U.S. patent application number 16/914682 was filed with the patent office on 2021-05-27 for active-electrode integrated biosensor array and methods for use thereof.
The applicant listed for this patent is Board of Regents, The University of Texas System. Invention is credited to Arjang Hassibi, Arun Manickam, Rituraj Singh.
Application Number | 20210156813 16/914682 |
Document ID | / |
Family ID | 1000005373684 |
Filed Date | 2021-05-27 |
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United States Patent
Application |
20210156813 |
Kind Code |
A1 |
Hassibi; Arjang ; et
al. |
May 27, 2021 |
ACTIVE-ELECTRODE INTEGRATED BIOSENSOR ARRAY AND METHODS FOR USE
THEREOF
Abstract
A method and device for performing DNA sequencing and extracting
structural information from unknown nucleic acid strands. The
device includes a microwell structure, where identical DNA strands
are immobilized within the microwell structure on a surface of a
micro-bead, an active electrode or a porous polymer. The device
further includes a CMOS-integrated semiconductor integrated
circuit, where the CMOS-integrated semiconductor integrated circuit
includes metal layers on a silicon substrate, where the metal
layers form an active electrode biosensor. In addition, a sensing
electrode is formed by creating openings in a passivation layer of
the CMOS-integrated semiconductor integrated circuit to hold a
single bead, on which the DNA strands are immobilized.
Inventors: |
Hassibi; Arjang; (Sunnyvale,
CA) ; Manickam; Arun; (Sunnyvale, CA) ; Singh;
Rituraj; (Saratoga, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Board of Regents, The University of Texas System |
Austin |
TX |
US |
|
|
Family ID: |
1000005373684 |
Appl. No.: |
16/914682 |
Filed: |
June 29, 2020 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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15097037 |
Apr 12, 2016 |
10739293 |
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16914682 |
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13527742 |
Jun 20, 2012 |
9341589 |
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15097037 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
C12Q 1/6874 20130101;
G01N 27/27 20130101; G01N 27/3276 20130101 |
International
Class: |
G01N 27/27 20060101
G01N027/27; C12Q 1/6874 20060101 C12Q001/6874; G01N 27/327 20060101
G01N027/327 |
Claims
1. A system for assaying for a presence of a nucleic acid sequence
in a nucleic acid sample, comprising: a chip comprising (i) a
plurality of wells, wherein during use, an individual well of said
plurality of wells comprises a porous polymer, a template nucleic
acid molecule derived from said nucleic acid sample and reagents
necessary for a polymerization reaction, and (ii) a sensor adjacent
to said individual well, wherein said sensor detects signals
indicative of said presence of said nucleic acid sequence in said
template nucleic acid molecule; and detection circuitry operatively
coupled to said sensor, wherein said detection circuitry (i)
detects said signals indicative of said presence of said nucleic
acid sequence in said template nucleic acid molecule when said
template nucleic acid molecule is subjected to said polymerization
reaction under conditions that are sufficient to yield a nucleic
acid strand that is complementary to said template nucleic acid
molecule, and (ii) provides an output that is indicative of said
presence of said nucleic acid sequence in said nucleic acid
sample.
2. The system of claim 1, wherein said sensor comprises a sensing
electrode.
3. The system of claim 2, wherein said sensor further comprises one
or more metal layers disposed in an insulating layer adjacent to
said sensing electrode.
4. The system of claim 1, wherein said detection circuitry
comprises an operational amplifier and a capacitor that is in a
parallel configuration with respect to said operational
amplifier.
5. The system of claim 1, wherein during use, said template nucleic
acid molecule is immobilized in said porous polymer.
6. The system of claim 1, wherein said signals are indicative of an
impedance at an interface between said sensor and said individual
well.
7. The system of claim 1, wherein said signals are indicative of a
localized ionic current in said individual well during said
polymerization reaction.
8. The system of claim 1, wherein said signals are detected within
a double layer in said individual well adjacent to said sensor.
9. The system of claim 1, wherein said signals are indicative of a
localized ionic current and interface impedance in said individual
well during said polymerization reaction.
10. The system of claim 1, wherein said plurality of wells
comprises an additional well, wherein during use, said additional
well comprises an additional template nucleic acid molecule derived
from said nucleic acid sample.
Description
TECHNICAL FIELD
[0001] The present invention relates generally to biosensors and
bioelectronics, and more particularly to a type of
electro-analytical biosensor, referred to herein as the
"active-electrode" biosensor, that is compatible with Very Large
Scale Integration (VLSI) manufacturing processes and is used in
genomics and proteomics applications.
BACKGROUND
[0002] Biosensors are devices that use biochemical reactions to
identify and detect various molecules and biochemical analytes.
Biosensors are widely used in different life-science applications,
ranging from environmental monitoring and basic life science
research to Point-of-Care (PoC) in-vitro molecular diagnostics.
Biosensors are known to be very sensitive and also extremely
versatile in terms of detection as they can detect a small number
of almost any kind of analyte, once a proper recognition molecule
is identified. Example analytes that have been detected using
biosensors include DNA and RNA strands, proteins, metabolites,
toxins, micro-organisms, and even explosives molecules.
[0003] All biosensors, independent of the analyte they are trying
to detect, include two key building blocks. One is the molecular
recognition layer which is responsible for identifying and/or
interacting with and/or reacting with and/or capturing the specific
target analyte from the sample. The other is the sensor apparatus
that detects and/or quantifies the interactions of the recognition
layer with the analyte and provides a measurable output signal,
generally in the form of an electrical signal. The molecular
recognition layer typically comprises of carefully engineered and
surface-assembled bio-molecules in the form of spotted or
synthesized DNA oligonucleotides, aptamers, and antibodies attached
to solid substrates, such as glass slides, micro-beads, electrodes,
semiconductor materials, or dense polymers while the sensor
includes optical-, MEMS- and/or electronics-based transducers
connected to a low-noise detection circuit.
[0004] So far, there have been many detection methods that have
been adopted in biosensor systems. A detection method is defined as
the specific type of physiochemical mechanism designed into the
molecular recognition layer, analytes, and the interaction
environments that make the identification of the specific target
analytes possible by the sensor. The most widely used detection
methods are different classes of optical (e.g., fluorescence,
bioluminescence) and electro-analytical (e.g., potentiometric,
amperometric, impedimetric). It is also common to classify
biosensors based on their detection method. For example, in
bioluminescence-based biosensors, the interaction of the analyte
and probes results in a bioluminescence phenomenon which is
detected by a specific sensor with a transducer sensitive to
bioluminescence signals.
[0005] Electro-analytical biosensors detect analytes by monitoring
different electronic changes in electrode-electrolyte transducers
that are specifically interfaced with a recognition layer. For
instance, in amperometric biosensors, low-frequency Faradaic
reduction-oxidation (redox) currents are used as an indicator for
analyte interactions with the recognition layer, whereas in
impedimetric biosensors, the changes in the electrode-electrolyte
impedance induced by the captured analyte are used as an indicator
of analyte interactions with the recognition layer.
[0006] Unfortunately, the existing state-of-the-art
electro-analytical biosensors are not compatible with semiconductor
Very Large Scale Integration (VLSI) manufacturing processes thereby
not being able to take advantage of the VLSI processes (e.g.,
highest level of integration, miniaturization, cost-efficiency, and
robustness).
SUMMARY
[0007] As discussed above, the existing state-of-the-art
electro-analytical biosensors are not compatible with semiconductor
Very Large Scale Integration (VLSI) manufacturing processes thereby
not being able to take advantage of the VLSI processes (e.g.,
highest level of integration, miniaturization, cost-efficiency, and
robustness). The principles of the present invention address this
impediment.
