U.S. patent application number 17/147212 was filed with the patent office on 2021-05-27 for method and device for detecting cellular targets in bodily sources using carbon nanotube thin film.
This patent application is currently assigned to UNIVERSITY OF LOUISVILLE RESEARCH FOUNDATION, INC.. The applicant listed for this patent is UNIVERSITY OF LOUISVILLE RESEARCH FOUNDATION, INC.. Invention is credited to Balaji Panchapakesan.
Application Number | 20210155475 17/147212 |
Document ID | / |
Family ID | 1000005373600 |
Filed Date | 2021-05-27 |
![](/patent/app/20210155475/US20210155475A1-20210527-D00000.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00001.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00002.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00003.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00004.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00005.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00006.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00007.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00008.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00009.png)
![](/patent/app/20210155475/US20210155475A1-20210527-D00010.png)
United States Patent
Application |
20210155475 |
Kind Code |
A1 |
Panchapakesan; Balaji |
May 27, 2021 |
METHOD AND DEVICE FOR DETECTING CELLULAR TARGETS IN BODILY SOURCES
USING CARBON NANOTUBE THIN FILM
Abstract
A device and method detect cellular targets in a bodily source
by utilizing a biofunctional pad comprised of a thin film of carbon
nanotubes (CNT's). When antibodies are absorbed by the CNT's,
cellular targets having markers matching the antibodies may be
detected in a bodily source placed upon the biofunctional pad by
measuring the conductivity of the thin film using conductive
contacts electrically coupled to the thin film, as the binding of
the receptors in the cellular targets to the antibodies changes the
free energy in the thin film. In many respects, the device
functions as a Field Effect Transistor (FET) with the bodily
source, e.g., blood, acting as a polyelectrolyte liquid gate
electrode to create a varying electrostatic charge or capacitance
in the thin film based upon the binding of cellular targets in the
source to the antibodies present on the biofunctional pad.
Inventors: |
Panchapakesan; Balaji;
(South Grafton, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
UNIVERSITY OF LOUISVILLE RESEARCH FOUNDATION, INC. |
Louisville |
KY |
US |
|
|
Assignee: |
UNIVERSITY OF LOUISVILLE RESEARCH
FOUNDATION, INC.
Louisville
KY
|
Family ID: |
1000005373600 |
Appl. No.: |
17/147212 |
Filed: |
January 12, 2021 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
16542973 |
Aug 16, 2019 |
10919759 |
|
|
17147212 |
|
|
|
|
15897851 |
Feb 15, 2018 |
10427938 |
|
|
16542973 |
|
|
|
|
13045135 |
Mar 10, 2011 |
9926194 |
|
|
15897851 |
|
|
|
|
61312913 |
Mar 11, 2010 |
|
|
|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
G01N 27/4146 20130101;
B82Y 15/00 20130101; Y10T 156/1039 20150115; G01N 27/4145 20130101;
B82Y 30/00 20130101 |
International
Class: |
B82Y 15/00 20060101
B82Y015/00; B82Y 30/00 20060101 B82Y030/00; G01N 27/414 20060101
G01N027/414 |
Goverment Interests
GOVERNMENT RIGHTS
[0002] This invention was made with Government support under Grant
No. ECCS 0853066 awarded by the National Science Foundation. The
Government has certain rights in this invention.
Claims
1. A method of fabricating a sensor for detecting cellular targets
in a bodily source, the method comprising: forming a thin film of
carbon nanotubes (CNT's) on a carrier using vacuum filtration;
mechanically bonding the thin film to a dielectric layer on a
semiconductor substrate; separating the thin film from the carrier;
patterning the thin film to form a biofunctional pad; depositing a
plurality of conductive contacts on the substrate, with at least a
portion of each conductive contact overlapping and electrically
coupled to the thin film.
2. The method of claim 1, further comprising annealing the
substrate after depositing the plurality of conductive contacts
thereon.
3. The method of claim 1, further comprising testing the IV
characteristics of the sensor.
4. The method of claim 1, further comprising depositing antibodies
on the biofunctional pad such that at least a portion of the
antibodies are absorbed by the thin film of CNT's.
5. The method of claim 4, wherein depositing the antibodies is
performed either during fabrication of the sensor or during
clinical use of the sensor.
6. The method of claim 1, wherein the cellular target comprises a
cancer cell, and wherein the bodily source comprises a drop of
blood.
7. The method of claim 6, further comprising depositing a layer of
antibodies on the biofunctional pad, wherein the layer of
antibodies is selected from the group consisting of IGF1R, Her2,
EpCAM, and EGFR, and wherein the cellular target comprises a breast
cancer cell.
8. The method of claim 1, wherein the biofunctional pad has a
hydrophobic surface, and wherein the conductivity of the thin film
is inversely proportional to the presence of the cellular target in
the bodily source.
9. The method of claim 1, further comprising altering a surface of
the biofunctional pad to provide the biofunctional pad with a
hydrophilic surface such that the conductivity of the thin film is
proportional to the presence of the cellular target in the bodily
source.
10. The method of claim 9, wherein altering the surface of the
biofunctional pad comprises annealing the substrate after forming
the thin film of CNT's.
11. The method of claim 1, wherein annealing the substrate
comprises annealing the substrate at a temperature of about 200
degrees Celsius to about 400 degrees Celsius.
12. The method of claim 1, wherein annealing the substrate
comprises annealing the substrate at a temperature of at least
about 300 degrees Celsius.
13. The method of claim 9, wherein altering the surface of the
biofunctional pad comprises applying infrared heat to the
biofunctional pad.
14. The method of claim 9, further comprising forming a second
sensor on the substrate that has a second biofunctional pad with a
hydrophobic surface, wherein altering the surface of the first
biofunctional pad is performed without altering the surface of the
second biofunctional pad.
15. A method of detecting cellular targets in a bodily source, the
method comprising: placing a bodily source on a biofunctional pad
comprising a thin film of carbon nanotubes (CNT's) upon which is
disposed antibodies associated with a cellular target; and
measuring the conductivity of the thin film using a plurality of
conductive contacts electrically coupled to the thin film, whereby
the conductivity of the thin film is indicative of the presence of
the cellular target in the bodily source.
