Development of Aluminum Oxide (Al2O3) Gate Dielectric
Protein Biosensor under Physiologic Buffer
THESIS
Presented in Partial Fulfillment of the Requirements for the Degree Master of Science in
the Graduate School of the Ohio State University
By
Fang Ren
Graduate Program in Electrical and Computer Engineering
The Ohio State University
2012
Master's Examination Committee:
Prof. Paul R. Berger, Advisor
Prof. George Valco
Copyright by
Fang Ren
2012
Abstract
Aluminum Oxide (Al2O3) is a high k dielectric material with promising biosensor
applications. The key feature of the Al2O3 device that allows its stable operation in high
salt buffers is the impermeability of the device to mobile buffer ions. Permeation of such
mobile buffer ions into traditional silicon-based device results in electrical instability of
the device of magnitudes sufficient to interfere with analyte sensing.
This thesis focuses on Al2O3 high k dielectric devices which could work in the
physiologic buffer solution (PBS). A low cost Al2O3 MOS capacitor was first fabricated
and tested as a tractable model for metal-oxide-semiconductor field effect transistor
(MOSFET). The MOS capacitors were dipped in sterile PBS solution for increasing
intervals of time starting from 30 mins upto 24 hours. Triangular voltage sweep (TVS)
method was used to characterize the Na+ ion penetration. No sodium ion (Na+)
penetration was observed for the Al2O3 capacitors. By contrast, the dose of Na+ ion
penetration into silicon-dioxide MOS capacitor increased with increasing soak time in the
PBS solution. Further, no Na+ ion response was observed for varying Al2O3 thickness of
10nm, 25nm, 50nm, 100nm.
A low cost Si based bioFET with 20nm high-k Al2O3 dielectric deposited by ALD
was fabricated and tested in the PBS solution later. After fully functionalizing the surface
with aminopropyldimethylethoxysilane (APDMES), the Al2O3 gate dielectric bioFET is
capable to work under PBS solution. The drain to source current decreased after detecting
ii
the streptavidin. This could be the mobile charge in the channel decreased after binding
of the protein. However, the decreased changes are very small and cannot be used as an
effective biosensor. To some extent, these changes could be noise because the common
source I-V measurements were done in solution which increases lots of unstable factors.
iii
Dedication
This document is dedicated to my family.
iv
Acknowledgments
I would like to thank my advisor, Prof. Paul Berger. During one and half years, Prof.
Berger gave me lots of insight guidance, effort and instruction on my research, as well as
financial support, without which it could never possible to finish this work.
I would also like to thank Prof. Valco who provided me lots of help in my curriculum and
research, and served as my committee member. I appreciate all the help of Mark Brenner,
without his constant assistance furnace part, I could not fabricate my device smoothly.
I want to express my gratitude to my group members: Anisha Ramesh., Minjae Kim,
Shang Wei, Tyler Growden, Ying Ding, Prof. Xiaona Li. Especially, I want to say thank
you to Anisha Ramesh for her generously shared her experiences with me so that I can
perform experiments along quickly and smoothly. I would also want to thank Patricia
Casal, and Andy Theiss for protein treatment.
Thank you to all the staffs in the Nanotechwest cleanrooms for maintaining the facilities,
especially to Derek Ditmer, Daniel Gallego-Perez, Paul D.Steffen.
Finally, I want to thank my husband Tianyou Kou. I cannot thank you enough for all the
support and love you have given me. I also want to say thank you to my parents, thank
v
you for being there and always supporting me. I would not be the person I am today
without your constant patience, love, and support.
vi
Vita
July, 2007 ................................................B.S. Applied Physics, Inner Mongolia University
July, 2010 ............................................M.E. Integrated Circuit Design, Beihang University
March, 2012 ....................Graduate Research Associate, ECE, The Ohio State University.
.
Publications
Fang Ren, Paul R. Berger et al., "Affordable in vivo Biosensors for Clinical Protein
Detection". IMR Meeting, Columbus, OH, Sep 2011.
Anisha Ramesh, Fang Ren, Paul R. Berger et al., Al2O3 Gate dielectric with ion
impermeability for in-vivo biosensor ", 2011 MRS Fall Meeting, Boston, MA, Nov. 28Dec 2, 2011.
Fang Ren and Jinming Dong, "Fast and Efficient Intra Mode Selection for H.264/AVC",
ICCMS2010 IEEE.
Fields of Study
Major Field: Electrical and Computer Engineering
vii
Table of Contents
Abstract ............................................................................................................................... ii
Dedication .......................................................................................................................... iv
Acknowledgments............................................................................................................... v
Vita.................................................................................................................................... vii
List of Tables ...................................................................................................................... x
List of Figures .................................................................................................................... xi
Chapter 1: Introduction ...................................................................................................... 1
1.1 Motivation and overview ....................................................................................... 1
1.2 Application of Biosensors ..................................................................................... 2
1.3 Classification of Biosensors .................................................................................. 4
1.4 Protein biosensor ................................................................................................... 7
1.5 Thesis Outline ........................................................................................................ 8
Chapter 2: Theoretical Model of BioFET ........................................................................... 9
2.1 Streptavidin-Biotin .............................................................................................. 10
2.2 Field effect transistor ........................................................................................... 11
2.3 Liquid water energy bands .................................................................................. 14
viii
2.4 BioFET Protein Sensor ........................................................................................ 16
Chapter 3: MOS Capacitor Study .................................................................................... 20
3.1 MOS Capacitor fabrication.................................................................................. 20
3.2 MOS Capacitor Ion Permeation Study ................................................................ 23
3.3 Result and Discussions ........................................................................................ 31
Chapter 4: BioField Effect Transistor Study ................................................................... 32
4.1 BioFET fabrication .............................................................................................. 32
4.2 Al2O3 BioFET Testing ........................................................................................ 50
4.3 SiO2 BioFET Testing .......................................................................................... 60
4.4 Result and Discussions ........................................................................................ 63
Chapter 5: Summary and further directions ..................................................................... 65
5.1 Summary.............................................................................................................. 65
5.2 Further Directions ................................................................................................ 66
References ......................................................................................................................... 69
ix
List of Tables
Table 1 Relationship between ∆ [Na+] and Time ............................................................ 27
x
List of Figures
Figure 1 Monomeric Streptavidin (displayed as ribbon diagram) with bound ................. 10
Figure 2 Average gate to channel potential for n-type FET ............................................. 12
Figure 3 I-V characteristics of ideal nMOSFET transistor ............................................... 14
Figure 4 Distribution of the electron state density of hydrated redox particles when
................................................................................................................. 16
Figure 5 A field effect transistor protein biosensor is presented. ..................................... 16
Figure 6 Typical structure of a MOS capacitor formed on a p-type Si substrate used in
this study. .......................................................................................................................... 20
Figure 7 Triangular voltage ramp used for detecting alkali ions ...................................... 24
Figure 8 Illustration of the peak obtained with low frequency C-V measurements due to
mobile ions and calculation of ion concentration ............................................................. 25
Figure 9 Triangular voltage sweep measurements of a typical thermal SiO2 MOS
capacitor at 250oC ............................................................................................................. 26
Figure 10 Relationship between increases in alkali ions concentration with increasing
soak times in PBS for the control SiO2 based capacitor. The line are joins of the measured
data. ................................................................................................................................... 27
Figure 11 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 100 nm oxide thicknesses. ........................................................... 28
xi
Figure 12 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 100 nm oxide thicknesses. ........................................................... 29
Figure 13 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 25 nm oxide thicknesses. ............................................................. 30
Figure 14 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 10 nm oxide thicknesses. ............................................................. 30
Figure 15 Oxide thickness versus oxidation time for silicon in H2O at 640 Torr[84] ...... 33
Figure 16 A p-type silicon wafer with oxide layer. .......................................................... 34
Figure 17 A schematic of phosphorus diffusion process. ................................................. 35
Figure 18 A schematic of drive-in diffusion. .................................................................... 36
Figure 19 A microscope image of Si substrate after phosphorus diffusion and drive-in
process. W:L=10:1 L=25µm ............................................................................................. 37
Figure 20 A microscope image of Si substrate with Al2O3 gate grown on phosphorus
diffusion area. W:L=10:1.L=25µm ................................................................................... 39
Figure 21 Oxide thickness as a function of time for dry oxidation of bare (100) silicon at
different temperature [85]. ................................................................................................ 40
Figure 22 A microscope image of Si substrate with SiO2 gate grown on phosphorus
diffusion area. W:L=10:1 25µm........................................................................................ 41
Figure 23 Mask level of FET with different gate electrodes: gate electrodes were
designed with (a) holes, (b) solid metal, (c) without metal and (d) slots. Gate W:L=10:1
........................................................................................................................................... 43
xii
Figure 24 Mask level of thick photoresist SPR-220 as reservoir (Green areas) for the
measurement solution above the senor surface. Gate W:L=10:1 ..................................... 44
Figure 25 Al2O3 gate BioFET with perforated gate metal (holes). Additionally the
reservoir is shown surrounding the gate test area. W:L 10:1 L=25
............................ 45
Figure 26 BioFET control device with solid metal gate electrodes and reservoir W:L 10:1
L=25
............................................................................................................................ 46
Figure 27 BioFET floating gate and reservoir. W:L 10:1 L=25
................................. 46
Figure 28 BioFET with slotted gate electrodes and reservoir. W:L 10:1 L=25
.......... 47
Figure 29 The surface silanization procedure for streptavidin immobilization atop on the
Al2O3 layer ........................................................................................................................ 48
Figure 30 The surface silanization procedure for streptavidin immobilization atop on the
SiO2 layer .......................................................................................................................... 48
Figure 31 Diagram of the chemical linking of the streptavidin protein binding to the
Al2O3 surface ..................................................................................................................... 49
Figure 32 Diagram of the chemical linking of streptavidin protein binding to the SiO2
surface ............................................................................................................................... 49
Figure 33 Schematic of cross-sectional view of an Al2O3 FET biosensor test under PBS
solution set up. .................................................................................................................. 50
Figure 34 Common-source I-V curve of Al2O3 gate FET with solid gate Al electrode.
W:L 10:1 L=25
............................................................................................................ 51
Figure 35 Common-source I-V of Al2O3 gate FET with slots gate Al electrode. W:L
10:1 L=25
.................................................................................................................... 51
xiii
Figure 36 Common-source I-V of Al2O3 gate FET with holes gate Al electrode. W:L
10:1 L=25
.................................................................................................................... 52
Figure 37 Common-source I-V of Al2O3 gate FET without gate Al electrode before apply
PBS. W:L 10:1 L=1000
............................................................................................... 52
Figure 38 Common-source I-V of Al2O3 gate FET without gate Al electrode after apply
PBS. (PBS was in the reservoir and a bias of 1-5 V was applied to gate area) W:L 10:1
L=1000
........................................................................................................................ 53
Figure 39 Id- VG curve of Al2O3 gate FET with solid gate Al electrode. (
) ...... 54
Figure 40 Square root Id- VG curve of Al2O3 gate FET with solid gate Al electrode.
(
) ....................................................................................................................... 55
Figure 41 Id- VG curve of Al2O3 gate FET with holes gate Al electrode (
) ...... 56
Figure 42 Square root Id- VG curve of Al2O3 gate FET with holes gate Al electrode.
