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Surface functionalization of nanomaterials and development of field effect transistor nanobiosensors
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SURFACE FUNCTIONALIZATION OF NANOMATERIALS
AND DEVELOPMENT OF FIELD EFFECT TRANSISTOR NANOBIOSENSORS
by
YAN SONG
A Dissertation Presented to the
FACULTY OF THE USC GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the
Requirements for the Degree
DOCTOR OF PHILOSOPHY
(CHEMISTRY)
August 2015
Copyright 2015 YAN SONG
! ii
EPIGRAPH
“We all have dreams. But in order to make dreams come into reality, it takes
an awful lot of determination, dedication, self-discipline, and effort.”
--Jesse Owens!
! iii
ACKNOWLEDGEMENTS
I would like to express my sincere gratitude to my PhD advisor, Professor Mark
Thompson, for his excellent guidance and patience throughout my time at USC. I am
very grateful and proud to work with this great chemist and kind mentor.
I would like to thank all the collaborating professors during my PhD study,
Professor Chongwu Zhou, Professor Richard Cote, Professor Ram Datar and Professor
Richard Roberts, for helping me develop my skills and knowledge out of the scope of
chemistry.
I would also like to thank all the committee members during my screening exam,
qualifying exam and defense, Professor Brent Melot, Professor Richard Brutchey,
Professor Edward Goo, Professor G. K. Surya Prakash and Professor Andrey Vilesov.
I would like to thank all the current and previous biosensing team members: Dr.
Marco Curreli, Dr. Fumi Ishikawa, Dr. Hsiaokang Chang, Dr. Rui Zhang, Dr. Xiaoli
Wang, Dr. Noppadol Aroonyadet and Qingzhou Liu, for all the things I learned and every
great discussions we had during my PhD research.
I would also like to thank all the friends I made during my PhD study at USC: Dr.
Bing Xu, Dr. Qiwen Zhong, Dr. Yifei Liu, Dr. Marcos Sainz and Dr. Yuan Ding for all
the support and good times we spent together.
In addition, I would like to give thanks to all the current and former Thompson
Group members, especially Judy Fong, for making the lab running so well and for the
valuable discussions and support.
Most importantly, I want to thank my family. I know I cannot be who I am today
or accomplish what I have accomplished without their love and support.
! iv
TABLE OF CONTENTS
EPIGRAPH ...................................................................................................................................... ii!
ACKNOWLEDGEMENTS ............................................................................................................ iii!
LIST OF FIGURES ........................................................................................................................ vi!
LIST OF TABLES ............................................................................................................................ x!
ABSTRACT ................................................................................................................................... xi!
1! CHAPTER ONE: Introduction ................................................................................................... 1!
1.1! Field-effect transistor (FET) nanobiosensors ...................................................................... 1!
1.1.1! Working Mechanism ................................................................................................... 2!
1.1.2! Device Fabrication ....................................................................................................... 5!
1.1.3! Current research trends in field-effect transistor nanobiosensors ............................... 7!
1.2! Surface functionalization of field effect transistor for nanobiosensors development ....... 10!
1.2.1! Surface functionalization of metal oxide semiconductor .......................................... 10!
1.2.2! Surface functionalization of Si materials ................................................................... 11!
1.2.3! Current research trends in surface functionalization of semiconductor materials ..... 13!
1.3! Toward prototype and commercialization of FET-based Nanobiosensors ....................... 16!
1.3.1! Current commercialized POCT devices .................................................................... 17!
1.3.2! Advantages and challenges of FET-based nanobiosensors as POCT devices ........... 18!
1.4! References ......................................................................................................................... 21!
2! CHAPTER TWO: Bottom-up Fabrication of In
2
O
3
Nanowire Field Effect Transistor
Nanobiosensors ............................................................................................................................... 26!
2.1! Introduction ....................................................................................................................... 26!
2.2! Results and Discussion ..................................................................................................... 27!
2.2.1! Synthesis and characterization of In
2
O
3
nanowires ................................................... 27!
2.2.2! Device fabrication and performances ........................................................................ 30!
2.2.3! pH sensing ................................................................................................................. 31!
2.2.4! Surface functionalization ........................................................................................... 32!
2.2.5! Real time biological molecule detection ................................................................... 34!
2.3! Chapter conclusion ............................................................................................................ 35!
2.4! Experimental section ......................................................................................................... 36!
3! CHAPTER THREE: Top-down Fabricated Polycrystalline Silicon Nanoribbon Field Effect
Transistor Nanobiosensors .............................................................................................................. 41!
3.1! Introduction ....................................................................................................................... 41!
3.2! Results and discussion ...................................................................................................... 43!
3.2.1! Device fabrication and performance .......................................................................... 43!
3.2.2! pH sensing ................................................................................................................. 45!
3.2.3! Surface functionalization of polycrystalline silicon with native oxide layer ............ 47!
3.2.4! Ovarian cancer biomarker sensing ............................................................................ 49!
3.2.5! Surface functionalization of polycrystalline silicon without native oxide layer ....... 50!
3.3! Chapter conclusion ............................................................................................................ 53!
3.4! Experimental section ......................................................................................................... 54!
3.5! References ......................................................................................................................... 57!
4! CHAPTER FOUR: Highly Scalable, Uniform, and Sensitive Biosensors Based on Top-Down
In
2
O
3
Nanoribbons and Electronic Enzyme-Linked Immunosorbent Assay .................................. 59!
4.1! Introduction ....................................................................................................................... 59!
4.2! Results and discussions ..................................................................................................... 61!
! v
4.2.1! Device fabrication and performance .......................................................................... 61!
4.2.2! Statistical analysis of the In
2
O
3
nanoribbon FETs ..................................................... 63!
4.2.3! pH sensing ................................................................................................................. 65!
4.2.4! Surface functionalization ........................................................................................... 66!
4.2.5! Real time detection of HIV1 p24 protein .................................................................. 68!
4.2.6! Electronic ELISA approach for HIV1 p24 protein detection .................................... 69!
4.3! Chapter conclusion ............................................................................................................ 76!
4.4! Experimental section ......................................................................................................... 77!
4.5! References ......................................................................................................................... 81!
5! CHAPTER FIVE: Surface Functionalization of In
2
O
3
Nanoribbon FETs with Histidine
Tagged Biological Molecules For Biosensing Application ............................................................ 83!
5.1! Introduction ....................................................................................................................... 83!
5.2! Results and discussion ...................................................................................................... 84!
5.2.1! Synthesis of N-imidazolylpropyl phosphonic acid molecule .................................... 84!
5.2.2! Surface functionalization of In
2
O
3
with histidine-tagged biomolecules .................... 85!
5.2.3! Streptavidin/Anti-Streptavidin sensing application ................................................... 88!
5.2.4! DNA methylation detection ....................................................................................... 89!
5.3! Chapter conclusion ............................................................................................................ 92!
5.4! Experimental section ......................................................................................................... 93!
5.5! References ......................................................................................................................... 95!
BIBLIOGRAPHY ........................................................................................................................... 97!
!
! !
! vi
LIST OF FIGURES
Figure 1.1 Typical Field-Effect Transistor (FET) sensor structure ................................... 2
!
Figure 1.2 Mechanism to modulate the conductance of a p-type nanomaterial-
based FET (holes as the main charge carriers). The source-drain current of
the channel is monitored against time. In route A, when a negatively charged
target binds to the receptor anchored on the nanomaterial, the charge carriers
will accumulate under the bound analyte, thus causing an increase in the
device conductivity and source-drain current. In route B, the binding of a
positively charged target leads to depletion of charge carriers beneath the
bound analyte, causing a decrease in conductivity and source-drain current. ............ 3
!
Figure 1.3 Simulation of drain current (I
ds
) against gate voltage (V
lg
) curves for
before (black) and after (red) protein attachment on an ambipolar carbon
nanotube FET sensor due to (a) mobility change, (b) dielectric change, and
(c) gate bias.
10
............................................................................................................. 5
!
Figure 1.4 Images of FETs fabricated in our research group. (a) Optical image of
an In
2
O
3
NW FET with interdigitated electrodes. (b) Scanning electron
microscope (SEM) image of the same FET at a higher magnification. (c)
SEM image of a group of 6 poly-silicon nanooribbon FETs (Inset: SEM
image of one poly-silicon nanoribbon FET at a higher magnification). (d)
Optical image of a group of 6 In
2
O
3
nanoribbon FETs (Inset: SEM image of
one In
2
O
3
nanoribbon FET). ....................................................................................... 6
!
Figure1.5 Mono-, bi-, or tridentate anchorage of a phosphonate ligand on a metal
oxide surface ............................................................................................................. 11
!
Figure 1.6 Chemical pathways used to anchor biological molecules to different
nanomaterial surfaces. (a) Si surface coated with native oxide. (b) H-
terminated Si surface functionalized with an Olefin. (c) H-terminated Si
surface functionalized with the chlorination/alkylation method. .............................. 13
!
Figure 1.7 Schematic representation of two pathways for surface click chemistry:
(a) azide terminated surface and (b) acetylene terminated surface. .......................... 14
!
Figure 1.8 Immobilization of His-tagged protein on Ni
2+
:NTA surface. A linear
sequence of 6–12 histidine residues can anchor a protein onto a surface
functionalized with Ni
2+
:NTA motif. Nitrogen lone pairs on each histidine
residue (orange pentagon) coordinate to the remaining two sites on
hexavalent Ni
2+
ions (red spheres) held by the NTA chelator group.
58
.................... 15
!
Figure 1.9 POCT device that consists of a bio-recognition layer on a transducer
attached to an analytical output.
63
............................................................................. 17
!
! vii
Figure 2.1 Scheme of Laser-Ablation Chemical Vapor Deposition System .................... 28
!
Figure 2.2 SEM images of as-synthesized In
2
O
3
nanowires. (a) SiO
2
/Si (100)
substrate and (b) Si (111) substrate ........................................................................... 30
!
Figure 2.3 (a) A photographic image of the fabricated devices in a wafer scale. (b)
I
ds
-V
ds
plots with V
gs
varying from 0.0 V to 0.5 V under liquid gate
condition. (c) I
ds
-V
gs
plot at V
ds
=200mV under liquid gate condition at linear
scale (blue) and log scale (red). ................................................................................ 31
!
Figure 2.4 pH sensing with bare In
2
O
3
nanowire FETs with pH value ranging
from pH=2 to pH=10 ................................................................................................ 32
!
Figure 2.5 (a) Scheme for surface functionalization of probe molecule on In
2
O
3
NW. (b) High-resolution XPS spectrum of P 2p
2/3
region. (c) Fluorescent
image of In
2
O
3
NW matt functionalized with biotin and Alexa Fluro 568-
streptavidin. (d) Fluorescent image of In
2
O
3
NW matt functionalized with
amine-PEG and Alexa Fluro 568-streptavidin. ......................................................... 34
!
Figure 2.6 Real time sensing signal of detection CA-125 protein with In
2
O
3
nanowire FETs biosensors. A non-target protein, BSA, was added to the
sensing media in the end to demonstrate the specificity of the biosensor ................ 35
!
Figure 3.1 (a)-(d) Scheme of top-down fabrication of polySi nanoribbon FETs. (e)
SEM image of a group of 6 polySi nanoribbon devices. (f) A photographic
image of the fabricated devices in a wafer scale. ...................................................... 44
!
Figure 3.2 (a) I
DS
-V
DS
plots under various VGS. (b) I
DS
-V
GS
plots under various
VDS ........................................................................................................................... 45
!
Figure 3.3 (a) Real time pH sensing ranging from pH=4 to pH=9. (b) Small range
of pH sensing from pH=7.2 to pH=8. ....................................................................... 46
!
Figure 3.4 (a) Surface functionalization scheme of polySi with native oxide. (b)
XPS high-resolution spectra of N 1s region. (c) XPS high-resolution of C 1s
region. (d) Stepwise XPS survey spectra on polySi: black trace(bare polySi);
red trace(polySi with silane monolayer); blue trace(polySi with carboxyl
terminal); green trace(polySi functionalized with biotin and treated with
streptavidin). ............................................................................................................. 48
!
Figure 3.5 Real time sensing of CA-125 detection. A non-target protein, BSA,
was added in the end to demonstrate the specificity of the sensors. ......................... 50
!
Figure 3.6 Surface functionalization scheme for polySi without native oxide. ................ 51
!
! viii
Figure 3.7 (a) High-resolution XPS spectrum of Si 2p region after HF treatment.
(b) High-resolution XPS spectrum of C 1s region after Grignard reaction. (c)
High-resolution XPS spectra of Br 3d region before and after bromine
treatment with the vinyl surface. (d) High-resolution XPS spectra of N 1s
region of two different routes creating amine surface. ............................................. 53
!
Figure 4.1 (a)-(d) shows the fabrication steps of In
2
O
3
nanoribbon FETs using
two-step photolithography. (e) A photographic image of a wafer scale
fabricated devices with a zoom-in photo of one single sensor chip. (f) SEM
image of two fabricated In
2
O
3
nanoribbon FETs. ..................................................... 62
!
Figure 4.2 (a) I
DS
-V
DS
family plots under various V
GS
. (b) I
DS
-V
GS
family plots
under various V
DS
. .................................................................................................... 63
!
Figure 4.3 (a)-(d) Statistic study of four FET key parameters: I
ON
(a); V
TH
(b);
gm(c); µ(d). (e)-(f) Comparison of I
ON
distribution for In
2
O
3
nanoribbon (e)
FETs and nanowire (f) FETs ..................................................................................... 65
!
Figure 4.4 Normalized current for pH sensing. (a) Sensing of pH 4 to pH 9 with
sensors of different In
2
O
3
nanoribbon thickness. (b) Normalized current
versus time with a pH step of about 0.3 in pH ranges of biological interest. ........... 66
!
Figure 4.5 (a) Surface functionalization scheme for In
2
O
3
nanoribbon. Fluorescent
images for biotin (b) and amine-PEG (c) as probe molecules after treatment
with Alexa Fluro-568 streptavidin. SEM images of In
2
O
3
nanoribbon for
biotin (d) and amine-PEG (e) as probe molecules after treatment with Au-
streptavidin. ............................................................................................................... 68
!
Figure 4.6 Real time detection of HIV1 p24 protein. PBS addition would cause
signal as shown in the inset. ...................................................................................... 69
!
Figure 4.7 (a) Schematic diagram of streptavidin electronic ELISA. (b)
Normalized real-time responses of 1 µM streptavidin electronic ELISA from
three In
2
O
3
nanoribbon devices monitored simultaneously. (c) Plot of
average normalized current responses and streptavidin concentration
calculated from three devices monitored simultaneously in each
concentration. (d) Plot of pH changes in the sensing chamber measured by a
commercial pH meter and streptavidin concentration. ............................................. 73
!
Figure 4.8 (a) Schematic diagram of electronic ELISA for HIV1 p24 detection.
(b) Real-time responses monitored from 3 In
2
O
3
nanoribbon devices
simultaneously at 20 fg/mL of p24 proteins in PBS. Conduction of all
devices decreased upon presence of urea in the sensing chamber. (c) Plot of
average normalized responses from 3 devices at each p24 concentration and
p24 concentration in pg/mL. (d) Plot of change in pH in the sensing chamber
! ix
measured from a commercial pH meter and p24 protein concentration in
pg/mL. ....................................................................................................................... 76
!
Figure 5.1 Synthetic scheme for N-imidazolylpropyl phosphonic acid. .......................... 85
!
Figure 5.2 Surface functionalization scheme of In
2
O
3
with hisitidine biomolecule. ........ 85
!
Figure 5.3 (a) Ni 2p
2/3
XPS spectra of samples functionalized with the linker
molecule and NiCl
2
(blue trace) and without the linker molecule but only
NiCl
2
(red trace). (b) N 1s XPS spectra of sample functionalized with only
histidine tagged streptavidin (black trace), sample functionalized with only
NiCl
2
and then histidine tagged streptavidin (red trace) and sample
functionalized with the linker molecule, NiCl
2
and histidine tagged
streptavidin. ............................................................................................................... 87
!
Figure 5.4 Fluorescent images of samples treated as the illustrated
functionalization schemes. ........................................................................................ 88
!
Figure 5.5 Real time detection of anti-SA with In
2
O
3
nanoribbon FETs biosensors
functionalized with histidine tagged streptavidin ..................................................... 89
!
Figure 5.6 DNA methylation detection with In
2
O
3
nanoribbon FETs. (a) I
DS
-V
GS
sweeps of In
2
O
3
nanoribbon FETs after adding in different concentration of
methylated DNA strands. Control experiment was performed in the end to
demonstrate the specificity. (b) I
DS
-V
GS
sweeps of sensors when performing
repeated washing after adding 100 nM methylated DNA strands. ........................... 92!
!
! !
! x
LIST OF TABLES
!
Table 1.1 Comparison between bottom-up and top-down fabrication approaches ............ 7
!
! !
! xi
ABSTRACT
Researchers have been developing Point-Of-Care Testing (POCT) devices over
the past decade to potentially substitute the current lab-based disease diagnostic
technologies with improved portability and efficiency. Nanomaterials-based field effect
transistor (FET) biosensors have become the hot spot for such research because of the
small device size comparable to the biological molecules and the instant electrical
response for easy electronic integration. Moreover, nanomaterials have a high surface-to-
volume ratio, which improves the sensitivity of the currently available POCT devices.
My PhD research is to develop FET nanobiosensor platforms to potentially become a
prototype for universal biomarker detection.
Chapter 1 briefly introduces the structure and working mechanism of FET
nanobiosensors. An overview of the surface functionalization approaches is also
included. Finally, the chapter discusses the situation of the current commercial POCT
devices and the potential to commercialize FET nanobiosensors.
Chapter 2 describes our first-generation FET nanobiosensor – In
2
O
3
nanowire
FET biosensor. In this chapter, a laser-ablation chemical vapor deposition system is
applied to synthesize In
2
O
3
nanowires and the FETs are fabricated with photolithography.
The pH sensing is performed to confirm the device sensitivity. Appropriate surface
chemistry is used to functionalize the In
2
O
3
nanowire surface for CA-125 biosensing.
Chapter 3 introduces our second-generation FET nanobiosensor – polycrystalline
Si nanoribbon FET biosensor. In this chapter, top down fabrication technique is
introduced to fabricate the polysilicon nanoribbon FETs. Two routes of surface chemistry
are explored to functionalize the polysilicon surface for biosensing. The pH sensing and
! xii
CA-125 sensing is performed to demonstrate the device sensitivity and the potential to
develop biosensors for other biomarkers.
Chapter 4 describes our newest generation FET nanobiosensor – In
2
O
3
nanoribbon FET biosensor. The sensor is fabricated via top-down technique with nearly
100% wafer-scale device yield and minimal device-to-device variation. A novel
electronic enzyme-linked immunosorbent assay (ELISA) combined with the In
2
O
3
nanoribbon FET biosensor is introduced in this chapter and this combined approach is
able to realize ultrasensitive HIV P24 detection with detection limit comparable to that of
PCR.
In the end, chapter 5 studies a newly designed surface functionalization route on
In
2
O
3
to anchor histidine-tagged biomolecules. This route consistently immobilizes
histidine-tagged biomolecules on the surface of In
2
O
3
for biosensing application, which
was demonstrated with streptavidin antibody sensing and DNA hypermethylation
detection. This is the first time report of biosensing application of FET using histidine
tagged biomolecule. With the prevalence of histidine-tagged biomolecule, this approach
can be a universal route to anchor such biomolecules on the surface for other
applications.
! 1
1 CHAPTER ONE: Introduction
1.1 Field-effect transistor (FET) nanobiosensors
Sensor technology has been an important part of many sectors of society ranging
from agricultural and energy to transportation security and medicine. The explosion of
nanotechnology within the last twenty years has pushed the boundary of response times,
detection limits, sensitivity, portability, etc. for sensor technology, particularly for
chemical and biological sensors. This is partly due to the fact that nanostructures that
have at least one dimension in the range of 1 to 100 nm have comparable sizes as many
of the chemical and biological species of interest, and are thus better for probing the
molecules. Another important feature of the nanostructures is their large surface-to-
volume ratio that allows their material properties to be strongly affected by their
environment. A diversity of sensor architectures has been designed and fabricated during
the past few decades that utilizes different nanomaterials as a sensing element
(cantilevers, quantum dots, nanotubes, NWs, nanobelts, nanogaps, and nanoscale films).