16. The method of claim 15, wherein the cellular target comprises a
cancer cell, and wherein the bodily source comprises a drop of
blood.
17. The method of claim 16, wherein the antibodies are selected
from the group consisting of IGF1R, Her2, EpCAM, and EGFR, and
wherein the cellular target comprises a breast cancer cell.
18. The method of claim 15, further comprising placing the bodily
source on a plurality of biofunctional pads, each comprising a thin
film of CNT's upon which is disposed antibodies selected from among
a plurality of antibody types associated with the cellular target
such that the bodily source may be tested against a plurality of
antibody types to detect different markers potentially associated
with the cellular target.
19. The method of claim 15, wherein the biofunctional pad has a
hydrophobic surface, and wherein the conductivity of the thin film
is inversely proportional to the presence of the cellular target in
the bodily source.
20. The method of claim 15, wherein the biofunctional pad has a
hydrophilic surface, and wherein the conductivity of the thin film
is proportional to the presence of the cellular target in the
bodily source.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is a Continuation of, and claims the filing
date benefit of U.S. application Ser. No. 13/045,135, entitled
METHOD AND DEVICE FOR DETECTING CELLULAR TARGETS IN BODILY SOURCES
USING CARBON NANOTUBE THIN FILM, filed on Mar. 10, 2011, which
claims benefit of U.S. Provisional Application No. 61/312,913
entitled METHOD AND DEVICE FOR DETECTING CELLULAR TARGETS IN BODILY
SOURCES USING CARBON NANOTUBE THIN FILM, filed on Mar. 11, 2010.
The contents of each of these related Applications are hereby
expressly incorporated by reference herein in their entireties.
FIELD OF THE INVENTION
[0003] The invention is generally related to detecting cellular
targets in bodily sources, e.g., detecting circulating cancer cells
in blood. The invention is also generally related to the use of
carbon nanotubes (CNT's) in medical diagnostic applications.
BACKGROUND OF THE INVENTION
[0004] Identification and quantitation of numerous biological
molecules to generate a complex molecular profile is required for
diagnosis, monitoring, and prognostic evaluation of complex
diseases such as cancer. Despite outstanding progress in the area
of cancer biology, significant challenges remain in translating
biological knowledge of cancer surface markers into clinically
relevant devices that could be used as diagnostic or monitoring
tools for cancer management. Developing high-throughput and low
cost diagnostic cell and tissue analysis for disease detection has
remained a challenge.
[0005] For example, breast cancer is the most diagnosed cancer in
women, and it has been found that developing breast cancers shed
transformed cells into the blood, with more malignant breast cancer
cells appearing in the blood in later stages. It is believed by
many that early detection of circulating breast cancer cells might
improve diagnosis of early breast cancer and ultimately reduce
breast cancer-related deaths. Therefore, significant efforts have
been made toward the development of methods and devices for
detecting circulating breast cancer cells in blood.
[0006] Circulating tumor cells (CTC's) have long been analyzed ex
vivo by flow cytometry and fluorescence microscopy to measure
characteristic cell surface markers, such as epithelial cell
adhesion marker (EpCAM), a general purpose epithelial cell marker
that is common to circulating tumor cells. Many of these
techniques, however, are expensive and time consuming, often
requiring several days to generate results.
[0007] More recently, it has been found that small bundles of
single wall carbon nanotubes (SWCNT), .about.10 nm diameter,
lithographically patterned between two electrodes, with adsorbed
monoclonal antibodies, will display a sensitivity to a single
cancer cell in 1 .mu.L of blood. Moreover, such devices have the
potential to detect the presence of cancer cells in blood in a
matter of minutes, rather than days as is often the case with other
methodologies. However, the use of single or small bundles of
SWCNT's presents challenges in a clinical setting due to the
difficultly in fabricating such single or small bundle SWCNT
samples, and ensuring that the cancer cells are bridging the
electrodes to achieve reliable detection. Furthermore, the use of
nanoscale devices precludes the use of large blood volumes that are
typically analyzed in a clinical setting.
[0008] Therefore, a need continues to exist in the art for an
improved methodology and device for detecting cancer cells and
other cellular targets in blood and other bodily sources.
SUMMARY OF THE INVENTION
[0009] The invention addresses these and other problems associated
with the prior art by providing a device and method of detecting
cellular targets in a bodily source utilizing a biofunctional pad
comprised of a thin film of carbon nanotubes (CNT's). When
antibodies are absorbed by the CNT's, cellular targets having
markers matching the antibodies may be detected in a bodily source
placed upon the biofunctional pad by measuring the conductivity of
the thin film using conductive contacts electrically coupled to the
thin film, as the binding of the receptors in the cellular targets
to the antibodies changes the free energy in the thin film. In many
respects, the device functions as a Field Effect Transistor (FET)
with the bodily source, e.g., blood, acting as a polyelectrolyte
liquid gate electrode to create a varying electrostatic charge or
capacitance in the thin film based upon the binding of cellular
targets in the source to the antibodies present on the
biofunctional pad.
[0010] Consistent with one aspect of the invention, a device for
detecting cellular targets in a bodily source includes a substrate;
a biofunctional pad comprising a thin film of carbon nanotubes
(CNT's) disposed on the substrate and adapted to receive antibodies
associated with a cellular target; and a plurality of conductive
contacts disposed on the substrate and electrically coupled to the
thin film. The plurality of conductive contacts are configured for
use in detecting the cellular target in a bodily source by
measuring a conductivity of the thin film when the antibodies are
received by the thin film and the bodily source is disposed on the
biofunctional pad and in contact with the antibodies, whereby the
conductivity of the thin film is indicative of the presence of the
cellular target in the bodily source.