[0008] In view of the limitations of biosensors currently
available, there is a need for improved biosensors and methods for
use thereof in fields, such as nucleic acid detection, nucleic acid
sequencing, proteomics, forensics, in-vitro diagnostics, medicine,
and the like. Needed improvements include reducing the cost,
increasing the throughput, and/or decreasing the size of biosensors
instrument; in order to advance the field of personalized medicine
for example.
[0009] Provided herein are Complementary Metal Oxide Semiconductor
(CMOS) biological sensors ("biosensors") fabricated using Very
Large Scale Integration (VLSI) manufacturing processes in various
applications, such as, for example, nucleic acid sequencing,
proteomics, and forensics. CMOS biosensors described in various
embodiments of the present invention can be used in a variety of
genomics and proteomics applications. In particular, they can be
used in nucleic acid sequencing, such as deoxyribonucleic acid
(DNA) sequencing, ribonucleic acid (RNA) sequencing, and forensics
analysis, such as short tandem repeat (STR) analysis.
[0010] In one embodiment of the present invention, a method for
assaying for a presence of a nucleic acid sequence in a nucleic
acid sample comprises providing a chip comprising (i) a plurality
of wells, where an individual well of the plurality of wells
comprises a porous polymer, a template nucleic acid molecule
derived from the nucleic acid sample and reagents necessary for a
polymerization reaction, and (ii) a sensor adjacent to the
individual well, where the sensor detects signals indicative of the
presence of the nucleic acid sequence in the template nucleic acid
molecule. The method further comprises subjecting the template
nucleic acid molecule to the polymerization reaction under
conditions that are sufficient to yield a nucleic acid strand that
is complementary to the template nucleic acid molecule. The method
additionally comprises using the sensor to detect the signals
indicative of the presence of the nucleic acid sequence in the
template nucleic acid molecule. Furthermore, the method comprises
based on the signals detected in (c), providing an output that is
indicative of the presence of the nucleic acid sequence in the
nucleic acid sample.
[0011] In another embodiment of the present invention, a system for
assaying for a presence of a nucleic acid sequence in a nucleic
acid sample comprises a chip comprising (i) a plurality of wells,
where during use, an individual well of the plurality of wells
comprises a porous polymer, a template nucleic acid molecule
derived from the nucleic acid sample and reagents necessary for a
polymerization reaction, and (ii) a sensor adjacent to the
individual well, where the sensor detects signals indicative of the
presence of the nucleic acid sequence in the template nucleic acid
molecule. The system further comprises detection circuitry
operatively coupled to the sensor, where the detection circuitry
(i) detects the signals indicative of the presence of the nucleic
acid sequence in the template nucleic acid molecule when the
template nucleic acid molecule is subjected to the polymerization
reaction under conditions that are sufficient to yield a nucleic
acid strand that is complementary to the template nucleic acid
molecule, and (ii) provides an output that is indicative of the
presence of the nucleic acid sequence in the nucleic acid
sample.
[0012] Additional aspects and advantages of the present disclosure
will become readily apparent to those skilled in this art from the
following detailed description, wherein only illustrative
embodiments of the present disclosure are shown and described. As
will be realized, the present disclosure is capable of other and
different embodiments, and its several details are capable of
modifications in various obvious respects, all without departing
from the disclosure. Accordingly, the drawings and description are
to be regarded as illustrative in nature, and not as
restrictive.
BRIEF DESCRIPTION OF THE DRAWINGS
[0013] The novel features of the present invention are set forth
with particularity in the appended claims. A better understanding
of the features and advantages of the present invention will be
obtained by reference to the following detailed description that
sets forth illustrative embodiments, in which the principles of the
invention are utilized, and the accompanying drawings of which:
[0014] FIG. 1 illustrates an active electrode circuit model in
accordance with an embodiment of the present invention;
[0015] FIG. 2 illustrates an active electrode system using a switch
capacitor amplifier sensor in accordance with an embodiment of the
present invention;
[0016] FIG. 3A illustrates the correlated double sampling method to
suppress the low frequency noise and offset of the amplifier within
the active-electrode sensor in accordance with an embodiment of the
present invention;
[0017] FIG. 3B is a timing diagram of the circuit of FIG. 3A in
accordance with an embodiment of the present invention;
[0018] FIG. 4 illustrates using active electrode sensors to measure
C.sub.D in accordance with an embodiment of the present
invention;
[0019] FIG. 5 illustrates using active electrode sensors to measure
the ionic currents within the electrolyte in accordance with an
embodiment of the present invention;
[0020] FIG. 6 illustrates hybrid electro-analysis using an active
electrode sensor in accordance with an embodiment of the present
invention;
[0021] FIG. 7 is a table illustrating the absorbed ions and net
free charge generated during a single nucleotide incorporation at
different pH levels in accordance with an embodiment of the present
invention;
[0022] FIGS. 8A-8C illustrate different embodiments for
immobilizing identical DNA in proximity of an active electrode in
accordance with an embodiment of the present invention;
[0023] FIGS. 9A-9B illustrate examples of CMOS-integrated sensing
electrodes in accordance with an embodiment of the present
invention;
[0024] FIGS. 10A-10C illustrate an integrated active electrode DNA
sequencing biochip and the basic layer structure of its pixels in
accordance with an embodiment of the present invention;
[0025] FIG. 11A illustrates the general architecture of the
CMOS-integrated active-electrode biosensor pixel for DNA sequencing
in accordance with an embodiment of the present invention;
[0026] FIG. 11B is a timing diagram of in-pixel and out-of-pixel in
accordance with an embodiment of the present invention;
[0027] FIG. 12 is a transistor-level schematic of a signal chain in
accordance with an embodiment of the present invention;
[0028] FIGS. 13A-13C are graphs illustrating electrical detection
performances in accordance with an embodiment of the present
invention;
[0029] FIGS. 14A-14C illustrate interface capacitance measurements
versus solution pH in accordance with an embodiment of the present
invention;
[0030] FIGS. 15A-15C illustrate DNA polymerization detection in
accordance with an embodiment of the present invention; and
[0031] FIG. 16 is a micrograph of the active-electrode CMOS biochip
for DNA sequencing in accordance with an embodiment of the present
invention.
DETAILED DESCRIPTION
Incorporation by Reference
[0032] All publications, patents, and patent applications mentioned
in this specification are herein incorporated by reference to the
same extent as if each individual publication, patent, or patent
application was specifically and individually indicated to be
incorporated by reference.
Principles of the Present Invention
[0033] While various embodiments of the present invention have been
shown and described herein, it will be obvious to those skilled in
the art that such embodiments are provided by way of example only.
Numerous variations, changes, and substitutions may occur to those
skilled in the art without departing from the present invention. It
should be understood that various alternatives to the embodiments
of the present invention described herein may be employed in
practicing the invention.
[0034] The principles of the preset invention relate to
electro-analytical biosensors. Generally speaking,
electro-analytical biosensors that detect analytes by monitoring
different electronic changes in electrode-electrolyte transducers
that are specifically interfaced with the recognition layers. For
instance, in amperometric biosensors, low-frequency Faradaic
reduction-oxidation currents are used as an indicator for analyte
interactions with the recognition layer, whereas in impedimetric
biosensors, the changes in the electrode-electrolyte impedance
induced by the captured analyte are used as an indicator of analyte
interactions with the recognition layer.
