Note: Descriptions are shown in the official language in which they were submitted.
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BIOSENSOR ASSAY DEVICE AND METHOD
Field of the Invention
The present invention relates to methods for
detecting or quantitating an analyte, and to devices for
carrying out the method.
Background of the Invention
l0
Many tools used for detecting or quantitating
biological analytes are based on analyte-specific binding
between an analyte and an analyte-binding receptor or
agent. Analyte/analyte binding pairs encountered commonly
in diagnostics include antigen-antibody, hormone-receptor,
drug-receptor, cell surface antigen-lectin, biotin-avidin,
and complementary nucleic acid strands.
A variety of methods for detecting analyte-binding
agent interactions have been developed. The simplest of
~0 these is a solid-phase format employing a reporter labeled
analyte-binding agent whose binding to or release from a
solid surface is dependent on the presence of analyte. In
a typical solid-phase sandwich type assay, for example,
the analyte to be measured is an analyte with two or more
~5 binding sites, allowing analyte binding both to a
receptor, e.g., antibody, carried on a solid surface, and
to a reporter-labeled second receptor. The presence of
analyte is detected (or quantitated) by the presence (or
amount) of reporter bound to solid surface.
30 In a typical solid-phase competitive binding
analyte analog for binding to a receptor (analyte -binding
agent) carried on a solid support. The amount of reporter
signal associated.with the solid support is inversely
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proportional to the amount of sample analyte to be
detected or determined.
The reporter label used in both solid-phase formats
is typically a visibly detectable particle or an enzyme
capable of converting a substrate to an easily detectable
product. Simple spectrophotometric devices allow for the
quantitation of the amount of reporter label, for
quantifying amount of analyte.
Detecting or quantitating analyte-specific binding
events is also important in high-throughput methods being
developed for combinatorial library screening. In a
typical method, a large library of possible effector
molecules (analytes) is synthesized. The library members
are then screened for effector activity by their ability
to bind to a selected receptor. The approach has the
potential to identify, for example, new oligopeptide
antigens capable of high-specificity binding to disease
related antibodies, or small-molecule compounds capable of
interacting with a selected pharmacological tar-get, such
as a membrane bound receptor or cellular enzyme.
High-throughput screening methods typically employ
simple analyte displacement assays to detect and
quantitate analyte binding to a receptor. Displacement
assays have the advantage of high sensitivity, e.g., where
the displaced analyte is radiolabeled, and also allow for
the determination of analyte-receptor binding affinity,
based on competitive displacement of a binding agent whose
binding affinity to the target receptor is known.
In both diagnostics and high-throughput screening,
there is increasing interest in developing biosensors
capable of detecting and quantifying analyte-receptor
binding events.
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One general type of biosensor employs an electrode
surface in combination with current or impedance measuring
elements for detecting a change in current or impedance in
°response to the presence of a ligand-receptor binding -
event. This type of biosensor is disclosed, for example,
in U.S. Patent No. 5,567,301.
Gravimetric biosensors employ a piezoelectric crystal
to generate a surface acoustic wave whose frequency,
wavelength and%or resonance state are sensitive to surface
mass on the crystal surface. The shift in acoustic wave
properties is therefore indicative of a change in surface
mass, e.g., due to a ligand-receptor binding event. U.S.
Patents Nos. 5,478,756 and 4,789,804 describe gravimetric
biosensors of this type.
Biosensors based on surface plasmon resonance (SPR)
effects have also been proposed, for example, in U.S.
Patents Nos. 5,485,277 and 6,492,840. These devices
exploit the shift in SPR surface reflection angle that
occurs with perturbations, e.g., binding events, at the
SPR interface. Finally, a variety of biosensors that
utilize changes in optical properties at a biosensor
surface are known, e.g., U.S. Patent No. 5,268,305.
The interest in biosensors is spurred by a number of
potential advantages over strictly biochemical assay
formats. First, biosensors may be produced, using
conventional microchip technology, in highly reproducible
and miniaturized form, with the capability of placing a
large number of biosensor elements on a single substrate
(e. g., see U.S. Patent Nos. 5,200,051 and 5,212,050).
Secondly, because small signals can be readily
amplified (and subjected to various types of signal
processing if desired), biosensors have the potential for
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measuring minute quantities of analyte, and
proportionately small changes in analyte levels.