(
) ....................................................................................................................... 56
Figure 43 Id- VG curve of Al2O3 gate FET with slots gate Al electrode (
) ....... 57
Figure 44 Square root Id- VG curve of Al2O3 gate FET with slots gate Al electrode.
(
) ....................................................................................................................... 58
Figure 45 Common-source I-V of Al2O3 gate FET before and after binding 50µg/ml
streptavidin in reservoir with PBS solution and a bias of 5 V was applied to gate area
W:L 10:1 L=1000
........................................................................................................ 60
Figure 46 Common-source I-V curve of SiO2 gate FET with solid gate Al electrode and a
bias of 5 V was applied to gate area. W:L 10:1 L=25
................................................. 61
Figure 47 Diode I-V measurement of drain of SiO2 FET. ............................................... 62
xiv
Figure 48 Diode I-V measurement of source of SiO2 FET. .............................................. 62
Figure 49 A field effect transistor Si incorporated with high-k dielectric stack layer
protein biosensor is presented (Multi-color layers in the gate imply different oxides
stacked together). .............................................................................................................. 66
xv
Chapter 1: Introduction
1.1 Motivation and overview
A biosensor is a device which could measure and convert a biological response into
an observed signal. It consists of three parts. The first one is a sensitive biological
element, for example biological material. The second one is a transducer or the detector
element which transforms a signal generated from the interaction of the analyte with the
biological element into another signal for further measurement and quantification. The
third part is a signal processor which is used to display results in a user-friendly way.
With the development of biosensor materials, precise detector platform and faster
hardware presently (for signal processing and display) breakthroughs make it possible to
achieve a high sensitivity of biosensor device. A motive of this thesis is to develop a
novel protein biosensor device with an affordable detector topology for protein detection.
The transducer platform is very important for a biosensor device because its intrinsic
operational properties, such as size, bio-electrical signal conversion form, and cost limit
the application environment. Another critical part is the interface between the transducer
platform and the biological agents (referred as bio-receptors hereon). Although different
interfaces between the transducer platform and bio-receptors have resulted in many kinds
of devices and techniques, the basic idea is to make bio-receptors immobile on the
1
surface of the transducer, such as an immunosensor assay [1] and electrophysiological
receptor [2].
There is a wide variety of structural and design features that determine operational
characteristics of biosensors. Bio-receptors primarily determine the selectivity, sensitivity,
and stability of the biosensor. Excellent bio-receptors are not only sensitive to analyte but
are also stable to heat and, light, and are able to be immobilized on surfaces. Sensor
materials are another critical part of biosensor. High-k materials are largely popular as
piezoelectric, electrochemical, optical, magnetic biosensors based on their unique nature
operation. The important role of high-k materials in biosensors is expounded. Lately,
there are many applications of novel functional high-k dielectric materials with biosensor,
such as Al2O3 [3], HfO2 [4] , AlGaN/GaN [5]. Research indicates that high-k dielectric
materials can provide tremendous value in enhancing performance of biosensors, such as
sensitivity, detection range, and repetitiveness [6].
1.2 Application of Biosensors
Biosensors have developed to be a new interdisciplinary frontier including chemistry,
biology, medical science and electronics dating back to the development of the first
glucose biosensor in 1967 [7]. Due to their simplicity, high sensitivity and potential
ability for real-time and on-site analysis, biosensors have been applied widely in various
fields including clinical diagnosis, environment monitoring and food control.
2
A. Clinical Diagnosis
Biosensors are very promising for clinical diagnosis, health care and other medical
applications, such as emergency-room screening, bedside monitoring, home self-testing
and ‘alternative-site’ testing (e.g the physician’s office) due to their highly accuracy and
sensitivity [8]. One successful example is glucose biosensors based on amperometric
enzyme electrodes [9-11]. The success of glucose blood meters has stimulated a number
of interests in in-vitro and in-vivo devices for monitoring other physiologically important
based new materials (membranes etc.) and concepts. Clinical diagnosis biosensors have
shown considerable economic prospects and fascinating research opportunities which
will benefit a wide variety of sensing applications.
B. Environmental Monitoring
The strong demand for precise and fast detection sensors for environment monitoring
has brought lots of opportunities for biosensor technologies development. The disposable
and self-contained integrated character of an environmental monitoring biosensor makes
them surpass conventional bioanalytical systems, which need additional processing steps,
such as a reagent addition [12]. For example, Choi and his research group [13] fabricated
a portable biosensor based on optical-fibers. The results show that the toxicity of a
sample could be detected by measuring the bioluminescence 30 min after addition to the
freeze-dried strains. This kind of device can be used for field sample analysis and the
monitoring of various water systems on-site.
3
C. Food Control
Biosensors can also be used as food quality control tools. It is estimated that more
than 36 million cases of illness occur annually because of food-and waterbome pathogens
resulting in more than 5000 deaths a year [14]. Many researches have been done in this
area, for examples, Mukhopadhyay and Gooneratne developed novel biosensor for
inspection of meat in a noninvasive and nondestructive way [15].
1.3 Classification of Biosensors
In the past two decades, research and development of biosensors and their
applications has undergone remarkable growth [16] because of their excellent prospects
for interfacing with biological specimens. Biosensors have been applied to different
target applications, including glucose [17], pH [18, 19], protein [20-23], and DNA [2426] detection and measurement. Basically, there are four types of biosensors: bioaffinity
biosensor, catalytic biosensor, optical biosensors and whole-cell biosensors.
A. Bioaffinity Biosensor
Affinity-based biosensors are analytical devices composed of biological recognition
elements such as an antibody, receptor protein, biomimetic material, or DNA interfaced
to a signal transducer, which together relates the concentration of an analyte to a
measurable electronic signal [27, 28]. The selectivity, sensitivity, and stability of
biosensor are mainly depending on bioaffinity elements. Affinity based biosensor have
been used in protein detection [29], antibody biological recognition [30] and
environmental monitoring [31]. The interface of signal transductors and bioaffinity
4
materials are critical to affinity-based biosensor. Usually, the technologies of bioaffinity
assays are very complex and require multiple components [32].
B. Catalytic Biosensor
Catalytic biosensors are kinetic devices that measure steady-state concentration of a
transducer-detectable species formed/lost due to a biocatalytic reaction. A sensor might
be described as a catalytic biosensor if its recognition element was comprised of an
enzyme or series of enzymes, a living tissue slice (vegetable or animal), or whole cells
derived from microorganisms such as bacteria, fungi, or yeast. Catalytic biosensors can
monitor rate of product formation, disappearance of a reactant and the inhibition of a
reaction. There are many kinds of biocatalyst biosensors, such as enzymes [33] and
microorganisms [34]. The working temperature is one of the limitation conditions for
catalytic biosensor. For example, a catalyst also plays a part in catalyzing side reaction,
especially at high temperature.
C. Optical Biosensor
Optical biosensors also have great potential in biomedical research, health care, and
environmental monitoring. Optical devices including laser diodes, LEDs and optical
fibers play increasingly important role in medical research. Generally, there are two
detection protocols that can be implemented in optical biosensors: fluorescence-based
detection and label-free detection. Fluorescence spectroscopy has become one of the
more useful bioanalytical and diagnostic tools in the past 30 years. In fluorescence-based
detection, either target molecules or biorecognition molecules are labeled with
fluorescent tags, such as dyes; the intensity of the fluorescence indicates the presence of
5
the target molecules and the interaction strength between the target and biorecognition
molecules [35]. The label-free techniques measure an inherent property of the query itself
(e.g. mass or dielectric property) thereby avoiding modifying interactors [36]. One of
problems of optical biosensor is that the fluorescent signal may not be strong enough and
cause in accurate sensing results.
D. Whole-cell Biosensor
The whole-cell applications of biosensors have been limited to very few specific
examples, such as protein, pH, and glucose measurement. Whole cell based biosensors
may extend the field of biosensor applications due to its biological recognition
characteristics. Recently, several researchers demonstrated that whole cell based
biosensors have potential applications to the field of drug discovery, clinical diagnostics
and for monitoring toxins, chemicals, heavy metals and so on [37]. There are several
advantages of using whole cell based biosensors. First, “group effects” such as toxicity,
mutagenicity, or pharmacological activity become accessible to measurements using
sensor technology. In addition, internal amplification cascades can be used to increase the
sensitivity of the device. Moreover, whole cells are the smallest biological entity which is
self-sustaining. If non-neural cells are used, the sensors are inexpensive because the
preparation costs of culture growth are low [38]. Two of the disadvantages of whole-cell
biosensor are the lack of genetic stability and short lifetime.
6
1.4 Protein biosensor
Proteins are biochemical compounds which are defined as naturally-occurring,
unbranched polymers in which the monomer units are amino acids [39]. Proteins can
also work together to achieve a particular function, and they often associate to form
stable protein complexes which are very important parts of organisms because they
participate in almost every process in cells.
Proteins perform many functions, such as catalysis, immune recognition, cell
adhesion, signal transduction, transport, movement and cell organization. Protein
biosensors are of particular interest and importance in modern medicine for their role in
the early detection and diagnosis of disease, for instance cancer
[40-43]. Different
approaches for protein biosensors based on different semiconductor materials have been
explored, such as Si [44-47], AlGaN/GaN [48-50], carbon nanotubes [51, 52], grapheme
[53, 54], etc. Compared to the alternative material platforms, Si-based protein biosensors
are low-cost and envisioned to be easily integratable onto a small chip atop a diagnostic
needle complete with readout circuitry. However, Si-based protein biosensors suffer from
long-term electrical drift and instability due to the diffusion of ions from high osmolarity
biological buffers into the gate oxides.
Long-term stable and low-cost Si-based in vivo protein biosensors are needed but
their translation to in vivo sensing applications introduces big challenges. Apart from
specific device designs, great attention should be paid to the physiological environment
and its interactions with the field effect transistor (FET). The typical in vivo physiological
7
environment contains Na+ and K+ ions that can be incorporated into the dielectric oxide
of the FET and contribute to mobile charge [55-57]. These mobile ions are more
deleterious than fixed charges due to gate oxide defects or interface charges, since the
mobile ions shift within the active device depending upon voltage, causing a variable
drift in the transistor threshold voltage, resulting in inaccurate protein sensing. Hence, a
key feature needed for in vivo biosensors is impermeability to mobile alkali ions with
stable transistor operation.
This thesis demonstrates the usage of aluminum oxide (Al2O3) as a replacement to
the conventional thermally grown silicon oxide (SiO2) as a viable alternative to obtain
stable transistor operation in physiological solutions allowing the future realization of
low-cost Si-based biosensors.
1.5 Thesis Outline
This thesis focuses on high-k dielectric biosensor for streptavidin detection. The
organization of the thesis is as follows. Chapter 2 focuses on the theoretical model of a
bioFET including the strong streptavidin-biotin bond, field effect transistor theory model,
and the working function of the bioFET. This chapter analyses the feasibility of a bioFET.
Chapter 3 demonstrates the fabrication process of the MOS capacitor and bioFET uesd in
this research. The fabrication processes leverage conventional silicon MOSFET processes.
This chapter documents all the fabrication processes and presents microscope images of
the most important aspects of the process flow. Chapter 4 shows the test methods and
results. Chapter 5 concludes the results and proposes new high-k materials as the gate
dielectric for bioFET for the further study.