[0035] It is noted that depending on the exact electrical
characteristic that is being probed (or monitored) in
electro-analytical biosensors, different instrumentation techniques
and electrode configurations are required. The principles of the
present invention described herein describe a specific type of
electro-analytical biosensor, called an active-electrode biosensor,
and the methods by which one can create high-performance and
highly-parallel DNA sequencing platforms employing this
biosensor.
[0036] One of the key advantages of active-electrode biosensors is
that they are compatible with semiconductor Very Large Scale
Integration (VLSI) manufacturing processes in general, and
Complementary Metal-Oxide-Semiconductor (CMOS) Integrated Circuits
(ICs) in particular. This means that all of the advantages of VLSI
processes (e.g., highest level of integration, miniaturization,
cost-efficiency, and robustness) can be applied to active-electrode
biosensors. Furthermore, the principles of the present invention
provide methods by which one can design arrays of active-electrode
biosensors using CMOS processes. While it is self-evident that the
present invention can be used for a variety of biosensing
applications, embodiments of the present invention described herein
are related to DNA sequencing applications.
[0037] There are many different techniques to perform DNA
sequencing and extract structural information from unknown nucleic
acid strands. Certain embodiments of the present invention may rely
on the Sequence-by-Synthesis (SBS) procedure. In SBS, individual
nucleotides (dATP, dCTP, dGTP and dTTP) are iteratively introduced
to a primed DNA complex in the presence of the DNA polymerase
enzyme while the occurrence of polymerization events is monitored.
Successful polymerization events at the 3'-terminus of the primer
suggest the presence of the complementary base on the template DNA
while the amplitude of the polymerization indicates the number of
consecutive identical bases. Different techniques have been
discussed in the art to detect the polymerization events to perform
the SBS procedure. Examples are bioluminescence-based enzymatic
cascade, fluorescent-label nucleotides, and pH-based. Embodiments
of the present invention described herein provide an alternative
electro-analytical technique which is based on using
active-electrode biosensors. The key advantages over the previous
methods are amenability to VLSI integration, miniaturization
capabilities, lower noise performance, and increased detection
dynamic range.
Definitions
[0038] An active electrode is defined as a highly conductive
material (i.e., an electrode) with its electrical potential,
.PHI..sub.0(t), set to V.sub.0(t), a defined voltage set by an
independent time-varying source, such that
.PHI..sub.0(t)=V.sub.0(t) at all time. In the context of
electro-analysis and the present invention, an active-electrode
system is defined as a conductive material (i.e., the electrode)
that is capacitively-coupled to an aqueous and
electrically-conductive solution (i.e., the electrolyte) such that
the capacitively-coupled electrode-to-electrolyte potential
difference, .PHI..sub.E(t), follows V.sub.0(t), such that
.PHI..sub.E(t)=V.sub.0(t) at all time.
[0039] An active electrode biosensor is defined here as an active
electrode system with an electrolyte containing the target analyte
in which by applying time-varying V.sub.0(t) and concurrently
monitoring the corresponding coupled charge (screening charge) on
the highly-conductive electrode, denoted by Q.sub.E(t), one may
infer information regarding the target analyte presence, and/or
abundance and/or molecular structure.
[0040] There are unique characteristics for the active electrodes
biosensor described herein. The first is that
"capacitively-coupled" means that the active-electrode is not
directly in contact with the electrolyte and that an electrically
insulating layer with a thickness between 5 nm to 100 nm is placed
between the electrode and the electrolyte. The second is that
Q.sub.E(t) is located in a very thin layer near the
electrode-insulator interface; however, -Q.sub.E(t), the opposite
screening charge on the electrolyte-insulator interface side, is
distributed non-uniformly within the electrolyte in a form
generally referred to in the art as the double layer (i.e., Helmotz
and diffusion layers). The third and final characteristic is that
V.sub.0(t) can change Q.sub.E(t) as well as the profile of the
charge within the double layer. It is known in the art that such
changes depend on the exact electrochemical characteristics of the
electrolyte (e.g., ionic species charges and their diffusion
coefficient), the electrolyte-insulator interface (e.g., surface
pKa and the concentration the surface traps for the ions in the
electrolyte), and the thickness as well as the material composition
of the insulator. In view of the foregoing, no net charge transfer
from the electrode to the electrolyte can occur (i.e., no DC
current can pass the interface); however, the ionic charges can
still electrostatically interact with the charge carriers within
the electrode.
[0041] Active-Electrode Sensor Circuit Architecture
[0042] A simple and widely accepted circuit model 100 for
active-electrode sensors is shown in FIG. 1 in accordance with an
embodiment of the present invention. Referring to FIG. 1, C.sub.1
and C.sub.D represent the insulator layer 101 and double layer 102,
respectively, while R.sub.EL represents the ohmic resistance of the
electrolyte from the electrode-insulator interface 103 to a counter
electrode 104 (e.g., Ag/AgCl electrode) in the electrolyte. For
typical insulating materials, such as SiO.sub.2, Si.sub.3N.sub.4,
TiO.sub.2, Al.sub.2O.sub.3O.sub.3 HfO.sub.2, one can safely assume
that C.sub.1 is a linear capacitor; however, C.sub.D is widely
known in the art to be inherently non-linear and function of the
voltage placed across it.
[0043] The fundamental sensor circuitry 200 in the present
invention for measuring Q.sub.E(t) is shown in FIG. 2 in accordance
with an embodiment of the present invention which is essentially a
Switched-Capacitor Charge Amplifier (SCCA). In this circuit, to
ensure .PHI..sub.E(t)=V.sub.0(t) at all times, one can take
advantage of an operational amplifier 201 with a capacitive
negative feedback that connects the (-) input of amplifier 201 to
its output, V.sub.OUT(t), while V.sub.0(t) is applied to its (+)
input. To detect Q.sub.E(t), the following steps occur:
[0044] (a) Reset step: First the feedback reset switch is activated
(connected) to discharge any accumulated charge on C.sub.f at t=0
such that V.sub.OUT(t)=V.sub.0(t).
[0045] (b) Read step: Subsequently, at t=.DELTA., the switch is
deactivated (disconnected) and V.sub.OUT(t) is read in real time
for t>.DELTA.. It can be shown that during this phase
V OUT ( t ) = 1 C f [ Q E ( t ) - Q E ( .DELTA. ) ] ( EQ 1 )
##EQU00001##
[0046] which means that V.sub.OUT(t) is effectively the amplified
version of the difference between the charge at time t compared
with the charge at t=.DELTA.. In most cases, the reset step
duration (.DELTA.) can be kept relatively small such that
Q.sub.E(.DELTA.).apprxeq.Q.sub.E(0), thereby resulting in
V OUT ( t ) .apprxeq. 1 C f [ Q E ( t ) - Q E ( 0 ) ] ( EQ 2 )
##EQU00002##
[0047] which indicate that V.sub.OUT(t) is the amplified version
difference between the charge at time t compared with the charge at
t=0.
[0048] (c) Iteration steps: Repeat steps (a) and (b) periodically
with interval T, i.e., activating the reset switch at t=T, t=2T,
t=3T, . . . and deactivating it at t=T+.DELTA., t=2T+.DELTA.,
t=3T+.DELTA., . . . .
[0049] (d) Constructing Q.sub.E(.DELTA.): Use the individual
measured values of each interval of step (c) to create the
Q.sub.E(t)-Q.sub.E(0) waveform sampled at the frequency 1/T.
[0050] The minimum detection level of the active-electrode sensor
system is limited by the inherent noise sources within SCCA,
particularly the noise contributed by the amplifier. Generally
speaking, high-gain amplifiers introduce a high level of
low-frequency noise (i.e., 1/f noise) and DC offset in the system.