A consequence of the features above is that a large
number of different analytes can be detected or
quantitated by applying a small sample volume, e.g., 10-5,0
~,1, to a single multi-sensor chip.
Heretofore, electrochemical biosensors have been more
successfully applied to detecting analytes that are
themselves electrochemical species, or can participate in
catalytic reactions that. generate electrochemical species,
than to detecting analyte-receptor binding events. This
is not surprising, given the more difficult challenge of
converting a biochemical binding event to an
electrochemical signal. One approach to this problem is to
provide two separate reaction elements in the biosensor: a
first element contains a receptor and bound enzyme-linked
analyte, and the second element, components for
enzymatically generating and then measuring an
electrochemical species. In operation, analyte displaces
the analyte-enzyme conjugate from the first element,
releasing the enzyme into the second element region, thus
generating an electrochemical species which is measured in
the second element.
Two-element biosensors of this type are relatively
complicated to produce, particularly by conventional
silicon-wafer methods, since one or more biological layers
and permseleCtive layers must be deposited as part of the
manufacturing process. Further, enzymes or receptors in
the biosensor can denature on storage, and the device may
have variable "wetting" periods after a sample is applied.
Biosensors that attempt to couple electrochemical
activity directly to an analyte-receptor binding event, by
means of gated membrane electrodes, have been proposed.
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For example, U.S. Patent Nos. 5,204,239 and 5,368,712
disclose gated membrane electrodes formed of a lipid
bilayer membrane containing an ion-channel receptor that
is either opened or closed by analyte binding to the
receptor. Electrodes of this type are difficult to make
and store, and are limited at present to a rather small
group of receptor proteins.
Alternatively, direct analyte/receptor binding may be
measured electrically by embedding the receptor in a thin
polymer film, and measuring changes in the film's
electrical properties, e.g., impedance, due to analyte
binding to the receptors. U.S. Patent No. 5,192,507 is
exemplary. Since analyte binding to the receptor will
have a rather small effect on film properties, and since
no amplification effect is achieved, the approach is
expected to have limited sensitivity.
PCT patent application PCT/CA97/00275, published
November 6, 1997, publication No. WO 97/41424, discloses a
novel electrochemical biosensor having a conductive
detection surface, and a hydrocarbon-chain monolayer
formed on the surface. Biosensor operation is based on
the flow of an ionized redox species across the monolayer,
producing a measurable current flow. In one embodiment of
the biosensor disclosed, binding of an analyte to its
opposite binding member attached to the surface of some of
the hydrocarbon chains increases measured current flow by
increasing the disorder of the monolayer, making it more
permeable to the redox species. In another general
embodiment, the opposite binding member is anchored to the
monolayer through a coiled-coil heterodimer structure,
allowing any selected binding member carried on one oc-
helical peptide to be readily attached to a "universal"
monolayer surface carrying the opposite a,-helical peptide.
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The previously disclosed biosensor is capable of
detecting and quantifying analyte-binding events and
characterized by: (i) direct electrochemical conversion of
the binding event to electrical signal; (ii) a high
electron flow "turnover" from each binding event; (iii)
adaptable to substantially any analyte, and (iv) good
storage characteristics and rapid wetting with sample
application.
Given these features, it would be desirable to
improve the operational characteristics and suitability of
the biosensor to a wide variety of analytes, as well as
the adaptability of the biosensor to multianalyte formats,
e.g., in a microfabricated form. The present invention is
designed to provide these advantages.
Summary of the Invention
The invention includes, in one aspect, a method for
detecting or quantitating an analyte present in a liquid
sample. The method includes reacting the liquid sample
with an analyte-reaction reagent, thereby generating a
solution form of a first coil-forming peptide having a
selected charge and being capable of interacting with a
second, oppositely charged coil-forming peptide to form a
stable a-helical coiled-coil heterodimer.
The coil-forming peptide is contacted with a
biosensor having~a detection surface with surface-bound
molecules of such second, oppositely charged coil-forming
peptide, under conditions effective to form a stable a-
helical coiled-coil heterodimer on the detection surface,
where the binding of the solution form of the coil-forming
peptide to the immobilized coil-forming peptide is
effective to measurably alter a signal generated by the
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biosensor, which is measured to determine whether such
coiled-coil heterodimer formation on said detector surface
has occurred.