1-
4
Some of these sensing devices, such as those based on cantilevers and quantum dots, are
highly specific, ultrasensitive, and have short response times. However, these devices
require integration with optical components in order to translate surface-binding
phenomena into a readable signal. The need for detection optics is expected to
significantly increase the cost of operation for such a device. In contrast, sensors
designed to operate like FET can directly translate the analyte–surface interaction into a
readable signal, without the need for elaborate optical components. These devices utilize
the electronic properties of the sensing element, such as its conductance, to produce the
signal output. Sensors based on FETs promise to revolutionize bioanalytical research by
! 2
offering the direct, real-time, highly specific, ultrasensitive, and label-free detection of
the desired biomolecule.
5-8
1.1.1 Working Mechanism
An FET sensor has the structure of a common three-electrode transistor, where
the source and drain electrodes bridge the semiconductor channel and the gate electrode
modulates the channel conductance. The typical structure of an FET sensor is illustrated
in Figure 1.1. In the case of FET nanosensors, the semiconductor channel is made of a
nanomaterial and is used as the “sensing” component of the device. Semiconductor
channels can be fabricated using several nanomaterials, including CNTs and NWs. In
order to provide selectivity toward a unique analyte, a specific recognition group (also
called a receptor, ligand, or probe) is anchored to the surface of the semiconductor
channel. This receptor is typically chosen to recognize its target molecule (also called
analyte) with a high degree of both specificity and affinity.
!
Figure 1.1 Typical Field-Effect Transistor (FET) sensor structure
The semiconductor channel has a uniform conductance determined by the main
carrier density in the nanomaterials (holes for a p-type semiconductor or electrons for an
n-type semiconductor). The carrier density is proportional to the conductance of the
channel, which can be determined from the source-drain current of the device. Any
Back Gate
S D
Nanomaterial
Dielectrical layer
Receptor
Target
! 3
change in the current can be related to a change in conductance of the channel. When a
charged analyte molecule binds to a receptor anchored on the nanomaterial, an electric
field created on the surface exerts an effect both inside and outside the semiconductor
channel.
9
If the bound analyte molecule carries a charge opposite to the main carriers in
the FET, then charge carriers will accumulate under the bound analyte, thus causing an
increase in the device conductivity. This mechanism is shown in Route A in Figure 1.2,
where a negatively charged molecule such as DNA binds to the p-type channel, causing a
buildup of hole carriers, thus resulting in an increase in conductivity. In contrast, analytes
with molecular charges same as that of the main carriers in the FET lead to depletion of
main carriers beneath the bound analyte, causing a decrease in conductivity. The latter
case is shown in Route B in Figure 1.2, where a positively charged molecule, such as a
protein below its isoelectric point, depletes the carriers upon binding to the channel.
!
Figure 1.2 Mechanism to modulate the conductance of a p-type nanomaterial-based FET
(holes as the main charge carriers). The source-drain current of the channel is monitored
against time. In route A, when a negatively charged target binds to the receptor anchored
on the nanomaterial, the charge carriers will accumulate under the bound analyte, thus
causing an increase in the device conductivity and source-drain current. In route B, the
S D
+ + + +
Route A
Route B
S D
+ + + + + +
# #
S D
+ +
++
Current
Time
Target0bind
Current
Time
Target0bind
! 4
binding of a positively charged target leads to depletion of charge carriers beneath the
bound analyte, causing a decrease in conductivity and source-drain current.
The mechanism of the conduction changing during molecular binding is also a
debated topic.
10-13
According to the ideal transistor linear (region often used for
biosensing) current equation
!
!"
= !"#!
!
!
!
!
!"
!
(!
!"
−!
!
)
While the transistor dimenions (A, d, and L) and the drain voltage (V
DS
) are constant, a
change in conduction current (I
DS
) can be caused by either a change in mobility (µ), a
change in capacitance due to the difference in the dielectric constant (ε
r
) of the sensing
environment versus the binding molecule, or a gating effect (V
GS
) caused by charges
from the binding molecule. These three situations are illustrated in Figure 1.3 by
comparing the I
ds
–V
lg
curves of an ambipolar FET device before and after protein
binding. Figure 1.3(a) shows that a decrease in the slope of the I
ds
–V
lg
curve after protein
binding also decrease the I
ds
at fixed V
lg
. A change in I
ds
due to the slope indicates a
reduction in mobility and transconductance inside the channel, possibly due to an uneven
electrostatic field distribution caused by random binding with charged biomolecules. In
Figure 1.3(b) the gate bias is shown to be less effective at inducing I
ds
. The current
reduction in this case can be attributed to a reduced gate capacitance caused by the low
permittivity of the bound biomolecule. Finally, Figure 1.3(c) shows an I
ds
change due to
electrostatic gating of the FET channel by charged target biomolecules. This type of
change causes a threshold voltage (V
T
) shift seeing in the figure.
! 5
!
Figure 1.3 Simulation of drain current (I
ds
) against gate voltage (V
lg
) curves for before
(black) and after (red) protein attachment on an ambipolar carbon nanotube FET sensor
due to (a) mobility change, (b) dielectric change, and (c) gate bias.
10
1.1.2 Device Fabrication
Once the nanomaterials have been prepared, the source, drain, and gate electrodes
are deposited to complete the structure of the FET. Our research group has been using Si
substrate as the back gate electrode. In the case of bottom-up NWs, the NWs are
randomly dispersed on the substrate and metal source and drain electrodes are deposited
on the insulating layer (SiO
2
of 500 nm) on top of the NWs to define the channel length
and width of the FET (shown in Figure 1.4(a) and 1.4(b)). It has been reported that the
device dimensionality directly affects the response time
14
and the sensitivity
15
of sensors.
A common channel length is on the order of 2-10 µm. In the case of top-down poly-
silicon and In
2
O
3
nanoribbons, ribbons with leads are patterned with uniform width and
length at designated locations and metal electrodes are deposited to create electrical
connections (Shown in Figure 1.4(c) and 1.4(d)).
! 6
!
Figure 1.4 Images of FETs fabricated in our research group. (a) Optical image of an
In
2
O
3
NW FET with interdigitated electrodes. (b) Scanning electron microscope (SEM)
image of the same FET at a higher magnification. (c) SEM image of a group of 6 poly-
silicon nanooribbon FETs (Inset: SEM image of one poly-silicon nanoribbon FET at a
higher magnification). (d) Optical image of a group of 6 In
2
O
3
nanoribbon FETs (Inset:
SEM image of one In
2
O
3
nanoribbon FET).
The comparisons between bottom-up and top-down fabrication techniques are
listed in the table below (Table 1.1). The bottom-up fabrication utilizes the random
assembly of the nanomaterials, mostly one-dimensional (1D) nanomaterials, and
electrodes are patterned on top of that. Interdigitated electrodes were used in our case of
In
2
O
3
NW FETs in order to increase the chance to bridge two electrodes with the
nanomaterials (Figure 1.4(a)). The top-down fabrication controls the position and the
dimensions of the nanomaterials used in FET channel and electrodes are defined based on
the location of the channel materials. The difference in controllability between two
fabrication techniques results in the significant difference in device yields and uniformity.
Because of the randomness of nanomaterials position for bottom-up fabrication, a wafer
scale device yield can only get as high as 74% in our In
2
O
3
NW FETs fabrication, with
50 µm
(a)
(b)
(c) (d)
In2O3 nanowire
In2O3 nanoribbon
4.µm
15.µm
! 7
noticeable device-to-device variation.
16
On the other hand, the top-down fabrication can
achieve almost 100% wafer scale device yield with minimal device-to-device variation,
thanks to the good controllability over not only the channel position but also the precise
dimensions.
17
Both fabrication techniques just apply the conventional photolithography,
which will keep the process low-cost and easy to handle.
Bottom-up Fabrication Top-down Fabrication
Controllability Low, random assembly High, precise control
Device-to-device variation High Low
Wafer-scale device yield As high as 74% As high as 100%
Fabrication procedures Easy, photolithography Easy, photolithography
Table 1.1 Comparison between bottom-up and top-down fabrication approaches
1.1.3 Current research trends in field-effect transistor nanobiosensors
Massive amount of researches have devoted to field-effect transistor
nanobiosensors over the past decade. Efforts have been made to produce ultrahigh
sensitivity, great specificity and minimal sample preparation process in order to apply
such nanobiosensors for Point-Of-Care (POC) settings. Currently there are several
popular research directions in this field to facilitate the widespread adoption of such
technology.
Device structure engineering Scientists and engineers are trying to create novel
device structure to achieve higher sensitivity and different functionality for the FET
nanobiosensors. Traditionally, an FET consists of three electrodes and one
semiconducting channel. Ahn, et. al., have added a secondary gate electrode to improve
the sensitivity of their FET nanobiosensors.
18
By means of the secondary gate, it is!easy
! 8
to control the carrier conduction paths, which critically affect the device parameters such
as the subthreshold slope, threshold voltage, and drain current. It was experimentally
observed by antibody-antigen interaction and theoretically supported by the
commercialized simulator that the nanowire structure with the double gate showed
improved sensitivity compared to that with a conventional single gate.
Multiplex sensing Due to the complexity of biological systems, especially the
human body, a single biomarker alone is not effective enough for accurate diagnosis.
Medical decision based on single biomarker usually has a high possibility of false
positive and false negative. Recent research shows that combination of multiple
biomarkers generates improved accuracy compared to single biomarker.
19, 20
This fact
brings up the importance of multiplexing assay of biomarkers. An ideal biosensing
technology should be capable of simultaneous detection of a combination of biomarkers.
To construct sensor arrays for multiplexed biosensing, the sensors must be
selectively functionalized with different capturing probes against their designated
analytes. Efforts have been made to achieve the selective functionalization of
nanomaterial-based devices, by using microfluidic chips, microspotting techniques
7, 21
and
electroactive monolayers.
22, 23
Lieber, et. al., have demonstrated the highest level of
multiplexing to date for NW-FET sensors in a simultaneous assay for three cancer
markers with a detection of 0.9 pg/mL in desalted but undiluted serum samples.
7
Monoclonal antibodies specific for each of the targets were spotted onto different NW-
FETs. Samples solutions were delivered through microfluidic channels and the electrical
signal from each FET was monitored in real time during exposure to each of the targets.
Compared to the microspotting technique, the use of electroactive monolayers possesses
! 9
a unique advantage because it is only limited by the ability to electronically address the
individual sensors.
23
In this method, the key step is to design a bifunctional molecule,
which bears a NW-anchoring group on one end and an electroactive moiety on the other.
After being covalently linked to the NWs, the molecule can be activated from the
chemically inert “OFF” state to its “ON” state by applying an external voltage to device
electrodes. In the “ON” state the electroactive moiety reacts with the desired capture
probe to covalently anchor it to the NW surface.
24
Physiological samples With today’s nanosensors, researchers claim that they are
able to detect proteins and DNAs down to femtomolar or even attomolar range with good
selectivity.
7, 25-28
However, these detections are performed in purified buffers with very
low ionic strengths. When it comes to clinical diagnosis, the sensitivity and selectivity of
a biosensor will be significantly suppressed due to the complexity of the sample
composition. Efforts have been made to address this problem by sample purifications and
novel surface modification approaches.
Mark Reed, et. al., have reported biomarker detection from whole blood samples
purified by a microfluidic purification chip (MPC).
29
The biomarkers spiked in a whole
blood sample were captured by the antibody-modified MPC and the antibody/antigen
complexes were released into 0.01X PBS buffer. The complex solution was then
delivered to Si NW-based sensors functionalized with a secondary antibody to perform
sensing. This study marks the first use of label-free nanosensors with physiological
solutions.
In order to overcome the complication caused by physiological samples, our
group developed a faster approach without requiring extra process for device fabrication.
! 10
By passivating the In
2
O
3
NW surface with tween-20, we successfully blocked the signal
induced by nonspecific binding when performing active measurement in whole blood.
30
The detection limit of tween-20 passivated sensors for biomarkers in whole blood was
enhanced to the level similar to the detection limit for the same analyte in purified buffer
solutions at the same ionic strength, suggesting minimal decrease in device performance
in the complex media.
1.2 Surface functionalization of field effect transistor for nanobiosensors
development
An as-fabricated FET device will not have the desired molecular recognition
properties. The surface of the sensing element (semiconductor nanomaterial) needs to be
modified so that the device acquires specific recognition toward a desired analyte. This
selectivity is typically achieved by anchoring a specific recognition group to the surface
of nanomaterials. A bifunctional linker molecule with two chemically different termini is
used to help anchor the receptor molecules to the nanomaterial surface. In my doctoral
research, I focused on the surface functionalization of metal oxide and Si materials.
1.2.1 Surface functionalization of metal oxide semiconductor
Metal oxide surface can be functionalized with a linker molecule that bears a
functional group capable of forming a nonhydrolizable conjugate, such as phosphonate or
siloxide. Phosphonic acids are found to bind strongly on the surface of In
2
O
3
and ITO.
28, 31
Silane molecules have been applied to functionalize ZnO and Fe
3
O
4
surfaces.
32, 33
Also
carboxylic acids, especially fatty acids, have been used to functionalize TiO
2
nanoparticle
surface.
34
The optimum linker molecule was found to be a phosphonate derivative, like 3-
! 11
phosphonopropanoic acid.
22
This phosphonate spontaneously self assembles on the
nanowires from aqueous solutions or polar solvents. A major feature of the attachment of
a phosphonate group on a metal oxide surface is that the anchorage can be mono-, bi-, or
tridentate (Figure 1.5). For example, a recent investigation of the binding of (11-
hydroxyundecyl)phosphonic acid or (12-carboxy-dodecyl)phosphonic acid on a SnO
2
surface by solid state
31
P NMR showed a bi- and tridentate attachment of phosphonate
ligands.
35
The multidentate attachment is another stabilizing factor for the modified
nanoparticles.
36, 37
In the case of the bifunctional (12-carboxy-dodecyl)phosphonic acid, it
is interesting to note that the phosphonate group and not the carboxylate group was
bonded to the stannia surface, which proves the phophonate group is more favored to
form the covalent bond compared with the carboxylate group.
!
Figure1.5 Mono-, bi-, or tridentate anchorage of a phosphonate ligand on a metal oxide
surface
1.2.2 Surface functionalization of Si materials
Si surface forms a thin layer (approximately 2 nm) of SiO
2
because of the
oxidation process when exposing the material in the air. The surface functionalization
schemes are dependent whether the surface oxide layer is removed or not. Alkoxysilane
derivatives, such as 3-(trimethoxysilyl)propyl aldehyde , 3-aminopropyltriethoxysilane
and 3-aminopropyldimethylethoxysilane are the most widely used linkers for the Si
P
O
O
HO
P
O
O
O
P
O
O
O
! 12
surface with the native oxide layer.
7, 12, 26, 38
The Si-methoxide or Si-ethoxide reacts with
the surface OH group, anchoring the linker molecule to the silicon oxide surface and
creating a monolayer terminated with aldehyde or amine groups. These groups can then
react with amine or carboxylic acid groups that are commonly present in biological
capture probes. As for Si surfaces without the native oxide layer, two methods have been
employed to functionalize the surface for further bioconjugation. Several research groups
use UV light to rapidly photo dissociate the Si-H bond to generate radical species on the
Si surface (Figure 1.6). These radicals can subsequently react with terminal olefin groups
on linker molecules, thus forming stable Si-C bonds at the Si surface.
27, 39, 40
The linker
molecules usually carry a protected amine terminal,! which can be used to attach
biological probes after deprotection. The other method, developed by Nathan Lewis, uses
a two-step chlorination/alkylation reaction to form Si-C bond on the surface.
41, 42
The Si-
H surface is first chlorinated to form Si-Cl bond and then the surface was treated by an
allyl Grignard. The resulted allyl surface can be used for further bioconjugation.
43
! 13
!
Figure 1.6 Chemical pathways used to anchor biological molecules to different
nanomaterial surfaces. (a) Si surface coated with native oxide. (b) H-terminated Si
surface functionalized with an Olefin. (c) H-terminated Si surface functionalized with the
chlorination/alkylation method.
1.2.3 Current research trends in surface functionalization of semiconductor materials
Scientists are trying to engineer the surface of materials with reactions that are
normally carried out in liquid phase. Organic synthesis has endowed researchers with a
significant amount of reactions to work with. Surface chemists start to involve some of
the “star” reactions that are highly yielding for surface functionalization.
Click Chemistry The click chemistry approach has attracted significant attention
in various parts of chemistry, such as in materials science and polymer chemistry during
the past years,
44, 45
since it was introduced in 2001 by Sharpless.
46
The concept addresses
several criteria. The reaction has to be modular and wide in scope, provides furthermore
very high yields, generates inoffensive byproducts, is stereospecific, can be performed
OH OH OH
Si
O
MeO
OMe
OMe O O O
Si
O
O O O
Si
HN
H H H
t-Boc
HN
H
HN
t-Boc
H
H
2
N
H
HN
O
H
HN
HN
H H H PCl
5
Cl Cl Cl
N N
N
! 14
under mild reaction conditions and with easily available starting materials.
46
The
purification can be preferably achieved by nonchromatographic methods. The most
perfect click reaction up to date is the Huisgen 1,3-dipolar cycloaddition of organic
azides and acetylenes.
47
Thereby, a mixture of 1,4- and 1,5-disubstituted 1,2,3-triazole
systems is formed. A variant of this reaction is the copper catalyzed coupling of azides
and terminal acetylenes, which selectively results in the formation of the 1,4-disubstituted
triazole.
48, 49
The general characteristics of the click chemistry approach fit very well with
the requirements of chemical reactions performed on surfaces. This is documented by a
large number of literature examples, which demonstrate the use of click chemistry for the
introduction of functional groups into the monolayer system on different substrates, e.g.
gold, silicon and glass.
50-52
Thereby, two synthetic preparation methods are used to
introduce 1,2,3-triazole moieties into the monolayer. Whereas the first method uses azide
terminated substrates for the coupling with functional acetylenes (Figure 1.7(a)), the
second modification sequence involves the generation of surfaces with terminal acetylene
moieties (Figure 1.7(b)).
!
Figure 1.7 Schematic representation of two pathways for surface click chemistry: (a)
azide terminated surface and (b) acetylene terminated surface.
H H H PCl
5
Cl Cl Cl
N N
N
OH OH OH
Si
MeO
OMe
OMe O O O
Si
Br
Br
O O O
Si
N
3
NaN
3
O O O
Si
N
N
N (a)
(b)
! 15
Binding biological molecules with affinity tag Traditional method to anchor
biomolecule is to use one of the functional groups that are intrinsic to the molecule, e.g.,
amine group or carboxyl group, etc., to couple with surface functional groups.
22, 28, 30
This
approach is very universal because of the functional groups used are present in most of
the biomolecules. However, there are possibilities that the functionality being altered
during the process. In particular, it has been demonstrated that the direct interface
between solid surfaces and proteins can affect the catalytic efficiency of an enzyme
towards its substrate
53
or antibody–antigen recognition.
54
Therefore, applying the affinity
tag to bind biomolecules on the surface appears to be an attractive strategy. The affinity
tag allows the biomolecules immobilized specifically on the substrate with a particular
orientation.
55
Histidine tag, commonly used for protein purification,
56, 57
allows a stable,
non-covalent interaction between pairs of histidine residues and a divalent metal ion such
as nickel (Ni
2+
), which is anchored to a surface through a chelator such as nitrilotriacetic
acid (NTA) (Figure 1.8
58
). The approach has been extensively applied to surface
functionalization of quantum dots,
59
Si materials,
58
gold substrate for SPR sensing,
60
magnetic particles for efficient protein purification,
61
etc.
!
Figure 1.8 Immobilization of His-tagged protein on Ni
2+
:NTA surface. A linear sequence
of 6–12 histidine residues can anchor a protein onto a surface functionalized with
Ni
2+
:NTA motif. Nitrogen lone pairs on each histidine residue (orange pentagon)
coordinate to the remaining two sites on hexavalent Ni
2+
ions (red spheres) held by the
NTA chelator group.
58
! 16
1.3 Toward prototype and commercialization of FET-based Nanobiosensors
The existing health care system is focused on treating diseases rather than
preventing them. Patients are generally not tested until physiological symptoms are
present. When they do get tested, the results often take several days and can be
inconclusive if the disease is at an early stage. In order to facilitate the diagnostics
process and make tests more readily available for patients, the concept of “point of care
testing” (POCT) has been brought up and developed in recent years.
62
POCT is defined
as the medical testing at or near the site of patient care. The major driving force behind
POCT is to provide convenient tests to patients and deliver timely results to the health
care team. This increases the chances that a disease is diagnosed at its early stage and
allows immediate clinical decisions to be made.
POCT is often accomplished through the use of low cost, portable, handheld
testing devices and kits. They are expected to make rapid and precise diagnosis using
only small amount of samples (e.g. a drop of blood for glucose test). A typical POCT
device normally contains three components (Figure 1.9):
63
(1) a biological recognition
element (e.g. an antibody), (2) a material or matrix that transduces the bio-recognition
events into detectable signals and (3) a detector. For a POCT device to be workable, there
are several requirements that need to be satisfied.