8
Chapter 2: Theoretical Model of BioFET
There has been a variety of research reports documenting the detectional proteins
which are central to biological processes. Generally, there are optical [58-61],
spectrometric [62-64], electrochemical [65-68], and surface plasmon resonance (SPR)
measurements [69-71] as the method to analyze biological molecules. However, all these
methods involve time consuming, and multi-stage processes that are expensive and
unsuitable for real time monitoring. The field-effect transistor (FET) is a transistor that
relies on an electric field to control the conductivity of a channel of one type of charge
carrier in a semiconductor material. Field effect transistors are widely used in everyday
electronics. Biosensors using the field effect mimics a conventional FET but with the
controlling gate electrode replaced by a biological solution. Similar to the conventional
FET, the gate dielectric layer is crucial in controlling the sensing behavior (i.e. gain), and
stability and reliability are central requirements for the gate dielectric. BioFET based
biosensors are a promising field of high-sensitivity, label-free and easy to fabricate
sensors [48, 72, 73].
9
2.1 Streptavidin-Biotin
Figure 1 Monomeric Streptavidin (displayed as ribbon diagram) with bound
The crystal structure of streptavidin with biotin bound was first solved in 1989 by
Hendrickson and his research group [74] (Figure. 1). The high affinity of the noncovalent
interaction between biotin and streptavidin forms the basis for many diagnostic assays
that require the formation of an irreversible and specific linkage between biological
macromolecules. Among the most common uses of streptavidin-biotin is the purification,
or detection, of various proteins. The strong streptavidin-biotin bond can be used to
attach various biomolecules to one another, or onto a solid support. Harsh conditions are
needed to break the streptavidin-biotin interaction, which often denatures the protein of
interest being purified. However, it has been shown that a short incubation in water above
70°C will reversibly break the interaction without denaturing streptavidin, allowing re10
use of the streptavidin solid support [75]. The strong affinity between these two
molecules, and their high degree of characterization, make it an ideal detection couple for
discovery of new bioFET platforms.
2.2 Field effect transistor
There have been various research reports into detecting proteins which are central
to biological processes. Generally, there are optical [58-61], spectrometric [62-64],
electrochemical [65-68], and surface plasmon resonance (SPR) measurements [69-71] as
the method to analyze biological molecules. However, all these methods involve time
consuming, and multi-stage processes that are expensive and unsuitable for real time in
vivo monitoring. The field effect transistor (FET) is a transistor that relies on an electric
field to control the shape and hence the conductivity of a channel of one type of charge
carrier in a semiconductor material. Field effect transistors are widely used in everyday
electronics. Biosensors using the field effect are mimics of conventional FET with the
controlling gate replaced by a biological solution. Similar to the conventional FET, the
gate dielectric layer is crucial in controlling the sensing behavior (i.e. gain), and stability
and reliability are central requirements for the gate dielectric. BioFET based biosensors
are a promising field of high-sensitivity, label-free and easy to fabricate sensors [48, 72,
73].
Figure 2 shows and n- channel field effect transistor. A FET has three modes of
operation [76-78]. There are the cutoff or subthreshold region, Linear or nonsaturation
region and saturation region
11
Gate
Drain
Source
+
N+
-
+
=
Channel
L
+
N+
P-type body
Figure 2 Average gate to channel potential for n-type FET
In the cutoff region, the gate to source voltage is less than the threshold voltage
(
). There is no electron inversion layer therefore in the channel and almost zero
current flows from drain to source. When
,
the gate attracts
electrons to form a channel and the FET works in the linear region. The electrons move
from source to drain at a speed proportional to the electric field in these regions. The
charge in the channel
Where
can be modeled in the follow equation.
is the capacitance of the gate to the channel.
The average voltage of the channel is
The gate can be modeled as a parallel plate capacitor.
12
. Therefore
where
is the permittivity
for SiO2 ,
permittivity of free space.
Where v is the average velocity ,
for Al2O3.
is the
F/cm.
is the mobility constant, and E is electric field.
is the current between source and drain. This equation describes the linear
region. When
,
, FET will work in a saturation region.
Increasing the drain voltage will not increase the current because the channel is no longer
inverted in the vicinity of the drain. The performance of a FET in saturation region can be
modeled in the following equations.
To sum up, an FET works in three regions and each of three regions can be
modeled by the following equations.
{
Figure 3 shows an example I-V curve of ideal nMOSFET.
13
Figure 3 I-V characteristics of ideal nMOSFET transistor
2.3 Liquid water energy bands
Since the protein biosensor is designed to work in a physiologic buffer solution, the
interface between the liquid and semiconductor as well as the Fermi level of hydrated
redox electrons are very important to the device performance. The electron energy bands
of hydrogen oxide are very similar to metal oxides but the band edges are indefinite due
to its amorphous structure [79]. There are two types of reaction in the liquid solutions,
oxidation and reduction.
Where RED is the reductant, OX is the oxidant, and
is the redox electron.
The Fermi level of hydrated redox electrons can be modeled by the following
equations.
14
√
√
where
and
is the electron state density in the donor and acceptor bands of
hydrated redox particles, respectively. where
and
densities of electron energy fluctuation. where
and
are the probability
are the concentrations and
where is electrons and ions fluctuation energy in liquid solution.
The sum of
and
is the total state density
.
=
When
, the level
is called the Fermi level of the redox
electron which can be derived into the following equation [80].
=
where
+kTln
=
is the standard Fermi level. Figure 4 shows a schematic distribution of
the electron state density of hydrated redox particles when
15
.
OX
RED
W( )
Figure 4 Distribution of the electron state density of hydrated redox particles when
2.4 BioFET Protein Sensor
Figure 5 A field effect transistor protein biosensor is presented.
16
Figure 5 shows a sensing channel connects the source (S) and drain (D) with a reference
electrode (RE). When the streptavidin protein binds to the receptor, it induces charges in
the substrate (electrons as pictured here), causing a change in the current flow between
the source and drain. Different kinds of FETs are made from many materials, such as Si,
SiGe, GaAs, InGaAs, and so on. No matter what type materials were applied, the basic
principle of the FET is similar. Conventional FETs consists of an active channel through
which electrons (or holes) flow from the source to the drain [81]. The conductivity of the
channel is modulated by a potential applied to the gate which results in the modulation of
the charge density flowing in the channel.
Additional voltage is applied between the drain and source electrodes which results
in a current flow through the modified channel now with its voltage induced conductivity,
thereby exhibiting gain in the drain current from the small gate voltage applied. For a
bioFET protein sensor, siliane linker layer and streptavidin layer are deposited on the gate
instead of a metal. The gate is not directly biased by an external voltage source but is
floating and in direct contact with the solution being tested. The gate area is then
modified by immobilized molecular receptors attached to the functionalized gate oxide
for the binding of charged analytes (protein to be detected). The surface result in the
charge induced in the channel, which manifests as a change in the output drain current.
Since a gate metal is absent, a voltage can be applied to the electrolyte through a
reference electrode (RE) to shift the baseline transistor bias condition towards maximize
transistor gain. When the gate-source voltage (VGS) is greater than the drain-source
17
voltage (VDS) the transistor operates in the linear region and the drain current-voltage
relationship is given by:
(
)
As the drain-source voltage is increased and exceeds VGS-Vt , the device enters
saturation and the drain current-voltage relation is given by:
where µ is the electron/hole mobility, Cox is the specific capacitance given by
, W and L are the width and length of the gate, is the oxide permittivity,
is oxide
thickness and Vt is the threshold voltage.
The threshold voltage is the minimum gate voltage to turn on the transistor and is
given by:
where
is the flat-band voltage. The flat-band voltage is defined as the applied gate
voltage such that there is no band bending in the semiconductor,
controlled by the doping density,
substrate doping concentration.
is a potential energy
is the semiconductor permittivity, and
is the
is the fixed oxide charge introduced in the oxide
during growth and is constant for a device.
is the mobile ion charge, which is of
primary concern in biosensor operation due to its manifestation as sensor drift. It is clear
that changes in
results in changes in device threshold voltage and hence output
18
current of the device. This will conflict with changes due to adsorbed protein analyte and
result in erroneous sensor operation.
19
Chapter 3: MOS Capacitor Study
This chapter demonstrates the usage of aluminum oxide (Al2O3) as a replacement to
the conventional thermally grown silicon oxide (SiO2) as a viable alternative to obtain
stable transistor operation in physiological solutions allowing the future realization of
low-cost Si-based biosensors.
3.1 MOS Capacitor fabrication
A tractable model for a metal-oxide-semiconductor field effect transistor (MOSFET),
is a simple MOS capacitor that can be effectively used to determine the presence of
mobile ions, such as sodium (Na+) ions, in the oxide. The typical structure of a MOS
capacitor is as shown in Fig. 6.
Figure 6 Typical structure of a MOS capacitor formed on a p-type Si substrate used in
this study.
20
3.1.1 Device Fabrication
In this study the traditional thermally grown silicon oxide (SiO2) or aluminum oxide
(Al2O3) deposited by the atomic layer deposition (ALD) process were used for the
dielectric oxide layer.
ALD is a layer-by-layer deposition method relying on self-limiting surface reactions
to obtain atomic layer control of deposition. An advantage of ALD is precise thickness
control at the Ångstrom or monolayer level. The self-limiting aspect of ALD leads to
excellent step coverage and conformal deposition on high aspect ratio structures [82].
The Al2O3 deposition was carried out with trimethylaluminum (TMA) and water as
the precursors deposited at 300oC using a Picosun SunaleTM reactor. The substrates used
are moderately doped (~ 1016 cm-3) p-type silicon wafers.
Prior to deposition, the wafers were cleaned using the standard clean process
consisting of RCA1 (1NH4OH: 1H2O2: 5 de-ionized (DI) H2O at 70oC for 10 minutes)
and RCA2 (1HCl: 1 H2O2: 5 DI H2O at 70oC for 10 minutes). This was followed by a 1
minute dip in 1HF:10 DI and a 1 minute DI H2O rinse. The ALD pulsing sequence for
one cycle was 0.1 sec. per TMA pulse, 4 sec. per N2 purge, 0.1 sec. per H2O pulse, and 4
sec. per N2 purge. Typical ALD deposition rates of 0.8 Å/cycle were obtained. Aluminum
metal was deposited on the topside and patterned by photolithography and lift-off to
obtain square electrodes with various areas of 275×275, 550×550, 1100×1100,
1650×1650 and 2200×2200 µm2. The square electrodes were designed additionally with
holes and slots to permit various levels of ion permeation and a control electrode was
21
included with no holes. Finally aluminum metal was deposited on the backside of the
wafer followed by a post-metallization anneal at 450oC for 10 min. in nitrogen ambient.
3.1.2 Oxide Characterization
For the comparative study between ALD Al2O3 and thermal SiO2, a target thickness
of 100 nm was chosen and followed by a 20 minute nitrogen anneal at 700℃. . As-grown
Al2O3 was measured to be 103 nm.
Thermally grown silicon oxide (SiO2) was used as the control sample. The sample
was prepared using the same p-doped substrate and wafer cleaning procedure as
described above for ALD Al2O3. Silicon dioxide was grown in an atmospheric tube
furnace at 1050oC with a dry oxygen ambient followed by a 20 minute nitrogen anneal at
the same temperature.