To suppress both the low-frequency noise and offset of the
active-electrode sensor, one can use different Correlated Double
Sampling (CDS) techniques which are widely used in the art. In FIG.
3A, an exemplary embodiment of circuitry 300 implementing a CDS
technique which requires three additional switches activated by
signals .PHI..sub.1, .PHI..sub.2, and .PHI..sub.3, and an offset
storage capacitor, C.sub.S, is shown in accordance with an
embodiment of the present invention. As shown in the timing diagram
301 of FIG. 3B, initially the charge across the C.sub.f is reset
and subsequently during the offset calibration phase (high
.PHI..sub.1 and .PHI..sub.3) the input referred offset (and noise)
of operational amplifier 201 is stored on C.sub.S. Finally, in the
readout phase, this stored voltage is subtracted from the input of
operational amplifier 201 (i.e., the offset and noise are
cancelled) and the output becomes independent of this stored value.
In this case, .DELTA. is defined as the duration of time between
the rising edge of the reset and the falling edge of
.PHI..sub.3.
[0051] Active-Electrode Biosensor System
[0052] In order to create a biosensor using an active electrode
sensor, one should devise methods to couple the measurable
Q.sub.E(t) to the biosensing interactions that occur between the
target analyte and the recognition layer. There are two general
approaches to carry this out this:
[0053] Method 1:
[0054] The first method is to incorporate the recognition layer
within the double layer of the active-electrode system, such that
analyte-recognition layer interactions directly affect the
distribution of -Q.sub.E(t) and therefore the value of C.sub.D. A
typical example application for this approach is label-free DNA
hybridization detection in which the capturing DNA strands (in the
recognition layer) are attached to the solid surface and are
physically immobilized within the interface double layer of the
active-layer. In this example, successful hybridization of the
target charged DNA molecule modifies the interface charge and
subsequently the C.sub.D. In FIG. 4, it is illustrated how this
type of biosensor 400 can be accommodated by the SCCA-based
active-electrode sensor in accordance with an embodiment of the
present invention. To measure C.sub.D, which is an indicator of
analyte-recognition layer interactions, one may apply a sinusoidal
signal 401 at frequency .omega. across the interface by applying
V.sub.0(t)=V.sub.0+V.sub.1 cos(.omega.t) (both V.sub.0 and V.sub.1
are constant values) and examine the content of V.sub.OUT(t) at
frequency .omega.. In this case, the phase vector of the output,
denoted by V.sub.OUT(.omega.), can be described by the following
formula:
V OUT ( .omega. ) = V 1 [ 1 + C D C I C f .times. 1 1 + j .omega. R
E L C D C I ] ( EQ 3 ) ##EQU00003##
[0055] Since C.sub.f, C.sub.1, and V.sub.1 are known values, one
can use (EQ 3) and the measurement at a plurality of frequencies to
estimate both C.sub.D and R.sub.EL. By measuring V.sub.OUT(.omega.)
at more than two frequencies, redundant information is created
which can be used to further improve the estimated C.sub.D and
R.sub.EL.
[0056] Method 2:
[0057] The second method is to create ionic currents within the
biosensing reaction volume (i.e., coordinated where R.sub.EL is the
dominant electrical element) that are indicative of
analyte-recognition layer interactions. An example of such a
configuration is common in electrochemical enzyme biosensors which
take advantage of electro-active enzymes (e.g., horseradish
peroxidase or glucose oxidase) attached to their detection
antibody. In FIG. 5, it is illustrated how the active electrode
sensor 500 can detect such currents in accordance with an
embodiment of the present invention. As illustrated in FIG. 5,
i(t), represents the ionic current within the solution which is
triggered at t=0. It is straightforward to show that the output of
the SCCA in this case, the sensor output for t>0 can be
formulated by:
V OUT ( t ) = e t R EL C D C I C f .intg. 0 t e - .alpha. R EL C D
C I i ( .alpha. ) d .alpha. ( EQ 4 ) ##EQU00004##
[0058] For systems in which i(t) changes occur at a much slower
rate compared to the sensor relaxation time, defined by
.tau.=R.sub.ELC.sub.D.parallel.C.sub.1, one can simplify and
rewrite (EQ 4) as
V OUT ( t ) = R EL C D C I C f i ( t ) ( EQ 5 ) ##EQU00005##
[0059] One critical issue here is that while (EQ 4) or (EQ 5) offer
a means to evaluate i(t) using the SCCA output; however, the exact
relationship between these two parameters relies on the values of
both R.sub.EL and C.sub.D which are known to be susceptible to
unwanted drifts during electro-analysis. This is a known and
widely-recognized problem in this field. R.sub.EL drifts generally
happen when the ionic content of the electrolyte is changed (e.g.,
by injecting in or washing away different reagents), or when the
reference electrode remains for a long time in the electrolyte and
ages. C.sub.D drifts are often a result of unwanted interaction of
the insulator surface with reactants and ions in the electrolyte
which slowly alter the charge distribution at the
electrolyte-insulator surface. If such drifts are not continually
monitored and effectively calibrated out the quality of the
measurements will be significantly degraded.
[0060] Hybrid Method:
[0061] In one embodiment, a method for implementing an active
electrode biosensor of the present invention is to concurrently use
method 1 and method 2 during electro-analysis. The basic idea is to
take advantage of method 1 for continual calibration and monitoring
of the surface and electrolyte (C.sub.D and/or R.sub.EL) and use
method 2 to monitor any ionic currents. To enable simultaneous
operation of both these methods, one should operate method 1 in
frequencies above the Nyquist bandwidth of i(t). For example, if
the informative frequency content of i(t) is within DC to 1 kHz,
method 1 is operated in frequencies higher than 1 kHz. This
approach essentially de-couples the operation of method 1 and
method 2 by separating their frequency operation.
[0062] In FIG. 6, we illustrate an exemplary embodiment of
circuitry 600 as to how the hybrid method can be implemented using
the SCCA in accordance with an embodiment of the present invention.
The general idea is to use a Low-Pass filter (LPF) and High-Pass
Filter (HPF) to isolate the output of each method from V.sub.OUT(t)
and analyze them independently.
[0063] Arrays of Active-Electrode Sensors
[0064] In some embodiments, an array of active-electrode sensors
are built on a common substrate, such as a semiconductor substrate
(e.g., CMOS) and by using VLSI fabrication processes. In some
cases, the number of the pixels within this array is greater than
10 and can be as large as 10.sup.8 per single substrate. Techniques
for selecting the row and column of the pixel to be interrogated
are widely known to those skilled in the art of design of sensor
array and image sensor arrays.
[0065] In some situations, a biosensor array includes at least 1,
2, 3, 4, 5, 6, 7, 8, 9, 10, 10.sup.2, 10.sup.3, 10.sup.4, 10.sup.5,
10.sup.6, 10.sup.7, 10.sup.8, 10.sup.9, or 10.sup.10 within a
cross-sectional area of at most about 1000 cm.sup.2, 100 cm.sup.2,
10 cm.sup.2, 1 cm.sup.2, 0.5 cm.sup.2, or 0.1 cm.sup.2.
[0066] In some embodiments in which an array of active-electrode
sensors is built in a semiconductor substrate, each pixel may have
part of the required circuitry to enable a SSCA-based active
electrode sensor, and, for example, the operational amplifier may
be shared by a plurality of pixels in, for example, all the pixels
within the column. This method of sharing the circuitry in the
signal path is a widely used method in CMOS sensor array and image
sensor arrays. In some embodiments, the shared circuits are placed
in the periphery of the array to minimize the size of the
individual pixels.