[0011] Consistent with another aspect of the invention, a method of
fabricating a sensor for detecting cellular targets in a bodily
source includes forming a thin film of carbon nanotubes (CNT's) on
a carrier using vacuum filtration; mechanically bonding the thin
film to a dielectric layer on a semiconductor substrate; separating
the thin film from the carrier; patterning the thin film to form a
biofunctional pad; and depositing a plurality of conductive
contacts on the substrate, with at least a portion of each
conductive contact overlapping and electrically coupled to the thin
film.
[0012] Consistent with yet another aspect of the invention, a
method of detecting cellular targets in a bodily source includes
placing a bodily source on a biofunctional pad comprising a thin
film of carbon nanotubes (CNT's) upon which is disposed antibodies
associated with a cellular target; and measuring the conductivity
of the thin film using a plurality of conductive contacts
electrically coupled to the thin film, whereby the conductivity of
the thin film is indicative of the presence of the cellular target
in the bodily source.
[0013] These and other advantages and features, which characterize
the invention, are set forth in the claims annexed hereto and
forming a further part hereof. However, for a better understanding
of the invention, and of the advantages and objectives attained
through its use, reference should be made to the Drawings, and to
the accompanying descriptive matter, in which there is described
exemplary embodiments of the invention.
BRIEF DESCRIPTION OF THE DRAWINGS
[0014] FIG. 1 is a top plan view of a cellular detection and
profiling sensor array consistent with the invention.
[0015] FIG. 2 is a top plan view of one of the sensors from the
sensor array of FIG. 1.
[0016] FIG. 3 is a cross-sectional view of the sensor of FIG. 2,
taken along lines 3-3.
[0017] FIG. 4 is a flowchart illustrating exemplary steps in
fabricating the sensor array of FIG. 1.
[0018] FIGS. 5A-5H are cross-sectional views illustrating the
fabrication of the sensor array of FIG. 1 during various of the
steps illustrated in FIG. 4.
[0019] FIG. 6 is a top plan view of an alternate sensor to that of
FIG. 2, incorporating separate drive and sensing contacts.
[0020] FIG. 7 is a flowchart illustrating exemplary steps in
detecting and profiling cellular targets using the sensor array of
FIG. 1.
[0021] FIG. 8 is a chart illustrating I-V characteristics of an
exemplary sensor consistent with the invention functionalized with
IGF1R antibodies for blood samples with varying numbers of MCF7
cells.
[0022] FIG. 9 is a chart illustrating a change of resistance in an
exemplary sensor consistent with the invention as a number of MCF7
cells in blood samples when functionalized with several different
antibodies.
[0023] FIG. 10 is a chart illustrating overexpression ratios
plotted as a function of antibody type, illustrating relatively
high overexpression in IGF1R and EpCAM antibodies, medium
overexpression in Her2 and EGFR antibodies, and low overexpression
in IgG and PSMA antibodies.
[0024] FIGS. 11A and 11B are scanning electronic microscope (SEM)
image of a CNT thin film surface, uncoated (FIG. 11A) and coated
with 10 .mu.L of 5 .mu.g/mL antibodies.
[0025] FIG. 12A is a representative I vs. Vds plots of experimental
stages of an anti-Her2 functionalized CNT thin film transistor.
[0026] FIG. 12B is a plot of I vs. time comparisons for 1 .mu.L of
anti-IGF1R, 1 .mu.L of pure blood, and 1 .mu.L of blood spiked with
5 MCF-7 cells.
[0027] FIG. 13 is a plot of normalized electrical signals as a
function of the number of MCF-7 cells for A) non-specific IgG
antibody B) non-specific PSMA antibody C) specific IGF1R antibody
D) specific HER2 antibody and E) specific EpCAM antibody
functionalized CNT films.
[0028] FIG. 14 illustrates changes in a normalized electrical
signal for devices functionalized with IgG, IGF1R, Her2, EpCAM, and
EGFR antibodies.
[0029] FIGS. 15A and 15B are respective plots of I-Vg for anti-IgG
and MCF7 cells interaction (FIG. 15A), and of I-Vg for anti-IGF1R
and MCF7 cells interaction (FIG. 15B).
[0030] FIG. 16 is a bar graph illustrating percentage changes in an
electrical signal applied across a CNT thin film for samples of
samples of ruby gold nanoparticles, isolated white blood cells,
isolated MCF7 cancer cells, blood and blood mixed with MCF7
cells.
[0031] FIGS. 17A and 17B are plots of current vs. time for blood
samples disposed on CNT thin films with hydrophobic (FIG. 17A) and
hydrophilic (FIG. 17B) surfaces.
DETAILED DESCRIPTION
[0032] Embodiments consistent with the invention use thin films of
carbon nanotubes (CNT's) for detecting surface receptors or markers
in cellular targets in bodily sources, e.g., cancer cells in blood.
As will become more apparent below, when blood mixed with cancer
cells is brought into contact with a thin film of CNT's that has
been functionalized with monoclonal antibodies, the conductivity of
the thin film changes, and typically does so in a manner that is
directly related to the number of cancer cells in blood. Therefore,
by applying a known voltage across the thin film, e.g., through a
pair of conductive contacts or electrodes electrically coupled to
the thin film, the presence of cancer cells can be determined from
the current sensed through the film, with the current decreasing a
function of the number of cells in the blood.
[0033] Among other benefits, the techniques described herein
provide devices that are readily adaptable to clinical
environments, and may have applicability in third world countries
or in other instances where access to health care facilities is
limited. Thin films of CNT's are readily adaptable to batch
fabrication techniques and CMOS/MEMS fabrication techniques. In
addition, as compared to conventional technologies, surface markers
can be detected in blood in a few minutes vs. a few days, and the
level of skill required of the technician may be substantially
reduced.
[0034] One application of the invention, for breast cancer
detection in blood, will hereinafter be the focus of the instant
application. It will be appreciated by one of ordinary skill in the
art that the invention may have applicability in connection with
the detection of other forms of cancer, e.g., prostate cancer, or
the detection of other cellular targets. In addition, the invention
may have applicability in connection with detecting cellular
targets in other bodily sources, e.g., other bodily fluids.