Reducing the oxide thickness further increases the capacitance and hence the
sensitivity of a potential biosensor. As described in section 2, the MOSFET channel
current is directly proportional to the oxide capacitance, Cox. Where
Thus,
increasing ε (using high-k dielectrics such as Al2O3) while concurrently reducing the
oxide thickness, tox, provides a large sensitivity boost, critical for biosensing applications.
MOS capacitors using Al2O3 as their dielectric and with reduced thicknesses were
obtained by repeating the ALD process and reducing the number of cycles to obtain
samples with target oxide thicknesses of 50, 25 and 10 nm, in addition to the 100 nm
sample. The measured oxide thickness values using ellipsometry were 52, 30 and 12 nm,
respectively.
22
3.2 MOS Capacitor Ion Permeation Study
The in vivo physiological environment can be simulated by conducting
experiments in physiological buffer solutions (pH 7.4, 0.15M Na+,K+). Natural in vivo
protein environments contain comparable concentrations of alkali ions at a similar pH.
Hence, impermeability of ions or immunity of transistor electrical response to these
environments serves as a viable proof of applicability of Si-based FET sensors.
3.2.1 Triangular Voltage Sweep Technique
Mobile ion contamination in semiconductor technology is generally performed on a
MOS capacitor using C-V bias temperature stressing (BTS) or the triangular voltage
sweep (TVS) technique. The TVS method has a higher sensitivity than BTS and is hence
the technique chosen in this work. TVS also leads to reduced measurement times since
measurements are done at high temperature, unlike BTS which requires sequential
heating, stress and cooling prior to the C-V measurement. Permeation of mobile charges
into the oxide can be quantified using the TVS method. The TVS technique is based upon
measuring the charge flow through the oxide at an elevated temperature in response to an
applied time-varying voltage [83]. The MOS sample is heated to a temperature (~250oC)
where the mobile ions have sufficient thermal energy, and thus mobility, to respond to an
applied bias. The MOS capacitor is stressed for 5 minutes at a voltage that generates
about 1 MV/cm electric field across the oxide. This moves all the mobile ions to the
capacitor plate charged with the opposite polarity. A triangular voltage ramp (Fig. 7) is
subsequently applied to the gate of the capacitor.
23
Voltage
Stress
+V
Voltage
Ramp
A+
0
Stress
-V
A+
Time
Voltage
Slope
Figure 7 Triangular voltage ramp used for detecting alkali ions
The ramp frequency should be slow enough so that the ions can drift through the
oxide. Hence, a quasi-static capacitance-voltage C-V measurement is performed. This
generates a displacement current in the capacitor. As the voltage crosses from positive to
negative or negative to positive, a peak in the measured capacitance is observed. The
capacitor is next stressed at an opposite polarity bias and a reverse voltage sweep is
applied. The capacitance is obtained by measuring the charge flow (ΔQ) through the
oxide when a time varying voltage is applied (ΔV) given by ΔQ/ΔV. The peaks in the
two sweep directions may not be identical since the ions are at different interfaces (metaloxide, oxide-semiconductor) after stressing at two different polarities.
Next, a high frequency C-V measurement is performed, where the ions do not have
sufficient time to respond, and no significant peak due to mobile ions is observed. Using
24
this as the baseline, the area between these two curves (high frequeny and low frequency)
is determined by integration to obtain the mobile ion charge density within the oxide (Fig.
8). Finally, MOS capacitors with ALD Al2O3 and thermal SiO2 gate dielectrics were
soaked in PBS solution for varying amounts of time and subsequently measured by the
TVS technique.
Figure 8 Illustration of the peak obtained with low frequency C-V measurements due to
mobile ions and calculation of ion concentration
25
3.2.2 Thermal Silicon Oxide versus ALD Aluminum Oxide
0.085
2
Capacitance Density (F/cm )
0.080
PBS0Min
PBS30Min
PBS60Min
PBS90Min
0.075
0.070
0.065
0.060
0.055
0.050
0.045
0.040
0.035
0.030
0.025
0.020
-4
-2
0
2
4
Voltage (V)
Figure 9 Triangular voltage sweep measurements of a typical thermal SiO2 MOS
capacitor at 250oC
Figure 9 shows the result of Triangular Voltage Sweep measurements for a typical
100nm SiO2 MOS capacitor at 250℃ and 1M Hz. The delay time is 5 min and the sweep
direction is from the positive bias to negative bias. After dipping in sterile physiological
buffer solution (PBS, 150 mM Na+) for 30 min, 60 min, 90 min, penetration by Na+ ions
was revealed for the 100nm SiO2 based MOS capacitors. The area between each curve
quantifies the increased mobile charge (Na+).
The Na+ density is calculated by the following equation:
26
Table 1 Relationship between ∆ [Na+] and Time
Time (min)
0
30
60
90
∆[Alkali ions]
0
1.77
3.69
10.87
(× 1010cm-2)
10
[Alkali ions] 10 cm
-2
12
10
Thermal 100nm SiO2
8
6
4
2
30
40
50
60
70
80
90
100
Time (min)
Figure 10 Relationship between increases in alkali ions concentration with increasing
soak times in PBS for the control SiO2 based capacitor. The line are joins of the measured
data.
Table 1 and Figure 10 demonstrate the increase in alkali ion penetration into SiO2
MOS capacitors with increasing soak times in PBS solution.
Figure 11 shows the result of the triangular voltage sweep measurements for the Si
based 100nm Al2O3 MOS capacitor sensor at 250℃ and 1 MHz. The delay time is 5
27
minutes. The sweep direction is from the positive bias to negative bias and then a return
to positive bias. The Al2O3 MOS capacitor was dipped in sterile PBS solution for
increasing intervals of time starting from 30 minutes up to 24 hours. No measurable
sodium ion (Na+) penetration was observable for any 100nm Al2O3 capacitors. The
hysteresis of each C-V curve for the Al2O3 MOS capacitance is due to the charge
Capacitance Density (nF/cm2)
injection into the oxide at high temperatures and is not related to mobile ion response.
90
100nm ALD Al2O3
80
70
0min
30min
60min
90min
120min
600min
1440min
60
50
-2
-1
0
1
2
Voltage (V)
Figure 11 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 100 nm oxide thicknesses.
28
3.2.3 Al2O3 Thickness Study
To increase the sensitivity of future bioFETS, 50nm, 25nm, 10nm Si based Al2O3
MOS capacitor were fabricated. Figure 12, Figure 13, and Figure 14 show that no
measurable sodium ion (Na+) penetration has been shown using Triangular Voltage
Sweep measurements at 250℃ and 1MHz for both positive bias to negative bias and
2
Capacitance Density (nF/cm )
negative bias to positive bias sweep.
160
50nm ALD Al2O3
140
120
100
80
-2
0min
30min
60min
90min
120min
600min
1440min
-1
0
1
2
Voltage (V)
Figure 12 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 100 nm oxide thicknesses.
29
2
Capacitance Density(nF/cm )
280
25nm ALD Al2O3
240
200
0min
30min
60min
90min
120min
600min
1440min
160
120
-2
-1
0
1
2
Voltage(V)
Figure 13
Triangular voltage sweep measurements of MOS capacitors with Al2O3
2
Capacitance Density (nF/cm )
dielectric at 250oC for 25 nm oxide thicknesses.
700
10nm ALD Al2O3
600
500
0min
30min
60min
90min
120min
600min
1440Min
400
300
200
100
-2
-1
0
1
2
Voltage (V)
Figure 14 Triangular voltage sweep measurements of MOS capacitors with Al2O3
dielectric at 250oC for 10 nm oxide thicknesses.
30
3.3 Result and Discussions
Atomic layer deposition of high-k Al2O3 barrier layers plays a very important role
for bio-devices. Before fabricating bioFETs, MOS capacitors were fabricated to quantify
alkali ion penetration from the PBS solution. MOS capacitors with different thicknesses
of ALD Al2O3 gate oxide were fabricated. The capacitors were dipped into a sterile
physiological buffer solution (PBS) solution for increasing intervals of time ranging from
30 minutes to 24 hours. The triangular voltage sweep (TVS) method was used to
characterize alkali ion penetration. No measurable alkali ion permeation was observed for
any Al2O3 capacitors, regardless of Al2O3 thickness. By contrast, the dose of alkali ion
penetration into silicon dioxide MOS capacitors increased with increasing soak times in
the sterile PBS solution. No evidence of alkali ion contamination was observed for
varying Al2O3 thicknesses ranging from 100 nm all the way down to only 10 nm (10nm,
25nm, 50nm, 100nm).
31
Chapter 4: BioField Effect Transistor Study
4.1 BioFET fabrication
Moderately doped (~ 1016 cm-3) p-type silicon wafers were used as bioFET
substrate. Five photomasks were used to pattern bioFET. SiO2 layer was grown to act as
the field oxide layer. The first mask used was to open the source and drain area.
Phosphorus pre-deposition followed by a drive-in process were used to obtain the n-type
dope. After the n-type diffusion, the second mask was used to open the gate area. Al2O3
and SiO2 were used as gate oxide separately. This were followed by annealed in N2
environment for 10min at 700 oC and 1050 oC separately. The third mask was used to
pattern via holes to allow electrical contact to the source and drain. Aluminum was
patterned using the fourth mask and the photolithographic lift-off method was used. A
post-metallization anneal was performed at 450oC for 10 min. Finally, the fifth mask was
used to pattern photoresist SPR-220 which was used to create a reservoir for the
measurement solution above the senor surface. The radio of width to length of gate area
is 10:1. FETs with various gate length of 25µm, 250µm, 1000µm were fabricated.
4.1.1 Field Oxidation Grown
A field oxide is used to define the source and drain regions of N-type diffusion.
Prior to field oxidation, the wafers were cleaned using the standard clean process
consisting of RCA1 (1NH4OH: 1H2O2: 5 de-ionized (DI) H2O at 70oC for 10 minutes)
and RCA2 (1HCl: 1 H2O2: 5 DI H2O at 70oC for 10 minutes). This was followed by a 1
32
minute dip in 1HF:1 DI and a 5 minute DI H2O rinse. There are two methods to grow
SiO2 layers, wet oxidation and dry oxidation.
Si(solid) + 2 H2O →SiO2(solid) Wet Oxidation
Si(solid) + O2 →SiO2(solid)
Dry Oxidation
The wet oxidation method was chosen to grow the field oxide because of the
quality of the dielectric properties of the field oxide are not very critical and a thicker
oxide is desired to avoid parasitic channels. Further, the growth rate of wet oxidation
growth is faster than the dry oxidation. Figure 15 shows oxide thickness versus oxidation
time for (100) and (111) silicon in H2O at 640 Torr partial pressure.
Figure 15 Oxide thickness versus oxidation time for silicon in H2O at 640 Torr[84]
33
After theoretical calculation, 50 minutes was chosen for wet oxidation of (100)
silicon wafers at 1100 ℃ .
A silicon dioxide layer (520nm) was formed after wet
oxidation. (Schematic figure 16)
Figure 16 A p-type silicon wafer with oxide layer.