[0067] Detecting DNA Polymerization Using Active-Electrode
Biosensors
[0068] One of the primary applications that is targeted using the
principles of the present invention is DNA sequencing. There are
many different techniques to perform DNA sequencing and extract
structural information from unknown nucleic acid strands. In some
embodiments, DNA sequencing is accomplished via
Sequence-by-Synthesis (SBS), in which individual nucleotides
(adenine, cytosine, guanine or thymine) are iteratively introduced
to a primed DNA complex in the presence of the DNA polymerase
enzyme. The occurrence of one or more polymerization events is
monitored, such as with the aid of a biochip described herein.
Successful polymerization at the 3'-terminus of the primer is
indicative of the presence of the complementary base on the
template DNA while the amplitude of the polymerization is
indicative of the number of consecutive identical bases. Different
approaches have been provided for detecting the polymerization
events and perform SBS. Examples are bioluminescence-based
enzymatic cascade (e.g., J. M. Rothberg and J. H. Leamon, "The
development and impact of 454 sequencing", Nature Biotech., vol.
23, no. 10, pp. 1117-1125, 2008), fluorescent-label nucleotides
(e.g., U.S. Pat. No. 7,835,871), and pH-based (e.g., U.S. Pat. No.
7,948,015). In the present invention, an alternative
electrochemical technique is provided which is not only amendable
to integration and miniaturization, but also offers a significantly
better noise performance and measurement robustness.
[0069] In the following, the details of how SBS can be enabled by
the active electrode biosensors are described. Initially, the
electronic characteristics of DNA polymerization, where its
detection is fundamental in SBS, is discussed followed by
discussing the multiple embodiments in which DNA polymerization can
be detected using the active-electrode biosensors.
[0070] In DNA polymerization, the 3'-terminus of the primer is
extended by the DNA polymerase enzyme which facilitates the
incorporation of individual nucleotides (deoxyribonucleotide
triphosphates, dNTPs) that are complementary to the template DNA
strand. A single nucleotide incorporation event that extends the
primer from length n to n+1 is best described by
##STR00001##
[0071] which states that the DNA-enzyme complex absorbs a single
dNTP molecule and releases a pyrophosphate (PPi) molecule. Since
all of the participating molecules in the catalytic reaction of (EQ
6) (including both the substrates and products) are essentially
charged species, it is feasible to setup certain conditions in
which DNA polymerization can result in a measurable electronic
parameter in this system. If this is done, then monitoring this
particular parameter can be used to detect (and quantify) DNA
polymerization and therefore can be used to perform SBS. All of the
DNA sequencing embodiments described herein operate according to
this particular principle.
[0072] Previously, Sakurai et al. ("Real-Time Monitoring of DNA
Polymerase Reactions by a Micro ISFET pH Sensor", Anal. Chem, vol.
64, pp. 1996-1997, 1992) demonstrated that a small pH change
induced by dNTP incorporation can be detected by a micro ISFET pH
sensor and used, in real-time, to detect DNA polymerization events.
Later, Pourmand et al., ("Direct electrical detection of DNA
synthesis", PNAS, vol. 102, no. 17, pp. 6866-6870, 2006) suggested
that the transient electrical signal generated by DNA immobilized
on a polarized gold electrode during polymerization can be sensed
in real-time using differential voltage and current amplifiers and
be used to sense dNTP incorporations. Recently, Rothberg et al.,
("An integrated semiconductor device enabling non-optical genome
sequencing", Nature, vol. 475, pp. 348-352, 2011) demonstrated how
an array of ISFET in conjunction with random bead arrays can be
used to create high-throughput parallel pH-based SBS arrays.
[0073] In the present invention, an alternative approach is
introduced which uses embodiments of an active-electrode sensor to
detect DNA polymerization to enable DNA SBS. The premise of this
system is the fact that nucleotide incorporation when the primed
DNA is immobilized (i.e., is attached to a solid surface), can
result in an imbalance between the absorbed and released free ionic
charges in the electrolyte near the DNA, which in turn results in
localized ionic currents that can be sensed by the active-electrode
biosensor. Unlike the previously reported system, active-electrode
arrays, rely neither on the concentration of protons and ISFET
structures, nor polarized gold electrodes to detect polymerization.
In addition, active-electrode sensors can also measure the surface
capacitance concurrently with polymerization detection (Hybrid
method), a unique feature that none of the aforementioned
references have.
[0074] Referring now to FIG. 7, FIG. 7 is a table 700 illustrating
the ionic species and the generated net free charge of a single
nucleotide incorporation event when DNA is immobilized in
accordance with an embodiment of the present invention. As
illustrated in FIG. 7, depending on the pH of the electrolyte and
isoelectric characteristics of the involved molecules, the
generated average net free negative charge (from PPi) and net free
positive charge (from protons, H.sup.+) are different. Yet, the net
generated charge (i.e., the sum of the positive and negative
charges) is always positive and equal to +1. This indicates that
DNA polymerization always creates a positive diffusion potential in
the electrolyte at the DNA polymerization coordinate which in turn
can create an outward going ionic current through free net charge
diffusive spreading.
[0075] Based on this observation, one can state that if an
active-electrode biosensor is placed in the electrolyte such that
it measures the ionic current that is generated by DNA
polymerization, one can detect DNA polymerization events. It is
important to realize that the DNA itself does not have to be in
intimate proximity of the electrode and as long as the current
reaches the electrode, one can detect polymerization. In practice
this means that the electrolyte-insulator surface should be within
a few diffusion lengths for the PPi and proton ions, which is
defined by the average distance that these ions can move the DNA
polymerization until they recombine with other ions in the
solution. Depending on the characteristics of the solution (e.g.,
salt concentration, buffer capacity and temperature), this distance
is typically between 1 nm and 10 nm.
[0076] It is noted herein that the embodiments of the present
invention are categorically different from all ISFET-based sensors
described in the art. ISFET devices and sensors implemented by
them, by their definition, include electrically-floating electrodes
where the electrode potential is undefined (unlike active-electrode
sensors which are always set to V.sub.0(t)) and can (and should)
change during measurements. In addition, unlike Pourmand et. al and
its derivatives, active-electrode sensors described herein do not
use clamp or voltage amplifiers and furthermore do not require the
DNA to be attached to the polarized metal electrode surface to
ensure that the DNA backbone protonization is directly shielded by
the electrode.
[0077] In some embodiments, the components and procedures to detect
DNA polymerization using active electrode biosensors are:
[0078] (a) Immobilize a plurality of identical primed DNA molecules
in a reaction chamber and interface an active-electrode biosensor
system to it;
[0079] (b) Introduce dNTPs into the reaction chamber in presence of
the DNA polymerase enzyme;
[0080] (c) Measure in real-time the ionic current, R.sub.EL, and
C.sub.D using the aforementioned hybrid biosensing method;
[0081] (d) Estimate i(t) by using the measured V.sub.OUT(t),
R.sub.EL, and C.sub.D; and
[0082] (e) Use the estimated i(t) to identify the occurrence and
amount of DNA polymerization events for the primed DNA molecules
interfaced to the active-electrode biosensor.
[0083] In some embodiments, to carryout a DNA SBS procedure and
identify the sequence of primed DNA molecules (e.g., a clonal
population thereof), the following steps are used:
[0084] (a) Immobilize a plurality of identical primed DNA molecules
in a reaction chamber and interface an active-electrode biosensor
system to it;
[0085] (b) Introduce a single type of nucleotide (dATP, dTTP, dCTP,
or dGTP) in the presence of DNA polymerase and monitor DNA
polymerization;
[0086] (c) Remove the remaining unreacted nucleotides from the
reaction chamber;
[0087] (d) Repeat step (b) through (c) for a different nucleotide;
and
[0088] (e) Use the occurrence and the quantity of DNA
polymerization events to find the sequence.