Tn one embodiment, the analyte is a ligand, and the
reacting includes mixing the analyte with a conjugate of
the first coil-forming peptide and the analyte or an
analyte analog, and reacting the analyte and conjugate
with an analyte-binding anti-ligand agent, such that the
amount of unbound conjugate generated is directly
proportional to the amount of analyte. The analyte-bound
agent is preferably immobilized. In another related
embodiment, the conjugate is bound to the analyte-binding
agent, and displaced from the binding agent in the
presence of analyte.
In still another embodiment, the analyte is an enzyme
and the reacting step is effective to enzymatically
release the first coil-forming peptide in soluble form in
the presence of analyte.
In one general embodiment, the biosensor is an
electrochemical biosensor that includes a conductive
detection surface, a monolayer composed of hydrocarbon
chains anchored at their proximal ends to the detection
surface, and the second charged coil-forming peptide also
anchored to said surface, where the binding of the first
peptide to the second peptide, to form such heterodimer,
is effective to measurably alter current flow across the
monolayer mediated by a redox ion species in an aqueous
solution in contact with the monolayer, relative to
electron flow observed in the presence of the second
peptide alone.
Where the redox ion species has the same charge as
said second coil-forming peptide, the binding of the first
peptide to the second peptide is effective to enhance ion-
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mediated current flow across said monolayer. Where the
redox ion species has a charge opposite that of said
second coil-forming peptide, the binding of the first
peptide to the second peptide is effective to reduce ion-
s mediated current flow across said monolayer.
In another aspect, the invention includes a
diagnostic device for detecting or quantitating an analyte
present in a liquid sample. The device includes a
reaction reagent effective to react with analyte to
generate a solution form of a first coil-forming peptide
having a selected charge and being capable of interacting
with a second, oppositely charged coil-forming peptide to
form a stable a-helical coiled-coil heterodimer.
A biosensor in the device has a detection surface
with surface-bound molecules of a second charged, coil-
forming peptide capable of interacting with the first
oppositely charged coil-forming peptide to form a stable
a.-helical coiled-coil heterodimer, where the binding of
the first peptide to the second peptide, to form such
heterodimer, is effective to measurably alter a signal
generated by the biosensor, which is measured by a
detector in the device.
The device may further include a substrate having
formed therein (i) a sample-introduction region, (ii) the
biosensor, and (iii) a sample-flow pathway between said
sample-introduction region and the biosensor. The
reaction reagent is disposed in the sample-flow pathway
and includes a conjugate of the first coil-forming peptide
and the analyte or an analyte analog, in a form releasable
into the sample liquid, and an analyte-binding agent. The
sample-flow pathway may include a mixing zone containing
the conjugate in releasable form, and a reaction zone
containing the analyte-binding agent in immobilized form.
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The device also preferably includes a background
control biosensor, and a control sample-flow pathway
connecting the sample-introduction region to the
background control biosensor, for measuring "baseline"
current. The control sample-flow pathway lacks the
analyte-conjugate.
In one general embodiment, the biosensor includes a
conductive detection surface, a monolayer composed of
hydrocarbon chains anchored at their proximal ends to the
detection surface, and the second charged coil-forming
peptide also anchored to the surface, where the binding of
the first peptide to the second peptide, to form such
heterodimer, is effective to measurably alter current flow
across the monolayer mediated by a redox ion species in an
l5 aqueous solution in contact with the monolayer, relative
to electron flow observed in the presence of the second
peptide alone.
The redox ion species may have the same charge as the
second coil-forming peptide, where the binding of the
first peptide to the second peptide is effective to
enhance redox ion-mediated current flow across the
monolayer. Examples are the redox ion species is Fe(CN)63-
if the charge of the second coil-forming peptide is
negative, and Ru(NH3)63~, if the charge of the second coil-
forming peptide is positive.
Alternatively, the redox ion species may have a
charge opposite to that of the second coil-forming
peptide, where the binding of the first peptide to the
second peptide is effective to reduce ion-mediated current
flow across Said monolayer. Examples are Fe(CN)63-, if the
charge of the second coil-forming peptide is positive, and
Ru(NH3)63+, if the charge of said second coil-forming
peptide is negative.
