: step1 is after 3G functionalization; step2 is after
probe binding; step3 is after specific hybridization in sites 1 and 2 while
is aspecific in site 3 and only buffer were added on site 4.
Chapter 9
EEPROM memory as DNA
In the last decade, genetic research has started to make large use of miniaturized devices (in general known as microarrays), consisting of glass
or quartz slides featuring a two-dimensional array of small sites, each
aimed at recognizing the presence of a specific base sequence within the
unknown DNA ”target” molecules to be analyzed/ recognized. To this
purpose each site is ”functionalized” with the immobilization of known
single-stranded base sequences, called probes, able to selectively bind (hybridize) with the complementary ones possibly present within the DNA
targets. In operation, all sites are simultaneously exposed to the material to be analyzed and the probes essentially work as ”selective glue”
for the target molecules. Thus, the recognition of the presence of specific
sequences within the DNA targets is transformed in the identification
of the sites where probe- target hybridization has taken place. Some of
these devices, featuring densities up to hundreds of thousands sites [4],
are able to perform tests on a whole genome scale. Moreover, low or
medium density arrays (featuring from ten [80, 81, 82] to forty thousand sites [83]) are sufficient for analysis of a restricted number of genes.
For the devices currently used to-day, the DNA strands to be identified
(hereafter referred to as targets) are preliminary marked with fluorescent
molecules and, when they hybridize with specific probes and are consequently immobilized in the corresponding sites of the array, their presence
is revealed by means of an optical scanner or a fluorescence microscope.
However, the high cost of the scanner, the unreliability due to the use of
markers (also called labels), and the complex data processing procedures
needed to extract useful information from these arrays pose critical limits
to the widespread use of these tools. In the attempt to overcome these
drawbacks alternative detection techniques are intensely investigated. In
this paper, we present a new approach based on measurements of UV
absorbance without the use of any labels. This technique is particularly
interesting because UV absorption by polynucleotide molecules is highly
specific and commonly employed as a tool in standard spectrophotometers [84, 85]. However, until very recently [86], it has not been used as
a basis of microfabricated DNA sensors/detectors. At system level, our
approach envisages an array of ”sensing” sites, each consisting of a biolayer of specific probes, positioned between an external UV source and
a UV detector. The bio-layer is conventional, in that it is obtained by
means of DNA probes, aligned with the underlying sensor and working as
”selective glue” with respect to the DNA target molecules. In operation,
all sites are simultaneously exposed to the targets molecules, that will
selectively hybridize only with complementary probes. Since DNA has a
significant UV absorption, the sites where hybridization has taken place
will present different UV absorption with respect to those where the same
amount of DNA is in the non hybridized state. The radiation transmitted
through such sites will be smaller in the first case, due to masking effect
of the hybridized DNA. As already explained, the recognition of this difference implies that of complementary molecules. A previous work [86]
presented an implementation of the concept described above consisting
of bio-functionalized quartz slides, UV sensors fabricated in amorphous
silicon and read-out electronics realized in the form of a PCB with standard components. The result is a low cost device, suitable for low-density
arrays and aimed at point-of-care applications. In this work, instead, we
present a complementary implementation of the same approach, aimed
at high performance devices, exploiting all the advantages of silicon integration. In particular, both the UV sensors and the electronics needed
to address and read the individual sites could be integrated on the same
chip and the resulting devices would be extremely compact, fast and
suitable for high-density arrays of sensing sites.
Devices and method
The key point of this work is the use of NV memory cells as UV sensors.