64
First, an easy-to-construct and reliable
interface is required to output reproducible signals. Second, the detection method needs
to be highly sensitive and specific. Third, the sensing element (transducing material) is
ideally to be cost-efficient and miniaturized. Finally, due to the complexity of biological
systems, especially the human body, clinical diagnosis usually requires simultaneous
detection of multiple disease markers in order to accurate results. Therefore, the next
! 17
generation of POCT devices will require the capability to be operated in a multiplexed
manner.
!
Figure 1.9 POCT device that consists of a bio-recognition layer on a transducer attached
to an analytical output.
63
1.3.1 Current commercialized POCT devices
The most successful POCT device on the market now is the glucose meter. It has
been developed for more than 50 years and was commercialized in the 1980s. The current
glucose meter delivers accurate, rapid test result with minimum sample volume and
simple procedures. However, glucose meter only detects one substrate and thus lacks the
versatility for a broader range of other substrates. Since it is a perfect platform for
accurate and rapid test, research has been around applying such a platform for a more
general spectrum of biomarkers. Xiang and Lu combined the glucose meter with a
separate DNA sensor and successfully extended the glucose meter to detect a variety of
target molecules, with decent detection limits and dynamic ranges.
65
Another POCT platform is the lateral-flow testing strip. The widely used
pregnancy test strip is based on such a platform. Currently on the market, the later-flow
strips are developed for a large variety of biomarkers. Although this platform is able to
deliver rapid qualitative test results, the relatively high detection limit and false positive
Biorecognition Events Transduction Data Processing
! 18
rate make the conventional lab test still a must for a more confirmative result. Moreover,
a much more complicated technology is required to conjugate with the lateral-flow assay
in order to obtain quantitative test results.
66
Examples are to use spectroscopy to read the
intensity of the sample colored line on the strip, which is similar to the ELISA process.
Theranos
TM
, a bay area-based biotech company, starts to provide services to detect
a large variety of analytes with only a finger prick of blood. The company is currently
pairing with doctors to deliver test results for certain analytes within hours, instead of
days for conventional test turn-around time. The analytes cover a large number of protein
biomarkers, different chemical elements, small molecules and blood cells. More
importantly, the company has started to work with Walgreens to bring the testing service
in Walgreens store for a more convenient experience for patients. Database will be
established for a certain patient to monitor one or several specific biomarkers chronically,
providing physicians a closer track of the health condition of the patient.
67
The
technology of the company, though not disclosed on their website, is mainly optical
sensing and ELISA-type sensing technology based on several of their issued patents.
68
They have engineered the sample delivery system and sensing assembly for faster testing
process and smaller sample volume.
69
So far, Theranos
TM
is the most successful company
to provide POCT services on the market and with the establishment of their Theranos
TM
Wellness Center in Walgreens, this service will surely become more prevalent and the
development of POCT devices will be even more demanding.
1.3.2 Advantages and challenges of FET-based nanobiosensors as POCT devices
POCT devices require rapid and accurate test results from minimum sample
volume and easy sample handling without well-trained personnel. FET is potentially a
! 19
favorable platform to develop reliable POCT devices. The fast response of the electrical
signal induced by the external electrical field on the transistor is instant, which is very
important for repaid sensing result delivery. Furthermore, the electrical signal can be
easily integrated with other electronics components for signal processing and readout.
Similar to the glucose meter, the use of electrical signal will enhance the portability for
the application of the FETs as POCT devices. In our FET nanobiosensors, we use
nanomaterials for the semiconductor channel. With the help of nanoscale size of the
channel materials, the high surface-area-volume (S/V) ratio will significantly improve the
sensitivity. A direct consequence of this high S/V ratio is that a large fraction of the
atoms in the material are located at or near the surface. This proximity causes the surface
atoms to play an important role in determining the physical, chemical, and particularly
electronic properties of the nanomaterials. This dependence on the properties of the
nanomaterial/surrounding interface makes nanomaterials highly sensitive in molecular
sensing applications. The small size of these nanomaterials is another important feature
that makes them ideal candidates for POCT devices. They are comparable in size with
most biological entities, such as proteins, nucleic acids, cells,!viruses, etc., making them
the ideal interface materials between biological molecules and scientific instruments.
Also, their extreme smallness would allow packing a huge number of sensing elements
into a small chip of an array device, which can be used in multiplexed sensing of a panel
of disease markers.
Although the advantages of the FET nanobiosensors are attractive, development
of such devices into commericalizable POCT devices are still challenging. Several
! 20
important issues need to be well explored and addressed before potentially
commercializing the technology.
Device fabrication cost One important factor for commercializing any
technology is the cost-efficiency. Device fabrication can be quite costly if the materials
are difficult to obtain and the processes are too complicated. In order to control the
fabrication cost, the semiconductor materials used need to be abundant materials or
materials easy to synthesize, e.g., Si materials, ZnO, In
2
O
3
, etc. During the fabrication
process, the most conventional photolithography is optimal because of the low-cost of
materials and simple process. However, photolithography can only define the dimensions
at the micrometer scale. Therefore, a good design of the device structure is desired to
relax the dimension requirement and still maintain the same nanoscale characteristics.
Moreover, large-scale fabrication capability is also a key feature to further reduce the
fabrication cost. And this can also be fulfilled by applying the CMOS-compatible
photolithography process during the fabrication.
Consistency A good product requires delivering consistent testing results under
any circumstances. Large-scale fabricated transistors need to maintain very similar if not
identical electrical performances among different devices. The low device-to-device
variation is one of the most important aspects to consider when designing the fabrication
process of the transistors. Low device-to-device variation also ensures the high device
yield. The almost 100% device yield saves time and labor for additional device screening
process before actually packaging the final product. Reliable and efficient surface
functionalization scheme for the semiconductor channel is another important feature to
provide testing result consistency.
! 21
After scrutinizing the challenges to further commercialize the FET
nanobiosensors technology, I devoted myself to conquering the issues during my PhD
study along with my colleagues in Department of Electrical Engineering. From the first
generation of In
2
O
3
nanowire devices to the latest generation of In
2
O
3
nanoribbon
devices, the journey of developing FET nanobiosensors is summarized in the following
four chapters, which are focused at device fabrication, semiconductor materials selection,
surface chemistry development and sensing sensitivity enhancement.
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!
! 26
2 CHAPTER TWO: Bottom-up Fabrication of In
2
O
3
Nanowire Field
Effect Transistor Nanobiosensors
2.1 Introduction
One-dimensional (1D) nanomaterial has been widely used for nanotechnology
applications in sensors,
1
electronics
2
and medical devices
3
because of its high aspect
ratio, small diameter and large surface-to-volume ratio. The synthesis of 1D
nanomaterials could be achieved via hydrothermal/solvothermal synthesis,
4
template
deposition,
5, 6
chemical vapor deposition (CVD) synthesis,
7
etc., among which CVD has a
better control over the dimensions of the nanowires (NWs) and gives better yield. Thus,
vapor deposition has become the most common method for synthesizing carbon
nanotubes,
8, 9
silicon NWs
10, 11
and metal oxide NWs.
12, 13
The application of 1D
nanomaterials in biosensors is of great interest because of the high surface-to-volume
ratio and size similarity between nanowires diameter and biomolecules. These advantages
provide the opportunities to achieve biosensors with ultrahigh sensitivity that could
compete with the current state-of-art diagnostic techniques.
In
2
O
3
is a wide band gap semiconducting material that has been studied to work
as gas sensors,
6, 14
optical sensors
6
and photocatalyst.
15
In
2
O
3
NWs are mostly
synthesized via vapor deposition methods. There are few studies focusing the synthesis of
In
2
O
3
NWs with solvothermal method or sol-gel synthesis. One example of the
solvothermal synthesis of In
2
O
3
used ether as the solvent at 200 °C creating an ultrahigh
pressurized vessel, which potentially would be dangerous to work with;
16
on the other
hand, the sol-gel template method yielded polycrystalline In
2
O
3
NWs,
17
and the grain
! 27
bounderies were not preferred to the transportation of charge carriers, which in turn will
affect the transistor performance.
Here in this chapter, we described a laser-ablation CVD system for In
2
O
3
NW
synthesis with high aspect ratio and decent yield. Using the as-synthesized In
2
O
3
NWs,
we fabricated an array of In
2
O
3
NWs field-effect transistors (FETs) with good transistor
performances. By applying appropriate surface chemistry on the In
2
O
3
NWs, we
established the interface between biological molecules and the electrical devices, from
which nanobiosensors could be achieved. In this chapter, we transformed the In
2
O
3
NWs
FETs into nanobiosensors for ovarian cancer diagnosis, detecting protein CA-125. The
detection limit was 1 U/mL, which is approximately one order of magnitude lower than
the clinical relevant level, 35 U/mL.
18
2.2 Results and Discussion
2.2.1 Synthesis and characterization of In
2
O
3
nanowires
Chemical vapor deposition (CVD) method for nanowire growth usually requires a
thin layer of Au nanoparticles as catalyst for confining nanowire diameter and initiating
the nanowire growth. Vapor-liquid-solid (VLS) mechanism of controlling nanowire
diameter distribution is applied in this method. In this project, we employed laser-
ablation CVD system to synthesize the In
2
O
3
NWs, as shown in Figure 2.1. We first
coated the surface of the selected substrate for NWs growth with poly-Lysine, following
with drop casting a thin layer of 10 nm Au nanoparticle gel solution. Au nanoparticles
were bound on the substrate with the help of poly-L-Lysine, and the uncovered poly-
Lysine was removed with O
2
plasma. The Au nanoparticle covered substrate was then put
at the downstream end of a tube furnace where the CVD took place. Indium vapor was
! 28
supplied from laser ablating on InAs target, which was placed at upper stream of the tube
furnace. When the vaporized indium reached the catalyst nanoparticles, the gold and
indium fused forming an In/Au alloy. Continuous feeding of such alloy with indium
caused the molten solution to reach saturation; and each crystalline indium oxide grew in
only one direction from the catalyst nanoparticle. The furnace was purged with Ar gas to
decrease the O
2
level before starting the laser ablation. Usually the O
2
level was
controlled under 40 ppm to create enough oxygen vacancy for the synthesized In
2
O
3
NWs to exhibit better semiconducting characteristic. During the laser ablation process,
the temperature of the furnace was maintained at 790 °C and the growth process normally
lasted for 50 min.
!
Figure 2.1 Scheme of laser-ablation chemical vapor deposition system.
Figure 2.2 shows the scanning electron microscopy (SEM) images of the as-
synthesized In
2
O
3
NWs on SiO
2
/Si (100) substrate and Si (111) substrate, respectively.
Our initial experiment started with SiO
2
/Si (100) substrate and the yield of In
2
O
3
nanowires was quite poor, as shown from the SEM images. There are short nanowires
InAs target
2”
10 nm Au
Indium vapor
High Temp
40 ppm O
2
Au-In Alloy Droplet
Ar
In
2
O
3
Nanowire
Au
! 29
(less than 2 µm) and some ribbon-like particles. However, after we switched our substrate
to Si (111) substrate, with the same parameters during nanowire synthesis, the length and
yield of In
2
O
3
NWs increased significantly. Note that we did not perform any
etching/cleaning of SiO
2
prior to the substrate preparation for the Si (111) wafer, thus
there was still native SiO
2
thin layer on the substrate surface. Therefore, the surfaces of
both substrates are amorphous SiO
2
, but with different thickness, 20 Å for Si (111)
substrate and 500 nm for SiO
2
/Si (100) substrate. Because the main purpose of this part
of the research is to synthesize In
2
O
3
NWs with high purity and yield, we did not look
into the reason why the difference of nanowires growth emerged. Our assumption was
that the surface hydrophilicity differed between two substrates and thus the nanowire
growth condition varied in each case. More experiments are desired to confirm this
assumption including the contact angle measurement and the surface analysis with XPS
and/or AFM. Previous study using the same synthesis parameters revealed the cubic
phase of the synthesized In
2
O
3
with XRD pattern and growth direction as [110] with
TEM and SAED pattern.
19
As shown in Figure 2.2(b), the diameter of the synthesized
In
2
O
3
nanowires on Si (111) substrates is well controlled to be around 10 nm, with a
narrow distribution. This was achieved by using monodispersed Au nanoparticles as the
catalyst. The idea is based on the VLS growth mechanism, where the In vapor first
diffuses into the gold catalytic particles, and grows out and reacts with O
2
to form In
2
O
3
once the In/Au alloy reaches supersaturation. Continued addition of In into the In/Au
nanoparticle feeds the In
2
O
3
growth and eventually the diameter of the In
2
O
3
nanowire is
directly linked to the catalytic particle size.
! 30
!
Figure 2.2 SEM images of as-synthesized In
2
O
3
nanowires. (a) SiO
2
/Si (100) substrate
and (b) Si (111) substrate.
2.2.2 Device fabrication and performances
In
2
O
3
nanowire field effect transistors are fabricated based-on a bottom-up
fabrication technique. The as-synthesized nanowires were first dispersed in isopropyl
alcohol with sonication and a SiO
2
/Si wafer was drop casted by the nanowires suspension.
In this step, we randomly distributed the nanowires on the surface of the wafer without
any location control. Then the metal electrodes were patterned on the wafer with one step
of photolithography followed by electron-beam evaporator deposition of 5 nm of Ti and
45 nm of Au. After the metal deposition, we applied PECVD for depositing a thin layer
of Si
3
N
4
on a selected area of metal contacts to prevent unnecessary current leakage
during sensing. The fabrication was finished by a lift-off process, leaving the fabricated
wafer as shown in Figure 2.3(a).
The as-synthesized In
2
O
3
NWs exhibit n-type semiconductor behaviors due to the
oxygen vacancies created during the synthesis. Figure 2.3(b) shows one regular In
2
O
3
NW FETs performance of the source-drain current vs. source-drain voltage (I
DS
-V
DS
) plot
and Figure 2.3(c) shows the source-drain current vs. gate voltage (I
DS
-V
GS
) plot under
liquid gate condition. From the I
DS
-V
GS
curve, the device got to be turned on at
(a)
(b)
! 31
V
TH
=180mV and the on/off current ratio is around 10
3
. The curve agrees with the n-type
semiconductor FETs behavior.
!
Figure 2.3 (a) A photographic image of the fabricated devices in a wafer scale. (b) I
ds
-V
ds
plots with V
gs
varying from 0.0 V to 0.5 V under liquid gate condition. (c) I
ds
-V
gs
plot at
V
ds
=200mV under liquid gate condition at linear scale (blue) and log scale (red).
2.2.3 pH sensing
In order to investigate the ion sensitivity of the In
2
O
3
NW FETs, we conducted
pH sensing on the as-fabricated devices. A mixing cell was mounted on top of our In
2
O
3
nanowire chip to confine the sensing solution at the channel area. Then we manually
exchanged different buffer solutions varying in pH in and out of the sensing chamber
using pipettes. The phosphate buffer solutions were made fresh each time we performed
the sensing experiments and the ionic strength were maintained at the same level for all
pH values so that the sensing signal would only be attributed to the change of pH. Figure
2.4 shows the real-time sensing response of a bare In
2
O
3
NW device to various pH buffer
solutions ranging from pH 2 to pH 10. The device shows an increase of conduction with
decreased pH, which agrees to the gate voltage modulation behavior of n-channel
transistors. In
2
O
3
nanowire surface is covered with hydroxyl groups exposed to the buffer
solution. Decreasing of the pH value will enhance the protonation of the surface hydroxyl
groups, and the positive charges on the surface will induce an accumulation of charge
V
g
= +0.5
V
(a)
(b) (c)
! 32
carriers in the In
2
O
3
channel. The charge carriers’ accumulation will generate an increase
of conduction as shown in the sensing signal. The device shows a constant increasing of
conductance over a wide dynamic range (from pH 2 to pH 10) and good reversibility
when changing back to higher pH solutions.
!
Figure 2.4 pH sensing with bare In
2
O
3
nanowire FETs with pH value ranging from pH=2
to pH=10
2.2.4 Surface functionalization
Surface functionalization with biomolecules is required to utilize the FETs as
nanobiosensors. A general scheme of the surface functionalization is shown in Figure
2.5(a). Phosphonic acid has been widely used as interface modifiers for ITO and other
metal oxides. We employed a phosphonic acid linker molecule with a carboxyl terminal
group, yielding a functional group that can be further coupled with biological molecules.
An overnight incubation of nanowires in a diluted phosphonic acid solution (0.1 mM)
followed by an overnight annealing at 120 °C under N
2
yielded a self-assembly
monolayer on the nanowire surface. Figure 2.5(b) shows the high resolution XPS
6
5
4
3
2
3
4
5
6
7
8
9
10
9
8
7
! 33
spectrum of P 2p
2/3
region for the sample after monolayer functionalization. The peak
position at 134 eV agrees with the peak position for (PO
3
)
-
group.
20
We further confirmed
the surface chemistry with contact angle measurement. Two samples were prepared
simultaneously, one of which was treated with 3-phosphonopropanoic acid (contact angle
65 º) and the other was treated with octylphosphonic acid (contact angle 105 º). The
ability to create hydrophilic and hydrophobic surfaces indicates that our surface
modification strategy is effective.
We chose the biotin/streptavidin as the model probe/target molecules pair for our
surface functionalization investigation. Amine-biotin was coupled with the carboxyl acid
group after the monolayer formation using the EDC/NHS coupling. A fluorescent-labeled
streptavidin protein solution was then introduced. For the control experiment, we used
amine-poly(ethylene glycol) (PEG) instead of amine-biotin as the probe molecule. When
we examined the surface of both samples under fluorescent microscope, only the one
functionalized with amine-biotin showed bright image, as shown in Figure 2.5(c and d),
indicating the success of the biotin anchoring on the surface.
! 34
!
Figure 2.5 (a) Scheme for surface functionalization of probe molecule on In
2
O
3
NW. (b)
High-resolution XPS spectrum of P 2p
2/3
region. (c) Fluorescent image of In
2
O
3
NW matt
functionalized with biotin and Alexa Fluro 568-streptavidin. (d) Fluorescent image of
In
2
O
3
NW matt functionalized with amine-PEG and Alexa Fluro 568-streptavidin.
2.2.5 Real time biological molecule detection
The demonstration of the potential usage of In
2
O
3
nanowire FETs as
nanobiosensors was fulfilled with the detection of CA-125 protein. CA-125 is one of the
biomarkers for ovarian cancer and the abnormally increasing level of CA-125 in female
serum indicates possible ovarian cancer. The devices were functionalized with CA-125
antibody to specifically recognize the target protein. During the sensing experiment, a
200 mV V
DS
and a 150 mV V
GS
were applied to the device and the current between
source and drain electrodes was monitored. CA-125 protein solutions with different
concentrations were added progressively into the sensing chamber.
The normalized current versus time for three monitored devices is plotted in
Figure 2.6. The device shows almost no response with small concentrations (0.1 U/mL),
while started to show decrease of current at CA-125 concentration of 1 U/mL and
In
2
O
3
nanowire
OH OH OH P
HOOC
O
HO OH
In
2
O
3
nanowire
O O O
P
HOOC
In
2
O
3
nanowire
O O O
P
C
H
2
N
NHS/EDC
NH
O
(a)
(b) (c)
(d)
! 35
progressively larger response with higher concentrations of CA-125 added. The detection
limit of In
2
O
3
nanowire FETs for CA-125 is 1 U/mL and this level is at least one order of
magnitude lower than the clinical relevant level for diagnosis (35 U/mL).
18
In order to
show the specificity of our nanobiosensors, a control experiment was performed with
bovine serum albumin (BSA). Device shows no responses to 150 nM of BSA, a
concentration that is much higher than the target molecules. Therefore, our In
2
O
3
nanowire FETs biosensor shows decent sensitivity and specificity toward the detection of
CA-125 protein and can be potentially used as nanobiosensors for other cancer or
infectious disease diagnosis.
!
Figure 2.6 Real time sensing signal of detection CA-125 protein with In
2
O
3
nanowire
FETs biosensors. A non-target protein, BSA, was added to the sensing media in the end
to demonstrate the specificity of the biosensor.
2.3 Chapter conclusion
In
2
O
3
nanowires were synthesized via laser-ablation chemical vapor deposition
system. The VLS mechanism controlled synthesis yields nanowires with narrow diameter
distribution and high aspect ratio. The nanowire yield was dramatically increased after
the change of synthesis substrate. In
2
O
3
nanowire FETs were then fabricated and the
! 36
devices showed n-channel transistor behaviors. Excellent sensitivity of the devices
towards ions was demonstrated with pH sensing over a wide range. Phosphonic acid
molecule was used to successfully functionalize the surface of In
2
O
3
and the FETs were
transformed into nanobiosensor by surface functionalization of CA-125 antibody. The
In
2
O
3
NW FETs exhibited detection limit of CA-125 protein at 1 U/mL, which is around
one order of magnitude lower than the clinical relevant level.