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In this general embodiment the electrode may have a
gold detection surface and the monolayer may be composed
of 8-22 carbon atom chains attached at their proximal ends
to the detection surface by a thiol linkage, at a
molecular density of about 3 to 5 chains/nm~.
The device may be designed for use in detecting or
quantitating a plurality of different selected analytes.
Here the device includes, for each analyte, (i) a separate
biosensor, and (ii) a separate sample-flow pathway
20 connecting the sample-introduction region to each
associated biosensor, where each sample-flow pathway
includes (i) a conjugate of the first coil-forming peptide
and one of the selected analytes or analog thereof, and
(ii) an associated selected analyte-binding agent.
Preferably each biosensor in this embodiment contains
substantially the same second charged, coil-forming
peptide, and the sample-introduction region is a single
port communicating with each of the sample-flow pathways.
The sample .introduction region, biosensors, and sample-
flow pathways may be microfabricated on the substrate.
Several embodiments of the invention, particularly
those listed just below, are practiced optionally without
the need for charged heterodimer subunits because of the
presence of an electro-active group on the end of the
first coil-forming peptide as a redox reactive agent.
The invention further provides for a method for
detecting or quantitating an analyte present in a liquid
sample by reacting said liquid sample with an analyte-
reaction reagent, by said reacting, generating a solution
form of a first coil-forming peptide capable of
interacting with a second coil-forming peptide to form a
stable cc-helical coiled-coil heterodimer, contacting said
first coil-forming peptide generated by said reaction with
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a biosensor having a detection surface with surface-bound
molecules of such second coil-forming peptide, under
conditions effective to form a stable a-helical coiled-
coil heterodimer on said detection surface, where said
binding of said solution form of said coil-forming peptide
to said immobilized coil-forming peptide is effective to
measurably alter a signal generated by said biosensor, and
measuring said signal generated by said biosensor, to
determine whether such coiled-coil heterodimer formation
on said detector surface has occurred. This method may
optionally include said first coil-forming peptide further
comprising an electro-active group attached to said first
coil-forming peptide where said group is available for
redox induction by an electrode adjacent to said detection
surface with said second coil-forming peptide attached
thereto when said first coil-forming peptide forms a
coiled-coil heterodimer with said second coil-forming
peptide.
The invention further provides for a diagnostic
device for use in detecting or quantitating an analyte
present in a liquid sample, comprising a reaction reagent
effective to react with analyte to generate a solution
form of a first coil -forming peptide capable of
interacting with a second coil-forming peptide to form a
stable a-helical coiled-coil heterodimer, a biosensor
having a detection surface with surface-bound molecules of
a second coil-forming peptide capable of interacting with
the first coil-forming peptide to form a stable a-helical
coiled-coil heterodimer, where the binding of the first
peptide to the second peptide, to form such heterodimer,
is effective to measurably alter a signal generated by the
biosensor, and a detector for measuring the change in a
signal generated by the biosensor, in response to
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conjugate binding to the second coil-forming peptide.
This device may optionally include said first coil-forming
peptide further comprising an electro-active group
attached to said first coil-forming peptide where said
group is available for redox induction by an electrode
adjacent to said detection surface with said second coil-
forming peptide attached thereto when said first coil-
forming peptide forms a coiled-coil heterodimer with said
second coil-forming peptide.
In another aspect, the invention provides a method
for detecting or quantitating an analyte present in a
liquid sample,
comprising the steps of:
(a) reacting the liquid sample with an analyte-
reaction reagent, where the reacting generates a
solution form of a first coil-forming peptide
capable of interacting with a second coil-forming
peptide to form a stable a-helical coiled-coil
heterodimer,
(b) contacting the first coil-forming peptide
generated by the reaction with a biosensor having
a detection surface with surface-bound molecules
of the second coil-forming peptide, under
conditions effective to form a stable a,-helical
coiled-coil heterodimer on the detection surface,
where the binding of the solution form of the
coil-forming peptide to the immobilized coil-
forming peptide is adapted to be detectable by a
detector, and
(c) detecting the binding to determine whether such
coiled-coil heterodimer formation on the detector
surface has occurred.
