U.S. patent number 8,017,938 [Application Number 11/723,369] was granted by the patent office on 2011-09-13 for apparatus for microarray binding sensors having biological probe materials using carbon nanotube transistors.
This patent grant is currently assigned to N/A, The United States of America as represented by the Department of Health and Human Services, University of Maryland, College Park. Invention is credited to Konrad Aschenbach, Michael Fuhrer, Romel Del Rosario Gomez, Javed Khan, Herman Pandana, Jun Stephen Wei.
United States Patent |
8,017,938 |
Gomez , et al. |
September 13, 2011 |
Apparatus for microarray binding sensors having biological probe
materials using carbon nanotube transistors
Abstract
A microarray apparatus is provided which contains at least one
chip having source and drain electrodes positioned on an array of
carbon nanotube transistors which allows for electronic detection
of nucleic acid hybridizations, thereby affording both increased
sensitivity and the capability of miniaturization.
Inventors: |
Gomez; Romel Del Rosario
(Silver Spring, MD), Khan; Javed (Derwood, MD), Pandana;
Herman (Lanham, MD), Aschenbach; Konrad (Laurel, MD),
Fuhrer; Michael (Hyattsville, MD), Wei; Jun Stephen
(Gaithersburg, MD) |
Assignee: |
The United States of America as
represented by the Department of Health and Human Services
(Washington, DC)
N/A (College Park, MD)
University of Maryland, College Park (N/A)
|
Family
ID: |
38328650 |
Appl.
No.: |
11/723,369 |
Filed: |
March 19, 2007 |
Prior Publication Data
|
|
|
|
Document
Identifier |
Publication Date |
|
US 20080035494 A1 |
Feb 14, 2008 |
|
Related U.S. Patent Documents
|
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
|
60743524 |
Mar 17, 2006 |
|
|
|
|
Current U.S.
Class: |
257/40; 977/702;
257/E51.04; 977/938; 977/920 |
Current CPC
Class: |
G01N
33/5438 (20130101); B82Y 30/00 (20130101); B82Y
15/00 (20130101); Y10S 977/702 (20130101); Y10S
977/938 (20130101); Y10S 977/92 (20130101) |
Current International
Class: |
H01L
51/10 (20060101); H01L 51/30 (20060101) |
Field of
Search: |
;257/9,20,40,E51.001,E51.002,E51.003,E51.004,E51.005,E51.006,E51.023,E51.024,E51.025,E51.026,E51.038,E51.039,E51.04
;977/700,701,702,703,704,705,706,785,920,921,922,936,938 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
|
|
|
|
|
|
|
WO 2005/000735 |
|
Jan 2005 |
|
WO |
|
WO 2006/024023 |
|
Mar 2006 |
|
WO |
|
Other References
International Search Report and Written Opinion of the ISA for
PCT/US2007/006809, mailed on Aug. 23, 2007. cited by other.
|
Primary Examiner: Gebremariam; Samuel
Assistant Examiner: Arena; Andrew O
Attorney, Agent or Firm: Polsinelli Shughart PC Scott, Jr.;
Teddy C. Jenny; Paul A.
Government Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH
The work leading up to the present invention was funded, at least
in part, by NSA under Grant H9823004C0470. As such, the U.S.
Government may have certain rights in the present invention under
the provisions of 35 U.S.C. 203 et seq.
Parent Case Text
CROSS REFERENCE TO RELATED CASES
This application claims the benefit of Provisional U.S. Application
Ser. No. 60/743,524, filed Mar. 17, 2006, which is incorporated by
reference herein in its entirety.
Claims
What is claimed is:
1. An apparatus comprising: one or more carbon nanotube transistors
on a silicon substrate, the one or more carbon nanotube transistors
each including a gate electrode, a source electrode, a drain
electrode, and a carbon nanotube channel bridging the source
electrode and the drain electrode, wherein the drain electrode and
the carbon nanotube channel are covered by insulating layer,
wherein the insulating layer insulates the drain electrode and the
carbon nanotube channel from an electrolyte solution, wherein the
gate electrode is configured to contact the electrolyte solution
comprising one or more target biological molecules, one or more
specific probe materials immobilized on the insulating layer, an
electronic circuitry to detect a change in electrical charge by
each of the one or more carbon nanotube transistors due to binding
of the one or more specific probe materials with the one or more
target biological molecules, a detector to detect the change in
electrical charge to directly quantify the amount of a bound target
biological molecules; and an automated sensing system for
determining a relative abundance of the specific target biological
molecule based on the change in electrical charge.
2. The apparatus of claim 1, wherein the one or more carbon
nanotube transistors are made by electrically contacting a nanotube
mat, and allowing an electrical current to flow through the
nanotube mat in proximity to a third electrode separated by an
insulating barrier where a voltage applied will cause an electric
field to affect the conductance of the one or more carbon nanotube
transistors.
3. The apparatus of claim 1, wherein the exposed metallic terminals
form the electrical connection to the one or more carbon nanotube
transistors.
4. The apparatus of claim 1, wherein the substrate is a glass or a
non-conducting polymer.
5. The apparatus of claim 1, wherein the thin insulating layer is
an oxide.
6. The apparatus of claim 1, wherein the one or more probe
materials are chosen from DNA, RNA, PNA, antibodies, modified
antibodies, proteins, peptides, aptamers, and peptide aptamers.
7. The apparatus of claim 1, wherein at least one of the exposed
metallic terminals is located on the back side of the
substrate.
8. The apparatus of claim 1, wherein the electronic circuitry is
external to the at least one of the one or more carbon nanotube
transistors.
9. The apparatus of claim 1, wherein the automated sensing system
is external to the at least one of the one or more carbon nanotube
transistors.
10. The apparatus of claim 1, wherein said at least one of the one
or more carbon nanotube transistors is comprised on an array of
about 20 to 50 carbon nanotube transistors.
11. The apparatus of claim 1, wherein the carbon nanotube channel
bridges a gap between the source and the drain electrodes.
12. The apparatus of claim 1, wherein the multiplicity of specific
probe materials comprise cell surface receptor sequences.
13. The apparatus of claim 1, wherein the drain electrode is
gold.
14. The apparatus of claim 1, wherein a separation distance between
the source and drain electrode is about 5 nm.
15. The apparatus of claim 1, further comprising: a gate oxide on
top of a drain source gap to avoid an electrolyte current leakage,
wherein the gate oxide is not insulated by the thin insulating
oxide or the nitride layer.
16. The apparatus of claim 15, wherein the gate oxide is an
aluminum oxide.
17. The apparatus of claim 16, wherein the aluminum oxide is about
100 nm thick.
18. The apparatus of claim 1, wherein the specific probe materials
are immobilized on the insulating oxide or the nitride layer by
silane functionalization.
19. The apparatus of claim 18, wherein the silane functionalization
comprises MPTMS coupled to a hydroxylated surface of the thin
insulating oxide or the nitride layer.
20. A method of electronically detecting a target biological
material in a mixture containing the target biological material and
other biological materials, which comprises: a) exposing the
mixture containing the target biological material and other
biological materials to the at least one chip of the apparatus of
claim 1; and b) determining an absolute amount of the target
biological material in the mixture by a change in electrical charge
due to binding of the target biological material to a probe
material on said at least one chip.
21. The method of claim 20, which further comprises electronically
determining a relative abundance of target biological materials in
a mixture comprising more than one target biological material.
22. The method of claim 20, wherein the apparatus comprises one
chip.
23. The method of claim 20, wherein the target biological material
is DNA.
24. The method of claim 20, wherein the target biological material
is RNA.
25. The method of claim 24, wherein the RNA is miRNA.
26. The method of claim 20, wherein the probe material is DNA, RNA
or PNA.
27. The apparatus of claim 1, wherein the insulating layer
comprises a thin insulating oxide or a nitride.
28. The apparatus of claim 27, wherein the thin insulating layer
includes exposed metallic terminals for providing electrical
conductivity.
Description
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates to an apparatus containing microarray
binding sensors having biological probe materials using carbon
nanotube transistors and various methods for detecting binding of
biological target materials thereto.
2. Description of the Background
DNA microarrays are powerful tools in molecular biology, and
generally contain an array of hundreds to tens of thousands of
genes spotted on a solid substrate, and which is used to identify
and quantify unknown gene samples. The microarray technique is
predicated upon the property that nucleic acid hybridization is
highly specific, i.e., cytosine binds only to guanine and thymine
to adenine. Thus, a specific sequence of nucleic acids, for
example, 5' ATCATC3,' will preferentially bind with its
complementary sequence, 3' TAGTAG5.'
DNA microarrays are invaluable techniques for high throughput
monitoring of gene expression at the transcription level,
determining genome wide DNA copy number changes, identifying
targets of transcription factors, sequencing and, more recently,
for profiling micro RNA (miRNA) levels in cancer. The central dogma
in molecular biology is that DNA is transcribed to ribonucleic acid
(RNA), and the information in the RNA is used to make proteins, by
a process called translation. Since the function and metabolism of
the cell is regulated by the protein produced in the cell, many
diseases caused by gene mutations, such as cancers, can be studied
by monitoring the gene expression. Thus, the identification and
quantification of genes is of particular interest. It is important
to know the particular gene or genes that contribute to a certain
phenotype, and also the amount of the gene that signifies the level
of the gene expression. There are diseases, however, which are not
necessarily caused by gene mutation or change in DNA sequence, but
which are caused by an abnormal amount of the gene or abnormal
level of gene expression. High throughput gene identification
enables researchers to quickly identify the genes that undergo
mutations in a certain disease. Comparative gene expression
compares the level of gene expression, between a cancerous cell and
a healthy cell, for example. In a typical DNA microarray experiment
that relies on fluorescent detection, comparative gene expression
is done by labeling the genes in one cell with one color of
fluorescent reporter molecules, and genes in the other cell with
another. The relative intensity of each color is a direct measure
of the abundance of the genes from the two cells. Given the
versatility of DNA microarrays, the impact thereof on healthcare is
expected to be quite significant if DNA microarrays can be deployed
widely and inexpensively. It will enable rapid diagnosis of
diseases, as well as enable drugs to be tailored to each patient to
achieve highest effectiveness.