4.1.2 Phosphorus Diffusion
Shipley 1813 was used as the photoresist to open windows for the phosphorus
diffused area. Wafers were spin coated with Shipley 1813 at 4000 rpm for 30sec and
followed by soft bake for 90 sec at 95℃. In order to obtain better photolithography
quality, wafers were baked at 180 ℃ for 20 min before applying photoresist to dehydrate
the surface. The coated wafers were exposed under UV light for 15 sec. Microposit MF319 developer was used to develop all samples for 45 sec. After development, all samples
were immersed in the buffered HF (BHF) solution for 4 min 25 sec. This was followed
by a 5 min rinse in DI H2O and blown dry with nitrogen gun. Windows for the
phosphorus were opened after this step. Before phosphorus diffusion, all photoresist was
removed from the wafer by soaking in acetone for 5 min, followed by soaking in
34
methanol for 5 min. All samples were rinsed in DI H2O for 5 min and blown dry with
nitrogen.
Prior to diffusion, the wafers were dipped in the DI:HF (10:1) solution for oxide
etch for about 10 sec. After that, RCA1 (1NH4OH: 1H2O2: 5 de-ionized (DI) H2O at 70oC
for 10 minutes) and RCA2 (1HCl: 1 H2O2: 5 DI H2O at 70oC for 10 minutes) were
applied to all wafers. This was followed by a 10 sec dip in 1HF:10 DI and a 5 minute DI
H2O rinse. Phosphorus predepostion diffusion was performed at 1050oC for 20 min.
Figure 17 shows a schematic of the phosphorus diffusion process.
Figure 17 A schematic of phosphorus diffusion process.
Phosphosilicate glass (PSG) was formed on the wafer during the phosphorus
predisposition. This layer will fix the surface concentration of the phosphorus doped
areas to the solid solubility limit. In order to remove PSG after the predeposition
diffusion the wafers were dipped in buffered HF (BHF) for 10 seconds. This was
followed by 5 min DI rinse and blown dry by nitrogen gun. Drive-in diffusion was
35
performed at 1100oC after etching PSG. A 17 min wet oxidation followed by 10 min dry
nitrogen was chosen for the drive-in process. Figure 18 shows a schematic of drive-in
diffusion process.
Figure 18 A schematic of drive-in diffusion.
After the drive-in process, the samples were ready for gate oxidation. Figure 19
shows a microscope image of the Si substrates after phosphorus diffusion and drive-in
process.
36
Figure 19 A microscope image of Si substrate after phosphorus diffusion and drive-in
process. W:L=10:1 L=25µm
4.1.3 Gate Oxidation
AZ5214 was used as the photoresist to open the area for gate oxidation. Wafers
were dehydration baked at 180 ℃ for 20 min in order to get better photolithography
quality. This was followed by spin coating with AZ5214 at 3000 rpm and soft bake for 60
sec at 90℃. The coated wafers were exposed under UV light for 0.5 sec then hard bake at
115℃ for 120 sec. After hard bake, an image reverse exposure was performed under UV
light for 10 sec. Microposit MF-319 developer was used to develop all samples for 60 sec.
After development, all samples were immersed in the buffered HF (BHF) solution for
7min 25 sec. This was followed by 5 min rinse in DI H2O and blown dry with nitrogen
gun. Windows for the gate oxidation were opened after this step. Before growing the
oxide gate, all photoresist was removed clean by soaking in acetone for 5 min and
37
followed by soaking in methanol for 5 min. This was followed by rinsing in DI H2O for 5
min and blown dry with nitrogen. In order to get better gate oxidation quality, the wafers
were cleaned prior to gate oxidation. All wafers were dipped in the DI:HF (10:1) solution
for oxide etch about 10 sec. After that, RCA1 (1NH4OH: 1H2O2: 5 de-ionized (DI) H2O
at 70oC for 10 minutes) and RCA2 (1HCl: 1 H2O2: 5 DI H2O at 70oC for 10 minutes)
were applied to all wafers. This was followed by a 10 sec dip in 1HF:10 DI and a 5
minute DI H2O rinse.
Al2O3 and SiO2 were chosen to be the gate oxidation, separately. The Al2O3
deposition was carried out with trimethylaluminum (TMA) and water as precursors at
300oC using a Picosun SunaleTM reactor. The ALD pulsing sequence for one cycle was
0.1 sec. per TMA pulse, 4 sec. per N2 purge, 0.1 sec. per H2O pulse, and 4 sec. per N2
purge. Typical ALD deposition rates of 0.8 Å/cycle were obtained. After 250 cycles,
20nm Al2O3 was grown on the gate area of samples. Figure 20 shows a microscope image
of Si substrate with Al2O3 ALD layer grow the devices.
38
Figure 20 A microscope image of Si substrate with Al2O3 gate grown on phosphorus
diffusion area. W:L=10:1.L=25µm
20nm SiO2 gate was grown as a control sample. Dry oxidation was used for
growing SiO2 because of it better oxidation quality.
Si(solid) + O2 →SiO2(solid)
39
Dry Oxidation
Figure 21 Oxide thickness as a function of time for dry oxidation of bare (100) silicon at
different temperature [85].
Figure 21 shows the relationship of oxide thickness as a function of time for wet
and dry oxidation of bare (100) silicon at different temperatures. After theoretical
calculation, 10 min SiO2 dry oxidation at 1100℃ was targeted for around 20 nm SiO2
of
gate oxide. Figure 22 shows a microscope image of the Si substrate with the SiO2 gate
grown on phosphorus diffusion area. SiO2 and Al2O3 were annealed in nitrogen ambient
40
for 10 min at 1050℃ and 700℃ separately.
Figure 22 A microscope image of Si substrate with SiO2 gate grown on phosphorus
diffusion area. W:L=10:1 25µm
4.1.4 Metal
Shipley 1813 was used as the photoresist to open contact holes. Wafers were spin
coated with Shipley 1813 at 4000 rpm for 30sec and followed by soft bake for 90 sec at
95℃. In order to obtain better photolithography quality, wafers were baked at 180 ℃ for
20 min before applying photoresist to dehydrate the surface. The coated wafers were
exposed under UV light for 15 sec. Microposit MF-319 developer was used to develop all
41
samples for 45 sec. After development, all samples were immersed in the buffered HF
(BHF) solution for 5 min 15 sec. This was followed by a 5 min rinse in DI H2O and
blown dry with nitrogen gun. Windows for the contact vias were opened after this step.
All photoresist was then removed from the wafer by soaking in acetone for 5 min,
followed by soaking in methanol for 5 min. All samples were rinsed in DI H2O for 5 min
and blown dry with nitrogen.
After etching the contact via holes, all samples were subjected to an Al metal lift
off process. AZ5214 was used as the photoresist to create an inverse pattern. Wafers
were baked at 180 ℃ for 20 min in order to get better photolithography quality. This was
followed by spin coatings with AZ5214 at 3000rpm and a soft bake for 60 sec at 90℃.
The coated wafers were exposed under UV light for 0.5 sec then hard baked at 115℃ for
120 sec. After hard bake, an image reversal exposure was performed under UV light for
10 sec. Microposit MF-319 developer was used to develop all samples for 60 sec. After
development, Al deposition was performed by electron beam evaporator. A thickness of
600nm aluminum was deposited by e-beam. The samples were soaked in the acetone for
about 10mins and followed by blown dry with nitrogen. In order to get better
performance FET, a backside aluminum deposition of 150nm was also processed. A postmetallization anneal was performed at 450oC for 10 min. The gate electrodes were
designed with holes, slots, solid metal and without metal as shown in the Fig. 23
42
Figure 23 Mask level of FET with different gate electrodes: gate electrodes were
designed with (a) holes, (b) solid metal, (c) without metal and (d) slots. Gate W:L=10:1
4.1.5 Reservoir
The thick photoresist SPR-220 was used to create a reservoir for the measurement
solution above the senor surface as shown in the Fig. 24.
43
Figure 24 Mask level of thick photoresist SPR-220 as reservoir (Green areas) for the
measurement solution above the senor surface. Gate W:L=10:1
Wafers were spin coated with SPR 220 at 1600 rpm for 60 sec and followed soft
bake for 120 sec at 115℃. In order to obtain better photolithography quality, wafers were
baked at 180 ℃ for 20 min before applying photoresist to dehydrate the surface. The
coated wafers were exposed under UV light for 16 sec. Post exposure bake was
performed at 115℃ for 90 sec. Microposit MF-319 developer was used to develop all
44
samples for 150 sec. After development, all samples were immersed in DI water for 5
min. In order to get better microfluid channel quality, a post bake at 115℃ for 5 min was
used. Firgure 25-28 shows the bioFET with different gate electrodes with reservoir
surrounding the gate.
Figure 25 Al2O3 gate BioFET with perforated gate metal (holes). Additionally the
reservoir is shown surrounding the gate test area. W:L 10:1 L=25
45
Figure 26 BioFET control device with solid metal gate electrodes and reservoir W:L 10:1
L=25
Figure 27 BioFET floating gate and reservoir. W:L 10:1 L=25
46
Figure 28 BioFET with slotted gate electrodes and reservoir. W:L 10:1 L=25
4.1.6 Bio-functionalized surface of the device
Before detecting Streptavidin, the gate areas of FETs were modified to grow an
immobilization layer respectively. First, the samples were rinsed in a DI water bath for 30
min
at
100 ℃
and
then
blown
dry
with
nitrogen
gun.
Second,
aminopropyldimethylethoxysilane (APDMES) was applied to the gate oxide surface of
the device for 30 min. Propanol solution were used to remove un-bound silane molecules,
and then baked 5 min at 120 ℃ (Figs. 29-30). 1mg/ml sulfo-NHS-biotin solution was
used as the receptor for the streptavidin and the silanized gate area. After applying
50ug/ml streptaviding solution to the whole device surface, the devices were incubated at
37℃ for 1 hour. PBS solution was used to rinse up all un-bounded Streptavidin (Figs. 3132).
47
Figure 29 The surface silanization procedure for streptavidin immobilization atop on the
Al2O3 layer
Figure 30 The surface silanization procedure for streptavidin immobilization atop on the
SiO2 layer
48
Figure 31 Diagram of the chemical linking of the streptavidin protein binding to the
Al2O3 surface
Figure 32 Diagram of the chemical linking of streptavidin protein binding to the SiO2
surface
49
4.2 Al2O3 BioFET Testing
PBS solution was applied in the reservoir and a metal probe was used to apply the
gate bias as shown in the Fig. 28.
measurements of Al2O3 FET biosensor with
gate metal were performed before applying PBS on the gate area.
measurements
of Al2O3 FET biosensor without gate metal were performed before and after apply PBS
on the gate area.
measurements were also studied before apply PBS solution.
Figure 33 Schematic of cross-sectional view of an Al2O3 FET biosensor test under PBS
solution set up.
4.2.1 Common-source I-V voltage study
Common source I-V measurements were performed on Al2O3 gate FETs with
three different gate electrodes as shown in Figs. 25, 26 and 27 before apply PBS in the
gate area. The results are shown in Figs. 34, 35 and 36.
50
Figure 34 Common-source I-V curve of Al2O3 gate FET with solid gate Al electrode.
W:L 10:1 L=25
Figure 35 Common-source I-V of Al2O3 gate FET with slots gate Al electrode. W:L
10:1 L=25
51
Figure 36 Common-source I-V of Al2O3 gate FET with holes gate Al electrode. W:L
10:1 L=25
As shown from Figs. 34-36, a series
curves measured on the FETs with
different gate electrodes patterns. Each of the FETs with different gate metal patterns
were functional at the same gate biases and drain voltages.