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In another aspect, the first coil-forming peptide of
the above method further comprises an electro-active group
attached to the first coil-forming peptide where the group
is available for redox induction by an electrode adjacent
to the detection surface with the second coil-forming
peptide attached thereto when the first coil-forming
peptide forms a coiled-coil heterodimer with the second
coil-forming peptide.
In yet another aspect, the first coil-forming peptide
of the method above further comprises a detectable
feature, the feature being detectable by the detector when
the first coil-forming peptide is bound to the second
coil-forming peptide.
In another aspect of the invention further provides
the detectable feature being indirectly detectable upon
binding of a detection moiety to the detectable feature,
and where the moiety is detectable by the detector. In
another aspect, the detectable feature is directly
detectable by the detector. In still another aspect, the
detectable feature is selected from the group consisting
of fluorophores, radioisotopes, and dyes. In another
aspect; the detectable feature is selected from the group
consisting of biotin, streptavidin, one component of an
antigen-antibody pair, one component of a ligand-receptor
pair, one component of a homologous nucleic acid duplex or
triplex, wherein the detectable feature is complementary
to a complementary detection moiety so that the
complementary detection moiety specifically binds the
detectable feature, the complementary detectable moiety
being adapted to be detectable by the detector.
The invention further provides for the detection
surface of the method above to be in electrical
communication with an amperometric detector so that when
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the electro-active group is adjacent the detection
surface, the detector detects current flow through the
detection surface.
Another aspect of the invention provides the method
above where the detector measures graviometric changes.
Another aspect of the invention provides the method
above where the detector measures surface plasmon
resonance changes.
The invention further provides for a diagnostic
l0 device for use in detecting or quantitating an analyte
present in a liquid sample, comprising
(a) a reaction reagent effective to react with
analyte to generate a solution form of a first
coil -forming peptide capable of interacting with
l5 a second coil-forming peptide to form a stable
cc-helical coiled-coil heterodimer, and
(b) a biosensor having a detection surface with
surface-bound molecules of a second coil-forming
peptide capable of interacting with the first
20 coil-forming peptide to form a stable a,-helical
coiled-coil heterodimer, where the binding of
the first peptide to the second peptide, to form
such heterodimer, is effective to generate a
signal, the signal being adapted for detection
25 by a detector for measuring the change in the
signal generated by the binding, in response to
the first coil-forming peptide binding to the
second coil-forming peptide.
The invention, in another aspect provides the device
30 above where the first coil-forming peptide further
comprises an electro-active group attached to the first
coil-forming peptide where the group is available for
redox induction by an electrode adjacent to the detection
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surface with the second coil-forming peptide attached
thereto when the first coil-forming peptide forms a
coiled-coil heterodimer with the second coil-forming
peptide.
The invention further provides the device above where
the first coil-forming peptide further comprises a
detectable feature, the feature being detectable by the
detector when the first coil-forming peptide is bound to a
second coil-forming peptide. Still further, the invention
provides the device above where the detectable feature is
indirectly detectable upon binding of a detection moiety
to the detectable feature, and where the moiety is
detectable by the detector. In another aspect, the
detectable feature is directly detectable by the detector.
In another aspect, the detectable feature is selected from
the group consisting of fluorophores, radioisotopes, and
dyes. In yet another aspect, the detectable feature is
selected from thee group consisting of biotin,
streptavidin,~one component of an antigen-antibody pair,
one component of a ligand-receptor pair, one component of
a homologous nucleic acid duplex or triplex.
The invention further provides for a preferred
modification of the hydrocarbon chain of the monolayer
where such monolayer further includes a hydrophillic
portion to reduce the hydrophobic nature of the
hydrocarbon thereby reducing the occurrence of non-
specific binding to the monolayer, the hydrophillic
portion, for example, being a hydroxyl, carbohydrate, or
hydrophillic group.
The invention further provides for the device above
where the detection surface is in electrical communication
with an amperometric detector so that when the electro-
active group is adjacent the detection surface, the
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detector detects current flow through the detection
surface.
Another aspect of the invention provides for the
device above where the detector measures graviometric
changes. Another aspect of the invention provides for the
device above where the detector measures surface plasmon
resonance changes.
These and other objects and features of the invention
will become more fully apparent when the following
detailed description of the invention is read in
conjunction with the accompanying drawings.