The first reported DNA microarray was fabricated on nylon membranes
using cDNA clones and utilized radioactively labeled targets for
detection. Since then, many large-scale DNA microarray platforms
have been developed, which have included, double-stranded cDNA,
single stranded short 25mers (Affymetrix), mid-sized 30mer
(Combimatrix) or long 50-70mers (Nimblegen or Agilent)
oligonucleotides. All of these methods rely upon various
combinations of enzymatic amplification of the nucleic acid and
fluorescent labeling of targets, hybridization, and amplification
of signal followed by detection by optical scanners.
In a microarray experiment, an array of known single stranded DNA
sequences, called probes, is immobilized on a substrate and later
exposed to an unknown set of target genes (or single stranded DNA
sequences) that have been chemically tagged with fluorescent
molecules. In places on the array where the probe and target
sequences are complementary, hybridization occurs and the locations
of these specific binding events are reported by the fluorescent
molecules.
A major hurdle of using DNA microarray as a clinical tool is that
the technique is laborious, requires complex protocols, requires
large amounts of reagents, and suffers from low signal to noise
ratio and rapid optical degradation. While significant strides have
been made in fluorescent-based DNA microarray technology, the
methodologies are often time-consuming and in addition rely on the
determination of fluorescence intensity and the sensitivity is thus
limited by the ability to detect small numbers of photons.
Moreover, fluorescent molecules suffer from photobleaching, which
means that the fluorescent molecule will stop to fluoresce after
receiving a certain amount of excitation.
A variety of DNA detection schemes has been reported in the
literature. The detection mechanisms involve detection of the
existence of the reporter molecules or tags, such as radioisotopes,
fluorophores, quantum dots, gold nanoparticles, magnetic
nanoparticles, or enzymes, for example. A brief survey of known
fluorescent based DNA microarray, and other label-free electronic
field effect DNA detection schemes is described below in
subsections 1) and 2).
1. Fluorescence-Based Microarrays
Typically, microarrays are microscope glass slides spotted with
thousands of different genes. The array does not have built-in
reader. The detection is performed using a fluorescence scanner
after hybridization with fluorescent-tagged target DNA. There are
two ways to make microarrays: (i) spotting cDNA or oligonucleotides
onto the substrate with a robotic spotter, or (ii) direct
oligonucleotide synthesis on the solid support. A robotic spotter
uses thousands of capillary pins dipped into wells containing
different kind of genes and transports the genes onto a
functionalized solid substrate to create gene spots. Another
approach, such as the one employed by affymetrix, uses direct
oligonucleotide synthesis on the substrate. The ingredients are
solutions of the four nucleotides: adenine, guanine, cytosine and
thymine which bear light sensitive protecting group. The process
starts with a quartz wafer that is coated with a light-sensitive
chemical compound that prevents coupling between the wafer and the
first nucleotide of the DNA probe being created. Lithographic masks
are used to either block or transmit light onto specific locations
of the wafer surface. The exposed spots are now ready to couple
with a nucleotide. The surface is then flooded with a solution
containing either adenine, thymine, cytosine, or guanine, and
coupling occurs only in those regions on the glass that have been
deprotected through illumination. The coupled nucleotide also bears
a light-sensitive protecting group, so the cycle of deprotection
and coupling until the probes reach their full length, usually 25
nucleotides.
2. Field Effect DNA Detection
In general, many field effect based biomolecule detection schemes
resemble the structure of ISFET (ion sensitive field effect
transistor), which was first introduced by Bergveld in 1970. IEEE
Transactions on Biomedical Engineering, 17(1): 70-71 (1970). ISFET
is similar to the conventional MOSFET (metal oxide semiconductor
field effect transistor), except that the metal layer is replace by
an ion-sensitive membrane, an electrolyte solution and a counter
electrode. EISFET (electrolyte-insulator-silicon FET) is another
acronym that refers to the same structure. The drain source current
is modulated by field effect from the ions that can reach the
oxide. ISFET technology has been so well-developed that it has made
its way to the market as pH meters. Souteyrand et al. is the first
to demonstrate label-free-homo-oligomer DNA (18-mer and 1000-mer of
poly(dA)DNA) hybridization detection using silicon ISFET. Journal
Physical Chemistry B, 101(15): 2980-2985 (1997). They observed a
shift in the flat-band potential of the underlying semiconductor in
response to the increase of surface charges induced by
hybridization between the complementary homo-oligomer strands.
Several other papers demonstrating successful field effect DNA
detection using silicon ISFET structure are mentioned below.
Pouthas et al. demonstrated field effect detection of 5 .mu.M, 10
.mu.M, 20 .mu.M of 20-mer oligonucleotide and emphasized the need
for low ionic buffer. Physical Review E, 70(3): 031906 (2004).
Fritz et al. were able to detect in real time as dilute as 2 nM of
12-mer oligonucleotide. Proceedings of the National Academy of
Science USA, 99(22): 14142-14146. They utilized poly L-lyssine
(PLL) to immobilize the probe DNA, and claimed that real time rapid
hybridization at low ionic buffer (23 mM phosphate buffer) was
enable by the positively charged PLL surface that compensated for
electrostatic repulsion between complementary DNA strands. Peckerar
et al. demonstrated detection of 1 fM 15-mer DNA. IEEE Circuits
& Davies Magazine 19(2): 17-24 (2003).
Thus, current methods for detecting DNA rely upon various
combinations of enzymatic amplification of nucleic acids and
fluorescent labeling of targets, which entail enzymatic
manipulation of the nucleic acid being tested and chemical
labeling, respectively. These methods are both time consuming and
afford limited sensitivity.
Further, while more recently, DNA microarray technology has been
deployed as a tool for monitoring gene expression patterns and
profiling of micro RNA (miRNA) in normal and cancerous tissue,
quantification of changes has typically been optically-based. While
this technique is highly sensitive, use of optical methods impedes
progress in both system miniaturization and in direct interfacing
with data collection electronics.
Hence, a need exists for a method of detecting DNA that overcomes
these disadvantages.
SUMMARY OF THE INVENTION
Accordingly, it is an object of the present invention to provide an
apparatus for microarray binding sensors having biological probe
materials using carbon nanotube transistors.
It is a more particular object of the present invention to provide
such an apparatus containing at least one chip, each being
positioned on an array of carbon nanotube transistors on an
insulating substrate and covered by a thin insulating oxide or
nitride with exposed metallic terminals.
It is, moreover, another object of the present invention to provide
a method of electronically detecting binding of biological probe
materials to target materials therefor.
It is also an object of the present invention to provide a method
for electronically detecting oligonucleotide-oligonucleotide
binding.
It is further an object of the present invention to provide a
method of forming iron nanoparticle catalysts for carbon nanotube
growth.
Additionally, it is an object of the present invention to provide a
process for preparing an insulating gate material to afford
transistors having improved conductance.
It is, moreover, an object of the present invention to provide a
method of oligonucleotide immobilization.
Further, it is an object of the present invention to provide a
method of electronically detecting biological materials using bound
aptamers.
It is, in addition, an object of the present invention to provide a
method of measuring signal variation as a function of target
material concentration, as well as a method of electronically
determining relative abundances of specific target materials.
The above objects and others are provided by an apparatus
containing:
(a) at least one chip, incorporating an array of carbon nanotube
transistors on an insulating substrate and covered by a thin
insulating oxide or nitride with exposed metallic terminals,
(b) a conducting metal to provide electrical conductivity to the
exposed transistor terminals,
(c) a multiplicity of specific biological probe materials attached
on the insulating oxide or nitride layer,
(d) a microfluidic channel above the insulating oxide or nitride
layer to direct flow of liquid solutions containing biological
material;
(e) electronic circuitry configured to detect change in electrical
charge due to binding of the biological probe materials to target
materials in the biological material of step d),
(f) means configured to quantitatively correlate detected change in
electrical charge from e) into an amount of bound target material,
and
(g) an automated sensing means configured to determine a relative
abundance of specific target materials using the electronic
circuitry of e) and the means of f).
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 illustrates the structure of double stranded DNA bonded by
hydrogen bonding. Electronic detection of DNA, for example, is
possible as the phosphate backbone of the DNA is charged when
ionized or dissolved in solution.
FIG. 2 is a schematic of a carbon nanotube transistor of the
present invention for liquid gating. The drain source current
(I.sub.DS) is modulated by field effect from the DNA charges
adsorbed on the oxide layer.
FIG. 3 is a photograph of a carbon nanotube transistor array of the
present invention. Nanotubes bridge the gap between source and
drain electrodes. (S) is the bonding pad for the common source
electrode, and (G) is the bonding pad for the gate counter
electrode. The square on the other end of (G) is in contact with
the solution. Unmarked squares are bonding pads for drain
electrodes. Shown are 38 transistors in a cell of 3 mm.times.3
mm.
FIG. 4 illustrates the geometry for determining line charge next to
a conducting cylinder.