[0035] Turning now to the Drawings, wherein like numbers denote
like parts throughout the several views, FIG. 1 illustrates a
device 10 consistent with the principles of the invention. Device
10 includes a sensor array of sensors 12 disposed on a substrate,
e.g., a silicon or other semiconductor wafer 14. In the illustrated
embodiment, each sensor 12 may be separately configured with
different antibodies to target different cellular targets, and it
may be desirable to provide multiple sensors 12 with the same
antibodies, e.g., to provide the ability to double check results,
or to test blood samples from different patients on the same
device. While device 10 is illustrated with a 3.times.4 array of
sensors 12, it will be appreciated that any number of sensors 12
may be disposed in a given device consistent with the invention.
For example, it may be desirable in some embodiments to utilize 200
sensors 12 so that a 1 mL blood sample may be analyzed in 5 .mu.L
drops applied to the 200 sensors.
[0036] FIG. 2 illustrates one of sensors 12 in greater detail,
while FIG. 3 illustrates a cross-section of one of sensors 12,
taken along lines 3-3 of FIG. 2. Sensor 12 in the illustrated
embodiment is formed on top of a dielectric layer 16 on substrate
16, e.g., a silicon dioxide layer. Sensor 12 includes a
biofunctional pad 18 formed from a thin film of CNT's and a pair of
contacts or electrodes 20 through which the conductivity of the
thin film may be measured. Contacts 20 may be formed of gold or
another conductive material, and partially overlap the thin film in
regions 22.
[0037] In use, biofunctional pad 18 is functionalized to detect a
particular marker on a cellular target by applying antibodies 24 on
a surface thereof. Then, a drop of blood or other bodily source,
illustrated at 26, is deposited on biofunctional pad 18, separated
from contacts 20, and a voltage is applied across contacts 20 to
generate a current that is measured to calculate the conductivity,
i.e., the IV characteristics, of the thin film forming the
biofunctional pad.
[0038] As noted above, each sensor 12 effectively functions as a
Field Effect Transistor (FET) with the bodily source, e.g., blood,
acting as a polyelectrolyte liquid gate electrode to create a
varying electrostatic charge or capacitance in the thin film based
upon the binding of cellular targets in the source to the
antibodies present on the biofunctional pad.
[0039] As discussed in the aforementioned paper, CNT's are
generally p-type materials, and as a result, applying a positive
gate voltage to the thin film of CNT's depletes the carriers and
reduces the overall conductance through the thin film. The
dependence of conductance on gate voltage is ideal for biosensing
applications, as the binding of charged species to the gate
dielectric is analogous to applying a voltage through a gate
electrode. Thus the conductance of a p-type CNT would decrease when
a protein with a positive surface charge binds to an antibody. It
has been found that blood spiked with cancer cells decreases or
increases the conductance of the sensor with increasing number of
cells depending on the net charge. This is a general pattern for
many antibodies, although other antibodies, e.g., EGFR, may
increase the conductance of the thin film with increasing number of
cells. The mechanism is therefore one of electrostatic gating of
the CNT thin film. It is believed that blood spiked with cancer
cells acts as a gate electrode. Varying positive (negative) voltage
at the gate electrode decreases (increases) the conductance of the
device. In this case, increasing the number of cells is equivalent
to increasing the voltage of the liquid gate. While this seems
simplistic, one can also look at the capacitance of the liquid gate
as a function of the Debye length to understand the reason for
excellent gate coupling of blood with increase in cancer cells.
[0040] The total gate capacitance, which determines the charging of
the CNT's under a certain gate voltage, consists of electrostatic
(Ce) and quantum (Cq) components. For back-gating devices, the
capacitance of the gate is given by
C.sub.bg=2.pi..epsilon..epsilon..sub.0/In (2 h/r), where
(.epsilon..epsilon..sub.0) is the gate material dielectric constant
(e.g., 3.9.times.8.85.times.10.sup.-12 F/m), h is the gate oxide
thickness (e.g., 500 nm), and r is the thickness of the CNT film
(e.g., 180 nm). The calculated capacitance per unit length of 500
nm silicon dioxide back gating is about 1.245.times.10.sup.-10 F/m.
For blood as a top gate one can approximate it as a liquid
electrolyte top gate with a capacitance given by: C.sub.liquid
gate=2 .pi..epsilon..epsilon..sub.0/In (r+.lamda..sub.D/r), where
.lamda..sub.D is the Debye length or the electronic screening
length resulting from ions. Now if one assigns a value of .about.10
nm for Debye length, the capacitance per unit length of the liquid
gate is about 6.445.times.10.sup.-8 F/m. Due to the higher
capacitance, better gate channel coupling is achieved. Further, it
can be seen that blood as a gate has a capacitance of two orders
better than a conventional back gated structure. Although the Debye
length varies with different salt concentrations, the estimated
Debye length used in the calculation is still valid because the
total capacitance variation caused by the change in the Debye
length is small as long as it is still on the order of a few
nanometers. Cell surface receptors with a net positive or negative
charged increase or decrease the current in blood depending on
their surface charge. Thus, by optimizing a sensor with thinner CNT
films, one can achieve excellent gating in a liquid environment
that can be used as a mechanism for sensing circulating cancer
cells.
[0041] In addition, it is believed that change in current in a
sensor consistent with the invention is also related to the
extracellular and intra-cellular potentials of the cellular
targets. This makes the sensor and method of its use highly
specific for specific cell types.
[0042] It should also be noted that, after etching and patterning
the thin film of CNT's, the surface of the thin film becomes highly
hydrophobic due to the nature of the single walled CNT's (SWCNT's).
The hydrophobic nature of the surface causes a deposited blood
droplet to remain on a specific spot on the biofunctional pad
without shorting the contacts. The hydrophobic nature of the
surface also causes the antibodies to diffuse slowly and arrange
themselves on the surface of the CNT's, and makes it possible for
the adsorbed antibodies to interact with the cell surface antigens
and create the change in conductivity.