As known, these cells can be erased by means of UV radiation, particularly if designed so that their Floating Gate (FG) can be directly exposed
to the incoming radiation. Since during erasing the cell threshold voltage shift (∆VT H = VT H0 − VT H , where VT H and VT H0 denote the actual
and initial value of the threshold voltage, respectively) increases with the
UV dose, NV cells represent almost ideal UV dosimeters that can be exploited for the purpose of this work. Therefore, the DNA chip envisaged
in this work features an array of suitable NV memory cells, all initially
programmed at the same threshold voltage VT H0 (¿¿ VT HN where VT HN
is the cell threshold voltage with no charge on the FG). Each cell (or
group of cells) is located below a specific bio-functionalized layer. After
exposure to the target DNA the whole array is exposed to UV radiation
for a fixed time. Consequently all cells are partially erased, but those
covered by hybridized targets receive less radiation, hence exhibit a final
higher value of VT H . (i.e a smaller ∆VT H ), compared with those where
non DNA binding has occurred. Of course, all the electronics needed to
individually select and read the cells in the array are essentially the same
as in (multi-level) memories, hence can be integrated on the same chip
as the cells. Moreover, since memory cells can be very small, a number
of them can be used for each sensing sites, thus allowing the possible implementation of useful checking algorithms. In this context, the target of
this work is the characterization of NV memory cells as DNA detectors, in
order to show that they are suitable for the task. For this purpose, liquid
DNA solutions in quartz containers will be ”superimposed” to memory
cells (fabricated ad hoc to be easily accessible and measurable). In this
way, not only a possible implementation of DNA sensors (featuring microfabricated wells as liquid containers) is directly explored, but DNA
concentrations can be controlled with great accuracy. The DNA ”sensing site” used in this work is shown in Fig. 9.1. A quartz container with
the DNA targets is placed between a UV lamp and the NV memory cell
used as a dosimeter for the impinging UV radiation. A mechanical shutter allows to switch on and off the incident UV radiation. The memory
cell is initially programmed at a fixed value (VT H0 ¿¿ VT HN ) of threshold voltage. Then, the UV radiation is switched on and is transmitted
through the DNA solution. Since the UV dose reaching the memory cell
depends on the absorption of the DNA solution (namely on DNA concentration and whether or not DNA is in the double or single strand state),
measuring ∆VT H after a fixed time provides information about the fact
that probe- target hybridization has or has not occurred. As for recognition of specific DNA sequences, the experiments of this work exploit
the so called hypochromic effect, namely the fact that for the same number of DNA bases (nucleotides), the UV absorption of double-stranded
molecules is smaller (by about 20%) than that of the single stranded form.
Of course, for the purpose of this work a major question concerns the
sensitivity of NV cells used as UV detectors and, from this point of view,
a suitable device has to be used. From this point of view, unfortunately,
standard double poly Flash cells are not particularly suitable because,
when exposed to UV, their FG is masked by the polysilicon Control Gate
(CG), that absorbs most of the radiation. For this reason, ”single-poly
EEPROM devices” have been chosen. As known, in this technology, the
CG is realized by means of a n+ diffusion under the FG, that extends
outside the channel area (Fig. 9.1). This feature has some very important benefits, namely: a) the FG is completely exposed to U.V. radiation
(i.e. no ”masking” CG is present); 2) the FG area is larger than normal (thus more UV radiation can be collected); 3) the cell can be easily
fabricated by CMOS technology. For the experiments of this work, test
chips has been designed and fabricated with 0.25 µm CMOS technology
at STMicroelectronics (Milan, Italy). The chips contains single memory
cells that can be directly contacted by means of suitable pins. The area
of the cells is approximately 20 µm2 but could easily be made larger to
improve the UV sensitivity.
Measurement setup
The experimental set-up used in this work is schematically represented
in Fig. 9.2. The source of UV radiation is a Xenon lamp characterized by
a spectral distribution with high emission values in the range of interest
(250 to 270 nm). The lamp filament is controlled by a special circuit
which maintains the supplied power constant. However, since this does
not guarantee radiation stability over long time periods, a specific circuit
has been used to measure the instantaneous radiation and compute the
total dose. This circuit is based on a photodiode (whose spectral sensitivity is particularly high in the range of interest) whose output current
is amplified by an I/V converter and sampled by means of an acquisition
board equipped with 16-bit converters. The photodiode is placed directly
after the UV filter needed to select the radiation range of interest and
before the quartz container: thus the UV dose is independent of the DNA
solutions placed after. The sample data are processed by a LabVIEW
program in order to evaluate the radiation dose of the experiments. The
same program controls the shutter and switches off the radiation once
the dose has reached a desired value. The EEPROM VT H is measured,
when the light source is disabled, by means of a HP4156 Semiconductor
Parameter Analyzer (SPA) driven by the control PC. VT H is conven-
tionally defined as the CG voltage needed to obtain a drain current of
2 µA when the drain is driven to 1 V and source and bulk are grounded.