2.4 Experimental section
Materials and instruments Poly-Lysine was purchased from Ted Pella. 10 nm
Au colloidal solution was purchased from BBI solutions. Amine-terminated biotin was
purchased from Pierce. Streptavidin-Alexa fluro 568 was purchased from Invitrogen. CA-
125 antibody and antigen were purchased from Fitzgerald. All other chemicals were
purchased from Sigma-Aldrich. XPS was performed on an M-probe surface spectrometer
(VG Instruments). Monochromatic Al KR X-rays (1486.6 eV) incident at 35 º from
horizontal were used to excite electrons from the sample, and the emitted electrons were
collected by a hemispherical analyzer at a takeoff angle of 35 º from the plane of the
sample surface (horizontal). Fluorescent images were taken by a Nikon Eclipse ME600
fluorescent microscope equipped with a Microfire digital color camera. The electrical
testing of devices was performed by an Agilent 4156B semiconductor analyzer. The real-
time sensing was done using an Agilent B1500 semiconductor analyzer.
Nanowire Synthesis Si(111) wafer was immersed in boiling acetone and
isopropyl alcohol for 5 min, respectively. After drying the surface with a stream of N
2
,
the surface was treated with UV/Ozone cleaner for 10 min. Poly-Lysine was dropped on
the wafer surface and was spread evenly. The wafer was dried by a stream of N
2
after the
! 37
poly-lysine sitting on the substrate for 6 min. Au colloidal solution was diluted 20x in
isopropyl alcohol and sonicated for 8 min. The Au nanoparticle solution was dropped on
the as-prepared wafer and spread evenly on the surface. After 5 min, the wafer was blow-
dried with a stream of N
2
. The wafer was then treated with O
2
plasma to remove the
excess poly-lysine. The substrate was loaded into a quartz tube at the down stream end of
a furnace, and an InAs target was placed at the upper stream of the furnace. The system
was first purged with Ar at a rate of 150 sccm until the oxygen level was below 40 ppm,
while the InAs target was ablated to supply the indium vapor. During the laser ablation
process, the chamber was maintained at 790 ºC. A pulsed Nd:YAG laser (Continuum)
(L= 532 nm) with repetition rate of 10 Hz, and a pulse power of 1.0 W was used. The
typical reaction time used was about 35 min. After the furnace cooled down to room
temperature, light gray materials were found on the surface of the substrate.
Device Fabrication The substrate with synthesized nanowires was sonicated in
isopropyl alcohol for 1 min. The In
2
O
3
nanowire suspension was then dropped and
spread on a SiO
2
/Si wafer cleaned by hot acetone and isopropyl alcohol. The suspension
was repeatedly applied on the wafer once the surface was dry. The substrate was rinsed
thoroughly with isopropyl alcohol after all the suspension was dropped on the substrate
surface. After blow-dried with N
2
, the wafer was patterned by photolithography to define
the metal contact. E-beam evaporation metal deposition was performed to deposit 5 nm
of Ti and 45 nm of Au for the metal electrodes. A passivation layer of Si
3
N
4
was then
deposited by PECVD at 100 ºC on top of the selected metal electrodes area.
pH sensing 10 mM phosphate buffer solutions at different pH values were freshly
made and the total ionic strength of all solutions was adjusted to 100 mM by adding NaCl.
! 38
A Teflon cell was mounted on the active channel area of In
2
O
3
nanowire FETs as the
sensing chamber. A V
DS
=200 mV and a V
GS
=150 mV were applied to the FETs. pH
sensing started with pH 6 buffer solution and was changed to different pH solutions with
pipettes manually.
Surface functionalization In
2
O
3
nanowire sensors or In
2
O
3
nanowire matt on Si
substrates were first cleaned by boiling in acetone and IPA for 5 minutes each, and then
treated by UV/O
3
for 10 minutes (2 minutes in the case of sensors). The cleaned sensors
or substrates were submerged in 0.1 mM aqueous solution of 3-phosphonopropioninc
acid for 16 hours before they were rinsed with deionized (DI) water to remove unbound
linkers. After immobilization of linker molecules, devices were annealed at 120 ºC in N
2
environment for 12 hours to dehydrate surface and to reinforce linkers on the In
2
O
3
nanoribbon surface. The functionalized In
2
O
3
was then subject to the standard NHS/EDC
coupling reaction to attach amine-terminated biotin or antibodies. For fluorescent
imaging experiments, the biotinylated sample was exposed to a solution of streptavidin-
Alexa Fluro 568 (in 1X PBS) for 1 hour and then rinsed by 1X PBS and then D.I. water
for 3 times, respectively.
Biomarker sensing In
2
O
3
nanowire sensors were functionalized with CA-125
antibody and immersed in 0.01X PBS. During the sensing experiment, a 200 mV V
DS
and
a 150 mV V
GS
were applied to the devices and the source-drain currents of three devices
were monitored simultaneously by an Agilent B1500 analyzer. CA-125 antigen solutions
(prepared in 0.01X PBS) were added into the mixing cell from low to high concentrations.
The non-target BSA solution (10 mg/ml in 0.01X PBS) was added at the end of the
sensing experiment to confirm the sensors’ selectivity.
! 39
2.5 References
1. Wan, Q.; Li, Q. H.; Chen, Y. J.; Wang, T. H.; He, X. L.; Li, J. P.; Lin, C. L.,
Fabrication and ethanol sensing characteristics of ZnO nanowire gas sensors. Applied
Physics Letters 2004, 84 (18), 3654-3656.
2. McAlpine, M. C.; Friedman, R. S.; Jin, S.; Lin, K.-h.; Wang, W. U.; Lieber, C.
M., High-Performance Nanowire Electronics and Photonics on Glass and Plastic
Substrates. Nano Letters 2003, 3 (11), 1531-1535.
3. Patolsky, F.; Zheng, G.; Lieber, C. M., Nanowire sensors for medicine and the life
sciences. Nanomedicine 2006, 1 (1), 51-65.
4. Joo, J.; Chow, B. Y.; Prakash, M.; Boyden, E. S.; Jacobson, J. M., Face-selective
electrostatic control of hydrothermal zinc oxide nanowire synthesis. Nat Mater 2011, 10
(8), 596-601.
5. Patrick, N.; Wendy, U. D.; Stefan, B.; Jörg, P. K.; Friedrich, C. S., Polyaniline
nanowire synthesis templated by DNA. Nanotechnology 2004, 15 (11), 1524.
6. Zheng, G.; Jinyun, L.; Yong, J.; Xing, C.; Fanli, M.; Minqiang, L.; Jinhuai, L.,
Template synthesis, organic gas-sensing and optical properties of hollow and porous
In
2
O
3
nanospheres. Nanotechnology 2008, 19 (34), 345704.
7. Liu, Z.; Zhang, D.; Han, S.; Li, C.; Tang, T.; Jin, W.; Liu, X.; Lei, B.; Zhou, C.,
Laser Ablation Synthesis and Electron Transport Studies of Tin Oxide Nanowires.
Advanced Materials 2003, 15 (20), 1754-1757.
8. Hata, K.; Futaba, D. N.; Mizuno, K.; Namai, T.; Yumura, M.; Iijima, S., Water-
assisted highly efficient synthesis of impurity-free single-waited carbon nanotubes.
Science 2004, 306 (5700), 1362-1364.
9. Javey, A.; Guo, J.; Wang, Q.; Lundstrom, M.; Dai, H. J., Ballistic carbon
nanotube field-effect transistors. Nature 2003, 424 (6949), 654-657.
10. Patolsky, F.; Zheng, G. F.; Lieber, C. M., Fabrication of silicon nanowire devices
for ultrasensitive, label-free, real-time detection of biological and chemical species. Nat
Protoc 2006, 1 (4), 1711-1724.
11. Tsakalakos, L.; Balch, J.; Fronheiser, J.; Korevaar, B. A.; Sulima, O.; Rand, J.,
Silicon nanowire solar cells. Applied Physics Letters 2007, 91 (23).
12. Chang, P. C.; Fan, Z.; Chien, C. J.; Stichtenoth, D.; Ronning, C.; Lu, J. G., High-
performance ZnO nanowire field effect transistors. Applied Physics Letters 2006, 89 (13).
13. Soci, C.; Zhang, A.; Xiang, B.; Dayeh, S. A.; Aplin, D. P. R.; Park, J.; Bao, X. Y.;
Lo, Y. H.; Wang, D., ZnO nanowire UV photodetectors with high internal gain. Nano
Letters 2007, 7 (4), 1003-1009.
14. Waitz, T.; Wagner, T.; Sauerwald, T.; Kohl, C.-D.; Tiemann, M., Ordered
Mesoporous In2O3: Synthesis by Structure Replication and Application as a Methane
Gas Sensor. Advanced Functional Materials 2009, 19 (4), 653-661.
15. Li, B.; Xie, Y.; Jing, M.; Rong, G.; Tang, Y.; Zhang, G., In
2
O
3
Hollow
Microspheres:" Synthesis from Designed In(OH)
3
Precursors and Applications in Gas
Sensors and Photocatalysis. Langmuir 2006, 22 (22), 9380-9385.
16. Yu, D.; Yu, S. H.; Zhang, S.; Zuo, J.; Wang, D.; Qian, Y. T., Metastable
Hexagonal In2O3 Nanofibers Templated from InOOH Nanofibers under Ambient
Pressure. Advanced Functional Materials 2003, 13 (6), 497-501.
! 40
17. Cao, H.; Qiu, X.; Liang, Y.; Zhu, Q.; Zhao, M., Room-temperature ultraviolet-
emitting In2O3 nanowires. Applied Physics Letters 2003, 83 (4), 761-763.
18. Nossov, V.; Amneus, M.; Su, F.; Lang, J.; Janco, J. M.; Reddy, S. T.; Farias-
Eisner, R., The early detection of ovarian cancer: from traditional methods to proteomics.
Can we really do better than serum CA-125? Am J Obstet Gynecol 2008, 199 (3), 215-23.
19. Li, C.; Zhang, D.; Han, S.; Liu, X.; Tang, T.; Zhou, C., Diameter-Controlled
Growth of Single-Crystalline In2O3 Nanowires and Their Electronic Properties.
Advanced Materials 2003, 15 (2), 143-146.
20. Viornery, C.; Chevolot, Y.; Léonard, D.; Aronsson, B.-O.; Péchy, P.; Mathieu, H.
J.; Descouts, P.; Grätzel, M., Surface Modification of Titanium with Phosphonic Acid To
Improve Bone Bonding:" Characterization by XPS and ToF-SIMS. Langmuir 2002, 18
(7), 2582-2589.
!
! 41
3 CHAPTER THREE: Top-down Fabricated Polycrystalline Silicon
Nanoribbon Field Effect Transistor Nanobiosensors
3.1 Introduction
Nanobiosensors based on nanostructured field effect transistors have become an
area of intense research because of the potential time and cost efficiency over the current
state-of-art diagnostic platform such as PCR or ELISA.
1-7
In one previous chapter, we
have described a “bottom-up” approach of fabricating such nanobiosensors, in which
nanomaterials are assembled to make devices. Ultrahigh detection sensitivity could be
achieved with this approach because bottom up synthesis yields high crystalline
nanowires with critical dimensions as small as a few nanometers.
1, 8
However, the major
challenge for bottom up technique is the less controllability of the nanomaterials
assembly, which can significantly limit the yield and uniformity of the devices. Although
researches have been conducted to overcome such obstacles, most of the bottom-up
fabricated nanosensors still lack controllability, reliability and scalability. Top-down
fabrication, on the other hand, uses external techniques to define the nanostructure, which
will add scalability and controllability. As a result, the top-down fabricated devices can
yield more uniform and reliable sensing signals. The challenge for such top-down
technique is to create nanostructures with a large surface-to-volume ratio that can be
equivalent to the synthesized nanowires. Recent researches have used e-beam
lithography,
9
imprint lithography
10
and spacer technique
11
to achieve dimensional control
and scalability with promising results. However, such techniques are extremely time-
consuming and cost-inefficient, which can limit the possible commercialization and
prevalence of the devices. In order to further reduce the cost for fabricating nanosensors,
! 42
one recent study demonstrates relaxed lateral dimensons with well-controlled critical
dimension, which is the channel depth, can still maintain a relative highly sensitive
nanosensors. Such devices can be achieved with precise selection of deposition
techniques for the critical channel depth at nanoscale and straightforward
photolithography to define the micrometers nanostructure channel length and width. The
larger area with the relaxed lateral dimensions provides more surface area for analytes to
bind while the nanoscale critical dimension provides the comparable surface-to-volume
ratio to that of bottom-up fabricated devices, which can potentially yield a high
sensitivity.
12, 13
Silicon has been the most widely used material in semiconductor industry. Top-
down fabricated silicon NWs FETs have been achieved on silicon-on-insulator (SOI)
substrates using e-beam lithography.
2, 11, 12
The Si-NWs FETs show good transistor
performance and have been transformed to chemical sensors and biosensors with high
sensitivity. However, cost-inefficient SOI wafer and e-beam writing technique will
significantly limit the commercial impact for the techniques. On the other hand,
polycrystalline silicon wafers can be achieved with relatively lower-cost than SOI wafers
and top-down fabricated nanoribbons FETs with photolithography further reduce the cost
for the sensors to be more economically friendly.
In this chapter, we described a top-down process for the fabrication of
polycrystalline silicon nanoribbon field-effect transistors. The poly-Si NR FETs show
uniform device performance and ~100 % device yield. The whole process is compatible
with conventional photolithography with only two masks, thus is time and cost efficient
and highly scalable. By performing pH sensing, the as-fabricated FETs are shown
! 43
sensitive to the change of pH over a wide dynamic range. With appropriate surface
functionalization of poly-Si, we can transform the FETs into biosensors for ovarian
cancer diagnosis. Therefore, this poly-Si NR FETs platform can be potentially applied for
highly scalable, cost-efficient, label-free biosensors for biomolecules, which will be
beneficial for developing new point-of-care diagnostic devices.
3.2 Results and discussion
3.2.1 Device fabrication and performance
The top-down fabrication started with a 3” SiO
2
/Si wafer. Oxide layer was
thermally grown on a Si wafer and the thickness was controlled at 500 nm that is selected
as the desired dielectric thickness. A thin layer (50 nm) of polycrystalline silicon was
deposited with low-pressure chemical vapor deposition (LPCVD) and then doped with
boron at a desired concentration that is determined by the FET performances. The first
photolithography step followed by a CF
4
dry etching was to define the active
semiconductor channel area and contact leads. The second photolithography step and e-
beam evaporator metal deposition was performed to define the source and drain
electrodes (5 nm Ti and 45 nm Au) for the transistors. Finally, a Si
3
N
4
deposition via
plasma-enhanced chemical vapor deposition (PECVD) was performed to passivate the
contact leads and selected metal electrodes area. The whole fabrication process is highly
scalable and efficient, and with only two photolithography steps involved, the fabrication
is compatible with most industrial fabrication facilities and can be achieved in a much
larger scale. Figure 3.1 shows the SEM images of a group of 6 fabricated poly-Si
nanoribbon devices and a photograph of a whole wafer of fabricated devices to show the
! 44
scalability of such top-down process. These devices bare high uniformity and device
yield due to the high controllability of our fabrication process.
!
Figure 3.1 (a)-(d) Scheme of top-down fabrication of polySi nanoribbon FETs. (e) SEM
image of a group of 6 polySi nanoribbon devices. (f) A photographic image of the
fabricated devices in a wafer scale.
The electrical performances of the FETs using the back-gate are shown as in
Figure 3.2. Doping process was necessary to generate better semiconducting behavior of
the deposited polySi film. We chose 1×10
17
/cm
3
as the doping concentration after trails
of experiments. FETs showed ohmic contact and p-type transistors behavior from the
source-drain current (I
GS
) – source-drain voltage (V
DS
) under various gate voltage (V
GS
)
plots shown in Figure 3.2(a) and I
GS
-V
GS
under different V
DS
plots shown in Figure 3.2(b).
! 45
From the plots, with doping concentration of 1×10
17
/cm
3
, polySi NR FETs exhibited
decent conductance and I
on
/I
off
ratio of 250 at VDS=1V.
!
Figure 3.2 (a) I
DS
-V
DS
plots under various V
GS
. (b) I
DS
-V
GS
plots under various V
DS
.
3.2.2 pH sensing
To demonstrate the ion sensitivity of our poly-Si nanoribbon FETs, we performed
pH sensing experiments under various conditions as shown in Figure 3.3. We
mechanically mount a mixing cell on top of our poly-Si nanoribbon chip to confine the
sensing solution at the channel area. Then we manually exchanged different buffer
solutions varying in pH in and out of the sensing chamber using pipettes. The phosphate
buffer solutions were made fresh each time we performed the sensing experiments and
the ionic strength were maintained at the same level for all pH values so that the sensing
signal would only be attributed to the change of pH. Figure 3.3(a) shows the real-time
sensing response of a 50 nm bare poly-Si nanoribbon devices to various pH buffer
solutions ranging from pH 4 to pH 10 and then back to pH 4. The device showed an
increase of conduction with increased pH, which agrees to the gate voltage modulation
behavior of p-channel transistors. Poly-Si surface is covered with a thin layer of SiO
2
with hydroxyl groups exposed to the buffer solution. Increasing of the pH value will
enhance the deprotonation of the surface hydroxyl groups, and the negative charges on
(a) (b)
! 46
the surface will induce an accumulation of charge carriers in the poly-Si channel. The
charge carriers’ accumulation will generate an increase of conduction as shown in the
sensing signal. The device showed a constant increasing of conductance over a wide
dynamic range (from pH 4 to pH 10) and good reversibility when changing back to lower
pH solutions. To test the sensitivity of nanoribbon devices to pH change in a more
relevant range for biosensing, we performed pH sensing from pH 7.2 to pH 8 with steps
of 0.2 in pH. With around 10% change of conductance for a variation of 0.2 in pH, the
poly-Si nanoribbon FETs show high sensitivity to ions, as shown in Figure 2.3(b).
!
Figure 3.3 (a) Real time pH sensing ranging from pH=4 to pH=9. (b) Small range of pH
sensing from pH=7.2 to pH=8.
pH 7.2 pH 7.6 pH 7.8
(b)
pH 7.4
pH 8
! 47
3.2.3 Surface functionalization of polycrystalline silicon with native oxide layer
An interface between an electronic device and a biomolecule needs to be
established before we can transform our FETs into biosensors. Silicon is known to grow a
thin layer of native oxide on the surface because of the oxidation process when exposed
in the air. Appropriate surface chemistry is required to anchor the specific probe molecule,
such as antibodies or oligonucleotides, on the active channel area (nanoribbon area) in
order to capture the analytes of interests. Numerous researches have been report to
functionalize surface of silicon, and the most effective way to anchor biomolecule is to
covalently bond the molecule via the surface chemistry between surface oxide layer and
alkoxysilane.
2, 14, 15
To explore the efficiency of surface chemistry, we chose
biotin/streptavidin as the model probe/target molecules pair. The surface
functionalization scheme is shown in Figure 3.4(a). First, a pre-clean polysilicon
substrate was functionalized with a linker molecule, 3-aminopropyldimethyethoxysilane
(APDMS) to bare an amine-terminated surface. The reason we chose this
monoethoxysilane molecule over the much more commonly used trialkoxysilane is to
prevent the formation of a thick layer of polymer film from self-polymerization and that
thick layer will significantly decrease the sensitivity of the FETs biosensors. The amine
terminated polysilicon surface was then converted to carboxyl groups via reaction
between amine and succinic anhydride in the presense of triethylamine. Amine-
terminated biotin molecules (NH
2
-PEG
3
-Biotin) are attached on the surface via
NHS/EDC-mediated coupling reaction.
To verify the success of yielding silane monolayer, the film thickness was
measured by ellipsometer and there was a gain of 7 Å after this step and the change of
! 48
contact angle from less than 10 º to 60 º also confirms the change of surface functional
groups from highly hydrophilic hydroxyl groups to amine groups. High-resolution XPS
spectra of N 1s region are plotted for the bare polysilicon and silane monolayer
functionalized polysilicon, as shown in Figure 3.4(b). The increase of the intensity
indicates the formation of the monolayer due to the amine groups from APDMS molecule.