FIG. 5 illustrates the geometry for determining capacitance between
a conducting cylinder and a plane.
FIG. 6 shows that back gating I-V characteristics improve upon
annealing: (left) before annealing, and (right) after annealing.
Annealing temperature was 500.degree. C. for one hour in
vacuum.
FIG. 7 illustrates a comparison of back-gating and
electrolyte-gating on the same transistor. Triangles of the markers
denote the gate voltage sweep direction (VDS=-0.1 v
FIG. 8 illustrates the coupling mechanism of MPTMS to hydroxylated
silicon oxide surface.
FIG. 9 illustrates an interface of an electrode in an electrolyte
showing positive surface charges as the electrode and the diffused
ions in the outer Helmholtz layer (x>.delta.).
FIG. 10 illustrates charge distribution among various
interfaces.
FIG. 11 schematically illustrates DNA adsorbed on the
oxide-electrolyte interface. The length of 15-mer DNA is about 51
{acute over (.ANG.)}, while the thickness of the inner Helmholtz
layer is typically about 5 {acute over (.ANG.)}. An appropriate
electrolyte concentration is assumed to ensure that the diffuse
layer is larger than the DNA.
FIG. 12 illustrates the case where the charge to be detected is
beyond the reach of the diffuse layers of the oxide-electrolyte
interface.
FIG. 13 illustrates transconductance curves of the carbon nanotube
transistor (VDS=-0.1 v): (ss) after treated with single-strand DNA;
(unmatched hyb) incubated with unmatched sequence of DNA; (ds
matched) hybridized with complete matched DNA. Triangles denote
sweep direction.
FIG. 14 illustrates a feedback circuit to fix I.sub.DS by adjusting
voltage applied to the gate.
FIG. 15 is a cross-sectional view of a carbon nanotube transistor.
A semiconducting carbon nanotube is contacted by two electrodes,
labeled (S) and drain (D) on opposite ends, and covered with an
insulting oxide barrier. The source and drain electrodes have
electrical connections (not shown) to bring signals from outside of
the oxide barrier.
FIG. 16 illustrates voltage connections for transconductance
measurements.
FIG. 17 a scheme for DNA-DNA hybridization. Functionalized probe
DNAs (pr_DNA) are immobilized on the oxide surface using
silane-acrydite binding. A solution containing non-modified target
DNA (tar_DNA) is introduced for complementary hybridization.
FIG. 18 illustrates a single well transistor setup for a single
oligomer sequence, which contains a redundant set of drain and CNT
connections.
FIG. 19 illustrates an array of transistor wells for implementing
hybridization experiments for multiple target sequences.
FIG. 20 illustrates a plot of maximum current versus concentration
for a single transistor sensor. 20 .mu.M probe DNA was used.
Monotonic increase of sensor current from 100 nM and higher. A
nitride-gated transistor was used.
FIG. 21 illustrates sensitivity testing whereby sensitivity of the
apparatus is estimated from the transistor characteristics.
FIG. 22 illustrates response with various buffers and treatments.
Increase in conductance is observed as a function of treatment. The
largest change in conductance is due to complementary binding which
increased conductance from 2 to 6.5. Subsequent washing and
denaturization reduced conductance significantly.
FIG. 23 illustrates characteristics of a transistor fabricated
using aluminum oxide as a gate material. Significantly improved
performance was noted.
FIG. 24 show CNTs grown from Fe catalysts prepared by depositing a
very thin layer (less than 1 nm) of Fe film by using ultra high
vacuum iron deposition techniques.
FIG. 25 shows CNTs grown out of Fe Pt particles prepared using
cluster fabrication technique.
FIG. 26 illustrates an apparatus of the present invention
including, by way of example, a 96 well assay plate.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
The present invention is predicated, at least in part, upon the
provision of a highly sensitive apparatus based on carbon nanotube
transistors for the electronic detection of biological probe-target
binding. In accordance with the present invention, a single carbon
nanotube transistor (CNT) is associated with a distinct biological
probe material, thereby far surpassing existing technology in
sensitivity, ease of use and capability of miniaturization.
Importantly, the present apparatus offers a significant advantage
in simplicity of protocol as the method used therewith does not
require chemical or enzymatic manipulation of the target being
detected.
Further, the present invention does not rely upon optical detection
means so that the present apparatus can be miniaturized. Rather,
the present invention provides, in part, a fabricated nanoplatform
using field-effect transistor (FET) sensing with a gate terminal of
the FET functionalized with an biological probe material of
interest. While either carbon nanotubes or silicon nanowires may be
used as FET, carbon nanotubes are preferred.
The present invention, thus, provides an apparatus for biological
target material detection which uses an array of carbon nanotube
transistors, with each being operated as a field effect transistor.
The current versus voltage characteristics or transconductance
between the source and drain electrodes is measured before and
after a binding event between the biological probe and target
materials. By using a mathematical relationship, the exact amount
of target binding can be extracted. The present apparatus employs
peripheral electronic networks and amplifiers and measurement
algorithms to perform a highly quantitatively measure of the amount
of binding of the biological probe-target materials. The data is
calibrated using well known techniques familiar to artisans in
molecular biology. Thus, the present invention may be used for the
same purposes as conventional DNA microarrays but with the
aforementioned advantages, such as increased sensitivity but
without chemical or enzymatic manipulation of the nucleic acids
being detected.
The difference in current versus voltage characteristics or
transconductance between the source and drain electrodes before and
after a binding, such as an hybridization event, arises from the
known fact that the binding of an oligonucleotide, for example,
such as a target DNA, to a complementary sequence of another chain
of nucleic acid, termed a probe, in a sequence specific manner
results in a total charge change on the probe. This may be
appreciated from FIG. 1, which shows the occurrences of hydrogen
bonds in base pairing (two hydrogen bonds in an adenine (A)-thymine
(T) pair, and three in a cytosine (C)-guanine (G) pair), and
negative charges carried by the phosphate backbone.
The apparatus of the present invention is advantageous as it can be
miniaturized and the methodologies of using the apparatus exhibit
at least five major advantages for detecting DNA, which are:
(i) since the apparatus detects charge, the sample does not need a
labeling step, whereby the process is less laborious;
(ii) since the methodology is label-free, the sensitivity is better
than the fluorescent detection scheme, because in fluorescent
detection schemes, the sensitivity depends both on the
photodetector and the completeness of the fluorescent tagging;
(iii) the methodology does not suffer from photobleaching, and the
apparatus can be used for detection many times if time averaging is
needed;
(iv) application of electric field can increase the hybridization
reaction rate, which increases the throughput of the assay; and
(v) the methodology does not use optical detection means, thus
facilitating miniaturization.
The method described in detail below allows for the sensitive and
specific detection of nucleic acid hybridization without the need
of for extensive chemical or enzymatic manipulation of the DNA or
RNA. The applications of the present apparatus and methodologies
using the same are extensive. A few may be mentioned: 1. Monitoring
Gene expression: research and for diagnostics and predicting
prognosis 2. Monitoring micro RNA (miRNA) expressions, (MicroRNAs
(miRNAs) are small, RNA molecules encoded in the genomes of plants
and animals. These highly conserved, -21-mer RNAs regulate the
expression of genes by binding to the anywhere along the mRNA but
particularly the 3'-untranslated regions (3'-UTR) of specific
mRNAs. They have been reported to be differentially expressed in
cancers of specific types and there is evidence that certain
profiles may predict the patient outcome in cancer. 3. Detecting
DNA copy number changes, performing electronic comparative genomic
hybridization, detecting deletions in chromosomal regions. 4.
Sequencing entire genes, may replace the current gel based
sequencing techniques. 5. Single Nucleotide Polymorphism detection.
6. Detecting pathogens in the air, blood and body secretions.
DEFINITIONS
As used herein, the following terms are defined as follows:
APTAMER: an oligonucleotide or peptide that binds a specific target
molecule. These oligonucletide and/or peptides have been engineered
through in vitro selection or equivalently SELEX (systematic
evolution of ligands by exponential enrichment) to bind to various
molecular targets such as small molecules, proteins, nucleic acids,
and even cells, tissues and organisms. Aptamers often molecule
recognition properties that rival antibodies. CARBON NANOTUBE: A
one-atom thick of graphite (called graphene) rolled up into a
seamless cylinder with a diameter on the order of a nanometer (nm).
The length-to-diameter ratio may be in excess of 10,000. Carbon
nanotubes (CNT) may be either single-walled (SWNT) or multi-walled
(MWNT). MICROARRAY: This refers to, in the case of DNA for example,
a collection of DNA spots attached to a solid surface, such as
glass, plastic or silicon chip. The affixed DNA spots are often
referred to as probes or reporters. Microarrays may be fabricated
using a variety of techniques, such as photolithography using
ink-jet printing. OLIGONUCLEOTIDE: A nucleotide sequence of either
DNA or RNA. The length of a base sequence is often denoted by
`mer`. Thus, a fragment of 15 bases called a 15-mer. MICROFLUIDIC
CHANNEL: A channel having at least one dimension of less than 1 nm.
Common fluids used in microfluidic devices and channels thereof are
blood samples, bacterial cell suspension, and protein or antibody
solutions, for example. The volume of fluids within these channels
is on the order of a few nanoliters (nl). PROBE (OR PROBE
MATERIAL): Any biological material having the ability to bind to a
target material. Examples include segments of DNA, RNA, and
oligonucleotides and polynucleotides, generally; cell receptors or
viruses. As most commonly used, the term refers specifically to
segments of single-stranded DNA or RNA having the ability to bind
or hybridize, and thereby detect, complementary sequences in the
presence of large amount of non-complementary DNA and RNA,
respectively. TARGET (or TARGET MATERIAL): Any biological material
having the ability to bind to a probe material by hybridization,
for example.