Resolution in VT H measurements is 1 mV . A LabVIEW program automatically controls the whole measurement. The operator sets the step
dose and the number of steps. The controller then starts a measurement
step consisting of: i) a first determination of VTH, ii) the application
of a step dose; iii) a subsequent re-determination of VT H . This step is
repeated a number of times selected by the user. After the last step the
operator manually resets the cell to the initial state (by means of an over
program operation followed by short UV erase steps to set the VT H to
VT H0 = 5.5 V within 1 mV error). The whole measurement is repeated
5 times for each DNA solution to evaluate the standard deviation, hence
the measurement reliability. Resulting data present very low standard
deviation, thus no error bars have been plotted.
Experimental results
results described in this Section are shown as a function of UV dose
(a parameter equivalent to, but more appropriate than exposure time).
Fig. 9.3 shows the typical behavior of ∆VT H as a function of the UV
dose and, as can be seen, a difference is measured for different values of
single-stranded DNA concentration. Fig. 9.4 shows the curves obtained
for the same concentration of DNA but after an annealing step allowing
the molecules to hybridize. The Fig. 9.3 and Fig. 9.4 show that the
difference between different samples increase with UV exposure; however
in the long run all cells would be completely erased, regardless of DNA
concentration. Thus a question is in order about the optimum UV dose
for maximum sensitivity. To investigate this problem, we introduce a
new parameter denoted as Under Erasure (U.E.) defined as
U.E. = ∆VT H−BF F ER − ∆VT H−DN A
where ∆VT H−BF F ER and ∆VT H−DN A represent ∆VT H measured in the
case where only the buffer or the buffer containing DNA is interposed between the UV source and EEPROM cells, respectively (in this definition,
of course, ∆VT H−BU F F ER merely represents a convenient common reference). The results indicate that an optimum dose is about 300 (in A.U.)
and this value is found to be essentially independent of DNA concentration (while it may probably depend on the EEPROM cell technology
and geometry). This value allows to obtain maximum measurement sensitivity. Taking this result into account, in our experiments UV dose of
225 is used to obtain a good trade-off between measurement signal and
time exposure. As for DNA recognition, relevant results are illustrated
in Fig. 9.6, where the experimental points are drawn in such a way that
each value of concentration on the x axis corresponds to the same number
of single-stranded DNA bases (i.e. the elements actually absorbing UV
radiation), so that the difference among the various curves is due only to
whether or not the DNA is in the bound (i.e. hybridized) state. In particular, the points indicated with squares and diamonds represents the
case where the same number (N) of DNA single strands are presented
in double and single-stranded form, respectively. As can be seen, the
measurements allow to easily recognize the case of hybridization, that is
the key for the DNA sensor envisaged in this work.
The results presented before clearly indicate that a DNA microarray
based on the approach suggested in this work and realized in a single
chip with standard CMOS technology seems possible, since (at least)
single-poly EEPROM cells represent a suitable UV detectors. Compared
with the microarrays in use to-day, such a microarray would benefit of
all advantages of integration. In particular: a) sophisticated electronics
for site addressing and electrical read-out could be easily integrated on
the same chip (in fact, these features are essentially the same as those
of Multi-Level NV memories); b) a large number of small cells could be
available, and this would allow to use a number of them for each sensing
sites (to implement checking schemes for error detection and correction;
c) sensors would be cost-effective (for mass production); d) devices would
be highly reliable, and could be easily made user-friendly e) no markers
are needed for the analysis In the case of fully integrated sensors, small
containers (wells) for DNA could be fabricated directly on top of the
IC passivating glass. Such containers would be much smaller than those
used in this work. However, the relevant parameter is the number of
DNA molecules along the optical path within the container, in practice
the average number of molecules along a vertical path from the liquid
upper level to the container bottom. This number decreases linearly
with the vertical dimension of the container, however the amount of
molecules along such a path can be easily brought back to an adequate
level increasing the DNA concentration. At this regard, the results of
Fig. 9.6 indicates that double and single-stranded molecules in 1 cm
path length can be distinguished starting from a minimum concentration
of 1.5 µM bases Finally, it is worth mentioning that the ratio of the
slopes of the lower to upper curve of Fig. 9.6 is 0.63, thus indicating
that hybridization leads to a 37% decrease in UV absorption, in fair
agreement with published data [87] on the hypochromic effect exploited
in the experiments.