High-resolution XPS spectra of C 1s region are also plotted in Figure 3.4(c). The
increasing of 289 eV peak from black trace (amine-terminated monolayer) to red trace
(carboxyl-terminated monolayer) indicates the success of conversion from amine
terminals to carboxyl terminals on the polysilicon surface because the peak at 289 eV
represents the C=O content that comes from the carboxyl groups. Survey scans of X-ray
photoelectron spectroscopy (XPS) spectra are plotted after each steps of functionalization
in Figure 3.4(d). The significant increase of N 1s content at the last step of anchoring
streptavidin indicates the success of capturing the target protein in our model surface
chemistry study.
!
Figure 3.4 (a) Surface functionalization scheme of polySi with native oxide. (b) XPS
high-resolution spectra of N 1s region. (c) XPS high-resolution of C 1s region. (d)
Stepwise XPS survey spectra on polySi: black trace(bare polySi); red trace(polySi with
silane monolayer); blue trace(polySi with carboxyl terminal); green trace(polySi
functionalized with biotin and treated with streptavidin).
OHOH
Si
H
2
N
OEt
O
Si
NH
2
O O
O
O
Si
HN
OH O
O
NHS/EDC
NH
2
O
Si
HN
H
N O
O
poly Si
poly Si poly Si poly Si
394 396 398 400 402 404 406
(Bare(Poly1Si
(Poly1Si(+(SAM
(Poly1Si(+(SAM(w/(COOH
(
(
Intensity((a.u.)
Binding(Energy((eV)
!Poly&S i!+!S AM
!Poly&S i!+!S AM!w/!COOH
!Poly&S i!+!S AM!+!Biotin!+!S A
278 280 282 284 286 288 290 292
!
!
Intensity!(a.u.)
Binding!Energy!(eV)
0 100 200 300 400 500 600 700
O*1s
N*1s
C*1s
Si*2s
*
*
Intensity*(a.u.)
Binding*Energy*(eV)
Si*2p
(b)
(c)
(d)
(a)
! 49
!
3.2.4 Ovarian cancer biomarker sensing
We chose a biomarker, CA-125, as the detection target to demonstrate the
potential application of polysilicon nanoribbon FETs as nanobiosensors. The polysilicon
nanoribbon was functionalized with CA-125 antibodies on the surface to specifically
recognize the target proteins. During the sensing experiment, a 200 mV V
DS
and a -200
mV V
GS
were applied to the devices and the current between source and drain electrodes
of three devices are monitored simultaneously. CA-125 protein solutions with different
concentrations were added progressively into the sensing chamber.
The normalized current versus time for three monitored devices is plotted in
Figure 3.5. All three devices show almost no response with small concentrations (0.01-1
U/mL), while start to show increase of current at CA-125 concentration of 10 U/mL and
progressively larger response with higher concentrations of CA-125 added. Note that all
three devices show uniform responses when different concentrations of CA-125 solutions
were added. The detection limit of polysilicon nanoribbon FETs for CA-125 is 10 U/mL
and this level is at least one order of magnitude lower than the clinical relevant level for
diagnosis (35 U/mL – 275 U/mL).
16
In order to show the specificity of our
nanobiosensors, a control experiment was performed with bovine serum albumin (BSA).
No devices show responses to 150 nM of BSA, a concentration that is much higher than
the target molecules. Therefore, our polysilicon nanoribbon FETs biosensors show decent
sensitivity and good specificity for CA-125 detection and thus can be potentially used as
a sensing platform for point-of-care detection of other diseases.
! 50
!
Figure 3.5 Real time sensing of CA-125 detection. A non-target protein, BSA, was added
in the end to demonstrate the specificity of the sensors.
3.2.5 Surface functionalization of polycrystalline silicon without native oxide layer
The insulating native oxide layer on the surface of silicon screens out some
charges gathered on the surface so the sensitivity of silicon FETs as biosensors is
negatively affected. In order to optimize the sensitivity for the polysilicon nanoribbon
FETs nanobiosensors, we explored surface functionalization approach without the native
oxide layer on polysilicon surface. The scheme is shown in Figure 3.6. First, polysilicon
substrate was dipped in 2% HF solution for 30s to remove the native oxide layer. The
surface Si-H was substituted into Si-Br when exposing the treated substrate to Br
2
with
the aid of UV illumination. With the generation of bromine radicals during UV
illumination, the radical substitution reaction was happened at polysilicon surface,
yielding Si-Br bond.
17
The surface bromine was then further substituted with Grignard
reagents via S
N
2 substitution reaction. Note that a mixture of methyl and olefin Grignard
reagents were used during the functionalization. The olefin reagent was used for further
! 51
functionalization and the methyl reagent was used to dilute the surface olefin groups,
making big biomolecules more approachable during functionalization. Exposing the
olefin terminated surface to Br
2
under dark condition yielded 1,2-dibromides at the
terminals.
18
Efforts were made to convert the terminal bromine into amine group with
two different approaches. The first attempt was carried using phthalimide to substitute the
terminal bromine followed by deprotection with methylamine to achieve amine
terminals.
19
The second approach started with substituting the surface bromine with azide
groups followed by reducing the azide into amine with LiAlH
4
.
20
!
Figure 3.6 Surface functionalization scheme for poly-Si without native oxide.
XPS spectra were used to verify each steps of the surface reaction. High-
resolution spectra of Si 2p region were plotted in Figure 3.7(a). The HF treated spectrum
OH OH OH
2% HF
Br
2
(gas)
UV
OH H H H H
Br Br Br Br
H
2
C=CHCH
2
CH
2
MgBr
CH
3
MgBr
CH
3
CH
3
Br
2
in DCM
dark
CH
3
CH
3
Br
Br
Br
Br
CH
3
Br
Br
Br
Br
CH
3
CH
3
N
Br
N
Br
CH
3
CH
3
N
3
Br
N
3
Br
CH
3
N
+
K
-
O
O
O
O O
O
CH
3
NH
2
CH
3
NH
2
Br
NH
2
Br
CH
3
NaN
3
LiAlH
4
NaOH
CH
3
NH
2
Br
NH
2
Br
CH
3
! 52
showed no peak at 102.5 eV region, indicating the success of removing native oxide layer.
High resolution spectrum of C 1s region (Figure 3.7(b)) confirms the success of
anchoring vinyl groups on the surface with the existence of peak at 286.5 eV,
correspondent with the binding energy of C=C bond.
21
In Figure 3.7(c), the peak of Br 3d
region at 70 eV increased significantly from black trance (vinyl terminal) to red trace
(after bromine addition), suggesting the conversion of the surface from vinyl groups to
1,2-dibromides. Two different attempts were carried out to yield amine groups for further
surface functionalization. High-resolution spectra of N 1s for both approaches were
plotted in Figure 3.7(d). Both traces showed almost no peak at 400 eV, indicating the
failure of converting bromo-terminated surface into amine terminals. More experiments
are necessary to explore the reason of the failure. Even though the surface chemistry
without native oxide layer cannot be fulfilled completely, with the success of the first two
steps, it is still highly possible to realize the anchoring of biomolecules. An improvement
of sensitivity is expected without the insulating native oxide layer, thus biosensors baring
higher sensitivity could be fulfilled.
! 53
!
Figure 3.7 (a) High-resolution XPS spectrum of Si 2p region after HF treatment. (b)
High-resolution XPS spectrum of C 1s region after Grignard reaction. (c) High-resolution
XPS spectra of Br 3d region before and after bromine treatment with the vinyl surface.
(d) High-resolution XPS spectra of N 1s region of two different routes creating amine
surface.
3.3 Chapter conclusion
Polycrystalline silicon nanoribbon FETs are fabricated with top-down process
with only two steps of photolithography. Devices fabricated with this process show
excellent yield, good uniformity and transistor behavior. The fabrication process is
highly-scalable, low cost and compatible with most semiconducting industry fabrication
facilities. Polysilicon nanoribbon devices exhibited good sensitivity in both wide range of
pH solution from pH 4 to 10 and physiological range between 7.2 and 8.0. Surface
(a)
(b)
(c)
(d)
! 54
functionalization with biomolecules is conducted with the native oxide layer and the
FETs have been successfully transformed into biosensors for CA-125 detection with
good sensitivity and specificity. Efforts are made to functionalize the surface without
native oxide layer and further experiments are necessary to improve the sensitivity of
polysilicon nanobiosensors.
3.4 Experimental section
Materials and instruments Spin-on dopant was purchased from Emulsitone
cooperation. APDMS was purchased from Gelast, Inc. Amine-terminated biotin was
purchased from Pierce. Streptavidin-Alexa fluro 568 was purchased from Invitrogen. CA-
125 antibody and antigen were purchased from Fitzgerald. All other chemicals were
purchased from Sigma-Aldrich. XPS was performed on an M-probe surface spectrometer
(VG Instruments). Monochromatic Al KR X-rays (1486.6 eV) incident at 35 º from
horizontal were used to excite electrons from the sample, and the emitted electrons were
collected by a hemispherical analyzer at a takeoff angle of 35 º from the plane of the
sample surface (horizontal). Fluorescent images were taken by a Nikon Eclipse ME600
fluorescent microscope equipped with a Microfire digital color camera. The electrical
testing of devices was performed by an Agilent 4156B semiconductor analyzer. The real-
time sensing was done using an Agilent B1500 semiconductor analyzer.
Device Fabrication 50 nm of polycrystalline silicon thin layer was deposited by
LPCVD on a Si (100) substrate with thermally grown SiO
2
(typically 500 nm) layer.
Spin-on dopant with desired concentrations was spun on the substrate followed by a
drive-in annealing process at 1100 ºC for 15 min under N
2
. BOE etching was performed
to remove the spin-on dopant. A first photolithography was performed followed by CF
4
! 55
dry etching to define the contact leads and active nanoribbon area of polysilicon. Then
the second photolithography was carried out to define the metal contact. E-beam
evaporation metal deposition was performed to deposit 5 nm of Ti and 45 nm of Au for
the metal electrodes. A passivation layer of Si
3
N
4
was then deposited by PECVD at 100 º
on top of the selected metal electrodes area.
pH sensing 10 mM phosphate buffer solutions at different pH values were freshly
made and the total ionic strength of all solutions was adjusted to 100 mM by adding NaCl.
A Teflon cell was mounted on the active channel area of polysilicon nanoribbon FETs as
the sensing chamber. A V
DS
=200 mV and a V
GS
=-200 mV were applied to the FETs. pH
sensing started with pH 4 buffer solution and was changed to different pH solutions with
pipettes manually.
Surface functionalization with native oxide Polysilicon wafers or sensors were
first cleaned by boiling in acetone and IPA for 5 minutes each, and then treated by UV/O
3
for 10 minutes (2 minutes in the case of sensors). The cleaned wafers or sensors were
immediately transferred to an APDMS solution in dry toluene (2% v/v) and incubated for
2 hours in N
2
. They were then washed by fresh toluene and methanol and annealed under
N
2
at 120 ºC for 12 hours. The carboxyl surface were generated by immersing annealed
polysilicon in a 5 mg/ml succinic anhydrous solution in dry THF containing 5% v/v
triethylamine for 4 hours under N
2
. The functionalized polysilicon was then subject to the
standard NHS/EDC coupling reaction to attach amine-terminated biotin or antibodies.
Surface functionalization without native oxide Polysilicon wafers were first
cleaned by solvent boiling and UV/O
3
as discribed above, and then dipped into a
deoxygenated HF solution (2%) for 30 seconds. The wafers were then quickly rinsed with
! 56
D.I. water and transferred into a pre-dried custom-made quartz tube. The tube was
immediately pumped down to vacuum to ensure an oxygen-free environment. After back
filling the tube with N
2
, a dry and deoxygenated Br
2
was transferred via cannula into the
tube. The polysilicon was fully immersed in Br
2
and a handheld UV light was used to
illuminate the quartz tube for 30 min. Br
2
was then pumped out from the quartz tube and
a mixture of Grignard reagents methylmagnesium bromide solution (1.4 M in THF:
Toluene) and 4-Pentenylmagnesium bromide solution (0.5 M in THF) was transferred
into the quartz tube via cannula at volume ratio of 1:3. The whole quartz tube was filled
with N
2
and kept at 60 °C overnight. The reaction was quenched by adding anhydrous
methanol into the system and the polysilicon substrate was sonicated in MeOH for 3
times, 5 min each time. Polysilicon was then immersed in Br
2
solution in dry DCM (v/v
10%) under N
2
for 16 h. The as-prepared polysilicon substrate was immersed in 18 mM
phthalimide potassium solution in dry DMF under N
2
at 60 °C for 3 h then washed with
dry DMF and MeOH 3 times, respectively. Then polysilicon substrate was immersed in
40% CH
3
NH
2
for 1 min and was washed with MeOH 3 times blew dry under stream of
N
2
. The other route was also performed to convert the 1,2-dibromides surface to amines.
Br
2
solution treated polysilicon was immersed in supersaturated solution of NaN
3
in DMF
for 24 h at room temperature and was washed with D.I. water 3 times. Then the substrate
was treated with 0.2 M LiAlH
4
solution in dry THF for 24 h at room temperature
followed by immersing in 5% HCl for 5h. The substrate was then washd with D.I water
and MeOH 3 times, respectively and blew dry under N
2
stream.
Biomarker sensing Polysilicon nanoribbon sensors with native oxide were
functionalized with CA-125 antibody and immersed in 0.01X PBS. During the sensing
! 57
experiment, a 200 mV V
DS
and a -200 mV V
GS
were applied to the devices and the
source-drain currents of three devices were monitored simultaneously by an Agilent
B1500 analyzer. CA-125 antigen solutions (prepared in 0.01X PBS) were added into the
mixing cell from low to high concentrations. The non-target BSA solution (10 mg/ml in
0.01X PBS) was added at the end of the sensing experiment to confirm the sensors’
selectivity.
3.5 References
1. Cui, Y.; Wei, Q.; Park, H.; Lieber, C. M., Nanowire nanosensors for highly
sensitive and selective detection of biological and chemical species. Science 2001, 293
(5533), 1289-92.
2. Stern, E.; Klemic, J. F.; Routenberg, D. A.; Wyrembak, P. N.; Turner-Evans, D.
B.; Hamilton, A. D.; LaVan, D. A.; Fahmy, T. M.; Reed, M. A., Label-free
immunodetection with CMOS-compatible semiconducting nanowires. Nature 2007, 445
(7127), 519-22.
3. Li, C.; Curreli, M.; Lin, H.; Lei, B.; Ishikawa, F. N.; Datar, R.; Cote, R. J.;
Thompson, M. E.; Zhou, C., Complementary detection of prostate-specific antigen using
In2O3 nanowires and carbon nanotubes. J Am Chem Soc 2005, 127 (36), 12484-5.
4. Ishikawa, F. N.; Chang, H. K.; Curreli, M.; Liao, H. I.; Olson, C. A.; Chen, P. C.;
Zhang, R.; Roberts, R. W.; Sun, R.; Cote, R. J.; Thompson, M. E.; Zhou, C., Label-free,
electrical detection of the SARS virus N-protein with nanowire biosensors utilizing
antibody mimics as capture probes. ACS Nano 2009, 3 (5), 1219-24.
5. Chang, H. K.; Ishikawa, F. N.; Zhang, R.; Datar, R.; Cote, R. J.; Thompson, M.
E.; Zhou, C., Rapid, label-free, electrical whole blood bioassay based on nanobiosensor
systems. ACS Nano 2011, 5 (12), 9883-91.
6. Gao, A.; Lu, N.; Dai, P.; Li, T.; Pei, H.; Gao, X.; Gong, Y.; Wang, Y.; Fan, C.,
Silicon-nanowire-based CMOS-compatible field-effect transistor nanosensors for
ultrasensitive electrical detection of nucleic acids. Nano Lett 2011, 11 (9), 3974-8.
7. Patolsky, F.; Zheng, G.; Hayden, O.; Lakadamyali, M.; Zhuang, X.; Lieber, C.
M., Electrical detection of single viruses. Proc Natl Acad Sci U S A 2004, 101 (39),
14017-22.
8. Cui, Y.; Lauhon, L. J.; Gudiksen, M. S.; Wang, J. F.; Lieber, C. M., Diameter-
controlled synthesis of single-crystal silicon nanowires. Appl Phys Lett 2001, 78 (15),
2214-2216.
9. Park, I.; Li, Z. Y.; Pisano, A. P.; Williams, R. S., Top-down fabricated silicon
nanowire sensors for real-time chemical detection. Nanotechnology 2010, 21 (1).
10. Wang, D. W.; Sheriff, B. A.; Heath, J. R., Silicon p-FETs from ultrahigh density
nanowire arrays. Nano Letters 2006, 6 (6), 1096-1100.
! 58
11. Hakim, M. M. A.; Lombardini, M.; Sun, K.; Giustiniano, F.; Roach, P. L.; Davies,
D. E.; Howarth, P. H.; de Planque, M. R. R.; Morgan, H.; Ashburn, P., Thin Film
Polycrystalline Silicon Nanowire Biosensors. Nano Letters 2012, 12 (4), 1868-1872.
12. Elfstrom, N.; Karlstrom, A. E.; Linnros, J., Silicon nanoribbons for electrical
detection of biomolecules. Nano Letters 2008, 8 (3), 945-949.
13. Vacic, A.; Criscione, J. M.; Stern, E.; Rajan, N. K.; Fahmy, T.; Reed, M. A.,
Multiplexed SOI BioFETs. Biosens Bioelectron 2011, 28 (1), 239-242.
14. Wang, W. U.; Chen, C.; Lin, K. H.; Fang, Y.; Lieber, C. M., Label-free detection
of small-molecule-protein interactions by using nanowire nanosensors. Proc Natl Acad
Sci U S A 2005, 102 (9), 3208-12.
15. Zheng, G.; Patolsky, F.; Cui, Y.; Wang, W. U.; Lieber, C. M., Multiplexed
electrical detection of cancer markers with nanowire sensor arrays. Nat Biotechnol 2005,
23 (10), 1294-301.
16. Nossov, V.; Amneus, M.; Su, F.; Lang, J.; Janco, J. M.; Reddy, S. T.; Farias-
Eisner, R., The early detection of ovarian cancer: from traditional methods to proteomics.
Can we really do better than serum CA-125? Am J Obstet Gynecol 2008, 199 (3), 215-23.
17. Linford, M. R.; Chidsey, C. E. D., Surface Functionalization of Alkyl Monolayers
by Free-Radical Activation:" Gas-Phase Photochlorination with Cl
2
. Langmuir 2002, 18
(16), 6217-6221.
18. Brozek, E. M.; Zharov, I., Internal Functionalization and Surface Modification of
Vinylsilsesquioxane Nanoparticles. Chem Mater 2009, 21 (8), 1451-1456.
19. Ofir, Y.; Zenou, N.; Goykhman, I.; Yitzchaik, S., Controlled Amine Functionality
in Self-Assembled Monolayers via the Hidden Amine Route:" Chemical and Electronic
Tunability. The Journal of Physical Chemistry B 2006, 110 (15), 8002-8009.
20. Lee, M. T.; Ferguson, G. S., Stepwise synthesis of a well-defined silicon
(oxide)/polyimide interface. Langmuir 2001, 17 (3), 762-767.
21. Adamkiewicz, M.; O’Hagan, D.; Hähner, G., Bis(trifluoromethyl)methylene
Addition to Vinyl-Terminated SAMs: A Gas-Phase C–C Bond-Forming Reaction on a
Surface. Langmuir 2014, 30 (19), 5422-5428.
!
! 59
4 CHAPTER FOUR: Highly Scalable, Uniform, and Sensitive
Biosensors Based on Top-Down In
2
O
3
Nanoribbons and Electronic
Enzyme-Linked Immunosorbent Assay
4.1 Introduction
Point-of-care testing (POCT) devices have drawn intensive attention recently.
Researches have been mostly focused to enhance the portability and accuracy while
reduce the cost for detection of various biomarkers and diseases.
1-4
The most successful
POCT device on the market is the glucose meter. It has been developed for over 50 years
and the relatively low-cost of the device makes it the most widely used in-house glucose
monitor. Efforts have been made to produce more POCT devices, however none of which
is comparable with the glucose meter. They either lack of sensitivity or cost-efficiency.
5, 6
Instead, patients still need to go to the hospital to get the test results, and the conventional
testing approaches, e.g. ELISA or PCR, require several days of turn-around time, well-
trained personnel and expensive facilities. POCT devices would be of great significance
in the emergency room setting where rapid test results are more beneficial for physicians
to deliver appropriate treatment plan. Moreover, it will be much useful for patients who
need to keep track of one or several biomarkers, such as glucose, constantly to ensure a
healthy condition.
Nanomaterials based field-effect transistors have attracted interests from scientists
and engineers to develop POCT devices.