[0043] With additional reference to FIGS. 5A-5H, a sensor
consistent with the invention may be fabricated using a process 50
shown in FIG. 4. Process 50 begins in block 52 by forming a CNT
wafer, e.g., by a vacuum filtration technique such as disclosed in
Lu et al., "Nanotube micro-optomechanical actuators," Applied
Physics Letters 88, 253107 (2006). The wafer includes a thin film
18 of CNT's, e.g., about 100 nm to about 150 nm in thickness,
deposited on a mixed cellulose ester (MCE) carrier or filter 28
(FIG. 5A).
[0044] Next, in block 54, the substrate is prepared by forming a
dielectric layer 16 on a silicon or other semiconductor wafer 14
(FIG. 5B), e.g., through oxidation of the silicon wafer to create a
layer of silicon dioxide.
[0045] Next, in block 56, the CNT wafer is bonded to the substrate
through mechanical compression while heating to about 75 degree
Celsius, bonding thin film 18 to dielectric layer 16 (FIG. 5C).
Thereafter, the CNT wafer filter 28 is removed in block 58 using an
acetone vapor bath to dissolve the filter away from the thin film
of CNT's (FIG. 5D).
[0046] Next, in block 60, the CNT thin film is etched by patterning
a photoresist mask 30 using a lithographic process to cover the
regions of each biofunctional pad, and then etching the remaining
CNT thin film using an etching technique such as deep reactive-ion
etching (DRIE) (FIG. 5E). Thereafter, the photoresist mask is
removed.
[0047] Next, in block 62, electrodes or contacts are deposited,
first by patterning a photoresist mask 32 using a lithographic
process to expose the regions of each contact (FIG. 5F) and then
depositing a conductive material such as gold, aluminum, copper,
platinum, or other conductive metal or alloy providing a low
contact resistance, e.g., via sputtering or other suitable
deposition technique. The contacts may be deposited, for example,
to a thickness of about 100 nm. Thereafter, the photoresist mask is
removed, resulting in contacts 20 being formed with overlapping
regions 22 (FIG. 5G).
[0048] Next, in block 64, the assembly is annealed, e.g., at about
150 to about 200 degrees Celsius in an Argon or other inert gas
atmosphere for about 20 minutes. Doing so improves the contact
between each contact or electrode and the thin film, thereby
lowering the contact resistance so that the bulk of the contact
resistance in the sensor is due to the binding of cellular targets
to the antibodies absorbed into the thin film of CNT's.
[0049] Next, in block 66, it may be desirable to test the sensors
on the wafer, e.g., by measuring the IV characteristics of the
sensors by applying a voltage across the contacts 20. In addition,
Raman spectroscopy may be performed to characterize the CNT's in
the thin film of each sensor and identify any potentially defective
films.
[0050] Subsequent to testing, it may also be desirable during
fabrication, as shown in block 68, to deposit antibodies 24 on the
biofunctional pad (FIG. 5H). Antibodies may be deposited, for
example, via drop coating of pure antibodies, or via covalent
bonding or other known techniques. In one suitable technique,
antibodies may be drop coated onto a biofunctional pad and allowed
to set for about 10 minutes to enable the antibodies to diffuse and
be absorbed into the thin film, and then wash away the remaining
liquid using deionized (DI) water. Depending upon the type of
antibodies and the environmental robustness thereof, it may also be
desirable to autoclave or freeze the sensors during the fabrication
process to preserve the antibodies on the surface of the
biofunctional pad until the sensors are ready to be used.
[0051] In an alternate embodiment, however, the sensors may be
fabricated without antibodies deposited thereon, requiring the
antibodies to be deposited immediately prior to use in a clinical
environment.
[0052] In the embodiment illustrated in FIGS. 1-5H, each sensor
includes a biofunctional pad 18 of about 1.5 mm.times.1.5 mm, with
each contact 20 being about 1 mm.times.1 mm and overlapping the
biofunctional pad in regions 22, sized about 0.1 mm.times.0.3 mm.
In this embodiment, it may be desirable to drop coat about 5 .mu.L
to about 10 .mu.L of antibodies at a concentration of about 5
.mu.g/mL, which results in an about 20 nm thick layer of antibodies
deposited on the biofunctional pad.
[0053] While a pair of contacts, disposed over two adjacent corners
of the biofunctional pad, is used in each sensor 12, other
configurations and numbers of contacts may be used in a sensor
consistent with the invention. For example, FIG. 6 illustrates an
alternate sensor 80 including a thin layer of CNT's 82 and four
contacts 84, 86 overlapping in regions 88. One pair of contacts 84
disposed on opposite corners of biofunctional pad 82 may be used as
drive pads, through which a current is passed, and the other pair
of contacts 86 may be used as sensing pads, through which the
resistance or conductivity of the biofunctional pad is measured. In
other embodiments, contacts may be disposed in other positions,
e.g., overlapping the edges, of a biofunctional pad.
[0054] FIG. 7 next illustrates a process 100 for testing a blood
sample using device 10 of FIG. 1. Process 100 begins in block 102
by optionally depositing the antibodies on the biofunctional pads
in the manner discussed above, if not already so done during
fabrication. Next, in block 104, blood droplets (e.g., about 5
.mu.L) are deposited on the biofunctional pads, without contacting
the contacts, either manually or via a robotic system. Then, in
block 106, the IV characteristics of the sensor are recorded over
time by applying a known voltage, e.g., up to about 25 mV across
the contacts thereof and measuring the current. In many instances,
each sensor will stabilize within several minutes, e.g., 5 minutes
or so, once all cellular targets in the blood bind with the
antibodies. Thereafter, in block 108, the results may be analyzed
to determine what cellular targets were found in the blood based
upon what markers were expressed with the different antibodies on
different sensors in the array. In addition, it may also be
possible to predict or determine the number of cellular targets
within each blood sample, e.g., the number of cancer cells
predicted in a given drop of blood, as the IV characteristics of
each sensor will change based upon the number of cells, and thus
the number of bindings that occur with the antibodies.