Figure 9.1: DNA detection principle.
Figure 9.2: Schematic representaion of the experimental set-up.
Buffer solution TE 1x (1,5 ml)
DNA - 900 nM single stranded
DNA - 1950 nM single stranded
DNA - 3600 nM single stranded
DNA - 5400 nM single stranded
ǻVth [V]
Figure 9.3: Variation of EEPROM cell threshold voltage shift as a function of the UV dose for the case of single-stranded (i.e. non hybridized)
Buffer solution TE 1x (1,5 ml)
DNA - 825 nM double stranded
DNA - 1800 nM double stranded
DNA - 3300 nM double stranded
DNA - 4950 nM double stranded
ǻVth [V]
Figure 9.4: Variation of EEPROM cell threshold voltage shift as a function of the UV dose for the case of double-stranded (i.e. hybridized)
ǻvth [mV]
Dose [A.U.]
Figure 9.5: Typical behavior of the Under Erasure (U.E.) as a function
of the UV dose (the DNA concentration here is 100 nM).
ǻVth for DNA single stranded
ǻVth for DNA double stranded
U.E. [mV]
Base concentracion [nM]
Figure 9.6: U.E. for the case of hybridized and non- hybridized DNA,
lower and upper curves, respectively.
Chapter 10
Conclusions and perspectives
During PhD I have studied the possibility to detect by mean of fully
electrical capacitance measurement DNA hybridization reaction on metal
surface. I started from literauture where evidence of this detection were
reported [40] but results were obtained with a laboratory setup. Therefore, I follow two ways in order to do a step in advance: i) low-cost board
based on commercial electronics componenent (e.g. µ-controller, ADC,
DAC,...); ii) integration of metal electrodes and electronic circuits for
measurement on the same chip. Results show that both ways are able to
detect DNA hybridization by means of capacitance measurements, therefore, different application can be satisfy by mean of different electronics.
For example, the first approach is more useful for a low parallell sites
detection and specific gene analysis (e.g. one or few viruses which determine a particular group of patologies), while the second one can be
addressed to thousands parallell analysis exploiting microfabrication process. Problems related to stable and reproducible measurements are still
open since detection reported in this thesis are obtained by mean of a differential approach, but as shown in chapter 8 new molecules can be used
to modified metal surface and stabilization of capacitance measurement
has been observed showing higher reproducility of capacitance measurement. Elegant research on new sensing concepts has opened the door
to a widespread clinical applications of electrochemical devices. Such
devices are extremely useful for delivering the diagnostic information in
a fast, simple, and low cost fashion, and are thus uniquely qualified for
meeting the demands of point-of-care cancer screening. The high sensitivity of modern electrochemical bioaffinity assays should facilitate early
detection and treatment of diseases and should lead to increased patient survival rates. The attractive properties of electrochemical devices
are thus extremely promising for improving the efficiency of diagnostic
testing and therapy monitoring, and for point-of-care cancer testing, in
general. The main challenge is to bring electrochemical techniques to the
patient’s side for use by non-laboratory personnel without compromising accuracy and reliability. The realization of decentralized electronic
testing of cancer would thus require additional extensive developmental
work. Special attention should be given to non-specific adsorption issues
that commonly control the detection limits of electrochemical bioaffinity
assays. It is expected that the creativity of electrochemists and material
scientists, coupled with proper resources, will revolutionize cancer diagnostics in a manner analogous to their current leading role in diabetes
monitoring. Disposable cartridges, containing electrode strips (coated
with numerous receptors) along with related sample processing, could
thus offer early screening of cancer in a point-of-care setting, by measuring abnormalities in protein profiles within few minutes. Accordingly,
there is no doubt that electrochemical biosensors will become a powerful
tool for cancer diagnostics in the near future.
Chapter 11
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