Terms in Figures:
In FIG. 9: x. horizontal axis of the figure, drawn perpendicularly
to the electrode-electrolyte, indicating distance away from the
interface. .rho.: the vertical axis of the figure to describe the
profile of the electrical charge density, arising from the
immobilized DNA molecules and rearrangement of ions in the buffer
solution 0: the origin of the axes, denotes oxide-electrolyte
interface .sigma..sub.s: the surface charge density of the
immobilized DNA molecules to the electrode-electrolyte interface.
.delta.: the closest distance from the electrode-electrolyte
interface a solvated ion can approach.
In FIG. 10: x. horizontal axis of the figure, drawn perpendicular
to the electrode-electrolyte interface, indicating distance away
from the interface; .rho.: the vertical axis of the figure to
describe the profile of the electrical charge density arising from
the immobilized DNA molecules and rearrangement of ions in the
buffer solution 0: origin of the x-axis, denoting oxide-electrolyte
interface. -h is the distance from the oxide-electrolyte interface
to the silicon-oxide interface h: denotes the thickness of the
oxide. .sigma..sub.s: the surface charge density of the immobilized
DNA molecules to the electrode-electrolyte interface. .delta.: the
closest distance from the electrode-electrolyte interface a
solvated ion can approach. .rho..sub.l: the charge density induced
in the carbon nanotube. .sigma..sub.2: the surface charge density
induced in the gate electrode. .phi..sub.ox: the voltage drop
across the oxide. .phi..sub.1: the voltage drop across the
oxide-electrolyte interface .phi..sub.2: the voltage drop across
the electrolyte-gate interface. V.sub.app: the applied voltage to
the gate and silicon, required to maintained a fixed charge density
on the carbon nanotube, such that a fixed electrical current is
flowing through the carbon nanotube.
In FIG. 11: .rho.: the vertical axis of the figure to describe the
profile of the electrical charge density, arising from the
immobilized DNA molecules and rearrangement of ions in the buffer
solution .rho..sub.l: the linear charge density induced in the
carbon nanotube. .rho..sub.DNA: the volume charge density of the
immobilized DNA, shown to span to a finite distance to the
electrolyte in the realistic situation, instead of being collapsed
to the interface and treated as a surface charge density in the
simplified model.
In FIG. 12: .rho.: vertical axis of the figure to describe the
profile of the electrical charge density; .rho..sub.ion volume
charge density of negative ions in solution; .rho..sub.S: volume
charge density of the positive ions in solution which surrounds and
screens .sigma..sub.s; .rho..sub.l: linear charge density induced
in the carbon nanotube; .sigma..sub.s: surface charge density of
the immobilized molecule of interest. The figure demonstrates that
the immobilized molecule must be situated close to the interface,
otherwise its change will be screened by the surrounding ions in
the electrolyte and cannot be detected by the device. Any change in
the charge .sigma..sub.s will be screened by the surrounding
diffuse layer and has effect to the transistor.
In FIG. 13: ID: drain current; VGS: voltage across source and gate;
ss: single stranded 15-base oligomer immobilized on the gate; ds:
double stranded DNA, achieved by exposing ss to its complementary
oligomer and hybridized at room temperature; unmatched hyb: ss
exposed to non-complementary sequence and hybridized under same
condition as in ds
In FIG. 14: V is voltage; R is resistance; S is source; D is drain;
I is current; V.sub.os is drain source voltage; and G is gate.
In FIG. 15: S is source; D is drain; G is gate; O.sub.x is
insulating oxide; and G-O.sub.x is gate oxide.
In FIG. 16: V.sub.ds is drain source voltage; and V.sub.gs is gate
source voltage.
THE APPARATUS OF THE PRESENT INVENTION
A. Overview of the Present Apparatus
Semiconducting carbon nanotubes function as channels in between two
conductors and respond to a gating field by modulating the channel
conductance. Carbon nanotubes are very sensitive charge detectors,
and thus are conducive to achieving high sensitivity DNA detection,
for example. The fabrication of a carbon nanotube transistor is
easy, simple and well-adaptable to a flexible substrate. The
cylindrical geometry of the nanotube allows for a less stringent
requirement on gate oxide thickness. In accordance with the present
invention, the nanotube is insulated with silicon oxide or nitride,
for example, and probe material is immobilized on the insulating
layer to avoid direct modification of the nanotube by the probe
material or the electrolyte buffer.
In accordance with another aspect of the present invention, iron
nanoparticles are used as a catalyst for effecting carbon nanotube
growth. In particular, sub-monolayer thin iron films are deposited
by thermal evaporation under less than 10.sup.-10 atm pressure.
Upon exposure to ambient atmospheric pressure, nanoparticles of
iron oxide are formed, and later reduced by high temperature
exposure to hydrogen during carbon nanotube growth.
B. Carbon Nanotube Transistors
The present apparatus contains an array containing up to hundreds
of thousands of carbon nanotube transistors, wire-bonded to a
circuit board that is easily fitted to a platform that houses
circuitry of switches that control the addressing signal to each
transistor. Each transistor is spotted with distinct biological
probe material, such as DNA, RNA, peptide or cell surface receptor
domain, so the device mimics a microarray of such probe material,
with an important exception, being among things, that the
transistor reader is built in. For example, since the DNA backbone
is negatively charged only if it is dissolved, the charge detection
measurement is done with the device in contact with the DNA
electrolyte solution. Thus, the device is well-insulated and
encapsulated in a robust package to prevent shorting among the
leads by the electrolyte solution.
A schematic of a simple apparatus of the present invention is
illustrated in FIG. 2 using DNA as a probe material. The drain (D),
source (S) electrodes and the carbon nanotube channel bridging the
two are insulated from the electrolyte solution by an oxide layer.
Only the gate (G) electrode is in contact with the electrolyte
solution. Probe DNA has been immobilized onto the oxide layer on
top of the carbon nanotube channel. The apparatus does not need to
be wet all the time. Only when charge measurement is performed does
the DNA need to be dissolved. The channel conductance is a function
of the field generated by the DNA charge adsorbed. As a prepared
carbon nanotube usually exhibits p-type conduction, and DNA carries
negative charges, the channel conductance increases upon
hybridization which causes an increase in the number of charges
adsorbed. If the drain-source potential is fixed, an increase in
channel conductance is manifested as a drain-source current
(I.sub.DS) increase.
The present apparatus or device operates in a feedback mode to fix
I.sub.DS by adjusting the gate-source potential (V.sub.GS) upon
hybridization. In this manner, the transistor action mechanism is
decoupled from the electrostatics of the surface charge adsorption.
The V.sub.GS shift reduces to a simple capacitance problem and is
proportional to the change of number of charge adsorbed. Thus, the
abundance of target DNA, for example, can be quantified by looking
at the V.sub.GS shift which indicates the number of hybridization
events. When it is desired to identify DNA, for example, the
present apparatus can achieve the two functionalities of DNA
microarray: (i) gene identification, by looking at which transistor
in the array show channel conductance increase; and (ii)
quantification, by looking at the amount of V.sub.GS shift.
C. Assembling the Present Apparatus
Carbon nanotubes (CNT) are grown on an insulating oxide substrate,
such as a silicon oxide substrate using chemical vapor deposition
(CVD). Generally, the substrate has several hundred nanometers (nm)
thickness of thermal oxide on a p-type silicon substrate. Catalyst
particles, such as iron particles, are deposited on the substrate
either by iron nitrate dripping or brief evaporation of iron to
generate iron dusting on the substrate. The growth is effected in a
furnace at a temperature in excess of about 750.degree. C.,
preferably about 900.degree. C., with appropriate flow of methane,
ethylene, hydrogen and argon gas. Carbon nanotubes grew out of the
catalyst and form a mat on the substrate.
From the prepared CNT mat on the silicon dioxide or nitride
substrate, the next step is to connect the nanotube channel to
source and drain electrodes, by depositing gold film. The substrate
is first spin-coated with photoresist, which is then patterned
using contact photolithography. A chromium adhesion layer and gold
layer are then deposited on the substrate by thermal evaporation;
followed by lift-off to remove the gold layer on top of the
photoresist. A typical source and drain separation or the channel
length is about 5 .mu.m. Unwanted nanotubes are etched out under
oxygen plasma.
FIG. 3 illustrates a transistor array of 38 transistors in one
cell, and all have a common source for economizing space. Each
electrode (drain or source) is connected to contact pad for wire
bonding. In this case, the contact pad is 200.times.200 .mu.m. The
device is intended for liquid gating operation since the DNA should
be dissolved in the buffer solution so there is also a common gate
counter electrode for applying gate voltage.
To provide isolation to the nanotubes from the electrolyte buffer,
a thin silicon oxide layer is evaporated on the device, followed by
the thick plasma enhanced CVD (PECVD) of oxide layer (up to
.about.0.5 .mu.m). The reasons for this redundancy is that the
plasma process destroys the nanotube mat. The oxide layer at the
active area is RIE (reactive ion etched) back again to get a
thickness of .about.100 nm, while leaving thick oxide on the other
part.
Further, the present invention also provides a design and process
for forming insulating gate material. The process entails
depositing aluminum gate oxide by atomic vapor deposition. A thin
layer of less than about 10 nm of silicon nitride is deposited on
the top layer to improve the protection of the underlying carbon
nanotubes from water. The resulting transistors have significantly
improved conductance curves, i.e., less hysterisis, and better
uniformity.