4, 7-9
The instant electrical signal produced by
transistors is essential for a rapid assay result. Nanoscale semiconductor channel
introduces a large surface-to-volume ratio, which is critical for ultrahigh sensitivity. Also,
the dimensions of the nanoscale materials are comparable with the size of biomolecules,
! 60
making them the ideal interface materials between biological molecules and scientific
instrument. Moreover, CMOS (Complementary-metal-oxide-semiconductor) compatible
fabrication process ensures the possibility of large-scale production of such devices with
the current industrial facilities, which can reduce the device cost and fabrication turn-
around time. Therefore, nanomaterials based FETs are good candidates to develop rapid,
ultrahigh sensitive and low-cost POCT devices.
Diagnosis of HIV-1 infection, the cause of acquired immune deficiency syndrome
(AIDS), relies on the detection of HIV-1 ribonucleic acid (RNA), capsid antigen p24, and
anti-HIV antibody.
10
Most common FDA-approved HIV-1 diagnostic assays target HIV-
1 RNA because of the ease of polymerase chain reaction (PCR) mediated amplification.
Although this method is very sensitive, it is expensive, requires well-trained staff, and
must be performed in well-equipped laboratories. RNA testing in rural or remote settings
is also a major challenge. Testing for the HIV-1 p24 antigen is a good alternative because
it can be done in resource-limited situations. Antigen p24 level is significantly high
during the early, acute phase of infection and the terminal stage of AIDS. It is also a
useful marker for predicting CD4+ T cell count decreases, disease progression for early
detection of HIV-1 infection, and patient prognosis. Early detection of HIV infection
helps to prevent HIV transmission and to prolong AIDS condition by receiving proper
treatment. HIV-1 p24 antigen is usually detected by enzyme-linked immunosorbent assay
(ELISA). However, the detection sensitivity of the conventional assay is less than
desirable: 10−20 pg/mL.
11
In this chapter, we introduce top-down fabricated In
2
O
3
nanoribbon field-effect
transistors with great uniformity and good transistor behavior. The as-fabricated devices
! 61
are then transformed into prototypes of POCT devices by combining this scalable sensing
platform with amplification from electronic enzyme-linked immunosorbent assay
(ELISA). The combined platform has shown great sensitivity comparing to conventional
ELISA with a much reduced turnover time. We choose HIV1 p-24 protein as the target
analyte and our sensing platform can achieve three orders of magnitudes lower detection
limit than commercial ELISA kit within several hours of sample preparation and
detection process. Therefore, this top-down fabricated In
2
O
3
nanoribbon FETs platform
has the potential for scalable, rapid and low-cost POCT devices for a variety of biological
analytes.
4.2 Results and discussions
4.2.1 Device fabrication and performance
The fabrication process requires two steps of photolithography, as shown in
Figure 4.1(a-d). First, 500 nm Si
3
N
4
was deposited on Si wafer substrates by low-pressure
chemical vapor deposition (LPCVD). Si
3
N
4
was chosen instead of SiO
2
to suppress the
competition binding of the target analytes on the surface of the substrates. The first mask
was used to define the metal contacts, followed by electron beam vapor metal deposition
of 5 nm Ti and 45 nm Au. After metal electrodes definition, a second photolithograph
step was introduced to define the dimension and position of the In
2
O
3
channel. In
2
O
3
nanoribbons were then deposited by RF sputtering at room temperature with thickness
targeted at 10 nm to 50 nm. Lift-off process was performed afterwards to realize the
In
2
O
3
nanoribbons, which also finalized the fabrication process. Note that the top-down
fabrication for In
2
O
3
nanoribbon FETs is different from the polysilicon FETs. In
2
O
3
nanoribbon, in this process, is never exposed to any photoresist or other additional layer,
! 62
thus the surface of the In
2
O
3
is left pristine for surface chemistry. Moreover, because the
sputtered In
2
O
3
is intrinsically n-type semiconductor, the device does not require any
further doping process. Figure 4.1(e) shows a wafer-scale photo of the fabricated In
2
O
3
nanoribbon FETs with 100% yield, and its inset shows a magnified optical image of one
nanoribbon chip, which contains 4 subgroups of 6 nanoribbon FET devices. A
comparison between the scanning electron microscopy (SEM) images of two nanoribbon
FETs in Figure 4.1(f) shows that the 50 µm by 2 µm channels are identical. This
uniformity in fabrication is expected to extend to the electrical properties of the devices.
!
Figure 4.1 (a)-(d) shows the fabrication steps of In
2
O
3
nanoribbon FETs using two-step
photolithography. (e) A photographic image of a wafer scale fabricated devices with a
zoom-in photo of one single sensor chip. (f) SEM image of two fabricated In
2
O
3
nanoribbon FETs.
Si
Si
3
N
4
PR
(c)
Si
Si
3
N
4
PR
(b)
Si
Si
3
N
4
Ti/Au
Ti/Au
(d)
(a)
Si
Si
3
N
4
In
2
O
3
(e)
(f)
50
µm
1.5 mm
1 cm
! 63
After fabrication, the devices are characterized by an Agilent semiconductor
analyzer 4156B using a back gate. Figure 4.2(a) and (b) respectively shows the family
plots of source-drain current versus source-drain voltage (I
DS
-V
DS
) and source-drain
current versus gate voltage (I
DS
-V
GS
) curves measured from one In
2
O
3
nanoribbon device.
Figure 4.2(a) indicates a good metal oxide field effect transistor behavior where I
DS
varies
linearly at lower V
DS
and starts to saturate at voltage above 20 V. Both Figure 4.2(a) and
(b) show increasing I
DS
with increasing V
DS
, which agrees with the trend of n-channel
transistor behavior. From figure y, the device is turned off at voltage below 0 V, with
on/off current ratio in the range of 10
5
to 10
6
.
!
Figure 4.2 (a) I
DS
-V
DS
family plots under various V
GS
. (b) I
DS
-V
GS
family plots under
various V
DS
.
4.2.2 Statistical analysis of the In
2
O
3
nanoribbon FETs
One of the main advantages of top-down fabrication process is the good
controllability, which achieves great device yield and excellent uniformity of the device
performance. We have shown, as in Figure 4.1(f), the uniformity of the channel ribbons
with SEM images. Here we performed statistical study with 50 In
2
O
3
devices for key
transistor characteristic factors. The average on-state current (I
ON
) measured at V
DS
= 600
mV and V
GS
= 30 V is 918.6 nA with a standard deviation of 36 nA, or 4% from the
0 10 20 30 40 50
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
(V
G
(=(100V
((((((((Step(=(10V
!
!
Drain(Current((µA)
Drain(Voltage((V)
!100 !50 0 50 100
0
30
60
90
120
150
180
210
*VD*=*1V
********step*=*0.2V
!
!
Drain*Current*(nA)
Gate*Voltage*(V)
(a)
(b)
! 64
average value, as shown in Figure 4.3(a). The V
TH
is extracted from the I
DS
−V
GS
curve
with VDS set to 600 mV, and the distribution is plotted in Figure 4.3(b) with an average
of 15.38 V and a standard deviation of 0.78 V, or 5% of the average. The electron
mobility (µ) was calculated from the relationship gm = µ(C/L
2
)V
DS
where gm is taken as
the maximum of the derivative of the I
DS
−V
GS
curve. The gate capacitance (C) was
calculated from the parallel plate model (C = εA/d), and the channel length (L) is 80 µm.
With a SiO
2
relative dielectric constant of 3.9, calculated plate area (A) of 460 µm
2
, and
SiO
2
dielectric thickness (d) of 50 nm, the gate capacitance is calculated to be 3.18 ×
10
−13
F. From these parameters, mobilities of all 50 devices are calculated and plotted in
Figure 4.3(c). The average electron mobility is 23.38 cm
2
/(V·s), and the standard
deviation is 1.42 cm
2
/(V·s) or 6% of the average. Lastly, Figure 4.3(d) shows the on/off
current ratios from 50 devices. Most of the In
2
O
3
nanoribbon devices demonstrate a good
on/off ratio between 10
5
to 10
7
. To bench mark the uniformity of our “top-down”
nanoribbon devices, the I
ON
of 50 In
2
O
3
nanoribbon FETs (Figure 4.3(e)) is compared to
that of 50 In
2
O
3
nanowire FETs shown in Figure 4.3(f). In
2
O
3
nanoribbon devices show
more uniform on-state current than nanowire devices due to their high dimensional
control.
! 65
!
Figure 4.3 (a)-(d) Statistic study of four FET key parameters: I
ON
(a); V
TH
(b); gm(c);
µ(d). (e)-(f) Comparison of I
ON
distribution for In
2
O
3
nanoribbon (e) FETs and nanowire
(f) FETs
4.2.3 pH sensing
To demonstrate the ion sensitivity of our In
2
O
3
nanoribbon FETs, we performed
pH sensing experiments under various conditions as shown in Figure 4.4. We
mechanically mount a mixing cell on top of our In
2
O
3
nanoribbon chip to confine the
sensing solution at the channel area. Then we manually exchange different buffer
solutions varying in pH in and out of the sensing chamber using pipettes. The phosphate
buffer solutions were made fresh each time we performed the sensing experiments and
the ionic strength were maintained at the same level for all pH values so that the sensing
signal would only be attributed to the change of pH. Figure 4.4(a) shows the real-time
sensing response of a 20 nm bare In
2
O
3
nanoribbon devices to various pH buffer
solutions ranging from pH 9 to pH 4. The device showed an increase of conduction with
decreased pH, which agrees to the gate voltage modulation behavior of n-channel
transistors. In
2
O
3
surface is covered with hydroxyl groups that are exposed to the buffer
solution. Decreasing of the pH value will enhance the protonation of the surface hydroxyl
! 66
groups, and the positive charges on the surface will induce an accumulation of charge
carriers in the In
2
O
3
channel. The charge carriers’ accumulation will generate an increase
of conduction as shown in the sensing signal. The device showed a constant increasing of
conductance over a wide dynamic range (from pH 9 to pH 4). To test the sensitivity of
nanoribbon devices to pH change in a more relevant range for biosensing, we performed
pH sensing from pH 6.7 to pH 8.2 with steps of 0.3 in pH. Figure 4.4(b) shows the
average of the normalized current responses from three In
2
O
3
nanoribbon FETs with 20
nm thickness. These devices show a 5 times decrease in conduction with a pH change of
1.5.
!
Figure 4.4 Normalized current for pH sensing. (a) Sensing of pH 4 to pH 9 with sensors of
different In
2
O
3
nanoribbon thickness. (b) Normalized current versus time with a pH step of about
0.3 in pH ranges of biological interest.
4.2.4 Surface functionalization
An interface between an electronic device and a biomolecule needs to be
established before we can transform our FETs into biosensors. We applied a phosphonic
acid molecule to covalently anchor biological molecules as probe molecules for specific
sensing. The same phosphonic acid to In
2
O
3
chemistry, along with the N-(3-
(dimethylamino)propyl)-N’-ethylcarbodiimide hydrochloride (EDC)/N-
4 5 6 7 8 9
0
10
20
30
40
50
60
70
80
90
100
!
!
10nm
20nm
30nm
40nm
50nm
I/I
0
pH
200 400 600 800
0
1
2
3
4
5
6
pH 6.72
pH 6.98
pH 7.32
pH 7.56
pH 7.85
Device A
Device B
Device C
I/I
0
Time (s)
pH 8.18
(a) (b)
! 67
hydroxysuccinimide (NHS) coupling to biomolecules, has been demonstrated in previous
studies on In
2
O
3
nanowire-based devices. Figure 4.5(a) shows the scheme for the surface
functionalization. Fluorescent experiments were carried out with biotin and fluorescent
dye-tagged streptavidin as the model probe/analyte system. The negative controls are
anchored with amine poly-(ethylene glycol) (PEG) as the probe instead of biotin. Figure
4.5(b) and (c) show that In
2
O
3
ribbons with biotin probes are bright while the ribbons
with amine-PEG probes are dark, confirming successful binding of probe biomolecules
using the phosphonic acid chemistry. Moreover, we also used gold nanoparticles
conjugated streptavidin protein to further confirm the efficiency of the surface chemistry.
After finishing the surface functionalization, both sample and control devices were taken
SEM images and the gold particle density on the one with biotin as probe molecule was
around 120 particles/µm
2
whereas the density on the control device was only as low as 2
particles/µm
2
. This significant difference of gold particle density again confirms the
success of the surface chemistry.
! 68
!
Figure 4.5 (a) Surface functionalization scheme for In
2
O
3
nanoribbon. Fluorescent
images for biotin (b) and amine-PEG (c) as probe molecules after treatment with Alexa
Fluro-568 streptavidin. SEM images of In
2
O
3
nanoribbon for biotin (d) and amine-PEG
(e) as probe molecules after treatment with Au-streptavidin.
4.2.5 Real time detection of HIV1 p24 protein
In order to demonstrate the application of In
2
O
3
nanoribbon FETs as biosensors,
we chose HIV1 p24 protein as the target molecules. The In
2
O
3
nanoribbon was
functionalized with HIV1 p24 antibodies on the surface to specifically recognize the
target proteins. During the sensing experiment, a 200 mV V
DS
and a 200 mV V
GS
were
In
2
O
3
nanoribbon
OH OH OH P
HOOC
O
HO OH
In
2
O
3
nanoribbon
O O O
P
HOOC
In
2
O
3
nanoribbon
O O O
P
C
H
2
N
NHS/EDC
NH
O
(a)
(b)
(c)
(d) (e)
! 69
applied to the devices and the current between source and drain electrodes of three
devices are monitored simultaneously. HIV1 p24 protein solutions with different
concentrations were added progressively into the sensing chamber.
The normalized current versus time for three monitored devices is plotted in
Figure 4.6. All three devices started to show decrease of current at p24 concentration of
20 pg/mL and progressively larger response with higher concentrations of p24 added.
Note that all three devices show uniform responses when different concentrations of p24
solutions were added. The detection limit of In
2
O
3
FETs for p24 is 20 pg/mL and this
level is already comparable with the detection limit of current commercially available
ELISA kit, which is as specified as 2-20 pg/mL per various venders.
!
Figure 4.6 Real time detection of HIV1 p24 protein. PBS addition would cause signal as
shown in the inset.
4.2.6 Electronic ELISA approach for HIV1 p24 protein detection
Direct electrical detection of biomolecules in their physiological environment is
often impeded by Debye screening from the high salt concentration in the sample
solution. The distance into the semiconducting nanoribbon at which surface charges are
1500 2000 2500 3000
0.5
0.6
0.7
0.8
0.9
1.0
1.1
2 µg/ml 200 ng/ml
20 ng/ml
2 ng/ml
200 pg/ml
20 pg/ml
Sensor 1
Sensor 2
Sensor 3
!
Normalized Current, I/I
0
Time (s)
350 400 450 500 550 600
0.90
0.95
1.00
1.05
1.10
!
!
I/I
0
Time (s)
PBS
0 500 1000 1500 2000
0.0
0.1
0.2
0.3
0.4
Sensor 1
Sensor 2
Sensor 3
!
ΔΙ/Ι
0
Concentration (ng/ml)
(a)
(b)
! 70
no longer felt is defined as the Debye length, !
!
= !!
!
!/(!
!
!), where !, k
B
, T, q, and
n stand for the permittivity (7.97x10
-13
F/cm for In
2
O
3
),
12
the Boltzmann's constant,
temperature, charge constant, and charge density, respectively.
13
To achieve good
sensitivity, the optimal nanoribbon thickness needs to be within the transistor Debye
length. The Debye length for 1x PBS buffer is calculated as 0.76 nm and that of 0.01x
PBS, which is the buffer we used during sensing, is 7.6 nm.
14
This size might be long
enough for the nanoribbon sensor surface to detect the charges brought by the
biomolecules, but the small buffer capacity of the diluted buffer solution may not be
efficient enough to maintain stable pH values for different concentrations of biomolecule
solutions. Moreover, to process physiological sample in our biosensors, the non-specific
interference from other proteins in the sample will be challenging. Tedious sample
preparation is commonly required, such as filtration,
9
purification,
8
etc. Sandwich
ELISA,
15
on the other hand, detects signals associated with the reactions between the
substrate solution and the conjugated enzymes on secondary antibodies instead of the
biomarker. The sandwich structure not only overcomes the Debye screening issue but
also incorporates an amplification scheme to lower the signal to noise ratio (SNR), which
can be much higher for direct analyte detection without amplification, especially when
the amount of analytes are small. Furthermore, sandwich ELISA, which involves several
rinsing steps, will effectively remove the unrelated proteins when dealing with
physiological samples, eliminating the sample preparation steps during clinical
application. In the following In
2
O
3
nanoribbon sensing experiments, we applied an
electronic ELISA technique that uses pH change due to urease enzyme activity as the
amplification signal. The urease enzymes are linked to the secondary antibody through
! 71
biotin and streptavidin. When a solution of urea is introduced to nanoribbon sensor
surface with the sandwich structure, the urea causes an increase in the pH of the solution
due to consumption of hydrogen ions according to the reaction:
16
!"#$+2!
!
!+!
!
!"#$%#
2!"
!
!
+!"#
!
!
The urease deprotonates free hydroxyl groups on the surface of In
2
O
3
nanoribbon, and the
pH increases due to the reduction of positive hydrogen ions and surface potential. The
increase in negative surface charges is responsible for the decrease in conduction of the
n-type In
2
O
3
nanoribbon FETs. The pH change is easily measured by the In
2
O
3
nanoribbon sensors because more charges are released during the pH increase than from
the direct binding between an analyte and a probe antibody. This allows the sensor to
detect very low concentrations of the analyte in physiological samples without the
limitation of the Debye screening effect.
16
We performed electronic ELISA sensing on streptavidin first so that the electrical
signal can be confirmed by the fluorescence signal from the binding between biotin and
streptavidin tagged with fluorescent dye. Figure 4.7(a) shows a schematic diagram of the
sensing setup and the sequence of molecule binding. Cleaned In
2
O
3
nanoribbon devices
were functionalized with 1 mM biotinylated phosphonic acid linkers in methanol for 5.5
hours. Devices were then immersed in a solution of streptavidin conjugated with red
fluorescent dye for 2 hours at room temperature before rinsing off excess streptavidin.
After attaching streptavidin, the devices were incubated in 100 µg/ml biotinylated
urease enzymes in 1xPBS for 2 hours at room temperature before rinsing off any excess
protein. The amount of urease enzymes was determined by streptavidin concentration
anchored on the nanoribbon surface. To perform the sensing, we immersed the devices in
! 72
a baseline solution of 0.01x PBS with pH 7.4. This was then replaced with 100 mM urea
in 0.01xPBS (pH = 6.61) to detect the presence of streptavidin. Figure 4.7(b) shows the
real-time responses when the urea solution was introduced into the sensing chamber. The
urease-urea interaction drastically reduced the device to 11.2% of the baseline signal. The
pH of the final solution in the sensing chamber was measured to be 8.68 by a commercial
Mettler Toledo pH meter. The 1.28 increase in pH is consistent with the decrease in
conduction of the In
2
O
3
nanoribbon device.
After testing with 1 µM streptavidin concentration, we repeated real-time sensing
for other streptavidin concentrations, namely 100 nM, 10 nM, 1 nM, 100 pM, 10 pM, and
1 pM. Figure 4.7(c) shows normalized steady state responses for each of the above
streptavidin concentrations. Each data point is an average of three sensors. The response
increases exponentially with increasing streptavidin concentration. The change in pH
between the final urea solution in the sensing chamber and baseline 0.01xPBS buffer is
plotted against streptavidin concentration as shown in Figure 4.7(d). This relationship
also shows the same exponential trend. With the pH amplification scheme enabled by the
electronic ELISA, our nanoribbon biosensors have detected a streptavidin concentration
that is 4 orders of magnitude lower than the 1 nM reported for In
2
O
3
nanowire sensors
with the same magnitude of the sensing response (about 2%).
9
! 73
Figure 4.7 (a) Schematic diagram of streptavidin electronic ELISA. (b) Normalized real-
time responses of 1 µM streptavidin electronic ELISA from three In
2
O
3
nanoribbon
devices monitored simultaneously. (c) Plot of average normalized current responses and
streptavidin concentration calculated from three devices monitored simultaneously in
each concentration. (d) Plot of pH changes in the sensing chamber measured by a
commercial pH meter and streptavidin concentration.