[0055] The types of antibodies used to test a given blood sample
may be different in different embodiments and applications. For
example, to test for the presence of breast cancer, it may be
desirable to utilize IGF1R, Her2, EpCAM, and EGFR antibodies on
different sensors, while to test for the presence of prostate
cancer, it may be desirable to utilize PSMA antibodies. It may even
be desirable to utilize antibodies that express for different types
of cancer on the same sensor array so that a single blood sample
may be tested for the presence of multiple types of cancer. Other
combinations of antibodies may also be used for other diagnostic
applications of the invention.
[0056] The provisional application cross-referenced herein
discusses test results performed with sensors fabricated in the
manner disclosed herein. A portion of these results are illustrated
in FIGS. 8-10. Detection and profiling of 10-300 MCF7 breast cancer
cells in 5 .mu.L aliquots of blood (the typical reported range for
circulating tumor cells in the blood of patients with metastatic
breast cancer) was performed. Incubation of blood spiked with
cancer cells resulted in unambiguous decreases in sensor
microjunction conductance, where pure blood resulted in higher
conductance (lower resistance) of the microjunctions compared to
blood spiked with cancer cells. The sensor was able to detect a
minimum of 10 MCF7 cells in 5 .mu.L of blood, as well as the
maximum number of 300 MCF7 cells in blood for several different
antibodies. FIG. 8, for example, illustrates the I-V
characteristics of sensors functionalized with IGF1R antibodies
using blood with 10-300 MCF7 cells, and shows a measurable decrease
in conductivity with an increase in the number of MCF7 breast
cancer cells in the blood.
[0057] FIG. 9 illustrates the change in resistance and conductance
of sensors vs. the number of MCF7 cancer cells in a 5 .mu.L blood
sample for various antibodies, including IGF1R, EpCAM, and Her2, as
well as non-specific IgG and PSMA antibodies, and shows that
anti-IGF1R, anti-EpCAM, and anti-Her2 showed a measurable
conductance change for MCF7 cells in blood as compared to the
non-specific antibodies.
[0058] One potential way to scale cellular measurements is to
determine a calibration curve between the cellular overexpression
of a surface antigen and the change in electrical signal. In
current clinical practice, diagnoses are mainly reported as the
presence or absence of malignant cells in the specimen. The
capability to quantify, profile, and stratify cancer cells would
likely improve diagnosis. A critical issue when screening cancer
cells is how to correlate the expression levels of tumor markers to
the number of malignant cells in a given sample. Without this
knowledge one could either measure high expression in relatively
few cells or low expression in many cells.
[0059] One can define an overexpression ratio as .DELTA.R.sup.Max#
of cells/.DELTA.R.sup.Min # of cells. This description is quite
appropriate as this ratio increases with overexpression. One of the
outcomes of results illustrated in FIG. 9 is that when one replots
the data as overexpression ratios, one finds that the ratio of
.DELTA.R.sup.300/.DELTA.R.sup.10 (IGF1R)=7.07,
.DELTA.R.sup.300/.DELTA.R.sup.10 (EpCAM)=5.4,
.DELTA.R.sup.300/.DELTA.R.sup.10 (Her2)=3.9, .DELTA.R.sup.300
/.DELTA.R.sup.10 (EGFR)=2.9, .DELTA.R.sup.300/.DELTA.R.sup.10
(IgG)=1.27, and .DELTA.R.sup.300/.DELTA.R.sup.10 (PSMA)=0.77. This
is shown in FIG. 10.
[0060] As can be seen in this figure, while the change in
resistance of IGF1R in FIG. 9 was lower than EpCAM and Her2, the
overexpression ratios were the highest for IGF1R. This shows that
the definition of overexpression is indeed valid, and is in fact
consistent with Western Blot analysis, which shows a similar
overexpression of IGF1R in MCF7 cells compared to Her2. Further,
the results also indicate specific numbers for EpCAM, EGFR and Her2
which are all valid surface markers for breast cancer. These
overexpression ratios from 1.0 to 7.0 may also be assigned
malignancy.
[0061] A ratio of 1.0 may be considered benign or negative for that
marker and 7.0 may be considered malignant or positive.
Furthermore, based on these numbers one can scale the number of
cells. From the aforementioned results, it is believed that
plotting the number of cells against their overexpression ratios
may give a linear change that can actually predict the number of
cells in blood. Furthermore, by using more markers, one may be able
to increase the accuracy of this technique.
[0062] It has also been found that, in some embodiments, it may
also be desirable to alter the hydrophobicity or hydrophilicity of
a biofunctional pad to alter the response characteristics of the
biofunctional pad. It has been found, in particular, that a CNT
thin film is typically hydrophobic in nature, and that the presence
of a cellular target in a bodily source disposed on a biofunctional
pad tends to decrease the conductivity of the CNT thin film in the
biofunctional pad such that the conductivity of the thin film is
inversely proportional to the presence of the cellular target in
the bodily source. However, by treating the biofunctional pad to
alter the physical structure of the CNT thin film, the
hydrophilicity of the CNT thin film may be increased, and notably,
the response of the biofunctional pad may be altered such that
conductivity increases, rather than decreases, in response to the
presence of a cellular target in a bodily source disposed on the
biofunctional pad, such that the conductivity of the thin film is
proportional to the presence of the cellular target in the bodily
source.
[0063] FIG. 11A, for example, illustrates an SEM image of the
entangled nature of CNT's in a thin film. This type of entangled
network presents ideal surface characteristics for cells to stick
to such a surface. It has been found that the entangled nature of
CNT's presents ideal surfaces for antibodies to stick to the
surface even for non-covalent functionalization methods. The CNT
surfaces show high degree of hydrophobicity due to the exposure of
carbon nanotubes in a oxygen plasma during device patterning
thereby creating rough surface. Adsorption of an antibody can
decrease the surface energy thereby decoupling its surface
wettability from bulk properties and enabling hydrophobicity. In
addition, in some embodiments, the surface chemistry can be
tailored with molecules such as silane to even create a
superhydrophobic surface with high contact angles. Such surfaces
are self-cleaning and therefore can enable variety of medical
related devices.