1. Design Parameters
(a) Oxide Thickness
There are two contradicting requirements for the gate oxide
thickness. On the one hand, a thick oxide is desired to minimize
electrolyte current leakage, but on the other hand, a thin oxide is
desired to maximize the gate coupling. The geometry of CNT reduces
the requirement for extremely thin oxides. A nanotube buried under
an oxide layer with surface charges on top of the oxide layer may
be modeled by considering the capacitance between an infinitely
long conducting cylinder and an infinite plane. Thus, can determine
the appropriate oxide layer thickness by considering a standard
textbook problem of an infinite cylinder and a grounded plane.
The problem of the capacitance between an infinitely long
conducting cylinder and an infinite plane is solved by the method
of image. This is an extension of the problem of an infinitely long
line charge .rho.i (C/m) located at a distance d from the axis of a
parallel conducting cylinder of radius .alpha.. One locates an
image line charge (.rho.i) that makes the cylinder surface an
equipotential surface, and the dimensions in the problem is shown
in FIG. 4.
Assigning .rho.i=d.sub.i=a.sup.2 d makes the cylinder surface
equipotential. To solve for the capacitance between a conducting
cylinder and a plane, another cylinder is added as shown in FIG. 5.
The original line change and the image line charge creates
equipotential cylinders around each line charge, with the axes of
the cylinders displaced by d.sub.i from the respective line charge.
If the plane is inserted right at the center between the two line
charges, it will be an equipotential plane, since each point on the
plane is equidistant from both line charges.
The potential difference between the cylinder surface (M) and the
plane (P) can be written as:
.rho..times..pi..times..times..times..times..times..times..times..times..-
times..times..times..times..times..times..times..times..times..times..time-
s..times..times..times..rho..times..pi..times..times..function..times..tim-
es..times..times..times..times..times..times..times..times..times..times..-
pi..function..times..times..times..times..times..times..times..function..t-
imes..times..times..times..times..times..times..times..times..times..times-
..times..times..times..times..times..times..times..pi..times..times..funct-
ion. ##EQU00001##
This simple model suggests that the capacitance is a slow function
of the oxide thickness (h), therefore, the requirement of
fabricating very thin gate oxide to provide optimal gate coupling
is largely alleviated. Thus, thick gate oxide may be used to avoid
electrolyte current leakage without drastic losses in gate
coupling. Typically, about 50 to 200 nm, preferably about 100 nm
thick of oxide, is used on top of the drain source gap in between
which carbon nanotube channels are bridging, and about 250-750 nm,
and preferably about 500 nm, thick of oxide everywhere else.
(b) Contact Resistance
While the carbon nanotube transistor mechanism of action remains
under investigation, there are rationales and evidence therefor
found in the literature suggesting I.sub.DS is controlled by the
gate voltage through: (i) charges induced in the nanotube; (ii)
modulation of Schottky barrier contact between the semiconducting
nanotube and the metallic source and drain electrodes; or (iii)
combination of both (i) and (ii). Unlike conventional MOSFET
(metal-oxide-semiconductor field-effect transistor) where the
source and drain are highly doped silicon, the material of drain
source electrodes for the carbon nanotube transistor used in the
present invention is gold, a different material from the nanotube
channel. Contact resistances between drain and channel junction and
between channel and source junction are therefore inevitable. In
practice, regardless of the physical mechanism behind the
transistor action, the preferred practice is to improve the
performance by annealing. FIG. 6 shows improvement of contact
resistance upon vacuum annealing of the device. The current level
is higher for the same drain-source voltage and the curve shows
less garble after annealing.
2. Device I-V Curve
As depicted in FIG. 2, the apparatus in operation has to be in
contact with the electrolyte solution. Gating is done through the
electrolyte which is in contact with the gate electrode where
V.sub.GS is applied. This is called electrolyte gating. However,
there is also another way of gating, which we call back gating,
i.e. through the back of the body of the silicon which is insulated
from the nanotube channel by 300 nm thick of thermal oxide. FIG. 7
shows comparison between back-gating and electrolyte-gating to the
same transistor. It is clear that electrolyte-gating requires less
voltage range to sweep the transistor on and off, and the
hysteresis effect is less prominent in the electrolyte-gating. The
hysteresis is undesired but unavoidable and caused by trapped
charges in the oxide layer that move around with the applied gate
voltage. In practice, the apparatus is should initially biased
towards one end to always choose the same hysteresis branch with
sweeping in only one direction.
D. Probe Material Immobilization and Hybridization
1. Immobilization
Probe material is immobilized on the insulating oxide or nitride
surface through silane functionalization. The insulting oxide or
nitride surface is exposed to brief oxygen plasma to generate
hydroxyl groups on the surface, on which
(3-mercaptopropyl)trimethoxysilane (MPTMS) can polymerize. The
coupling of MPTMS to hydroxylated silicon oxide surface is shown in
FIG. 8. Making aqueous solution of MPTMS substitutes the methoxy
groups to hydroxyl groups. A water molecule is released during the
coupling reaction, so it is important to perform the coupling in
dry environment. But the polymerization of silane molecule with
other silane molecules is inevitable, leading to formation of large
globule of polymers that induce roughness and heterogeneity on the
surface. After the surface is functionalized with mercaptan groups,
the acrydite-modified probe oligonucleotides, for example, react
readily with the mercaptan groups of the silane to form covalent
bonds by overnight incubation of the probe oligonucleotides. We
have used 15-mer oligonucleotide, for example. The surface is then
treated with 100 mM of sodium acrylate for 15 minutes to passivate
unbound MPTMS.
2. Hybridization
Hybridization with unlabeled or untagged target oligonucleotide is
done under normal hybridization condition, i.e. 10 mM phosphate
buffer solution pH 7, 0.3 M NaCl. Salt is very important to reduce
electrostatic repulsion among two complementary strands to achieve
hybridization. But, high salt or ionic concentration limits the
apparatus sensitivity. Hence, repeated washing steps are necessary
to reduce the salt without causing dehybridization. Washing is done
at least three times with gradual decrease of salt concentration
0.3M, 0.1M, 10 mM, and finally the device is washed with 0.3M
ammonium acetate pH 7, which is know to eliminate salt effectively.
Electrolyte gating measurements are taken before and after
hybridization and are always done under 1 mM phosphate buffer pH
7.
It has been suggested that target-DNA hybridization onto
preimmobilizied probe-DNA on solid substrate follows the Langmuir
adsorption model, which predicts that at high bulk concentration of
the adsorbate, the surface will be fully covered by the
adsorbate:
.GAMMA..GAMMA..beta..times..times..beta..times..times.
##EQU00002##
Where I' is the DNA surface coverage, I'.sub.max is the maximum DNA
surface coverage, C is the concentration of DNA in the bulk
electrolyte, and .beta. is usually extracted from experiment and is
typically in the range of 10.sup.7 M.sup.-1 from fluorescence or
surface plasmon resonance experiment.
E. Electrical Measurements
1. Device Electrostatic Model
(a) Electrolyte Capacitance (Gudy-Chapman-Stern Model)
When an electrode carrying surface charges is immersed in an
electrolyte solution, ionic space charges of opposite sign will
build up in the electrolyte solution. Ions in the space charge
cannot approach the electrode closer than the inner Helmholtz
layer, thus they are called out Helmholtz layer or diffuse layer.
Only chemically specific adsorbed molecules or ions can reside in
the inner Helmholtz layer.
Ions can move around in the electrolyte. The flux of ions consists
of diffusion of ions due to concentration gradient and drift of
ions due to an electric field. One can imagine that ionic space
charges build up close to the charged electrode, and decay with
distance away from the electrode. In the case of thermal
equilibrium where there is no net flux of ions in the solution, a
potential difference is setup to semiconductor p-n junction.
Using the flux equation j=DVc=qc.mu.(-.DELTA.o)=0, and Poisson
equation .differential..sup.2o/.differential.x.sup.2=-p/.epsilon.,
one can write down several important results in one dimensional
case for 1:1 electrolyte (e.g. sodium chloride, which ionizes into
the same amount of Na.sup.+ and CI in the solution).
The relationship between the electric field and the potential at
any arbitrary position in the electrolyte is:
.differential..PHI..differential..times..times..times..function..times..t-
imes..PHI..times..times..times..times..times..times..times..times..times..-
times..times..times..times..times..times..times..times..times..times.>.-
infin..times..times..times..times..times..times..times..times..times..time-
s..times..times..times..times..times..times. ##EQU00003##
.times..times..times..function..times..times..PHI..function..times..times-
..times..times..PHI..times..times..times..times..times..times..times..time-
s..times..times..times..times..times..times..times..times..times..times.&g-
t;.infin..times..times..times..times..times..times..times..times..times..t-
imes..times..times..times..times..times..times. ##EQU00003.2##
.times..times..times..function..times..times..PHI..function..times..times-
..times..times..PHI..times..times..times..times..times.
##EQU00003.3## where o.sub.0 is the potential at x=0.
For small argument of hyperbolic tangent, it can be approximated as
tan h(x).apprxeq.x. We Let x=e {square root over
(2c/k.sub.BT.epsilon.)}, then .phi.(x)=.phi..sub.0 exp(-xx). Since
o.sub.0 is the potential at x=0, and the potential at x=.infin. is
taken to be zero, than o.sub.0 is the potential drop across the
electrolyte.