0.10
0.08
Ids (A)
0.06
0.04
0.02
0.00
0
1
2
3
4
5
6
Vds (V)
Figure 37 Common-source I-V of Al2O3 gate FET without gate Al electrode before apply
PBS. W:L 10:1 L=1000
52
Vgs=0V
Vgs=1V
Vgs=2V
Vgs=3V
Vgs=4V
Vgs=5V
0.7
0.6
0.5
Ids (mA)
0.4
0.3
0.2
0.1
0.0
-0.1
0
1
2
3
4
5
Vds (V)
Figure 38 Common-source I-V of Al2O3 gate FET without gate Al electrode after apply
PBS. (PBS was in the reservoir and a bias of 1-5 V was applied to gate area) W:L 10:1
L=1000
Figure 37 and Fig. 38 shows Al2O3 FET without gate metal
measurements before and after apply PBS solution in the gate area. Before apply gate
bias, the Al2O3 FET was turned off due to no gate bias because there is no metal or
conductor applied gate bias voltage. After applying physiological buffer solution (PBS)
on the gate area, the solution which full of mobile ion works as a gate electrode (Fig. 38).
The Al2O3 FET was turned on even without gate bias. This indicates the PBS solution
induced and inversion layer which allowed the flow of electrons between the source and
drain. With the gate biased increase beyond 3V to 5V, the drain current is not increased
but decreased.
53
4.2.2 Threshold voltage study
The threshold voltage of a MOSFET is the gate voltage which forms the inversion
layer that allows the flow of electrons between the source and drain. Figure 39 shows the
results of drain current vs gate voltage measurements for a solid gate electrode Al2O3
MOSFET before apply PBS solution.
8
7
6
Ids (mA)
5
4
3
2
1
0
-1
-6
-4
-2
0
2
4
6
Vgs (V)
Figure 39 Id- VG curve of Al2O3 gate FET with solid gate Al electrode. (
54
)
0.10
0.08
Equation
Adj. R-Square
y = a + b*x
0.9989
B
B
Intercept
Slope
Value
0.00702
0.01597
Standard Error
1.35376E-4
5.76983E-5
Squart Id
0.06
0.04
0.02
0.00
-6
-4
-2
0
2
4
6
Vg (V)
Figure 40 Square root Id- VG curve of Al2O3 gate FET with solid gate Al electrode.
(
)
The length of the smallest Al2O3 FET is 25µm and the width is 250µm so it can be
represented by the long-channel MOSFET model.
The drain current in saturation is given by
I D  k (VGS  VT ) 2 
k
WCox  nc
(VGS  VT ) 2
2L
WCox  nc
2L
Figure 40 shows the square root drain current vs the gate voltage measurements. We
can see that √
and
ℎ
√
.
=-
7
6
2
7
55
Similar analysis was done for the Al2O3 FET with slots gate electrode and holes gate
electrode separately.
5
Ids (mA)
4
3
2
1
0
-6
-4
-2
0
2
4
6
Vgs(V)
Figure 41 Id- VG curve of Al2O3 gate FET with holes gate Al electrode (
)
0.07
Equation
y = a + b*x
Adj. R-Square
0.06
0.99874
Value
0.05
Standard Error
D
Intercept
0.00494
1.24253E-4
D
Slope
0.01302
4.66176E-5
Sqrt Id
0.04
0.03
0.02
0.01
0.00
-0.01
-6
-4
-2
0
2
4
6
Vgs(V)
Figure 42 Square root Id- VG curve of Al2O3 gate FET with holes gate Al electrode.
(
)
56
The drain current in saturation is given by
I D  k (VGS  VT ) 2 
k
WCox  nc
(VGS  VT ) 2
2L
WCox  nc
2L
Figure 42 shows the square root drain current vs the gate voltage measurements. We
can see that √
and
ℎ
.
7
=-
0.3794
2
6
√
7
6
Ids (mA)
5
4
3
2
1
0
-6
-4
-2
0
2
4
6
Vgs (V)
Figure 43 Id- VG curve of Al2O3 gate FET with slots gate Al electrode (
57
)
0.08
Equation
y = a + b*x
Adj. R-Square
0.99917
Value
Sqrt Id
0.06
Standard Error
D
Intercept
0.00488
1.15738E-4
D
Slope
0.01511
4.43835E-5
0.04
0.02
0.00
-6
-4
-2
0
2
4
6
Vgs(V)
Figure 44 Square root Id- VG curve of Al2O3 gate FET with slots gate Al electrode.
(
)
The drain current in saturation is given by
I D  k (VGS  VT ) 2 
k
WCox  nc
(VGS  VT ) 2
2L
WCox  nc
2L
As shown from Figure 44 shows the square root drain current vs the gate voltage
measurements. We can see that √
ℎ
and
7
=-
.
0.323
2
√
58
4.2.3 Streptavidin detection
Common-source I-V measurements were performed before and after streptavidin
detection. The Al2O3 gate FET without gate metal was used to detect streptavidin protein.
After fully functionalizing the surface of the FET by APDMES, the gate of the Al2O3
FET was exposed to streptavidin. This was performed by applying a solution with
50µg/ml streptavidin to the gate area and then incubating at room temperature for an hour.
After that, the device was rinsed in the physiological buffer solution (PBS) three times in
order to clean unbound streptavidin.
Figure 45 shows the
characteristics of 20nm Al2O3 gate BioFET before
and after applying 50µg/ml streptavidin in the sterile physiological buffer solution (PBS,
150 mM Na+). The first
reservoir. The second
curve is tested with only PBS in the biofunctionalized
curve is measured after binding protein. A gate bias of 5V
was used based on Fig. 38.
59
Binding with SA in PBS
Before Binding in PBS
0.5
IDS (mA)
0.4
0.3
0.2
0.1
0.0
0.0
0.2
0.4
0.6
0.8
1.0
1.2
1.4
1.6
1.8
2.0
VDS (V)
Figure 45 Common-source I-V of Al2O3 gate FET before and after binding 50µg/ml
streptavidin in reservoir with PBS solution and a bias of 5 V was applied to gate area
W:L 10:1 L=1000
As shown in the Figure 45, the drain to source current decreased slightly after
exposure to streptavidin. However, the changes for this device are very small and cannot
be used as an effective biosensor. Improvements in fabrication could address some of the
losses in sensitivity and improve future biosensors.
4.3 SiO2 BioFET Testing
Common-source I-V measurements were also attempted on the 20nm SiO2 gate
bioFETs as the control samples before apply apply PBS solution. However, after trying
60
almost all the FETs on the wafer, we can only get common source I-V curves like Fig. 46.
The current is negative and very small.
0.00E+000
-2.00E-011
Ids (A)
-4.00E-011
-6.00E-011
-8.00E-011
-1.00E-010
0
1
2
3
4
5
Vds (V)
Figure 46 Common-source I-V curve of SiO2 gate FET with solid gate Al electrode and a
bias of 5 V was applied to gate area. W:L 10:1 L=25
As shown from Fig. 46, the negative and small current means that the FET
doesn’t turn on after apply 5v drain voltage and 5v gate bias. In order to find the problem,
we test the source and drain to substrate junctions as diodes separately (Figs. 47,48)
61
2.50E-012
2.00E-012
1.50E-012
Id (A)
1.00E-012
5.00E-013
0.00E+000
-5.00E-013
-1.00E-012
-1.50E-012
-2
-1
0
1
2
Vd (V)
Figure 47 Diode I-V measurement of drain of SiO2 FET.
2.50E-012
2.00E-012
1.50E-012
Is (A)
1.00E-012
5.00E-013
0.00E+000
-5.00E-013
-1.00E-012
-1.50E-012
-2
-1
0
1
2
Vs (v)
Figure 48 Diode I-V measurement of source of SiO2 FET.
62
As shown from Figs. 47 and 48, the I-V curves are similar to resistor I-V curves
instead of diode which could be due to the following reason. After etching the via holes,
the SiO2 FET samples were exposed to air for about two days before e-beam metal
deposition. The native oxide that formed in the contact vias interfered with formation of
ohmic contacts between the aluminum and source and drain on these samples. The 10min
anneal after metal deposition was not sufficient to penetrate the native oxide. Further
evidence in support of this is that an a few of the source or drain contacts, repeated
measurements from -40V to 40V were able to breakdown the native oxide at the interface,
resulting in a diode-like I-V characteristic. But we ere not able to obtain such an
improvement to both the drain and source contacts on an any of the individual FETs on
the wafer.
4.4 Result and Discussions
20nm Al2O3 gate bioFETs were fabricated and tested before and after applying
PBS solution. The results show that 20nm Al2O3 bioFET without gate metal can operate
under PBS after full functionalization by APDMES. The drain to source current
decreased slightly after detecting the streptavidin. However, the decreased changes are
very small and cannot be used as an effective biosensor. To some extent, these changes
could be some noise because common source I-V measurements were done in solution
which increases lots of unstable factors.
As the control sample, 20nm SiO2 gate FETs were fabricated and tested. However,
the 20nm SiO2 gate FETs doesn’t work because the contact between diffusion area and
metal is not good. The reason of this problem may be that SiO2 FET samples were
63
exposed to air for about two days after etch the contact holes. The Si on the via hole areas
was oxidized very quickly. The deposited metal was insulated by the thin oxidate layer
and could not contact with diffusion area therefore could not turn on the FET.
Another problem is that some Al electrodes were etched by APDMES and PBS
solution which may make devices unstable. Gold metal deposition is a better choice for
future device fabrication due to its inert properties.
64
Chapter 5: Summary and further directions
5.1 Summary
Silicon in vivo protein biosensors suffer from long-term electrical drifting and
instability due to the contamination of alkali ions from high osmolarity biological buffers.
Their long-term stability and biocompatibility is of great concern which requires
significant improvements for clinical use.
In this thesis, a low-cost Si based MOS capacitor with a high-k Al2O3 dielectric
deposited by ALD has been fabricated. High-k dielectric layers not only prevent alkali
ions diffusion from high osmolarity biological buffers into the gate oxides but also result
in enhanced device sensitivity due to increased electrostatic coupling. Si-based ALD
Al2O3 MOS capacitors show no measurable peak before and after soaking in the PBS
solution indicating no alkali ions penetration for various oxide thicknesses of 100nm,
50nm, 25nm, 10nm. In contrast, alkali ion penetration into SiO2 MOS capacitors
increased with increasing soak times in PBS solution.
A low cost Si based bioFET with high high-k Al2O3 dielectric deposited by ALD was
fabricated and tested in high osmolarity environment. The results show that 20nm Al2O3
bioFET without gate metal can operate under PBS after full functionalization by
APDMES. The drain to source current decreased after detecting the streptavidin. This
could be the mobile charge in the channel decreased after binding of the protein.
However, the decreased changes are very small and cannot be used as an effective
65
biosensor. To some extent, these changes could be noise because the common source I-V
measurements were done in solution which increases lots of unstable factors.