[0064] In one experimental implementation, for example, the
electrical responses (I vs. Vds) of CNT thin film devices
functionalized with specific antibodies were recorded in order to
determine if different electrical signals were produced by
antibody, blood, or blood with MCF-7 cells. FIG. 11B, for example,
illustrates an SEM image of a CNT thin film functionalized with 10
.mu.L of 5 .mu.g/mL antibodies.
[0065] Upon adding biological components, noticeable changes in
conductance were observed, as shown in a typical electrical
measurement for an anti-HER2 coated device in FIG. 12A. The
reduction in conductance was negligible for a phosphate buffered
saline (PBS) wash. However, a 50% drop in the current of the device
was observed after the adsorption of 5 .mu.L of anti-HER2. After
antibody adsorption, the addition of blood mixed with cancer cells
resulted in an additional .about.30% decrease in device
conductance. This result is further observed from a real time
current measurement (I vs. T), shown in FIG. 12B. Current decreased
.about.10% after adding antibodies, .about.25% for a blood control
sample and .about.60% for blood mixed with cancer cells.
[0066] The electrical behaviors of CNT thin films functionalized
with specific (anti-IGF1R, anti-HER2 and anti-EpCAM) and
non-specific (anti-IgG) or non-cognate (anti-PSMA) antibodies were
measured in order to determine whether specific detection of MCF-7
cells was possible in a sample of unaltered blood. For an initial
study, 5 .mu.L of anti-IGF1R or anti-IgG were immobilized on the
surface of the CNT networks followed by the addition of blood
samples with a ramp of MCF-7 cell concentrations. FIG. 13 shows
that devices printed with IgG experienced less than a .about.10%
change in conductivity while devices printed with IGF1R exhibited a
.about.60% drop in conductivity with increasing number of MCF-7
breast cancer cells in blood.
[0067] A summary of these specificity studies are presented in FIG.
13. Here it can be observed that for specific antibodies such as
anti-IGF1R, anti-HER2 and anti-EpCAM, the electrical signal
(resistance changes between current baseline and after adding MCF-7
cells with blood) increased as a function of increasing number of
MCF-7 cells in blood samples. However, the same is not true for
non-specific IgG and non-cognate PSMA. The electrical signatures
remained the same despite the addition of blood mixed with MCF-7
cells. The specific interaction between antibodies and receptors on
the cell surface may be defined in the form of a ratio called the
overexpression ratio. The overexpression ratio is the ratio of
change in electrical signal for the maximum number of cells spiked
in blood (300 cells) to the change in electrical signal for the
minimum number of cells in blood (10 cells). This relates to the
specificity of the sensor. The ratio is highest for IGF1R
(.about.7.0), EpCAM (.infin.6.0), Her2 (.about.3.6) and IgG
(.about.0.8). The ratios give some degree of specificity based on
the binding of the antibodies to the receptors in cells. The number
of binding sites for EpCAM and Her2 surface markers and their ratio
have been shown in 9 different cancer cell lines using standard
titration methods. For MCF7 cells the EpCAM expression was reported
as 222.1 (713.7).times.103 binding sites while Her2 expression was
25.2 (71.6).times.103 binding sites respectively. Comparing these
measurements to the overexpression ratios, the EpCAM over
expression ratio was higher (.about.6.0) than the ratio for Her2
(.about.3.6) for MCF7 cells in blood, which suggests that the
CNT-antibody array data gives similar results to standard titration
methods for the number of binding events. In other words, it can be
inferred that the change in electrical signal arises from the
number of cooperative binding events happening on the surface of
the device.
[0068] In another experimental implementation, I-V plots were
recorded on 5 .mu.L stabilized blood samples from three patients
with metastatic breast cancer, the results of which are shown in
Table I below:
TABLE-US-00001 TABLE 1 Molecular analysis of metastatic breast
cancer patients Veridex CTC, CNT-mAb CTC, Patient ER PR Her2 in 7.5
mL in 5 .mu.L.sup..dagger. A + + - ND* ~50 B + + - 97 ~50 C - - + 1
~300 *deceased before Veridex available .sup..dagger.based on
overexpression ratio
[0069] In this implementation, the CNT thin film devices were
spotted with antibodies against IGF1R, HER2, EpCAM, EGFR, and
nonspecific IgG, and tested with patient blood samples. The change
in conductivity was almost 10-fold greater for specific antibodies
over non-specific IgG.
[0070] Electrical measurements equivalent to those performed from
controlled blood samples were recorded for the three metastatic
breast cancer patients. When I-V characteristics of patient C blood
interacting with all the different antibodies was compared with
immunohistochemical analysis done on Patient C, it was found that
the cells were Her2-positive in both cases. Additionally, Veridex
CellSearch analysis was performed for the patients. It was observed
from both Patient B and C in Table 1 that when the cells were
Her2-positive (Patient C, 1 cell detected), the Veridex gave low
cell numbers compared to Her2-negative (Patient B, 97 cells
detected). It is believed that the low cell numbers associated with
Her2 status may indicate that EpCAM targeting alone cannot capture
all CTC's in a blood sample. The surfaces of CTC's are heterogenous
and therefore many different types of markers may be necessary for
accurate capture, profiling, and enumeration of CTC's. FIG. 14
illustrates the changes in a normalized electrical signal for
Patient B for devices functionalized with IgG, IGF1R, Her2, EpCAM,
and EGFR antibodies.
[0071] In another experimental implementation, the surface
interactions occurring in devices based on CNT thin film
transistors were studied using liquid gated CNT FET's. The goal of
these experiments was to identify the interactions between the
CNT's and the antibody-receptor binding that lead to the charge
carrier depletion or decrease in current. The transfer
characteristics (I vs. Vg) of liquid gated transistors were
monitored upon the addition of 1 .mu.L of antibodies and 1 .mu.L of
5 MCF-7 cells mixed with blood in order to identify the
electrostatic interactions taking place between CNT's and the
binding of surface receptors with antibodies. For these
experiments, only one single concentration of MCF-7 cells (5 MCF-7
cells/.mu.L) and blood was used. The transfer characteristics were
recorded for devices functionalized with nonspecific IgG and
anti-IGF1R antibodies as shown in FIGS. 15A and 15B (the anti-IgG
and anti-IGF1R plots are above the MCF-7 plots in these figures).