[0089] In the embodiments of the present invention where the
integrated active-electrode biochip array can carries parallel
(multiplexed) DNA sequencing, one has:
[0090] (a) An immobilized primed DNA array created on the surface
of a CMOS-integrated semiconductor chip where identical DNA strands
are located in distinct coordinates within the array, referred to
herein as the pixels;
[0091] (b) An array of active-electrode biosensors built in a CMOS
chip in the form of an integrated circuit in which individual
pixels of the DNA array have one active-electrode biosensor;
[0092] (c) A reaction chamber which can contain aqueous solutions
on top of the CMOS chip where the DNA array is located;
[0093] (d) A fluidic single-directional flow-through system which
enables controlled injection and removal of different aqueous from
the reaction chamber including, but not limited to, dNTPs, DNA
polymerase enzyme, wash buffer, dNTPase, and pH standard
buffers;
[0094] (e) A system to monitor and control the temperature of the
reaction chamber; and
[0095] (f) A data acquisition and processing device which can
extract, read, and process the output of each integrated active
electrode biosensor.
[0096] Preparing DNA for Sequencing
[0097] The nucleic acid being sequenced is referred to herein as
the target nucleic acid. Target nucleic acids include, but are not
limited to, DNA, such as but not limited to, genomic DNA,
mitochondrial DNA, cDNA and the like, and RNA, such as but not
limited to, mRNA, miRNA, and the like. The nucleic acid may be from
any source including naturally occurring sources or synthetic
sources. The nucleic acids may be Polymerase Chain Reaction (PCR)
products, cosmids, plasmids, naturally occurring or synthetic
libraries, and the like. The present invention is not to be limited
in this regard. The methods provided herein can be used to sequence
nucleic acids of any length. The following is a brief description
of examples of these methods.
[0098] In some embodiments, target nucleic acids are prepared using
any manner known in the art. As an example, genomic DNA may be
harvested from a sample according to techniques known in the art
(see for example Sambrook et al. "Maniatis"). Following harvest,
the DNA may be fragmented to yield nucleic acids of smaller length.
The resulting fragments may be on the order of hundreds, thousands,
or tens of thousands nucleotides in length. In some embodiments,
the fragments are 200-1000 base pairs (bp) in size, or 300-800 bp
in size, although they are not so limited. Nucleic acids may be
fragmented by any means including but not limited to mechanical,
enzymatic or chemical means. Examples include shearing, sonication,
nebulization and endonuclease (e.g., Dnase I) digestion, or any
other technique known in the art to produce nucleic acid fragments,
optionally of a desired length. Fragmentation can be followed by
size selection techniques which can be used to enrich or isolate
fragments of a particular length or size. Such techniques are also
known in the art and include, but are not limited to, gel
electrophoresis or Solid-Phase Reversible Immobilization
(SPRI).
[0099] In some embodiments, the size selected target nucleic acids
are ligated to adaptor sequences on both the 5' and 3' ends. These
adaptor sequences comprise amplification primer sequences to be
used in amplifying the target nucleic acids. One adaptor sequence
may also comprise a sequence complementary to the sequencing
primer. The opposite adaptor sequence may comprise a moiety that
facilitates binding of the nucleic acid to a solid support, such as
but not limited to, a bead. An example of such a moiety is a biotin
molecule (or a double biotin moiety, as described by Diehl et al.
Nature Methods, 2006, 3(7):551-559) and such a labeled nucleic acid
can therefore be bound to a solid support having avidin or
streptavidin groups. The resulting nucleic acid is referred to
herein as a template nucleic acid. The template nucleic acid
comprises at least the target nucleic acid and usually comprises
nucleotide sequences in addition to the target.
[0100] Immobilization of DNA
[0101] There are different known methods in the art that one can
use to immobilize primed DNA strands near an on the
active-electrode biosensor. FIGS. 8A-8C illustrates different
embodiments for immobilizing clonal DNA in proximity of an active
electrode in accordance with an embodiment of the present
invention. Referring to FIGS. 8A-8C, FIGS. 8A-8C illustrate how
identical DNA strands 801 are physically placed near the integrated
active biosensors that are embedded in a CMOS chip 802 (includes a
silicon substrate 809 and metal layers 810). In one embodiment, DNA
strands 801 are immobilized within microwell structures 803
covalently using linkers or through base pairing (hybridization) on
the surface of functionalized micro-beads 804, active electrodes
805, or porous polymers 806, referred to as solid support herein.
In alternative embodiments, microwell structure 803 may not be
present and DNA strands 801 are immobilized covalently using
linkers, or through base pairing (hybridization) on the surface of
the electrolyte-insulator interface 807 (insulator identified as
element 808), or porous polymers 806. The size of the pixels
(parameter X) is preferably between 0.1 .mu.m to 50 .mu.m, while
the aspect ratio of individual microwells 803 (i.e., X/Y) varies
between 0.6 to 3.
[0102] In some embodiments, a linker (or spacer) is specifically
used to distance the template nucleic acid (and in particular the
target nucleic acid sequence comprised therein) from the solid
support. This can facilitate sequencing of the end of the target
closest to the surface. Examples of suitable linkers are known in
the art (see Diehl et al. Nature Methods, 2006, 3(7):551-559) and
include, but are not limited to, carbon-carbon linkers, such as,
but not limited to, iSp18.
[0103] The beads used in the present invention can be made of any
material including, but not limited to, cellulose, cellulose
derivatives, gelatin, acrylic resins, glass, silica gels, PolyVinyl
Pyrrolidine (PVP), co-polymers of vinyl and acrylamide,
polystyrene, polystyrene cross-linked with divinylbenzene or the
like (see, Merrifield Biochemistry 1964, 3, 1385-1390),
polyacrylamides, latex gels, dextran, crosslinked dextrans (e.g.,
Sephadex.TM.), rubber, silicon, plastics, nitrocellulose, natural
sponges, metal, and agarose gel (e.g., Sepharose.TM.). In one
embodiment, the beads are streptavidin-coated beads. The bead
diameter can depend on the density of the well array used with
larger arrays (and thus smaller sized wells) requiring smaller
beads. Generally, the bead size may be about 0.1 .mu.m-10 .mu.m, or
1 .mu.m-5 .mu.m. In an example, the beads are about 5.91 .mu.m in
diameter. In another example, the beads are about 2.8 .mu.m in
diameter. It is to be understood that the beads may or may not be
perfectly spherical in shape. It is to be understood that other
beads may be used and other mechanisms for attaching the nucleic
acid to the beads may be utilized.
[0104] In some embodiments, a homogeneous population of amplified
nucleic acids is conjugated to one or more beads with the proviso
that each bead will ultimately be bound to a plurality of identical
nucleic acid sequences. The degree of loading of nucleic acid
templates onto beads will depend on a number of factors including
the bead size and the length of the nucleic acid. In most aspects,
maximal loading of the beads is desired. Amplification and
conjugation of nucleic acids to solid support, such as beads, may
be accomplished in a number of ways, including, but not limited to,
emulsion PCR as described by Margulies et al. Nature 2005
437(15):376-380 and accompanying supplemental materials. In some
embodiments, the amplification is a representative amplification. A
representative amplification is an amplification that does not
alter the relative representation of any nucleic acid species.