5.2 Further Directions
One of the key elements that allows for successful scaling of biosensors is the
electrical properties of the gate dielectric through the equation, C = εA/d, indicating that a
higher permittivity, ε, or a thinner dielectric thickness, d, will elevate the capacitance, C,
which means increased biosensitivity. The output may be normalized per unit area, A,
rendering that parameter less critical. However, further scaling done of the gate thickness
may raise the leakage current flowing through the bioFET by a quantum mechanical
tunneling mechanism [86-90].
Figure 49 A field effect transistor Si incorporated with high-k dielectric stack layer
protein biosensor is presented (Multi-color layers in the gate imply different oxides
stacked together).
66
As shown in the Fig. 49, a sensing channel connects the source (S) and drain (D) with a
reference electrode (RE). When a target protein (Steptavidin) binds to the silane linker
layer, it induces charges in the substrate (electrons as pictured here), causing a change in
the current flow between the source and drain. In order to improve the performance and
make ultrathin biosensors, a variety of topologies can be used, including an exploration
of the viability of multilayer high-k dielectric stacks coated on the surface of the silicon,
such as combinations of Al2O3, Ta2O5, TiO2, HfO2 etc. deposited by atomic layer
deposition (ALD) creating ultrathin alternating layers, preferably toggling between
materials to provide the maximum of chemical potential for trapping the unwanted ions
and high permittivities.
The high-k material which will be used as a gate for the biosensor should satisfy a
long list of the requirements.
 Good thermal stability in contact with Si, preventing the formation of a parasitic SiOx
interfacial layer leading to lower “effective” permittivity, nor the formation of
undesired silicide layers;
 Low density of intrinsic defects at the Si/dielectric interface and in the bulk of the
material, providing high mobility of charge carriers in the channel and sufficient gate
dielectric lifetime;
 A sufficiently large energy band gap, providing high energy barriers at the
Si/dielectric and metal gate/dielectric interfaces, in order to reduce the leakage current
flowing through the structure;
67
Preliminary results indicate that high-k dielectric layers, such as Al2O3, prevent alkali
ions diffusion from high osmolarity biological buffers into the gate oxides.
68
References
[1]
[2]
[3]
[4]
[5]
[6]
[7]
[8]
[9]
[10]
[11]
[12]
[13]
C. A. Rowe, S. B. Scruggs, M. J. Feldstein, J. P. Golden and F. S. Ligler, "An array
immunosensor for simultaneous detection of clinical analytes," Anal. Chem., vol.
71, pp. 433-439, 1999.
D. Leech, "Affinity biosensors," Chem. Soc. Rev., vol. 23, pp. 205-213, 1994.
Z. Liu, B. Liu, M. Zhang, J. Kong and J. Deng, "Al2O3 sol-gel derived
amperometric biosensor for glucose," Anal. Chim. Acta, vol. 392, pp. 135-141,
1999.
I. S. Wang, W. Y. Chung, T. F. Lu, H. C. Chuang, C. E. Lue, C. M. Yang, D. G.
Pijanswska, J. C. Wang and C. S. Lai, "Urea Biosensor with Nitrited-ALD-HfO2
Membrane by Remote NH3 Plasma Based on EIS Structure," .
B. Kang, H. Wang, F. Ren and S. Pearton, "Electrical detection of biomaterials
using AlGaN/GaN high electron mobility transistors," J. Appl. Phys., vol. 104, pp.
031101-031101-11, 2008.
G. Steinhoff, O. Purrucker, M. Tanaka, M. Stutzmann and M. Eickhoff, "AlxGa1–
xN—A New Material System for Biosensors," Advanced Functional Materials, vol.
13, pp. 841-846, 2003.
S. Updike and G. Hicks, "Reagentless substrate analysis with immobilized
enzymes," Science, vol. 158, pp. 270, 1967.
J. Wang, "Amperometric biosensors for clinical and therapeutic drug monitoring: a
review," J. Pharm. Biomed. Anal., vol. 19, pp. 47-53, 1999.
L. Gorton, G. Bremle, E. Csöregi, G. Jönsson-Pettersson and B. Persson,
"Amperometric glucose sensors based on immobilized glucose-oxidizing enzymes
and chemically modified electrodes," Anal. Chim. Acta, vol. 249, pp. 43-54, 1991.
H. Tang, J. Chen, S. Yao, L. Nie, G. Deng and Y. Kuang, "Amperometric glucose
biosensor based on adsorption of glucose oxidase at platinum nanoparticle-modified
carbon nanotube electrode," Anal. Biochem., vol. 331, pp. 89-97, 2004.
A. Guerrieri, G. De Benedetto, F. Palmisano and P. Zambonin, "Electrosynthesized
non-conducting polymers as permselective membranes in amperometric enzyme
electrodes: a glucose biosensor based on a co-crosslinked glucose
oxidase/overoxidized polypyrrole bilayer," Biosensors and Bioelectronics, vol. 13,
pp. 103-112, 1998.
S. Rodriguez-Mozaz, M. J. Lopez de Alda and D. Barceló, "Biosensors as useful
tools for environmental analysis and monitoring," Analytical and Bioanalytical
Chemistry, vol. 386, pp. 1025-1041, 2006.
S. H. Choi and M. B. Gu, "A portable toxicity biosensor using freeze-dried
recombinant bioluminescent bacteria," Biosensors and Bioelectronics, vol. 17, pp.
433-440, 2002.
69
[14] A. Rasooly and K. E. Herold, "Biosensors for the analysis of food-and waterborne
pathogens and their toxins," J. AOAC Int., vol. 89, pp. 873-883, 2006.
[15] S. C. Mukhopadhyay and C. P. Gooneratne, "A novel planar-type biosensor for
noninvasive meat inspection," Sensors Journal, IEEE, vol. 7, pp. 1340-1346, 2007.
[16] D. Grieshaber, R. MacKenzie, J. Voeroes and E. Reimhult, "Electrochemical
biosensors-Sensor principles and architectures," Sensors, vol. 8, pp. 1400-1458,
2008.
[17] G. Piechotta, J. Albers and R. Hintsche, "Novel micromachined silicon sensor for
continuous glucose monitoring," Biosensors and Bioelectronics, vol. 21, pp. 802808, 2005.
[18] Y. Chen, X. Wang, S. Erramilli, P. Mohanty and A. Kalinowski, "Silicon-based
nanoelectronic field-effect pH sensor with local gate control," Appl. Phys. Lett., vol.
89, pp. 223512, 2006.
[19] O. H. Elibol, B. Reddy Jr and R. Bashir, "Nanoscale thickness double-gated field
effect silicon sensors for sensitive pH detection in fluid," Appl. Phys. Lett., vol. 92,
pp. 193904, 2008.
[20] H. Ouyang, L. A. DeLouise, B. L. Miller and P. M. Fauchet, "Label-free
quantitative detection of protein using macroporous silicon photonic bandgap
biosensors," Anal. Chem., vol. 79, pp. 1502-1506, 2007.
[21] A. Star, J. C. P. Gabriel, K. Bradley and G. Grüner, "Electronic detection of specific
protein binding using nanotube FET devices," Nano Letters, vol. 3, pp. 459-463,
2003.
[22] R. Gabl, H. D. Feucht, H. Zeininger, G. Eckstein, M. Schreiter, R. Primig, D. Pitzer
and W. Wersing, "First results on label-free detection of DNA and protein
molecules using a novel integrated sensor technology based on gravimetric
detection principles," Biosensors and Bioelectronics, vol. 19, pp. 615-620, 2004.
[23] T. Nicholson, S. Gupta, X. Wen, H. Wu, R. Anisha, P. Casal, K. Kwak, B. Bhushan,
P. Berger and W. Lu, "Rational enhancement of nanobiotechnological device
functions illustrated by partial optimization of a protein-sensing field effect
transistor," Proceedings of the Institution of Mechanical Engineers, Part N: Journal
of Nanoengineering and Nanosystems, vol. 223, pp. 149-161, 2010.
[24] M. Calleja, M. Nordström, M. Alvarez, J. Tamayo, L. M. Lechuga and A. Boisen,
"Highly sensitive polymer-based cantilever-sensors for DNA detection,"
Ultramicroscopy, vol. 105, pp. 215-222, 2005.
[25] Z. Li, Y. Chen, X. Li, T. Kamins, K. Nauka and R. Williams, "Sequence-specific
label-free DNA sensors based on silicon nanowires," Nano Letters, vol. 4, pp. 245247, 2004.
[26] Y. H. Kim, K. S. Shin, J. Y. Kang, E. G. Yang, K. K. Paek, D. S. Seo and B. K. Ju,
"Poly (dimethylsiloxane)-based packaging technique for microchip fluorescence
detection system applications," Microelectromechanical Systems, Journal of, vol.
15, pp. 1152-1158, 2006.
[27] A. P. F. Turner, "Current trends in biosensor research and development," Sensors
and Actuators, vol. 17, pp. 433-450, 1989.
70
[28] K. R. Rogers, "Principles of affinity-based biosensors," Mol. Biotechnol., vol. 14,
pp. 109-129, 2000.
[29] P. Darbon, V. Michel, F. Math, H. Giorgi and F. Machizaud, "Immunoelectrodes in
protein detection: comparison between glassy carbon and a semimetallic Ni/P thin
film as binding support. Biological applications," Anal. Chem., vol. 70, pp. 50725078, 1998.
[30] S. S. Iqbal, M. W. Mayo, J. G. Bruno, B. V. Bronk, C. A. Batt and J. P. Chambers,
"A review of molecular recognition technologies for detection of biological threat
agents," Biosensors and Bioelectronics, vol. 15, pp. 549-578, 2000.
[31] K. Rogers, "Recent advances in biosensor techniques for environmental
monitoring," Anal. Chim. Acta, vol. 568, pp. 222-231, 2006.
[32] K. R. Rogers, "Principles of affinity-based biosensors," Mol. Biotechnol., vol. 14,
pp. 109-129, 2000.
[33] I. Willner, E. Katz and B. Willner, "Electrical contact of redox enzyme layers
associated with electrodes: routes to amperometric biosensors," Electroanalysis, vol.
9, pp. 965-977, 1997.
[34] L. D. Mello and L. T. Kubota, "Review of the use of biosensors as analytical tools
in the food and drink industries," Food Chem., vol. 77, pp. 237-256, 2002.
[35] X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter and Y. Sun, "Sensitive
optical biosensors for unlabeled targets: A review," Anal. Chim. Acta, vol. 620, pp.
8-26, 2008.
[36] P. Y. Li, B. Lin, J. Gerstenmaier and B. T. Cunningham, "A new method for labelfree imaging of biomolecular interactions," Sensors Actuators B: Chem., vol. 99, pp.
6-13, 2004.
[37] L. Bousse, "Whole cell biosensors," Sensors Actuators B: Chem., vol. 34, pp. 270275, 1996.
[38] C. Ziegler, "Cell-based biosensors," Fresenius J. Anal. Chem., vol. 366, pp. 552559, 2000.
[39] H. S. Stoker, General, Organic, and Biological Chemistry. pp. 616, 2009.
[40] K. W. Wee, G. Y. Kang, J. Park, J. Y. Kang, D. S. Yoon, J. H. Park and T. S. Kim,
"Novel electrical detection of label-free disease marker proteins using piezoresistive
self-sensing micro-cantilevers," Biosensors and Bioelectronics, vol. 20, pp. 19321938, 2005.