Brief Description of the Drawings
Fig. 1A is a simplified, partly schematic perspective
view of a microfabricated substrate used in a single-
analyte detection device in accordance with one aspect of
the invention;
Fig. 1B is an exploded view of an embodiment of a
detection device in accordance with the invention.
FIG. 1C is a simplified, partly schematic perspective
view of another embodiment of a device in accordance with
the invention;
FIG. 1D is a schematic and enlarged diagram of a side
view of a section of the device of FIG. 1C taken along the
arrows 1D;
Fig. 2A is a partly schematic view of the mixing zone
in a biosensor device;
Fig. 2B is a partly schematic view of the mixing zone
50 in a biosensor device indicating the migration of test
sample as a cross-hatched area;
Fig. 3A is a partly schematic view of the reaction
zone in a biosensor device;
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Fig. 3B is a partly schematic view of the reaction
zone in biosensor device indicating the migration of
a
test sample as a cross-hatched area;
Fig. 4 is a cross-sectional view taken in the
direction of arrows 4-4 in Fig. 1A of an amperometric
biosensor constructed in accordance with one embodiment
of
the invention;
Fig. 5A is an enlarged view of a region of the
electrode in the biosensor shown in Fig. 4;
10Fig. 5B is an enlarged view of a region of the
electrode in the biosensor shown in Fig. 4, in the
presence
of first
coil-forming
peptide-analyte
conjugate;
Fig. 6 is a simplified, partly schematic perspective
view of microfabricated substrate used in a Pseudomonas
a
15PAK pilin peptide-detection device in accordance with
one
aspect the invention;
of
Fig. 7A is an enlarged view of a region of the
electrode in the biosensor shown in Fig. 6 taken in the
direction of arrows 7-7;
20Fig. 7B is an enlarged view of a region of the
electrode in the biosensor shown in Fig. 6 taken in the
direction of arrows 7-7, in the presence of first coil-
forming ptide-PAK peptide conjugate;
pe
Figs. 8A and 8B are idealized linear (8A) and semilog
25(8B) plotsshowing change in biosensor signal, measured
in
nA, as unction of PAK pilin peptide added to the
a f
biosensor device, where the second a-helical peptide
has
the same harge as the redox ion species;
c
Fig. 9 shows elements of a gravimetric biosensor
30constructed in accordance with an embodiment of the
invention;
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Fig. 10 shows elements of a surface plasmon resonance
biosensor constructed in accordance with an embodiment of
the invention;
Fig. 11 shows elements of an optical biosensor
constructed in accordance with an embodiment of the
invention;
Fig. 12 is a simplified, partly schematic perspective
view of a mufti-analyte detection device in accordance
with one aspect of the invention;
Fig. 13 is a view of a region of the device of Fig.
12 taken in the direction of arrows 13-13.
FIG. 14 is a plot showing change in biosensor signal,
measured in nA*sec, as a function of PAK pilin peptide
level in a competitive antibody binding assay;
FIG. 15 illustrates the structure of an electrode
monolayer having an embedded K-coil peptide subunit;
FIG. 16 illustrates the structure of an electrode
monolayer having an embedded K-coil peptide subunit, in
the presence of E-coil/PAK peptide conjugate.
FIGS. 17A - 17B depict the structures of the first
coil-forming peptide (17A) and the second coil-forming
peptide (17B).
FIG. 18 depicts an electrochemical reaction of the
DNP group that results in a detectable signal.
FIG, 19 is a graphical depiction of square-wave
voltammograms comparing the effect of the presence of
dissolved oxygen with the effect of purging the biosensor
system with argon.
FIGS. 20A - 20B depict model coiled-coil heterodimers
attached to a surface adjacent electroactive groups.
FIGS. 21A - 21C depict the detection of the
electroactive group embodiments of the invention upon
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their contact with an electrode surface (21A - 21B), and
the resulting signal (21C).
Detailed Description of the Invention
Fig. 1A is a simplified perspective view of a
diagnostic device 10 in accordance with one embodiment of
the present invention. The apparatus includes a substrate
12, sample introduction region 14, sample-flew pathway 16,
biosensor 20, and a detector for measuring a change in a
signal generated by the biosensor 22. The device
optionally includes a signal-responsive element for
recording the output signal, e.g., a visually readable
l5 output 24. Details of each element are further described
hereinbelow. The device preferably includes a control
sample-flow pathway 26 and background control biosensor
28.