Distinct differences between the electrical characteristics of
non-specific and specific interactions of MCF-7 surface receptors
were observed. For a device printed with IGF1R antibodies, there
was a shift in the threshold voltage (.about.250 mV), whereas
devices printed with anti-IgG showed no distinguishable shift
although conductance was reduced for negative gate voltage. The
current decreased for both non-specific and specific antibodies in
the negative gate voltage region. However, there was a shift in the
gate voltage for the specific antibody in the positive side,
suggesting that geometric deformations occur around the cellular
interactions giving rise to a stress, leading to scattering sites
on a CNT, and thus to reduced conductance. At the same time the
device characteristic is modified only for negative gate voltages,
leaving the transconductance in the positive gate voltage region
unaffected.
[0072] For this latter experiment, an alternate device
implementation may be used, where a localized liquid gate
configuration modulates current in a conducting channel. The CNT
FET's were scaled down to a smaller film area of .about.0.008
mm.sup.2 with only .about.10-100 .mu.m gap between patterned
electrodes in order to observe any charge transfer and minimize the
diffusive behavior of charged particles. Electrical currents were
measured for specific and non-specific antibody-cells surface
marker interactions. In addition, the relationship between electric
current as a function of number of CTC's in a sample was explored.
The change in signal level was related to the overexpression of
targeted cell surface antigens.
[0073] As noted above, it may also be desirable in some embodiments
to alter the surface of a CNT thin film to change the
hydrophobicity or hydrophilicity of the surface. In one exemplary
embodiment, for example, annealing may be performed, e.g., as
discussed above in connection with block 64 of FIG. 4, but at a
higher temperature than described in connection with this figure.
It is believed that annealing at a higher temperature, e.g., about
200 to about 400 degrees Celsius, or about 300 degrees Celsius or
higher, increases the hydrophilicity of a CNT thin film surface by
"burning" the CNT's on the surface of the thin film, causing the
CNT's to curl or curve, and effectively reducing the density of
CNT's at the surface of the film. It is believed that by doing so,
the continuity of the surface is interrupted, exposing holes or
pits in the surface that receive liquid and thus increase the
hydrophilicity of the surface. In one exemplary implementation, for
example, it was found that a CNT thin film transforms from
hydrophobic to hydrophilic at about 300 degrees Celsius.
[0074] It will be appreciated that altering the surface of a CNT
thin film may be performed in a number of manners consistent with
the invention. In addition to annealing a wafer after deposition of
the thin film and electrodes, annealing may be performed at other
points in the fabrication process, e.g. prior to electrode
deposition. In addition, heat may be applied to the thin film in
other manners, e.g., via infrared heating, etc. In addition, other
surface treatments may be performed, including, for example,
chemical treatment, oxygen plasma treatment, etc. In general, any
treatment that lowers the density of CNT's on a thin film surface
and increases hydrophilicity may be used consistent with the
invention.
[0075] It has been found that altering the surface of a CNT thin
film to render the surface hydrophilic causes the conductivity of a
biofunctional pad functionalized with an antibody to increase in
response to the presence of a biological target, which is opposite
to the response of a CNT thin film with a hydrophobic surface. As
such, in implementations where it is desirable to utilize a
positive conductivity relationship with the presence of a
biological target, treating the surface of the CNT thin film may be
desirable.
[0076] In addition, the ability to selectively alter the surface of
only some of the biofunctional pads on a wafer, e.g., as might be
performed using infrared heating, provides the ability to provide
both hydrophilic and hydrophobic biofunctional pads on the same
wafer. In some embodiments, for example, the outputs of hydrophilic
and hydrophobic biofunctional pads functionalized with the same
antibody may be combined to increase the sensitivity and/or signal
to noise ratio of a sensor, or to provide a reconfigurable
functionalized surface for a sensor.
[0077] In still other embodiments, different surfaces may be
combined with different antibodies to provide a more thorough
analysis of the types of biological targets present in a bodily
source. Different types of CTC's, for example, may exhibit
different responses to different antibodies, so that not only the
presence of a CTC in a blood sample, but the type of CTC, may be
detected through the analysis of the conductivity of different
biofunctional pads functionalized with different antibodies.
[0078] In an additional experimental implementation, a device
comprising a CNT thin film with a hydrophobic surface was
fabricated spanning between a pair of gold electrodes deposited on
a glass substrate. 5 .mu.L samples of ruby gold or gold
nanoparticles of 3-5 nm were tested along with 5 .mu.L samples of
isolated white blood cells, isolated MCF7 cancer cells, blood and
blood mixed with MCF7 cells, and the change in an electrical signal
applied across the electrodes was measured for each sample. As
shown in FIG. 16, the gold nanoparticles showed the smallest
percentage change, followed by isolated white blood cells. Isolated
MCF7 cells exhibited a larger percentage change; however, blood and
blood mixed with MCF7 cells exhibited unique conductivity on the
CNT surface, with the blood mixed with MCF7 cells exhibiting the
highest conductivity.
[0079] In addition, as shown in FIG. 17A, a measurement of current
over time for the aforementioned device illustrates the decrease in
conductivity over time seen in a hydrophobic CNT thin film upon
which 5 .mu.L of blood has been adsorbed. FIG. 17B, in contrast,
illustrates a measurement of current over time for a similar device
for which the CNT thin film is treated to render the surface
hydrophilic, and with 5 .mu.L of blood adsorbed thereon. In
contrast with the hydrophobic surface, the plot of current vs. time
for the hydrophilic surface exhibits an increase in conductivity
over time. In each of FIGS. 17A and 17B, the plots of two blood
samples are shown but are not normalized to one another, as the
purpose of these figures is merely to illustrate the relative
changes in conductivity that occur over time with CNT films having
hydrophobic and hydrophilic surfaces.
[0080] Various additional modifications beyond those discussed
herein will be apparent to one of ordinary skill in the art.
Therefore, the invention lies in the claims hereinafter
appended.
* * * * *