According to the Stern model, ions cannot go arbitrarily close to
the electrode. The ions have a finite size, they are probably
solvated, or a layer of solvent might separate the ions from the
electrode surface. Imagine that the ions can only go as close as
.delta. to the electrode surface. In order to determine the
relation of the potential drop across the electrolyte and the
surface charge on the electrode surface, from Gauss' law, we can
write:
.sigma..times..times..function..function..differential..PHI..differential-
..times..times..times..times..times..times..times..times..times..times..ti-
mes..times..times..times..times..times..times..times..times..times..times.-
.delta..times..differential..PHI..differential..differential..PHI..differe-
ntial..PHI..function..delta..PHI..delta..times..times..differential..PHI..-
differential..times..times..times..function..PHI..function..delta..times..-
times..times..sigma..times..times..times..function..times..times..times..P-
HI..sigma..times..delta..times..times..times..times..times..times..times..-
times..times..times..times..times..times..times..times..times..PHI..times.-
.times..times..function..sigma..times..times..times..times..times..times..-
sigma..times..delta. ##EQU00004##
The voltage drop derived in this section is based on thermal
equilibrium assumption, which is attainable when good insulation
between the nanotubes and the electrolyte exists. Leakage current
is ignored in this analysis, however electrolyte leakage current is
undesirable, because it means that the electrodes have degraded by
Faraday process.
An applied voltage between the source and the gate counter
electrode is the sum of potential drop across the nanotube channel,
the oxide layer, the oxide-electrolyte interface, and the
electrolyte-gate counter electrode interface, or V
app=o.sub.transistor+o.sub.ox+o.sub.1+o.sub.2. Thermal equilibrium
is assumed where there is no net ionic flux in the electrolyte. The
change of the potential drop across the nanotube channel can be
estimated.
The voltage drop between the electrolyte and the gate counter
electrode from Guoy-Chapman-Stern model derived in the preceding
section is given by:
.PHI..times..times..times..function..sigma..times..times..times..times..s-
igma..times..delta. ##EQU00005##
where now .sigma..sub.2 is the surface charge density on the
counter electrode. This potential drop is expected to be about the
same before and after DNA hybridization, for example, because the
surface charge density depends on the conditions of the electrolyte
and the counter electrode, which are not changed upon
hybridization.
Now looking at the inner Helmholtz layer, to which probe DNA is
attached, for example, Gauss' law is applied at the interface:
.epsilon..sub.cu(-E.sub.cu)+.epsilon..sub.liqE(x=.delta.')=.sigma..sub.s.
where .sigma..sub.s is the surface charge density of the
immobilized DNA, which will increase and at most double upon
hybridization. .delta. is in the order of several Angstrom and
denotes the boundary of inner and outer Helmholtz layers.
The voltage drop across the oxide and electrolyte interface is
then:
.PHI..times..times..times..function..times..times..times..function..delta-
..function..delta..delta..times..times..PHI..times..times..times..function-
..sigma..times..times..times..times..sigma..times..times..delta..times..ti-
mes..times..times..times..times..times..times..times..times..times..times.-
.times..times..times..times..times..times..times..times..times..times..tim-
es..times..times..times..times..times..times..times..times..times..times..-
times..times..rho..times..pi..times..times..times..times..times..times..ti-
mes..times..times..times..times..times..times..times..times..times..times.-
.times..times..times..times..times..times..PHI..times..times..rho..times..-
function..times..pi..times..times..times..times..times..times..times..time-
s..times..times..times..times..times..times..times..times..PHI..PHI..times-
..times..PHI..PHI. ##EQU00006##
(b) Voltage Shift as a Function of Adsorbed Charge
It is assumed that the apparatus is operating in a feedback mode to
keep I.sub.DS constant, so that electrostatics of the adsorbed
charges can be decoupled from the carbon nanotube transistor action
mechanism.
By fixing IDS, is fixed; therefore o.sub.transistor and o.sub.ox
are also fixed. The feedback circuit only needs to compensate for
the change of o.sub.1 due to hybridization. Plugging in typical
numbers: h=100 nm, a=2 nm, .epsilon..sub.ox=3.9 .epsilon.0,
.epsilon..sub.liq=80 .epsilon..sub.0 T=300 K, and .delta.=5 {acute
over (.ANG.)}, can write:
.PHI..function..times..function..sigma..function..times..times..times..ti-
mes..times..times..times..PHI..function..times..function..times..times..si-
gma..function..times..times..times..PHI..times..times..function.
##EQU00007##
Typically all voltages involved are in the order of one or two
volts. If the ionic strength of the electrolyte is c=1 mM, then the
denominator inside the bracket of sin h.sup.-1 is
3.79.times.10.sup.-3, so the first term is the dominant term.
Probe material, such as DNA, cannot reside completely within the
inner Helmholtz layer. Typically the inner Helmholtz layer is in
the order of .delta.=5 .ANG. thick, but the length of 15-mer DNA is
51 .ANG.. The more realistic picture of the interface should look
more like the one in FIG. 11. The free ion concentration is
Boltzmann-distributed in energy. We can write down the Poisson's
equation as:
.differential..times..PHI..differential..times..times..function..infin..t-
imes..function..times..PHI..function..times..rho. ##EQU00008##
.differential..times..PHI..differential..times..times..times..function..P-
HI..function..times..times..function..PHI..function..times..rho..times..fu-
nction..PHI..times..rho. ##EQU00008.2##
Unfortunately, the above equation cannot be solved for a closed
form expression. Typically two extreme cases are solved in the
textbook: (i) quasi-neutrality approximation, and (ii) depletion
approximation.
In quasi-neutrality approximation,
.differential..sup.2o/.differential.x.sup.2 is assumed to be very
small, or the charge is nearly neutral. The positive ions
compensate for the negative DNA charges: c
exp(-e.phi./k.sub.aT).times.|.rho..sub.DNA|
.phi..apprxeq.k.sub.BT/eln(|.rho..sub.DNA|/c) Or
.phi..varies.ln(|.rho..sub.DNA|) (5)
In the other extreme approximation, the depletion approximation,
where the free ions are depleted and cannot compensate for the DNA
charges, we can write:
.differential..times..PHI..differential..rho. ##EQU00009##
.differential..PHI..differential..rho..times. ##EQU00009.2##
(c) Device Sensitivity
Simple calculations can be performed to project the minimum
concentration of the 15-mer target DNA needed to induce appreciable
voltage shift. Let us assume we can detect a change of 51 mV, which
is the prefactor of the sin h.sup.-1 term. It is a safe assumption
indeed, because 51 mV is 2 k8 T/e. The denominator inside the sin
h.sup.-1 bracket is 3.79.times.10.sup.-3, for c=1 mM. If the change
is .sigma..sub.s is equal to 3.79.times.10.sup.-3, then the
o.sub.ox term can be neglected, and the o.sub.1 change is
approximately 51 mV.
To achieve 51 mV in the o.sub.1 change, we need a surface coverage
change of 3.79.times.10.sup.-3 coul/m.sup.2. For 15-mer
oligonucleotide, we assume that each base carries one electron
charge in solution, so each molecule carries
15.times.1.6.times.10.sup.-19 coul. The surface coverage needed is
then 1.57.times.10.sup.15 molecule/m.sup.2. This surface
concentration imposes the requirement for the minimum surface
density of probe-DNA immobilization. That is, if the probe DNA is
not dense enough, the change in surface charge upon hybridization
to target DNA cannot yield the desired voltage change. The maximum
DNA surface coverage is achieved when all of the DNA strands fill
up the surface in upright strand orientation. Assuming the strand
radius is 6 .ANG. gives:
.GAMMA..sub.max=(.pi.r.sup.2).sup.-1=8.8.times.10.sup.17
molecule/m.sup.2.
Using the Langmuir adsorption model, and .beta. value of 10.sup.7
M.sup.-1, we obtain that the target concentration in the bulk
needed to achieve the desired voltage change is .about.20 nM.
Although the sensitivity predicted by the model is only in the
order of 20 nM, but it has been shown experimentally that
sensitivity down to fM is achievable with field effect detection.
This suggests that the assumption of .beta. may actually
underestimate the sensitivity. We project that our approach can
also achieve fM sensitivity.
But the electric field
.differential..PHI..differential..rho..times. ##EQU00010##
.PHI..function..rho..times..times..times..PHI. ##EQU00010.2##
The potential drop across the length (l) of the DNA, is
.phi..sub.D-.phi.(x)=E.sub.0l-|.rho..sub.DNA|l.sup.2/2.epsilon. Or
.phi..varies.|.rho..sub.DNA| (6)
The quasi neutrality approximation predicts that the potential drop
is proportional to the logarithm of the DNA charges, while the
depletion approximation predicts that the potential drop is
linearly proportional to the DNA charges. The realistic case may
lie in between the two approximations. This is actually consistent
with the behavior of the sin h.sup.-1 function previously derived
from the potential drop, because sin h.sup.-1(x).apprxeq.x for
small x, and sin h.sup.-1(x).apprxeq.ln(2x) for large x.
FIG. 12 illustrates a situation where the charge to be detected is
beyond reach of the diffuse layer of the oxide-electrolyte
interface. Any change in the charge .sigma..sub.s will be screened
by the surrounding diffuse layer and has no effect on the
transistor.
Low ionic salt concentration is used in the electrolyte solution
during the measurement, to ensure that the field from the charge to
be detected can reach the carbon nanotube channel. Consider in FIG.
12, where the charge to be detected is farther than the Debye
length. The increase in .sigma..sub.s has no effect on the
transistor because it only modulates the diffuse layer around it,
and the transistor does not feel any electric field from
.sigma..sub.s since the field has been screened by the diffuse
layer. So it is very important that the diffuse layer of the
oxide-electrolyte interface overlaps with the charged biomolecules.