Before proceeding further with the description of the
specific embodiments of the present invention, a number of
terms will be defined.
"Analyte" is defined as the compound or composition
to be measured, which is a member of a specific binding
pair (sbp) and may be a ligand, which is mono- or poly-
valent, usually antigenic or haptenic, a single or
plurality of compounds which share at least one common
epitopic or determinant site, or a receptor.
"Analyte-binding agent" is defined as any compound or
composition capable of recognizing a particular spatial
~0 and polar organization of an analyte molecule, e.g.,
epitopic or determinant site. The device of the present
invention can be used in detecting the presence or amount
in a sample of an analyte which forms with an analyte-
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binding agent, an analyte-analyte-binding agent pair.
Non-limiting examples of such pairs include antigen-
antibody, hormone-receptor, drug-receptor, cell surface
carbohydrate-lectin, biotin-avidin, and complementary
nucleic acids. Numerous examples of such pairs are known
(e.g., as described in U.S. Patent No. 5,716,778 (1998) to
Unman) .
"Analyte analog" is defined as a modified analyte
that can compete with the analogous analyte for a
l0 receptor, the modification providing means to join an
analyte analog to another molecule.
"Bibulous material" is defined as a porous material
having pores of at least 0.1 ~.un, preferably at least 1.0
Vim, which is susceptible to traversal by an aqueous medium
l5 in response to capillary force. Such materials are
generally hydrophilic or are capable of being rendered
hydrophilic and include inorganic powders such as silica,
magnesium sulfate, and alumina~ natural polymeric
materials, particularly cellulosic materials and materials
20 derived from cellulose, such as fiber containing papers,
e.g.; filter paper, chromatographic paper, etc.; synthetic
or modified naturally occurring polymers, such as
nitrocellulose, cellulose acetate, poly (vinyl chloride),
polyacrylamide, cross linked dextran, agarose,
25 polyacrylate, etc.; either used by themselves or in
conjunction with other materials; ceramic materials; and
the like. The bibulous material can be attached to a
support. The bibulous material may be polyfunctional or
be capable of being polyfunctionalized to permit covalent
30 bonding of compounds as described hereinbelow.
Binding of molecules to the bibulous material may be
accomplished by well-known techniques, commonly available
in the literature. See, for example, "Immobilized
CA 02420743 2003-02-26
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Enzymes," Tchiro Chibata, Halsted Press, New York (1978)
and Cuatrecasas, et al., J. Bio. Chem., 245:3059 (1970).
The bibulous material can be a single structure such
as a sheet cut into strips or it can be particulate
material bound to a support or solid surface. In a
preferred embodiment, the material is applied using a
screen printing technique, such as described in U.S.
Patent No. 5,736,188.
The substrate 12 is a generally planar solid support
for the device, preferably composed of an electrically
insulating, non-porous, rigid, moisture impermeable
material. A wide variety of organic and inorganic
materials, both natural and synthetic, and combinations
thereof, may be employed provided only that the support
l5 does not interfere with the capillary action of the
bibulous material, or non-specifically bind assay
components, or interfere with the signal producing system.
Tllustrative polymers include polyethylene, polypropylene,
poly(4-methylbutene), polystyrene, polymethacrylate,
polyethylene terephthalate), nylon, polyvinyl butyrate),
glass, and ceramics. The device preferably includes a
cover 30 which can be transparent.
The sample introduction region 14 provides a site for
application of a liquid sample containing analyte 34
(Figs. 2A-2B, 3A-3B). As used herein, "liquid sample"
typically refers to a naturally occurring or artificially
formed liquid test medium suspected of containing the
analyte of interest. The liquid sample may be derived
from a wide variety of sources such as physiologic fluid
illustrated by blood, serum, plasma, urine, ocular lens
fluid, saliva, amniotic and spinal fluid, etc., food
products such as milk or wine, chemical processing
streams, or waste water, etc. The volume of sample can
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vary between 1-200 ~,Z. In a preferred embodiment, the
sample introduction region is an inlet port which is in
liquid communication with the with sample flow pathway 16.