[41] Y. Arntz, J. Seelig, H. Lang, J. Zhang, P. Hunziker, J. Ramseyer, E. Meyer, M.
Hegner and C. Gerber, "Label-free protein assay based on a nanomechanical
cantilever array," Nanotechnology, vol. 14, pp. 86, 2003.
[42] K. Martin, T. H. Steinberg, L. A. Cooley, K. R. Gee, J. M. Beechem and W. F.
Patton, "Quantitative analysis of protein phosphorylation status and protein kinase
activity on microarrays using a novel fluorescent phosphorylation sensor dye,"
Proteomics, vol. 3, pp. 1244-1255, 2003.
[43] D. W. Abbott, A. Wilkins, J. M. Asara and L. C. Cantley, "The Crohn's disease
protein, NOD2, requires RIP2 in order to induce ubiquitinylation of a novel site on
NEMO," Current Biology, vol. 14, pp. 2217-2227, 2004.
71
[44] M. Veiseh, Y. Zhang, K. Hinkley and M. Zhang, "Two-dimensional protein
micropatterning for sensor applications through chemical selectivity technique,"
Biomed. Microdevices, vol. 3, pp. 45-51, 2001.
[45] J. J. Saarinen, J. E. Sipe, S. M. Weiss and P. M. Fauchet, "Optical sensor based on
resonant porous silicon structures," in Conference on Lasers and Electro-Optics,
2005, .
[46] H. Ouyang, L. A. DeLouise, B. L. Miller and P. M. Fauchet, "Label-free
quantitative detection of protein using macroporous silicon photonic bandgap
biosensors," Anal. Chem., vol. 79, pp. 1502-1506, 2007.
[47] Y. Wang, Y. Zhou, J. Sokolov, B. Rigas, K. Levon and M. Rafailovich, "A
potentiometric protein sensor built with surface molecular imprinting method,"
Biosensors and Bioelectronics, vol. 24, pp. 162-166, 2008.
[48] S. Gupta, M. Elias, X. Wen, J. Shapiro, L. Brillson, W. Lu and S. C. Lee,
"Detection of clinically relevant levels of protein analyte under physiologic buffer
using planar field effect transistors," Biosensors and Bioelectronics, vol. 24, pp.
505-511, 2008.
[49] B. Kang, H. Wang, F. Ren and S. Pearton, "Electrical detection of biomaterials
using AlGaN/GaN high electron mobility transistors," J. Appl. Phys., vol. 104, pp.
031101-031101-11, 2008.
[50] B. S. Kang, F. Ren, L. Wang, C. Lofton, W. W. Tan, S. Pearton, A. Dabiran, A.
Osinsky and P. Chow, "Electrical detection of immobilized proteins with ungated
AlGaN∕ GaN high-electron-mobility Transistors," Appl. Phys. Lett., vol. 87, pp.
023508, 2005.
[51] J. J. Gooding, R. Wibowo, J. Liu, W. Yang, D. Losic, S. Orbons, F. J. Mearns, J. G.
Shapter and D. B. Hibbert, "Protein electrochemistry using aligned carbon nanotube
arrays," J. Am. Chem. Soc., vol. 125, pp. 9006-9007, 2003.
[52] K. Besteman, J. O. Lee, F. G. M. Wiertz, H. A. Heering and C. Dekker, "Enzymecoated carbon nanotubes as single-molecule biosensors," Nano Letters, vol. 3, pp.
727-730, 2003.
[53] C. Liu, "Recent developments in polymer MEMS," Adv Mater, vol. 19, pp. 3783–
3790, 2007.
[54] C. Wang, L. Zhang, Z. Guo, J. Xu, H. Wang, K. Zhai and X. Zhuo, "A novel
hydrazine electrochemical sensor based on the high specific surface area graphene,"
Microchimica Acta, vol. 169, pp. 1-6, 2010.
[55] G. Derbenwick, "Mobile ions in SiO2: potassium," J. Appl. Phys., vol. 48, pp.
1127-1130, 1977.
[56] M. Kuhn and D. Silversmith, "Ionic contamination and transport of mobile ions in
MOS structures," J. Electrochem. Soc., vol. 118, pp. 966, 1971.
[57] E. Snow, A. Grove, B. Deal and C. Sah, "Ion transport phenomena in insulating
films," J. Appl. Phys., vol. 36, pp. 1664-1673, 1965.
[58] G. H. Cross, A. A. Reeves, S. Brand, J. F. Popplewell, L. L. Peel, M. J. Swann and
N. J. Freeman, "A new quantitative optical biosensor for protein characterisation,"
Biosensors and Bioelectronics, vol. 19, pp. 383-390, 2003.
72
[59] P. R. Edwards, A. Gill, D. V. Pollardknight, M. Hoare, P. E. Buckle, P. A. Lowe
and R. J. Leatherbarrow, "Kinetics of protein-protein interactions at the surface of
an optical biosensor," Anal. Biochem., vol. 231, pp. 210-217, 1995.
[60] K. P. S. Dancil, D. P. Greiner and M. J. Sailor, "A porous silicon optical biosensor:
detection of reversible binding of IgG to a protein A-modified surface," J. Am.
Chem. Soc., vol. 121, pp. 7925-7930, 1999.
[61] A. J. Haes and R. P. Van Duyne, "A nanoscale optical biosensor: sensitivity and
selectivity of an approach based on the localized surface plasmon resonance
spectroscopy of triangular silver nanoparticles," J. Am. Chem. Soc., vol. 124, pp.
10596-10604, 2002.
[62] E. Katz and I. Willner, "Probing Biomolecular Interactions at Conductive and
Semiconductive Surfaces by Impedance Spectroscopy: Routes to Impedimetric
Immunosensors, DNA‐Sensors, and Enzyme Biosensors," Electroanalysis, vol. 15,
pp. 913-947, 2003.
[63] M. C. Rodriguez, A. N. Kawde and J. Wang, "Aptamer biosensor for label-free
impedance spectroscopy detection of proteins based on recognition-induced
switching of the surface charge," Chemical Communications, pp. 4267-4269, 2005.
[64] R. W. Nelson, D. Nedelkov and K. A. Tubbs, "Biosensor chip mass spectrometry:
A chip‐based proteomics approach," Electrophoresis, vol. 21, pp. 1155-1163,
2000.
[65] A. N. Asanov, W. W. Wilson and P. B. Oldham, "Regenerable biosensor platform:
a total internal reflection fluorescence cell with electrochemical control," Anal.
Chem., vol. 70, pp. 1156-1163, 1998.
[66] G. S. Bang, S. Cho and B. G. Kim, "A novel electrochemical detection method for
aptamer biosensors," Biosensors and Bioelectronics, vol. 21, pp. 863-870, 2005.
[67] J. Wang, "Electrochemical biosensors: towards point-of-care cancer diagnostics,"
Biosensors and Bioelectronics, vol. 21, pp. 1887-1892, 2006.
[68] K. Ikebukuro, C. Kiyohara and K. Sode, "Novel electrochemical sensor system for
protein using the aptamers in sandwich manner," Biosensors and Bioelectronics, vol.
20, pp. 2168-2172, 2005.
[69] V. Nanduri, A. K. Bhunia, S. I. Tu, G. C. Paoli and J. D. Brewster, "SPR biosensor
for the detection of L. monocytogenes using phage-displayed antibody," Biosensors
and Bioelectronics, vol. 23, pp. 248-252, 2007.
[70] J. S. Yuk, H. S. Kim, J. W. Jung, S. H. Jung, S. J. Lee, W. J. Kim, J. A. Han, Y. M.
Kim and K. S. Ha, "Analysis of protein interactions on protein arrays by a novel
spectral surface plasmon resonance imaging," Biosensors and Bioelectronics, vol.
21, pp. 1521-1528, 2006.
[71] V. Nanduri, A. K. Bhunia, S. I. Tu, G. C. Paoli and J. D. Brewster, "SPR biosensor
for the detection of L. monocytogenes using phage-displayed antibody," Biosensors
and Bioelectronics, vol. 23, pp. 248-252, 2007.
[72] T. Windbacher, V. Sverdlov and S. Selberherr, "Biotin-Streptavidin Sensitive
BioFETs and Their Properties," Biomedical Engineering Systems and Technologies,
pp. 85-95, 2010.
73
[73] K. Park, M. Kim, K. Park and S. Choi, "Fabrication of BioFET sensor for
simultaneous detection of protein and DNA," .
[74] W. A. Hendrickson, A. Pähler, J. L. Smith, Y. Satow, E. A. Merritt and R. P.
Phizackerley, "Crystal structure of core streptavidin determined from
multiwavelength anomalous diffraction of synchrotron radiation," Proceedings of
the National Academy of Sciences, vol. 86, pp. 2190, 1989.
[75] . Holmberg, A. Blomstergren, O. Nord, M. Lukacs, J. Lundeberg and M. Uhlén,
"The biotin ‐ streptavidin interaction can be reversibly broken using water at
elevated temperatures," Electrophoresis, vol. 26, pp. 501-510, 2005.
[76] N. H. E. Weste and D. Money, Cmos Vlsi Design. Pearson/Addison Wesley, 2005.
[77] C. T. Sah, "Characteristics of the metal-oxide-semiconductor transistors," Electron
Devices, IEEE Transactions on, vol. 11, pp. 324-345, 1964.
[78] W. Shockley, "A unipolar field-effect transistor," Proceedings of the IRE, vol. 40,
pp. 1365-1376, 1952.
[79] N. Sato, Electrochemistry at Metal and Semiconductor Electrodes. Elsevier Science,
1998.
[80] H. Gerischer and C. W. Tobias, "Advances in Electrochemistry and
Electrochemical Engineering. Vol. 1," John Wiley and Sons, pp. 446, 1961.
[81] U. K. Mishra and J. Singh, Semiconductor Device Physics and Design. Springer
Verlag, 2007.
[82] S. M. George, "Atomic layer deposition: An overview," Polymer, vol. 1550, pp.
111-131, 2010.
[83] D. K. Schroder, Semiconductor Material and Device Characterization. Wiley-IEEE
Press, 2006.
[84] S. A. Campbell, The Science and Engineering of Microelectronic Fabrication.
Oxford University Press New York, 2001.
[85] S. K. Gandhi, "VLSI Fabrication Principles, Silicon and Gallium Arsenide," 1983.
[86] M. Houssa, High-k Gate Dielectrics. Taylor & Francis, 2004.
[87] E. Gusev and C. D’Emic, "Charge detrapping in HfO high-κ gate dielectric stacks,"
Appl. Phys. Lett., vol. 83, pp. 5223, 2003.
[88] R. Chau, S. Datta, M. Doczy, B. Doyle, J. Kavalieros and M. Metz, "High-κ/metalgate stack and its MOSFET characteristics," Electron Device Letters, IEEE, vol. 25,
pp. 408-410, 2004.
[89] M. Passlack, R. Droopad, K. Rajagopalan, J. Abrokwah, R. Gregory and D. Nguyen,
"High mobility NMOSFET structure with high-κ dielectric," Electron Device
Letters, IEEE, vol. 26, pp. 713-715, 2005.
[90] G. Wilk, R. Wallace and J. Anthony, "High-κ gate dielectrics: Current status and
materials properties considerations," J. Appl. Phys., vol. 89, pp. 5243, 2001.
74