The requirement that .kappa..sup.-1 has to be large imparts
limitation to the ionic strength (c) of the electrolyte, since
.kappa..sup.-1=1/e {square root over (k.sub.BT.epsilon./2c)} is
inversely proportional to square root of c. If we take for example
a 15-mer oligonucleotide and the length of one monomer is 3.4
.ANG., then k.sup.-1>51 .ANG., or c<3.6 mM.
.kappa..sup.-1=1/e {square root over (k.sub.BT.epsilon./2c)}
2. Change of I-V Curve Upon Hybridization: Preliminary Proof of
Principle
FIG. 13 shows the transconductance curve of single strand DNA,
unmatched hybridization, and double strand DNA after completely
matched hybridization. As prepared, carbon nanotube transistors
generally show p-type conduction. Negative charges from DNA
adsorbed on the gate serve as extra negative bias voltage applied
that would increase the conductance of nanotube channel. It is
clear that the current level was boosted after the matched
hybridisation due to the increase number of charged adsorbed. A
slight increase in current is still observed after unmatched
hybridization due to non-specific binding. All of the three
transconductance curves were measured from the same transistor and
taken sequentially from single strand (ss):
5'/Acrydite//Spacer18/ATC CTT ATC AAT ATT -3' (SEQ ID NO:1),
hybridization with unmatched sequence: 5'/ATC CTT ATC AAT ATT -3'
(SEQ ID NO:1) (unmatched hyb), and hybridization with matched
sequence: 5'/AAT ATT GAT AAG GAT -3' (SEQ ID NO:2) (ds
matched).
FIG. 13 illustrates transconductance curves of the carbon nanotube
transistor (VDS=-0.1V): (ss) after treated with single-strand DNA;
(unmatched hyb) incubated with unmatched sequence of DNA; (ds
matched) hybridized with complete matched DNA.
Triangles denote sweep direction.
F. Summary of the Technique for Microarray Application
The aforementioned discussion relates the expected change in the
gate voltage as a function of the amount of charge deposited on the
gate. From a device perspective, one can approximate the behavior
of the CNT transistor as a PMOS operating under the so-called
triode region. Under this assumption, the drain current is modeled
as: i.sub.d=K[2(v.sub.gs-v.sub.t)v.sub.ds-v.sub.ds.sup.2)], (7)
where K is a parameter that determines the sensitivity, vt is the
effective threshold voltage, vds is the applied potential between
drain and source, and vgs is the applied potential between the gate
and source. K and vt are intrinsic properties specific to each
transistor and may vary from transistor to transistor. In the
presence of hybridization, the net effect is the change in the
threshold voltage v.sub.t which will be compensated by a
corresponding change in v.sub.gs
.delta..times..times..differential..differential..times..times..times..de-
lta..times..times..differential..differential..times..delta..times..times.-
.times..times..delta..times..times..times..times..delta..times..times.
##EQU00011##
If we employ a feedback circuit such that the current is kept
constant, then we can (7) to zero to obtain. In other words, the
change in v.sub.gs is equal to, .delta.v.sub.gs=.delta.v.sub.t
(9)
This relationship is independent of the transistor property.
Furthermore, the change in vgs can be used as a direct measure of
the charge of the bound DNA molecules given by equations (5) and
(6) for quasineutral or depletion approximations respectively.
Thus, the measurement of the change in gate voltage after
hybridization can provide a means to directly quantify the amount
of bound charge. This approach is a fundamental enabler for gene
expression experiments without the need for labeling.
The general methodology for microarray use as follows. A chip
containing an array of carbon nanotube transistors is fabricated on
a suitable substrate. The drain, source and gate electrodes are
exposed for electrical contact, and an external multiplexing
circuit is developed for each transistor. A microfluidic channel
made of PDMS or suitable polymer is fabricated on the chip surface
to direct the flow of probe and target oligonucleotide, such as
DNA, solutions. Prior to DNA exposure, a preliminary scan of the
transconductance curves will be measured for each transistor and
will serve to identify working transistors as well as establish the
baseline characteristics. Next, specific DNA solutions with
complementary sequences to the target genes and with appropriate
terminal modification for substrate binding are immobilized on the
transistors. An automated spotter immobilizes different
oligonucleotide sequences throughout the array in complete analogy
with existing fluorescent-based microarrays. Hybridization is then
performed on a target solution, and the voltage shift at constant
current will be measured for each transistor using appropriate
protocols. The operating point of the current is set at the
steepest point of the transconductance curve to simultaneously
increase the sensitivity and decrease the data acquisition time. In
order to implement the equivalent of competitive gene expression
experiments a second chip is fabricated and the same complementary
DNA solutions, for example, are immobilized on the second chip.
Hybridization and measurements of the voltage shifts are conducted
as in the previous case, except that the target solution is
prepared using the second sample. The concentration normalization
chips corresponding to DNA sequences that are known to be conserved
in both samples.
G. Testing Sensitivity of the Apparatus from Transistor
Characteristics
Referring to FIG. 21, the following procedure is used to test the
sensitivity of the present apparatus from transistor
characteristics. Transconductance slope=1.93 .mu.A/V Capacitance=59
fF, area=132 .mu.m.sup.2 To induce 1V of VGS change we need a
surface charge density of:
.times..times..times..times..times..times..times..times..times..times..ti-
mes..times..times..times..times..times..times..times..times.
##EQU00012## Using Langmuir adsorption model to relate surface
density to volume density:
.GAMMA..GAMMA..times..times. ##EQU00013##
.GAMMA..sub.max (maximum surface coverage) is obtained by assuming
the DNA to be a rod like structure with a base of 6 {acute over
(.ANG.)} radius and a height of 3.4 {acute over (.ANG.)} per base.
Then .GAMMA..sub.max.apprxeq.9.times.10.sup.13 molecules/cm.sup.2,
if all DNA stand upright, or
.GAMMA..sub.max.apprxeq.3.times.10.sup.13 molecules/cm.sup.2 if all
15-mer DNA lie down horizontally. Typical K.sub.A is
6.times.10.sup.7 M.sup.-1.
Assuming worst case scenario, that all DNA lie down, we use
.GAMMA..sub.max.apprxeq.3.times.10.sup.13 molecules/cm.sup.2. The
needed volume concentration to achieve
.GAMMA.=3.times.18.6.times.10.sup.9 molecules/cm2 in order to
induce 3 V change or 3.times.1.93 .mu.A increase in drain source
current is: .apprxeq.31 pM.
H. Various Uses of the Present Apparatus
The present apparatus may be used in a large variety of fields and
to accomplish diverse objectives. For example, the apparatus may be
used as a point-of-care clinical tool to examine certain forms of
cancer wherein the knowledge of the expression of certain genes in
the mRNA and miRNA can provide the basis for diagnosis, prognosis
and potential treatment. More significantly, the invention is
potentially applicable towards the sensing of a plethora of
biological materials such as certain proteins, metabolic by
products, drugs and even whole cells. The system is sufficiently
adaptable to incorporate artificial nucleic acids, commonly
referred to as aptamers, that are sensitive to the aforementioned
biological materials. Thus, in general the invention can be used in
the following areas:
Medical diagnosis and Preventive screening,
Drug discovery
Genetics and genetic screening
Bioagent detection for Homeland Security & Fighting Forces
Forensics and law enforcement
Although the apparatus of the present invention may be constructed
with one chip having from one to up to about 100 CNTs, the
inclusion of a plurality of chips in an apparatus is explicitly
contemplated. For example, apparatii having several, dozens,
hundred or thousands of chips are contemplated. See FIGS. 18 and
19, for single chip and multiple chip arrays, respectively.
For example, it is explicitly contemplated to provide 600-700 chips
in an array with from about 25 to 50 CNTs per chip, whereby each
CNT is spotted with a single human gene. Thereby, genes of the
known human genome may be accommodated in order to detect
mutations. This includes genes presently know as well as those yet
to be defined.
Further, it is explicitly contemplated to use cell receptor surface
sequences as probes for viruses. It is well known, for example,
that viral envelope glycoproteins bind to certain cell surface
receptors.
Thus, in accordance with the present invention, any biological
probe material may be used in the microarray to detect target
material or materials in a sample. For example, nucleotide probes
may be used which are complementary to characteristic pathogen
products. See U.S. Pat. No. 4,358,535, which is incorporated herein
by reference in the entirety. As another example, probes may be
used having cell surface receptor domains. See U.S. Pat. Nos.
5,861,479; 6,093,547 and 6,905,685, all of which are incorporated
herein by reference in the entirety. Moreover, aptamer probes may
be used for detection of various drugs. See U.S. Pat. No.
5,789,163, which is incorporated herein reference in the
entirety.
Additionally, as noted above, any conventional method of chip
printing or spotting may be used to prepare the microarrays of the
present invention: See U.S. Pat. Nos. 5,556,752; 6,953,551;
6,656,725; 6,544,698 and 6,594,432, all of which are incorporated
herein by reference in the entirety.
Reference will now be made to certain Examples which are provided
solely for purposes of illustration and which are not intended to
be limitative.
EXAMPLE 1
Use of Catalysts in Forming CNTs
At least two methods may be employed for creating Fe nanoparticles
that act as catalysts for CNT growth: a). Ultra high vacuum
deposition of Fe at a thickness of less than 0.5 nanometers; and
b). Formation of FePt particles using "nanocluster gun" technique
of Ping Wang at Univ. of Minnesota. FIGS. 24 and 25 show the growth
of CNT for each catalyst. Both produce CNTs, but longer and
sparsely distributed CNT's are formed on FePT clusters. Any
conventional method for forming CNTs may then be used with the
Fe-containing catalysts.