The ultra low limit of detection demonstrated by our nanoribbon platform is
advantageous for detecting biomarkers like the HIV p24 protein, whose presence even at
an extreme low level can indicate early stage of HIV infection. To perform electronic
ELISA detection for the p24 protein, we functionalized our devices with the phosphonic
linker molecules and used EDC/NHS chemistry to anchor the HIV1 p24 antibodies on the
surface of nanoribbons, as mentioned in the Experimental section. Known concentrations
of the p24 proteins in 1xPBS buffer were then introduced to the sensor for antigen-
antibody binding. Next, the secondary biotinylated HIV1-p24 antibodies were anchored
on the captured proteins by incubation at room temperature for 4 hours before rinsing
10
-3
10
-2
10
-1
10
0
10
1
10
2
10
3
0
20
40
60
80
100
ΔΙ/Ι
0
(%)
SA Concentration (nM)
0 1000 2000 3000 4000 5000
0.0
0.3
0.6
0.9
1.2
Device 1
Device 2
Device 3
I/I
0
Time (s)
0.01x
PBS
100 mM Urea
in 0.01x PBS
10
-3
10
-2
10
-1
10
0
10
1
10
2
10
3
0.0
0.3
0.6
0.9
1.2
1.5
ΔpH
SA Concentration (nM)
(b)
(c)
(d)
[SA] = 1 µM
(a)
! 74
extensively with 1xPBS. After that, devices were incubated in 100 pM streptavidin in
1xPBS to provide binding sites for 0.1 mg/ml biotinylated urease enzymes, which is the
last step before the sensing. Urease enzymes can also be directly linked to the secondary
antibody to reduce the number of binding steps. Figure 5a shows the schematic diagram
for the above sequence of molecules in the electronic ELISA for p24 protein detection.
Figure 5b shows normalized real-time electronic-ELISA sensing responses to
HIV1 p24 at 20 fg/ml monitored from three In
2
O
3
nanoribbon devices simultaneously.
The conduction of the devices was reduced by about 35% when they were exposed to 100
mM urea solution because of the increase in the pH of solution in the sensing chamber
that was induced by reaction between immobilized urease enzymes and urea solution.
Figure 5c shows a plot of average normalized responses from three devices monitored
simultaneously at each different p24 protein concentration from 20 fg/ml to 20 pg/ml.
Responses from electronic-ELISA show exponential relation with p24 concentration as
shown in Figure 5c. Figure 5d shows a plot of pH changes in the sensing chamber before
and after sensing measured by a pH meter and p24 concentration. It also shows an
exponential relationship, which agrees with the trend of changes in conduction in Figure
5c.
In addition, we have spiked known concentrations of p24 proteins in human
serum as the target analytes in several electronic ELISA experiments in order to
demonstrate the capability of our devices to selectively perform sensing in the
physiological solutions. We observed similar changes in electrical conduction and in the
pH of the sensing solution to what we had obtained from p24 sensing in PBS, despite the
fact that blood serum is composed of numerous competing proteins such as human serum
! 75
albumins and human serum globulins. The PBS sensing data are shown as black
rectangles, and human serum data are shown as red triangles in Figures 5c and d,
respectively. The matching of the buffer and the serum data is a good indicator that the
signals from both media are attributed to mainly the p24 proteins and not the competing
proteins in the physiological fluid. These results serve as a good evidence for the
selectivity of our sensors in the complex media because our devices can selectively detect
p24 proteins in human serum. As a result, we can use this electronic ELISA approach in
different kinds of physiological solutions without complicated sample preparation steps,
as competing proteins/biomolecules in the fluids are washed off, leaving only target
analytes immobilized by capture probes. From our approach, we could detect HIV-1 p24
proteins about 3 orders of magnitude lower than limit of detection of the conventional
ELISA approach.
11, 17
! 76
Figure 4.8 (a) Schematic diagram of electronic ELISA for HIV1 p24 detection. (b) Real-
time responses monitored from 3 In
2
O
3
nanoribbon devices simultaneously at 20 fg/mL
of p24 proteins in PBS. Conduction of all devices decreased upon presence of urea in the
sensing chamber. (c) Plot of average normalized responses from 3 devices at each p24
concentration and p24 concentration in pg/mL. (d) Plot of change in pH in the sensing
chamber measured from a commercial pH meter and p24 protein concentration in pg/mL.
4.3 Chapter conclusion
We have demonstrated a top-down approach for In
2
O
3
nanoribbon FET
fabrication using two photolithographic masks to define the position and the dimensions
of the metal electrodes and the nanoribbons. Devices fabricated using this approach show
good, uniform electrical performance without requiring doping or post-process annealing.
The fabrication is highly scalable, low cost, and a low temperature process that is
compatible with the CMOS fabrication facilities. In
2
O
3
nanoribbon devices exhibited
good sensitivity in both wide range of pH solution from pH 4 to 9 and physiological
(b)
(c) (d)
(a)
0 1000 2000 3000 4000 5000
0.6
0.7
0.8
0.9
1.0
1.1
Device 1
Device 2
Device 3
I/I
0
Time (s)
100mM Urea
in 0.01xPBS
[HIV1 p24] = 20 fg/ml
0.01 0.1 1 10
0.0
0.5
1.0
1.5
2.0
p24 in PBS
p24 in Human Serum
ΔpH
HIV p24 Concentration (pg/ml)
0.01 0.1 1 10
0
20
40
60
80
100
p24 in PBS
p24 in Human Serum
ΔI/I
0
(%)
HIV p24 Concentration (pg/ml)
! 77
range between 6.7 and 8.2. Streptavidin-biotin has been chosen to demonstrate signal
amplifying electronic ELISA with picomolar sensitivity showing 15% changes in
normalized current. We demonstrated electronic ELISA for detection of HIV p24
proteins at concentration about 20 fg/ml or 250 viruses/ml, which is about 3 orders of
magnitude lower than the commercial ELISA kit on the market. Depending on choice of
capture probes, our uniform, scalable, sensitive, top-down In
2
O
3
nanoribbon biosensor
platform integrated with the electronic ELISA technique can be utilized for diagnosis of
other infectious diseases and cancers. We believe that our In
2
O
3
nanoribbon platform can
be applied to other biological and medical applications.
4.4 Experimental section
Materials and instrument Amine-terminated biotin was purchased from Pierce.
Streptavidin-Alexa fluro 568 was purchased from Invitrogen. Biotinylated HIV1 p24
antibody, regular HIV1 p24 antibody and antigen were purchased from Fitzgerald.
Biotinylated Urease was purchased from GeneTex. All other chemicals were purchased
from Sigma-Aldrich. XPS was performed on an M-probe surface spectrometer (VG
Instruments). Monochromatic Al KR X-rays (1486.6 eV) incident at 35° from horizontal
were used to excite electrons from the sample, and the emitted electrons were collected
by a hemispherical analyzer at a takeoff angle of 35° from the plane of the sample surface
(horizontal). Fluorescent images were taken by a Nikon Eclipse ME600 fluorescent
microscope equipped with a Microfire digital color camera. The electrical testing of
devices was performed by an Agilent 4156B semiconductor analyzer. The real-time
sensing was done using an Agilent B1500 semiconductor analyzer.
! 78
Device Fabrication 500 nm of Si
3
N
4
thin layer on a SiO
2
/Si wafer was deposited
by LPCVD at 835 degree. A first photolithography was carried out to define the metal
contact. E-beam evaporation metal deposition was performed to deposit 5 nm of Ti and
45 nm of Au for the metal electrodes. Then the second photolithography was performed
to define the pattern for In
2
O
3
deposition of contact leads and active nanoribbon area.
In
2
O
3
nanoribbons were sputtered at room temperature by Denton II sputtering system.
After sputtering, unwanted In
2
O
3
was removed by the lift-off process yielding the pristine
surface.
pH sensing 10 mM phosphate buffer solutions at different pH values were freshly
made and the total ionic strength of all solutions was adjusted to 100 mM by adding NaCl.
A Teflon cell was mounted on the active channel area of In
2
O
3
FETs as the sensing
chamber. A V
DS
=200 mV and a V
GS
=200 mV were applied to the FETs. The pH sensing
started with pH 4 buffer solution and was changed to different pH solutions with pipettes
manually.
Surface Functionalization After device fabrication, In
2
O
3
nanoribbon FETs were
submerged in boiling acetone and isopropyl alcohol for 5 minutes before dried in N
2
stream. To generate hydroxyl groups on the surface of In
2
O
3
nanoribbon to accommodate
phosphonic acid linker molecules, devices were treated by O
2
asher at 100 W 150 mTorr
for 40 s. For biotinylated linker molecules, devices were incubated 1 mM biotinylated
phosphonic acid in methanol for 5.5 hours. After biotinylated liker molecule incubation,
devices were extensively rinsed with methanol. For the normal linker, devices were
submerged in 1 mM aqueous solution of 3-phosphonopropioninc acid for 5.5 hours
before they were rinsed with deionized (DI) water to remove unbound linkers. After
! 79
immobilization of linker molecules, devices were annealed at 120 °C in N
2
environment
for 12 hours to dehydrate surface and to reinforce linkers on the In
2
O
3
nanoribbon surface.
Anchoring of Amine Probe Molecules After devices were functionalized with 3-
phosphonopropioninc acid linker, devices were treated with mixture of 20 mM EDC and
5 mM NHS in DI water for 1 hour to convert from carboxylic acid functional groups to
NHS ester groups for covalently binding of amine molecules on the nanoribbon surface.
After 1 hour, devices were rinsed with DI water before devices were incubated in
solution of amine molecules (100 µg/ml of HIV1 p24 antibodies in 10 mM phosphate
buffer saline (PBS) solution pH 7.4 for 4 hours at room temperature to allow amine
functional groups to react with NHS ester functional groups on the nanoribbon surface.)
Synthesis of Biotinylated Phosphonic Acid Linker The process started with
synthesis of 6- (diethoxyphosphoryl)hexyl 5-(2-oxohexahydro-1H-thieno[3,4-d]imidazol-
4-yl) pentanoate (product 1) as shown in Figure 4.9. A stirred solution of diethyl (6-
hydroxyhexyl)phosphonate (0.952g, 4 mmole) in dry Dimethylformamide (DMF) (30 mL)
was mixed with biotin powder (0.813g 3.3 mmole), EDC•HCl (766 mg, 4 mmole) and 4-
Dimethylaminopyridine (DMAP) (488 mg, 4 mmole). The reaction mixture was stirred at
room temperature under inert atmosphere for 2 hours and poured into water. The aqueous
layer was extracted with ethyl acetate. The organic layer was washed with water, 1 M
sodium hydroxide (NaOH) aqueous solution, 1 M hydrocholic acid (HCl) aqueous
solution and brine, dried over Na
2
SO
4
and concentrated in vacuo to give the crude oil.
The crude residue was purified by flash chromatography (Dichloromethane/Methanol
alcohol, DCM/MeOH, 10:1) to give the purified of product 1 (0.777g, 1.6 mmole, 50%).
Synthesis of (6-((5-(2-oxohexahydro-1H-thieno[3,4-d]imidazol-4-yl)pentanoyl)oxy)
! 80
hexyl) phosphonic acid (product 2) started with a stirred solution of product 1 (0.1 g, 0.2
mmole) in dry DCM (10 mL). Bromotrimethylsilane (0.113 g, 0.75 mmole) was added
into a stirred solution under inert atmosphere. The reaction mixture was stirred at room
temperature for overnight and volatiles were evaporated in vacuo. 10 mL methanol was
added into the residue and the mixture was stirred for 2 hours at room temperature. The
reaction mixture was evaporated under in vacuo to give product 2 (quantitative) as a
hygroscopic solid.
Biomarker sensing In
2
O
3
nanoribbon sensors were functionalized with HIV1 p24
antibody and immersed in 0.01X PBS. During the sensing experiment, a 200 mV V
DS
and
a 200 mV V
GS
were applied to the devices and the source-drain currents of three devices
were monitored simultaneously by an Agilent B1500 analyzer. HIV1 p24 antigen
solutions (prepared in 0.01X PBS) were added into the mixing cell from low to high
concentrations. The non-target BSA solution (10 mg/ml in 0.01X PBS) was added at the
end of the sensing experiment to confirm the sensors’ selectivity.
Electronic ELISA sensing In
2
O
3
nanoribbon sensors were functionalized with
HIV1 p24 antibody and immersed in 1x PBS buffer or human serum solutions of
P
O
EtO
EtO
OH
O
S
NH HN
O
HO
EDC.HCl/DMAP/DMF
P
O
EtO
EtO
O
O
S
NH HN
O
BrSiMe
3
/DCM
P
O
HO
HO
O
O
S
NH HN
O
Product 1
Product 2
! 81
different concentrations of HIV1 p24 antigen for 4 hour each time. After the incubation,
the devices were rinsed by 1x PBS buffer solution vigorously in order to remove the
unbound antigen and other unrelated proteins from human serum. Then the devices were
exposed to 1uM biotinylated secondary HIV1 p24 antibody solution in 1x PBS for 10
min. After vigorous rinsing, 1uM solution of streptavidin in 1x PBS buffer was
introduced in the sensing chamber and the devices were incubated for 10 min. 100 nM
biotinylated urease solution in 1x PBS was added to the sensing chamber after vigorous
rinsing of the previous step. During the sensing experiment, a 200 mV V
DS
and a 200 mV
V
GS
were applied to the devices and the source-drain currents of three devices were
monitored simultaneously by an Agilent B1500 analyzer. 100 µM urea solution in 0.01x
PBS was added to the sensing chamber and the change of current of monitored. After the
current was stabilized, the final pH of the solution inside the sensing chamber was
measured by a pH meter. The sensing step was repeated with different concentrations of
p24 antigen solutions added to the sensing chamber.
4.5 References
1. Gooding, J. J., Electrochemical DNA hybridization biosensors. Electroanalysis
2002, 14 (17), 1149-1156.
2. He, Z.; Gao, N.; Jin, W., Determination of tumor marker CA125 by capillary
electrophoretic enzyme immunoassay with electrochemical detection. Analytica Chimica
Acta 2003, 497 (1-2), 75-81.
3. Tansil, N. C.; Xie, F.; Xie, H.; Gao, Z., An ultrasensitive nucleic acid biosensor
based on the catalytic oxidation of guanine by a novel redox threading intercalator.
Chemical Communications 2005, (8), 1064-1066.
4. Zheng, G.; Patolsky, F.; Cui, Y.; Wang, W. U.; Lieber, C. M., Multiplexed
electrical detection of cancer markers with nanowire sensor arrays. Nature Biotechnology
2005, 23 (10), 1294-1301.
5. Wang, J., Nanomaterial-based electrochemical biosensors. Analyst 2005, 130 (4),
421-426.
! 82
6. Wilson, M. S., Electrochemical immunosensors for the simultaneous detection of
two tumor markers. Analytical Chemistry 2005, 77 (5), 1496-1502.
7. Bunimovich, Y. L.; Shin, Y. S.; Yeo, W. S.; Amori, M.; Kwong, G.; Heath, J. R.,
Quantitative real-time measurements of DNA hybridization with alkylated nonoxidized
silicon nanowires in electrolyte solution. J Am Chem Soc 2006, 128 (50), 16323-31.
8. Stern, E.; Klemic, J. F.; Routenberg, D. A.; Wyrembak, P. N.; Turner-Evans, D.
B.; Hamilton, A. D.; LaVan, D. A.; Fahmy, T. M.; Reed, M. A., Label-free
immunodetection with CMOS-compatible semiconducting nanowires. Nature 2007, 445
(7127), 519-522.
9. Chang, H.-K.; Ishikawa, F. N.; Zhang, R.; Datar, R.; Cote, R. J.; Thompson, M.
E.; Zhou, C., Rapid, Label-Free, Electrical Whole Blood Bioassay Based on
Nanobiosensor Systems. ACS Nano 2011, 5 (12), 9883-9891.
10. Fiebig, E. W.; Wright, D. J.; Rawal, B. D.; Garrett, P. E.; Schumacher, R. T.;
Peddada, L.; Heldebrant, C.; Smith, R.; Conrad, A.; Kleinman, S. H.; Busch, M. P.,
Dynamics of HIV viremia and antibody seroconversion in plasma donors: implications
for diagnosis and staging of primary HIV infection. AIDS 2003, 17 (13), 1871-9.
11. Sutthent, R.; Gaudart, N.; Chokpaibulkit, K.; Tanliang, N.; Kanoksinsombath, C.;
Chaisilwatana, P., P24 Antigen detection assay modified with a booster step for diagnosis
and monitoring of human immunodeficiency virus type 1 infection. J Clin Microbiol
2003, 41 (3), 1016-22.
12. Bellingham, J. R.; Phillips, W. A.; Adkins, C. J., Electrical and optical properties
of amorphous indium oxide. Journal of Physics: Condensed Matter 1990, 2 (28), 6207.
13. Elfstrom, N.; Karlstrom, A. E.; Linnros, J., Silicon nanoribbons for electrical
detection of biomolecules. Nano Lett 2008, 8 (3), 945-9.
14. Noémie, L.; Rune, S. F.; Thor, C. M.; Nathalie, I. R.; Shivendra, U.; Luca De, V.;
Jan, H. J.; Jesper, N.; Karen, L. M., Effects of buffer composition and dilution on
nanowire field-effect biosensors. Nanotechnology 2013, 24 (3), 035501.
15. Butler, J. E., Enzyme-linked immunosorbent assay. J Immunoassay 2000, 21 (2-
3), 165-209.
16. Stern, E.; Vacic, A.; Li, C.; Ishikawa, F. N.; Zhou, C.; Reed, M. A.; Fahmy, T.
M., A nanoelectronic enzyme-linked immunosorbent assay for detection of proteins in
physiological solutions. Small 2010, 6 (2), 232-8.
17. Tang, S.; Hewlett, I., Nanoparticle-based immunoassays for sensitive and early
detection of HIV-1 capsid (p24) antigen. Journal of Infectious Diseases 2010, 201
(Supplement 1), S59-S64.
!
! 83
5 CHAPTER FIVE: Surface Functionalization of In
2
O
3
Nanoribbon
FETs with Histidine Tagged Biological Molecules For Biosensing
Application
5.1 Introduction
Surface functionalization of FETs semiconductor with biomolecules is essential to
fulfill functional biosensors. Most biomolecules are anchored on the surface with amine
coupling.
1
Although the method has been often used for biomolecules, there are
possibilities that the functionality being altered during the process. In particular, it has
been demonstrated that the direct interface between solid surfaces and proteins can affect
the quality of the detected molecular interactions (e.g. the catalytic efficiency of an
enzyme towards its substrate
2
, or antibody–antigen recognition
3
). Self-assembly of
biomolecules to the surface using affinity tags thus appears as an attractive strategy for
the design of FET nanobiosensors. The tags allow a specific immobilization of the
protein and reduce the probability of protein denaturation upon surface immobilization.
Moreover, they can confer a particular orientation to the biomolecules.
4
Histidine tagged
protein is widely applied in protein purification because of the strong affinity between the
histidine tag and divalent metal ions such as Ni
2+
.
5
Histidine tag usually consists 6-10
histidine residues, which is an amino acid that bares an imidazole group. The lone pair
electrons from the N atom of the imidazole moiety coordinate to the empty d orbit of
divalent metal ions, creating coordinate bonding. Therefore, in order to specifically bind
the histidine-tagged biomolecules on FETs, surface of the metal oxide channel needs to
be engineered to have the divalent metal ions exposed for chelating the histidine tag.
! 84
There have been several approaches to anchor divalent metal ions on the surface
of different substrates. The most commonly used reagent is nitrilotriacetic acid (NTA).
3, 6,
7
NTA is a tetradentate chelating agent that could capture the divalent metal ion, leaving
another two vacant binding sites at the metal center for histidine residues to coordinate.
In order to generate a metal oxide surface with NTA moiety, one possible approach is to
synthesize a molecule that bares both phosphonic acid group for surface covalent binding
and NTA moiety for divalent metal coordinating. A study described the synthesis of such
molecule.
8
The synthetic route is tedious and cost-inefficient with poor overall yield.
Moreover, when anchoring NTA-phosphonic acid on metal oxide surface, the
competition binding between carboxylic acid and phosphonic acid to the In
2
O
3
surface
could lead to insufficient amount of free NTA to capture metal ions. Recently, Zhang, et
al., have reported Fe
3
O
4
magnetic particles with imidazole surface for histidine-tagged
protein purification.
9
Herein, we synthesized a phosphonic acid molecule with an
imidazole moiety. After the formation of self-assembly monolayer (SAM) on In
2
O
3
oxide
surface, the imidazole terminal chelates Ni
2+
ions and then successfully captures histidine
tagged streptavidin and histidine tagged MBD2 polypeptide for biosensing application.
The proposed surface functionalization scheme is the first time use of histidine tagged
biomolecules on field effect transistor surface for biosensing.