[0105] Before and/or while in the wells of the flow chamber, the
beads are incubated with a sequencing primer that binds to its
complementary sequence located on the 3' end of the template
nucleic acid (i.e., either in the amplification primer sequence or
in another adaptor sequence ligated to the 3' end of the target
nucleic acid) and with a polymerase for a time and under conditions
that promote hybridization of the primer to its complementary
sequence and that promote binding of the polymerase to the template
nucleic acid. The primer can be of virtually any sequence provided
it is long enough to be unique. The hybridization conditions are
such that the primer will hybridize to only its true complement on
the 3' end of the template. Suitable conditions are disclosed in
Margulies et al. Nature 2005 437(15):376-380 and accompanying
supplemental materials.
[0106] Reaction Temperature
[0107] The sequencing reaction can be run at a range of
temperatures. Typically, the reaction is run in the range of
30-60.degree. C., 35-55.degree. C., or 40-45.degree. C. In some
embodiments, it is preferable to run the reaction at temperatures
that prevent formation of a secondary structure in the nucleic
acid. However, this is balanced with the binding of the primer (and
the newly synthesized strand) to the template nucleic acid and the
reduced half-life of Pyrophosphatase at higher temperatures. In one
embodiment, a suitable temperature is about 41.degree. C. The
solutions including the wash buffers and the dNTP solutions are
generally warmed to these temperatures in order not to alter the
temperature in the wells. The wash buffer containing
Pyrophosphatase, however, is preferably maintained at a lower
temperature in order to extend the half-life of the enzyme.
Typically, this solution is maintained at about 4-15.degree. C.,
and more preferably at about 4-10.degree. C.
[0108] Electrode Fabrication
[0109] Referring now to FIGS. 9A-9B, FIGS. 9A-9B illustrate
exemplary CMOS-integrated sensing electrodes 901, 902 in accordance
with an embodiment of the present invention. The metal of choice in
Integrated Circuits (ICs) is generally aluminum with certain amount
of impurities. Accordingly, in the CMOS-integrated embodiments of
the present invention, one may use aluminum metal layers 903 as the
active electrode. The top metal layer in CMOS-integrated sensing
electrode 901, 902 is the optimal metal layer since it is the
closest to the surface of the chip which can be coupled to the DNA
array. The metal layer in CMOS-integrated sensing electrode 901 is
generally covered by a thick passivation layer (typically made of
durable oxides such as SiO.sub.2 and Si.sub.3N.sub.4) to protect
the metal chemically and mechanically from the external
environment. Neither bare aluminum nor aluminum covered by a thick
dielectric layer is an optimal embodiment for the sensing electrode
of active-electrode biosensors. As a result, an insulating layer
904 should be formed. The sensing electrodes 901, 902 of the
present invention in are formed by first creating openings in the
passivation layer of the CMOS chip 905 (comprised of metal layers
906 and silicon substrate 907) and exposing the top metal layer. In
one embodiment, the top metal is subsequently covered by a blanket
layer of an insulator 904 (thickness varies between 5 nm to 2
.mu.m) using conventional thin-film oxide deposition techniques and
the microwells (if any) will be created on its top. Example
materials include SiO.sub.2, Ti.sub.3N.sub.4, Si.sub.3N.sub.4,
TiO.sub.2, Al.sub.2O.sub.3, and HfO.sub.2. The oxide or nitride
layer can be formed by various deposition techniques, such as
Chemical Vapor Deposition (CVD), Atomic Layer Deposition (ALD), or
Physical Vapor Deposition (PVD, such as, e.g., sputtering). In
alternative embodiments, a noble metal layer 908 (e.g., Pt or Au)
is first deposited over the exposed metal by using conventional
thin-film metal deposition techniques (e.g., evaporation or
electroplating), and afterwards, the blanket oxide layer is placed
on top of it. In one embodiment, the thickness of noble layer 908
can vary between 5 nm to 1 .mu.m.
[0110] Integration and SBS Arrays
[0111] In a preferred DNA SBS embodiment of the present invention,
the active-electrode biosensor array is built on the semiconductor
substrate of a CMOS process fabricated using VLSI fabrication
processes. In some cases, the number of the pixels within this
array is greater than 10 and can be as large as 10.sup.8 per single
substrate. Integrated circuit design techniques for selecting the
row and column of the pixel to be interrogated are widely known to
those skilled in the art in the design of CMOS sensor arrays and
image sensor arrays.
[0112] In some situations, the SBS array includes at least 1, 2, 3,
4, 5, 6, 7, 8, 9, 10, 10.sup.2, 10.sup.3, 10.sup.4, 10.sup.5,
10.sup.6, 10.sup.7, 10.sup.8, 10.sup.9, or 10.sup.10 within a
cross-sectional area of at most about 1000 cm.sup.2, 100 cm.sup.2,
10 cm.sup.2, 1 cm.sup.2, 0.5 cm.sup.2, or 0.1 cm.sup.2.
[0113] In some embodiments in which a SBA array is built in a
semiconductor substrate, each biosensing pixel may have part of the
required circuitry to enable a SSCA-based active electrode
biosensing, and, for example, the operational amplifier may be
shared by a plurality of pixels in, for example, all the pixels
within a column of the array. This method of sharing the circuitry
in the signal path is a widely used method in CMOS sensor arrays
and image sensor arrays. In preferred embodiments, the shared
circuits are placed in the periphery of the SBS array to minimize
the size of individual biosensing pixels.
Example Embodiment
[0114] In this section, it is described herein an exemplary
embodiment of the present invention. This system has been
successfully reduced to practice using the 0.18 .mu.m CMOS
fabrication process offered by Taiwan Semiconductor Manufacturing
Company (TSMC). It is noted for clarity that the principles of the
preset invention are not to be limited in scope (such as the scope
of its applications) to the below described details of this
embodiment.
[0115] Referring now to FIGS. 10A-10C, FIGS. 10A-10C illustrate an
integrated active electrode DNA sequencing biochip and the basic
layer structures of its pixels in accordance with an embodiment of
the present invention. Specifically, referring to FIGS. 10A-10C in
conjunction with FIGS. 8A-8C, FIGS. 10A-10C illustrate the basic
structure of the active-electrode CMOS DNA sequencing biochip and
further illustrates the scanning electron microscope image of the
CMOS chip surface with and without DNA-bead complexes. Each pixel
1001 is 16 .mu.m.times.16 .mu.m and the total array size is
90.times.90. The electrode (layer M6) 805 is made of aluminum and
the insulating layer is aluminum oxide (Al.sub.2O.sub.3). Each
pixel 1001 has a shallow microwell (Y=2 .mu.m) built on its top
with a 10 .mu.m.times.10 .mu.m openings 1003 to hold a single 10
.mu.m bead 804, on which DNA strands 801 are immobilized. In this
CMOS process, 6 metal layers (M1-M6) are used and the capacitors,
required for SCCA operation, are created using the
metal-insulator-metal (MIM) layers that are offered in this CMOS
process.
[0116] Referring now to FIGS. 11A-11B, FIG. 11A illustrates the
general architecture of the CMOS-integrated active-electrode
biosensor pixel 1101 for DNA sequencing in accordance with an
embodiment of the present invention and FIG. 11B is a timing
diagram 1102 of in-pixel 1101 and out-of-pixel 1103. The goal is to
measure i(t) by implementing a SSCA with CDS to reduce the effect
of amplifier 1/f noise and offset. The negative feedback of the
amplifier keeps the electrode potential constant and equal to
V.sub.0(t) at all times. The main capacitors in this topology are
the feedback capacitor (C.sub.F=90 fF), the offset storing
capacitor (C.sub.S=105 fF), and the electrode-electrolyte interface
capacitor (1 pF<C.sub.1.parallel.C.sub.D<10 pF).