EXAMPLE 2
Gold Deposition to Connect NT Channel to S and D Electrodes
Gold contracts are made using standards contact photolithography
techniques and depositing gold to the defined area using physical
vapor deposition technique. Briefly, standard contact
photolithography involves coating the substrate/sample with a
light-sensitive material, or a photoresist, which dissolves or
hardens depending on the type upon exposure to light and subsequent
developer solution. Patterns from a pre-drawn mask are transferred
to the photoresist by exposing it to an ultraviolet light through
the mask. A physical vapor deposition technique involves vaporizing
solid material of interest, which is gold for this particular
purpose, through various methods such as: resistive heating,
electron beam heating, plasma sputtering, such that the vapor
condenses on the substrate/sample thereby creating a thin film on
the substrate/sample. Gold adhesion towards the oxide
substrate/sample is improved by first depositing a thin coating of
chromium or titanium to act as the wetting layer.
EXAMPLE 3
Preparation of Aluminum Oxide Gate Material
A preferred method for creating oxides, such as aluminum or hafnium
oxides is by atomic layer deposition. Atomic layer deposition (ALD)
is a known deposition method in which a film is built up one atomic
layer at a time by saturating the functional group of the surface
with a suitable precursor. The ALD cycle/steps to deposit aluminum
oxide starts by saturating the surface, which inevitably contains
hydroxyl group due to air moisture, with trimethyl aluminum (TMA).
After excess/unreacted trimethyl aluminum is removed, water vapor
is introduced to convert the methyl group of TMA to hydroxyl group,
releasing methane as the byproduct. The newly converted hydroxyl
group is now ready to react with another cycle of TMA exposure.
Cycles of introduction of TMA and water vapor are repeated until
the desired thickness is achieved.
EXAMPLE 4
Probe Attachment to the Insulating Oxide Layer
DNA probes are attached to the insulating oxide layer by a strong
covalent bond between attachment molecule. There are a number of
protocols and chemistry employed for this process. In one approach,
an Acrydite.TM. molecule is added to the probe DNA sequence by
linking it to the 5' terminal of each DNA during DNA synthesis. In
other preparations, a spacer molecule such as a chain or carbon is
included between the probe DNA and Acrydite. Concurrent with DNA
synthesis, the oxide surface is functionalized with a thiol
(sulfur) group of 3-mercaptopropyltrimethoxysilane (MPTMS).
Functionalization of MPTMS to the insulating oxide layer is done in
vapor phase, where a small volume of MPTMS solution (.about.0.1 mL)
is placed at the bottom of glass container and the sample to be
functionalized is mounted faced down at the top of the glass
container. The bottom of the container is then heated
(.about.60.degree. C.) to drive the MPTMS vapor towards the sample
for about 10 minutes. Excess/unreacted chemicals from the sample
are driven off by means of thermal gradients, i.e., the bottom part
of the container is cooled while the top part is heated. DNA probe
attachment is done by pipetting and incubating the solution of the
probe of interest which contains the attachment molecule to the
MPTMS functionalized oxide layer. The chemical constitutes are
shown below.
##STR00001## Since the chemical attachment is dependent upon the
substrate surface, other methods, such as the functionalized of the
surface with amine groups may also be used.
EXAMPLE 5
Exemplary Circuitry Means for Detecting Change in Charge and on
Automated Sensing System
The electric charge of the biomolecule is detected as an apparent
threshold voltage shift in the sensor. An increase in the negative
charge of DNA hybridizing, for example, near the electrolyte-oxide
interface will cause a decrease in drain current of an n-channel
field-effect transistor. A feedback circuit is employed to keep the
drain current constant by adjusting the voltage applied to the
electrolyte (top gate). In the case of DNA, the electrolyte
potential increases in response to an increase of hybridized DNA
and serves as the electrical signal of interest.
Lab-on-a-chip: A self-contained automated DNA hybridization
detection system, for example, employs a custom-design CMOS
(complementary metal-oxide-semiconductor) integrated circuit to
periodically select and measure the sensor signals sequentially at
each element in the microarray. Initially, under buffer electrolyte
in the absence of biological analyte, the system scans through the
entire array, addressing each element via row and column decoders.
After introduction of the analyte and following of the appropriate
hybridization protocol, the signals are measured periodically. With
proper calibration, a greater magnitude of signal change from one
sensor to another indicates a greater degree of sequence expression
in the input sample. Given a profile of the levels of expression of
particular sequences that fingerprint a known bacterium, virus, or
genetic disease, the system then provides an assessment, which may
also be a medical diagnosis of genetic disease or infection by
pathogen.
EXAMPLE 6
The Detection Algorithm
The computer follows a control sequence embedded in permanent
memory once the user pushes a START button underneath the LCD
readout. The computer sequentially records values of the
electrolyte potential unique to each sensor as follows:
The computer selects an array element, RC, by outputting a row word
R and column word C. After a transient delay, the decoder outputs
logic 1's only on lines X(R) and Y(C), turning on transistors
connected to the lines. This allows current to flow through only
sensor RC and into the negative terminal (-) of the current-input
amplifier. The current through Sensor RC is compared with a
reference current generated by diode-connected transistor Mref. If
the current through Sensor RC is higher (lower) than the reference
current, the amplifier reduces (increase) the electrolyte
potential, which will reduces (increase) the current through the
sensor until it equals the reference current. After a programmed
delay to allow for this equilibrium, the computer will record the
electrolyte potential as a digital number, which is generated by
the analog-to-digital converter (ADC). An automated baseline
measurement is made for each transistor prior to exposure to target
analyte and this data is stored on the computer. Hence, by
comparing the baseline with the detection signal, the effective
threshold voltage shift of the transistor at fixed current due to
target capture is accurately measured.
In a similar fashion, the computer sequences through all 96 wells,
A1 to H12. The first time the user presses the START button, the
information acquired by the computer is stored as reference values.
The user then applies a biological protocol to the assay plate and
presses START again. The newly acquired measurements are compared
with the references values. The LCD readout indicates which wells
show an increase in measured values. The readout is compared with a
reference chart generated by the user indicating which wells will
undergo immobilization.
An analog to digital converters (ADC) is an electronic circuit that
converts continuous signals to discrete digital numbers. ADS
circuits are well known. See, for example, U.S. Pat. Nos. 6,407,692
and 6,456,223, each of which is incorporated herein in the entirety
by reference.
EXAMPLES 7
Extracting Concentrations of Bound Targets
The shift in electrolyte potential is an indicator of how much
target molecule has adsorbed on the sensor. In practice, the
relation of this potential shift with the concentration of target
molecule is obtained empirically through sensor calibration
experiments as mentioned above. To understand the response of the
present apparatus, a very simple model is offered that correlates
the potential shift to the electrical charges adsorbed through
total gate capacitance, which is a series of gate oxide
capacitance, and electrolyte double layer capacitance. For the
particular geometry of carbon nanotube transistor in the present
apparatus, the total capacitance is dominated by the gate oxide
capacitance, which has a typical value of 50 fF over an active area
of about 5.times.50 .mu.m.sup.2. This simple capacitance model then
predicts that 1 mV potential shift corresponds to a surface
concentration of 8.3.times.10.sup.6 molecules/cm.sup.2 for 15 bases
DNA. Assuming that the adsorption of the target molecule to the
sensor follows Langmuir model, which relates the surface density to
volume density through the following formula:
.GAMMA..function..GAMMA..GAMMA..GAMMA..times..GAMMA. ##EQU00014##
where .GAMMA. is surface concentration, .GAMMA.max is surface
concentration of maximum coverage, and K.sub.A is empirical
proportional constant. In theory
.GAMMA..sub.max.apprxeq.6.times.10.sup.13 molecules/cm.sup.2 for 15
bases DNA, also theoretical .GAMMA..sub.max is usually much larger
than .GAMMA., and K.sub.A is typically taken to be 6.times.107
M.sup.-1, and K.sub.A is typically taken to be 6.times.10.sup.7
M.sup.-1. Using these numbers, we estimate that the sensitivity of
our apparatus is 2.3 picomolar/mV.
EXAMPLE 8
Hand Held Device or Medical PDA
The apparatus of the present invention is also miniaturizable as a
medical personal digital assistant (PDA). From FIG. 26, a wireless
channel is used between the computer and a remote PDA. Thus, the
data acquired from the present apparatus at a lab site, for
example, may be sent to an appropriate PDA.
In additional to the probe biological materials listed above,
peptide nucleic acids (PNA) may also be used. PNA are widely used
in diagnostic assays and antisense therapies. PNAs are advantageous
in that due to the higher binding strength of PNA/DNA strands (both
single and double) than DNA/DNA strands, it is not necessary to
design long PNA oligomers for such uses, whereas oligonucleotide
probes of 20-25 mer are commonly required therefore. See U.S. Pat.
Nos. 5,582,985; 5,773,571; 6,015,710; 5,786,461 and 6,472,209, each
of which is incorporated herein in the entirety by reference.
While the foregoing specification teaches the principles of the
present invention, with examples provided for the purpose of
illustration, it will be appreciated by one skilled in the art from
reading this disclosure that various changes in form and detail can
be made without departing from the true scope of the invention.
SEQUENCE LISTINGS
1
2115DNAArtificial Sequencechemically synthesized oligonucleotide
probe 1atccttatca atatt 15215DNAArtificial Sequencechemically
synthesized oligonucleotide probe 2aatattgata aggat 15
* * * * *