5.2 Results and discussion
5.2.1 Synthesis of N-imidazolylpropyl phosphonic acid molecule
The N-imidazolylpropyl phosphonic acid linker was synthesized according to the
previous reported procedures
10
as shown in Figure 5.1. Briefly, 3-bromopropyl
diethylphosphonate was prepared via Arbuzov reaction with decent yield. Sodium
! 85
hydride deprotonated imidazole and enabled the nucleophilic attack on the electrophilic
3-bromopropyl diethylphosphonate molecule via S
N
2 pathway, yielding N-
imidazolypropyl diethylphosphonate. The diethylphosphonate was deprotected with
TMSBr and then MeOH to achieve the desired N-imidazolylpropyl phosphonic acid
molecule. The overall synthetic route was straightforward with decent yield.
Figure 5.1 Synthetic scheme for N-imidazolylpropyl phosphonic acid.
5.2.2 Surface functionalization of In
2
O
3
with histidine-tagged biomolecules
In order to specifically anchor histidine-tagged biomolecules, In
2
O
3
surface
should be covered by divalent metal ions, ready for histidine coordination. The procedure
for functionalizing In
2
O
3
surface with histidine-tagged biomolecules was illustrated as in
Figure 5.2.
Figure 5.2 Surface functionalization scheme of In
2
O
3
with hisitidine biomolecule.
OHOHOH OH OH OH
O O O O O O
P P
N N
N
N
100 mM NiCl
2
O O O O O O
P P
N N
N
N
Ni
2+
His-tagged Biomolecule
O O O O O O
P P
N N
N
N
Ni
2+
N
H
O
N
N
N
N
O
R
NH
R'
In
2
O
3 In
2
O
3
In
2
O
3 In
2
O
3
P N
O
HO
OH
N
Br
P(OEt)
3
150
o
C
P Br
O
EtO
OEt
N
HN
60%
NaH
0
o
C
N
N
-
Na
+
P Br
O
EtO
OEt
55
o
C
P N
O
EtO
OEt
N
BrSiMe
3
, MeOH
P N
O
HO
OH
N
70%
! 86
We chose to use imidazole as the functional groups to create Ni
2+
abundant In
2
O
3
surface. Previous study has shown the formation of well-structured SAM on In
2
O
3
surface with phosphonic acid group. Ni
2+
ion has a strong binding affinity with imidazole
groups, and it is rational that we can immobilize Ni
2+
ions onto the imidazole phosphonic
acid functionalized In
2
O
3
surface to anchor histidine-tagged biomolecules. X-ray
photospectrscopy (XPS) spectra were recorded to confirm the immobilization of Ni
2+
ions as well as histidine-tagged biomolecules, as shown in Figure 5.3. The high-
resolution Ni 2p
2/3
spectra (Figure 5.3(a)) illustrate the Ni amount on the surface with
(blue trace) and without (red trace) the imidazole linker molecule. The clean In
2
O
3
surface absorbed small amount of Ni
2+
even when there was no linker molecule present.
However, the amount of Ni increased significantly with the aid of the imidazole linker
molecule. In fact, from the integration area of the Ni 2p
2/3
peak, imidazole linker
molecule captured 16 times more Ni than the pure absorption on the substrate. The
ultimate goal for the surface functionalization strategy is to immobilize histidine-tagged
biomolecules. We applied histidine-tagged streptavidin on the Ni
2+
treated surface.
Figure 5.3(b) shows the traces of high-resolution spectra of N 1s region for different
samples. The significant increase of peak intensity from the black trace (bare In
2
O
3
with
His-tagged streptavidin) to the blue trace (imidazole linker and NiCl
2
treated In
2
O
3
with
His-tagged streptavidin) suggests the success of immobilizing histidine-tagged
streptavidin, as N component is one of the main elements in a protein molecule.
! 87
Figure 5.3 (a) Ni 2p
2/3
XPS spectra of samples functionalized with the linker molecule
and NiCl
2
(blue trace) and without the linker molecule but only NiCl
2
(red trace). (b) N 1s
XPS spectra of sample functionalized with only histidine tagged streptavidin (black
trace), sample functionalized with only NiCl
2
and then histidine tagged streptavidin (red
trace) and sample functionalized with the linker molecule, NiCl
2
and histidine tagged
streptavidin.
More research has been focused on engineering protein mimic peptides to achieve
similar protein activity with much smaller sizes and more versatile functionality.
11-13
Here,
we used histidine tagged MBD2-MBD polypeptide produced from human MBD2 protein
to verify a broader application of the proposed surface functionalization scheme. The
detail of the histidine-tagged polypeptide production procedure can be found in previous
reported literature.
13
We applied the same surface chemistry scheme to anchor the His-
tagged MBD2-MBD polypeptide on the surface of In
2
O
3
substrate. After
functionalization, a 6x-His Epitope Tag Antibody FITC conjugate was applied on the
surface. Negative control was performed without adding NiCl
2
during the
functionalization steps. Figure 5.4 shows the fluorescent images of the sample and the
control substrates. Following the proposed functionalization steps, the surface exhibits
strong fluorescence intensity, suggesting the success of anchoring the His-tagged
polypeptide on In
2
O
3
. On the other hand, the negative control shows an almost dark
(a) (b)
! 88
background with scattering of green dots on the surface. The bright dots were generated
from the non-specific absorption of either the polypeptide or the antibody on the substrate.
Figure 5.4 Fluorescent images of samples treated as the illustrated functionalization
schemes.
5.2.3 Streptavidin/Anti-Streptavidin sensing application
Anchoring histidine tagged biomolecules on In
2
O
3
substrate enables the
possibility to utilize such surface chemistry scheme transforming the In
2
O
3
nanorribon
FETs into nanobiosensors. The device was treated with N-imidazolylpropyl phosphonic
acid molecule followed by NiCl
2
solution. Then the mixing cell was mounted on the
active channel area. Histidine-streptavidin solution in 1x PBS was dropped in the
chamber and the devices were incubated for 2 hour before being rinsed by 1x PBS
vigorously. During the sensing experiment, a 200 mV V
DS
and a 200 mV V
GS
were
applied to the devices and the current between source and drain electrodes of three
devices were monitored simultaneously. Streptavidin antibody (anti-SA) solutions with
different concentrations were added progressively into the sensing chamber.
O O O O O O
P P
N N
N
N
In
2
O
3
100 mM NiCl
2
O O O O O O
P P
N N
N
N
Ni
2+
His-tagged MBD2-MBD
O O O O O O
P P
N N
N
N
Ni
2+
N
H
O
N
N
N
N
O
R
NH
R'
In
2
O
3
In
2
O
3
FITC
6x#His'Epitope'Tag'
Antibody'FITC'conjugate
O O O O O O
P P
N N
N
N
His-tagged MBD2-MBD
O O O O O O
P P
N N
N
N
N
H
O
N
NH
N
NH
O
R
NH
R'
In
2
O
3
In
2
O
3
FITC
6x#His'Epitope'Tag'
Antibody'FITC'conjugate
(a)
(b)
!
!
!
200 µm
!
200 µm
! 89
The normalized current versus time for three monitored devices is plotted in
Figure 5.5. All three devices showed almost no response when adding 0.01x PBS, while
started to show decrease of current at anti-SA concentration of 1 ng/mL (7 pM) and
progressively larger response with higher concentrations of anti-SA added. Note that all
three devices showed uniform responses when different concentrations of anti-
streptavidin solutions were added. The detection limit of the In
2
O
3
nanoribbon FET
biosensors is at 7 pM for streptavidin antibody. This is the first biosensing data that has
been achieved via histidine-tagged biomolecules with field effect transistor.
Figure 5.5 Real time detection of anti-SA with In
2
O
3
nanoribbon FETs biosensors
functionalized with histidine tagged streptavidin
5.2.4 DNA methylation detection
DNA methylation at the 5-position of cytosine in CpG dinucleotides is an
important aspect of physiological processes including embryonic development, X
chromosome inactivation, imprinting and transcriptional regulation.
14-17
While CpG
1000 1500 2000 2500 3000 3500
0.5
0.6
0.7
0.8
0.9
1.0
1.1
1 µg/ml
7nM
100 ng/ml
700pM
10ng/ml
70pM
1ng/ml
7pM
A (I
0A
=1.07µA)
B (I
0B
=1.06µA)
C (I
0C
=1.13µA)
I/I
0
Time (s)
0.01xPBS
! 90
dinucleotides are generally methylated throughout the genome of normal somatic cells,
CpG islands (CGIs), clusters of CpG dinucleotides in gene regulatory regions, are usually
unmethylated.
18
Abberant hypermethylation of CpG island sequence is almost a universal
somatic genome alteration in cancer.
19, 20
Most of the current DNA methylation detection
stratigies use sodium bisulfite to deamine cytosine into uracil, while leaving 5-
methylcytosine intact.
21
All the bisulfite-based strategies are all quite cumbersome and
involve time- and labor-intensive chemical treatments. Moreover, it also requires a
polymerase chain reaction (PCR) process after the bisulfite treatment, which is usually
expensive and time-consuming. Therefore, rapid and sensitive detection of methylated
DNA is of great significance to aid in cancer diagnosis and risk stratification.
Methylated-CpG Binding Domain (MBD) protein binds specifically to methylated
chromosomal DNA in mammalian cells
22
and thus can be used as the probe molecule for
DNA methylation detection with In
2
O
3
nanoribbon FETs. Note that the MBD protein
only binds to hybridized DNA strands with methylated-CpG islands on both sequences.
13
Moreover, DNA strands are highly charged in buffer solution, which is essential to obtain
detection signal with FET biosensors. The MBD2-MBD polypeptide was anchored on the
surface of In
2
O
3
via the surface chemistry described in the previous section. Two
complimentary sequences that are derived from the promoter of P16 gene, a tumor
suppressor gene, were chosen for our experiment. Methylation of CpG islands in the p16
promoter is significantly associated with lung cancer.
23
The sequence we used is shown
as follows:
Probe: 5’-CGGAGGTTGCAGTGAGC/i-MedC/GAGAT/i-MedC/G/i-MedC/GCCA-3’
Target: 5’-TGG/i-MedC/G/i-MedC/GATCT/i-MedC/GGCTCACTGCAACCTCCG-3’
! 91
Because of the nature of the MBD-MBD2 polypeptide, DNA strands need to be
hybridized in advance before they get to be captured by the peptide. We performed the
DNA hybridization by mixing the two strands in 1x PBS buffer solution and diluted to
different concentrations for sensing experiments.
During the sensing experiment, histidine-tagged MBD2-MBD polypeptide was
anchored on the surface. We then introduced different concentration of hybridized
methylated DNA strands to the sensing system. A V
DS
=200 mV and a V
GS
=150 mV was
applied to the In
2
O
3
nanoribbon FETs. The I
DS
-V
GS
sweep was taken after exchanging the
DNA strands solution with clean 0.01x buffer solution after 10 min incubation. We
started the experiment with DNA strands solution at 1 fM and started to increase the
concentration progressively to 100 nM. The conductance started to decrease once the
DNA strands was added, indicating the negatively charged DNA strands were captured
by the polypeptide. The progressive decreasing of the conductance when increasing the
DNA strands concentration up to 100 nM shows the wide dynamic range the sensor
covers. In order to demonstrate the specificity of the sensor, we performed a negative
control. A non-methylated target DNA strand was hybridized with the probe strand
beforehand and the hybridized strands were added at the end of the experiment. Since the
MBD2-MBD polypeptide only binds to DNA strands with methylated CpG islands on
both strands, the negative control DNA strands should not be captured by the peptide,
thus no decreasing of conductance being noticed. As shown in Figure 5.6(a), when
adding the control DNA strands, no obvious conductance change showed. Another
experiment was carried out to prove the binding between the polypeptide and methylated
DNA strands. 100 nM methylated DNA strands were added to a newly functionalized
! 92
sensor and a change of conductance was observed. By repetitive washing of the sensor
with 1x PBS buffer solution, no significant change of the conductance was shown from
the I
DS
-V
GS
sweep, as plotted in Figure 5.6(b). This experiment further confirmed the
interaction between peptide and the DNA strands was strong and the signal was
generated from such interaction.
Figure 5.6 DNA methylation detection with In
2
O
3
nanoribbon FETs. (a) I
DS
-V
GS
sweeps
of In
2
O
3
nanoribbon FETs after adding in different concentration of methylated DNA
strands. Control experiment was performed in the end to demonstrate the specificity. (b)
I
DS
-V
GS
sweeps of sensors when performing repeated washing after adding 100 nM
methylated DNA strands.
5.3 Chapter conclusion
Imidazole terminated monolayer was applied on metal oxide FETs surface for the
first time to immobilize histidine tagged biomolecule for biosensing application. We
synthesized N-imidazolylpropyl phosphonic acid molecule and successfully achieved to
create a Ni
2+
ion abundant surface on In
2
O
3
substrate. Histidine tagged streptavidin and
MBD2-MBD polypeptides were anchored on the surface of In
2
O
3
substrate as proved by
high-resolution XPS spectra and fluorescent images. With the anchoring of histidine
tagged streptavidin on the In
2
O
3
nanoribbon FETs surface, we were able to detect anti-
streptavidin down to 7 pM. DNA hypermethylation detection was also achieved after
(a)
(b)
! 93
immobilizing histidine tagged MBD2-MBD polypeptide on the transistor surface. The
device showed a decrease of conductance from I
DS
-V
GS
sweep at methylated DNA
strands concentration of 1 fM with good specificity. We believe this surface
derivertization strategy provides an alternative approach to anchor histidine protein on
metal oxide surface. The biosensing application with such surface functionalization
scheme demonstrates the possibilities to utilize such approach for general biomarker
detection.
5.4 Experimental section
Materials and instrument Streptavidin antibody, histidine tagged streptavidin
and were purchased from Fitzgerald. 6x-His Epitope Tag Antibody FITC conjugate was
purchased from Pierce. N-imidazolylpropyl phosphonic acid was synthesized following
published procedures. MBD2-MBD polypeptide was generously provided by our
collaborator from John Hopkins University. DNA strands (methylated and unmethylated)
were purchased from Integrated DNA Technologies. All other chemicals were purchased
from Sigma-Aldrich. XPS was performed on an M-probe surface spectrometer (VG
Instruments). Monochromatic Al KR X-rays (1486.6 eV) incident at 35° from horizontal
were used to excite electrons from the sample, and the emitted electrons were collected
by a hemispherical analyzer at a takeoff angle of 35° from the plane of the sample surface
(horizontal). Fluorescent images were taken by a Nikon Eclipse ME600 fluorescent
microscope equipped with a Microfire digital color camera. The electrical testing of
devices was performed by an Agilent 4156B semiconductor analyzer. The real-time
sensing was done using an Agilent B1500 semiconductor analyzer.
! 94
Surface functionalization In
2
O
3
substrate or In
2
O
3
nanoribbon FETs were first
cleaned by boiling acetone and isopropyl alcohol for 5 min, respectively. Then the In
2
O
3
sensors or the substrates were subjected to O
2
asher at 100 W 150 mTorr for 40 s Devices
were then quickly immersed into 1 mM imidazole linker molecule solution for 5.5 hr. 100
mM NiCl
2
solution was added to the surface of In
2
O
3
and the incubation last for 1 hr.
Histidine tagged biomolecules were then introduced to the surface of In
2
O
3
after NiCl
2
treatment. After 1 hr incubation, In
2
O
3
was rinsed by 1x PBS and then DI water
vigorously to remove the unbound biomolecules.
Streptavidin antibody real time sensing In
2
O
3
nanoribbon sensors were
functionalized with histidine tagged streptavidin and immersed in 0.01X PBS. During the
sensing experiment, a 200 mV V
DS
and a 200 mV V
GS
were applied to the devices and
the source-drain currents of three devices were monitored simultaneously by an Agilent
B1500 analyzer. Anti-streptavidin solutions (prepared in 0.01X PBS) were added into the
mixing cell from low to high concentrations.
DNA strands hybridization Two DNA strands with complimentary sequences
were diluted into 1 µM with 1x PBS. Then DNA solutions were immersed in 90 °C hot
water for 1 min separately. After cooling at room temperature for 1 min, the two strands
were mixed well with equal volume by vortex. Then the mixed solution was kept at 4 °C
overnight to finish the hybridization process. The stock solution was stored at -20 °C
after hybridization.
DNA methylation detection In
2
O
3
nanoribbon sensors were functionalized with
histidine tagged MBD2-MBD polypeptide and immersed in 0.01X PBS. During the
sensing experiment, a 200 mV V
DS
and a 200 mV V
GS
were applied to the devices and
! 95
the source-drain currents of three devices were monitored simultaneously by an Agilent
B1500 analyzer. Hybridized DNA strands solutions (prepared in 1X PBS) were added
into the mixing cell from low to high concentrations. The devices were incubated with
the DNA strands solution for 10 min before being exchanged by 0.01x PBS buffer
solution. The I
DS
-V
GS
sweep was then recorded with 0.01x PBS buffer solution. The
process was repeated with increasing concentrations of DNA strands solutions. In the end,
control experiment was carried out by adding unmethylated hybridized DNA strands
solution at 100 nM to the devices. Repetitive washing was carried out after adding 10 nM
methylated DNA strands solution. 0.01x PBS buffer solution was exchanged repetitively
and I
DS
-V
GS
sweeps were recorded each time.
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!
Abstract (if available)
Abstract
Researchers have been developing Point‐Of‐Care Testing (POCT) devices over the past decade to potentially substitute the current lab‐based disease diagnostic technologies with improved portability and efficiency. Nanomaterials‐based field effect transistor (FET) biosensors have become the hot spot for such research because of the small device size comparable to the biological molecules and the instant electrical response for easy electronic integration. Moreover, nanomaterials have a high surface‐to‐volume ratio, which improves the sensitivity of the currently available POCT devices. My PhD research is to develop FET nanobiosensor platforms to potentially become a prototype for universal biomarker detection. ❧ Chapter 1 briefly introduces the structure and working mechanism of FET nanobiosensors. An overview of the surface functionalization approaches is also included. Finally, the chapter discusses the situation of the current commercial POCT devices and the potential to commercialize FET nanobiosensors. ❧ Chapter 2 describes our first‐generation FET nanobiosensor—In₂O₃ nanowire FET biosensor. In this chapter, a laser‐ablation chemical vapor deposition system is applied to synthesize In₂O₃ nanowires and the FETs are fabricated with photolithography. The pH sensing is performed to confirm the device sensitivity. Appropriate surface chemistry is used to functionalize the In₂O₃ nanowire surface for CA-125 biosensing. ❧ Chapter 3 introduces our second-generation FET nanobiosensor—polycrystalline Si nanoribbon FET biosensor. In this chapter, top down fabrication technique is introduced to fabricate the polysilicon nanoribbon FETs. Two routes of surface chemistry are explored to functionalize the polysilicon surface for biosensing. The pH sensing and CA-125 sensing is performed to demonstrate the device sensitivity and the potential to develop biosensors for other biomarkers. ❧ Chapter 4 describes our newest generation FET nanobiosensor—In₂O₃ nanoribbon FET biosensor. The sensor is fabricated via top‐down technique with nearly 100% wafer‐scale device yield and minimal device‐to‐device variation. A novel electronic enzyme‐linked immunosorbent assay (ELISA) combined with the In₂O₃ nanoribbon FET biosensor is introduced in this chapter and this combined approach is able to realize ultrasensitive HIV P24 detection with detection limit comparable to that of PCR. ❧ In the end, chapter 5 studies a newly designed surface functionalization route on In₂O₃ to anchor histidine‐tagged biomolecules. This route consistently immobilizes histidine‐tagged biomolecules on the surface of In₂O₃ for biosensing application, which was demonstrated with streptavidin antibody sensing and DNA hypermethylation detection. This is the first time report of biosensing application of FET using histidine‐tagged biomolecule. With the prevalence of histidine‐tagged biomolecule, this approach can be a universal route to anchor such biomolecules on the surface for other applications.
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Song, Yan
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Core Title
Surface functionalization of nanomaterials and development of field effect transistor nanobiosensors
School
College of Letters, Arts and Sciences
Degree
Doctor of Philosophy
Degree Program
Chemistry
Publication Date
06/19/2015
Defense Date
05/20/2015
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field effect transistor,nanobiosensors,nanomaterials,OAI-PMH Harvest,surface functionalization
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Thompson, Mark E. (
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songyan@usc.edu,sy1987@gmail.com
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University of Southern California Dissertations and Theses
(collection)
Access Conditions
The author retains rights to his/her dissertation, thesis or other graduate work according to U.S. copyright law. Electronic access is being provided by the USC Libraries in agreement with the a...
Repository Name
University of Southern California Digital Library
Repository Location
USC Digital Library, University of Southern California, University Park Campus MC 2810, 3434 South Grand Avenue, 2nd Floor, Los Angeles, California 90089-2810, USA
Tags
field effect transistor
nanobiosensors
nanomaterials
surface functionalization