[0117] FIG. 12 illustrates the transistor-level schematic of the
signal chain in accordance with an embodiment of the present
invention. Referring now to FIG. 12, during the calibration phase,
the amplifier offset is stored onto C.sub.S. Subsequently during
the readout phase, the voltage stored on C.sub.S cancels the offset
and low frequency fluctuations while reducing the amplifier gain
error.
[0118] In order to minimize the pixel circuitry, only the switches
and the differential pair (M1 and M2) of the SCCA are integrated
in-pixel (pixel (i,j)) 1201 as shown in FIG. 12. The rest of the
transistors including the tail current source (M3) are shared at
the column level 1202 (comprised of column analog bus (column (j)
analog bus) 1203 and column amplifier (column (j) amplifier) 1204).
Both C.sub.F and C.sub.S are MIM capacitors and are placed on top
of the active circuitry and below the sensing electrode (see FIGS.
10A-10C). During readout, the i.sup.th row of the array is
activated by ROW[i], which connects the circuitry in the column to
the pixel. The outputs of the SCCAs are available at the column
level and are multiplexed to provide a single buffered output 1205
for the chip. The total consumed power in this chip is
approximately 13 mW using a 3.3V supply.
[0119] FIGS. 13A-13C illustrates the electrical performance of the
SCCAs in accordance with an embodiment of the present invention.
FIG. 13A includes a graph 1301 depicting the frequency versus the
input-referred noise. FIG. 13B includes a graph 1302 depicting the
bandwidth versus integrated noise and FIG. 13C includes a graph
1303 depicting the input voltage (V.sub.IN) and the output voltage
(V.sub.OUT). In the frequencies below 10 kHz, the noise power
spectral density (PSD) is dominated by 1/f noise of the operational
amplifier which can be suppressed by CDS. As evident, by using a 5
kHz CDS, the input-referred noise can be suppressed to
.about.10.mu. Vrms for a 100 Hz bandwidth. The 1 dB compression
point for this amplifier measured at a gain of two is 340 mV which
corresponds to a detection dynamic range of 90 dB.
[0120] FIGS. 14A-14C illustrate the interface capacitance
measurements versus the solution pH in graph 1400 in accordance
with an embodiment of the present invention. Referring to FIG. 14C,
FIG. 14C provides the C.sub.D.parallel.C.sub.1 measurement as the
solution pH is changed. As illustrated in FIG. 14A, a 10 mV
amplitude sinusoidal voltage with a frequency of 5 kHz is applied
to the reference electrode immersed in the solution. The generated
V.sub.OUT(t) (FIG. 14B) is then used to evaluate the SCCA gain and
subsequently C.sub.D.parallel.C.sub.1.
[0121] FIGS. 15A-15C illustrate the DNA polymerization detection in
accordance with an embodiment of the present invention. Referring
now to FIGS. 15A-15C, FIG. 15C provides the measurement results of
real-time DNA polymerization detection. In this experiment, as
shown in FIG. 15A, self-primed biotinylated DNA strands 1501 (SEQ
ID NO:1) are immobilized on streptavidin coated magnetic beads
1502. Deoxynucleotide triphosphates (dATP, dCTP, dGTP, and dTTP
nucleotides) 1503 are added sequentially to trigger polymerization
and perform SBS as shown in FIG. 15B. As is evident in the
experimental results as shown in FIG. 15C, the measured transient
current is large (in the order of 100s of fA) only when the correct
nucleotide is added (dTTP 1504 in this case) and negligible change
can be observed during the control experiment when the unmatched
nucleotides (dATTP, dCTP, and dGTP) are introduced into the
solution.
[0122] FIG. 16 illustrates the micrograph of the active-electrode
CMOS biochip for DNA SBS in accordance with an embodiment of the
present invention. As illustrated in FIG. 16, the 90.times.90
sensor array is placed in the middle of the chip and the electronic
input-output (I/O) of the chip is taken at the periphery. The total
chip size is 2.5 mm.times.2.5 mm.
[0123] The experimental protocol for the measurements is:
[0124] Reagents:
[0125] Biotinylated ss-DNA strand (5'-Biotin-CCTCTGAGTCAAAAAA
[0126] AAGCCGTCGTTATACAACGGAACGTTGTATAACGACGGC-3') from Integrated
DNA Technologies (IDT), USA; Streptavidin coated magnetic beads of
size 8-10 .mu.m from Spherotech, USA; DNA polymerization enzyme
used is Klenow fragment in a 10.times. reaction buffer (500 mM
Tris-HCl (pH 8.0 at 25.degree. C.), 50 mM MgCl.sub.2, 10 mM DTT)
from Fermentas, USA; Deoxynucleotide triphosphates (dNTP) from
Qiagen, USA; Standard buffers.
[0127] DNA Hybridization onto the Micro-Beads:
[0128] Single stranded biotin-DNA in Tris Buffer (20 mM Tris-HCl
buffer, 140 mM NaCl, 20 mM KCl, pH 7.5) was heated at 90.degree. C.
for 5 minutes and slowly cooled down to 25.degree. C. at
0.1.degree. C./min. The product was self-primed DNA.
[0129] 250 .mu.l of 1 mg/50 .mu.l streptavidin-coated magnetic
beads was washed with 1 mL of 20 mM Tris-Buffer and then dispersed
in 250 .mu.l of the same buffer. Next, 10 nmol of biotin-DNA was
added to the solution, followed by incubation on a rotator at room
temperature for 2 hours. The excess DNA not bound to the beads was
washed away with 500 .mu.l of Tris-buffer for five times. The
highest capacity of the beads for biotin DNA is 0.3 nmol/1 mg. At
last, the beads were dispersed in 250 .mu.l of 1.times. Klenow
reaction buffer to make a 1 mg/50 .mu.l beads solution.
[0130] DNA Polymerization Protocol:
[0131] 100 .mu.L of magnetic beads (1 mg/50 .mu.L) were deposited
in the reservoir on top of the CMOS chip. 5 .mu.L of Klenow enzyme
(5 units/.mu.L) was added. The Ag/AgCl reference electrode was
dipped into the reservoir. The CMOS chip was subsequently activated
and the real time data were captured on a PC when dNTPs were added
one at a time.
[0132] Measuring Interface Capacitance Vs. pH:
[0133] 150 .mu.L of Tris-HCl buffer (pH 8.0) was deposited in the
reservoir on top of the CMOS chip. The reference electrode
(Ag/AgCl) was dipped into the reservoir. Sinusoids of amplitude 10
mV and frequency ranging between 1-10 kHz were applied to the
reference electrode. The pixel output amplitude was measured, which
was then used to estimate the interface capacitance C.sub.1. The pH
was changed in steps by adding 1-5 .mu.L quantities of 120 mM
HCl
[0134] The descriptions of the various embodiments of the present
invention have been presented for purposes of illustration, but are
not intended to be exhaustive or limited to the embodiments
disclosed. Many modifications and variations will be apparent to
those of ordinary skill in the art without departing from the scope
and spirit of the described embodiments. The terminology used
herein was chosen to best explain the principles of the
embodiments, the practical application or technical improvement
over technologies found in the marketplace, or to enable others of
ordinary skill in the art to understand the embodiments disclosed
herein.
Sequence CWU 1
1
1155DNAArtificial SequencePrimer 1cctctgagtc aaaaaaaagc cgtcgttata
caacggaacg ttgtataacg acggc 55
* * * * *