The sample-flow pathway 16 is a channel or conduit
between the sample-introduction region 14 and the
biosensor 20. In one embodiment, the sample-flow pathway
is an essentially unblocked passage of a diameter suitable
fox conveying the liquid sample from the sample
introduction region 14 to the biosensor 20. Such a
microfluidic trench can be formed by micromachining
substrate material, e.g., as described in U.S. Patent No.
5,194,133 or by injection molding. Liquid sample migrates
in the sample flow pathway by capillary action or can be
driven by a micro-pump or by electroosmosis. As described
hereinbelow, assay reagents are bound either in a
releasable form or immobilized within the pathway 16. For
example, the reagents can be.bound to the walls of the
pathway 16. In a preferred embodiment, along the length
o.f the sample-flow pathway 16 is an insoluble bibulous
material 38 for conveying the sample by capillarity and
for binding assay reagents either in a releasable form or
an immobilized form.
Downstream of the sample introduction region 14, the
sample flow-pathway includes a mixing zone 40 as shown in
Fig. 1A and Figs. 2A-2B. The mixing zone contains a
conjugate 42 consisting of the analyte (or analyte analog)
34 linked to a charged coil-forming peptide 44. Peptide 44
is selected for forming a heterodimer with an oppositely
charged coil-forming peptide which is anchored within the
biosensor 20, as described hereinbelow. Conjugate 42 is
provided in mixing zone 40 in a form which is releasable,
i.e., diffusible, into the sample liquid when sample is
drawn, e.g., by capillarity, into the mixing zone. For
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example, conventional methods are used for releasably
binding the conjugate to the bibulous material, such as
the spot and dry method as described in U.S. Patent No.
5,580,794 (which is incorporated by reference in its
entirety herein) or using an applicator such as the Bio
Dot dispenser (Bio Dot, Inc. Irvine, CA).
Downstream of the mixing zone 40 is reaction zone 48
(Fig. 1A and Figs. 3A-3B) in which analyte-binding agent
50 is immobilized within the pathway 16. In one
embodiment, the analyte-binding agent is an antibody to
the analyte. The immobilization of proteins onto glass
and other surfaces, are known (e. g., as discussed in U.S.
Patent No. 5,192,507, which is incorporated by reference
in its entirety herein). Immobilization of molecules to a
bibulous material is performed using conventional methods
such as a soak and dry immobilization method, or by
immobilizing the protein to latex microparticles of about
5-20 ~,un and drawing these modified microparticles into the
membrane matrix using vacuum pressure.
During the passage of liquid sample through the
reaction zone 48, analyte 34 and conjugate 42 react with
the binding agent 50 under conditions effective to
immobilize analyte or conjugate so bound to the binding
agent. In a preferred embodiment of the invention, the
device further includes a control sample-flow pathway 26,
in liquid communication with the sample introduction
region 14 and a background control biosensor 28.
Preferably, the control sample-flow pathway lacks the
conjugate in the control mixing zone, but is otherwise
similar to the sample-flow pathway 16.
It will be appreciated from the discussion
hereinabove that due to competitive binding in reaction
zone 48, there will be an inverse relationship between the
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amount of analyte in a liquid sample and the amount of
conjugate 42 bound in zone 48. Sample liquid flowing
downstream from zone 48 contains conjugate 42 not bound in
zone 48. Thus, there is a positive relationship between
the level of conjugate in the liquid sample liquid
emerging downstream from zone 48 and the level of analyte
in the original liquid sample. The amount of reagents,
such as binding agent 50 and conjugate 42, provided within
the sample-flow pathway is predetermined by optimization
methods well known in the art. For example, in the
absence of analyte, the amount of binding agent 50 in zone
48 preferably is just sufficient to bind all of conjugate
42.
In general, the biosensor of the invention has a
detection surface with surface-bound molecules of a second
charged, coil-forming peptide capable of interacting with
a first oppositely charged coil-forming peptide to form a
stable a,-helical coiled-coil heterodimer. The two
oppositely charged peptides spontaneously self-assemble
into a heterodimer complex. In one embodiment, the
invention employs an electrochemical biosensor which
measures current flow across a hydrocarbon-chain
monolayer, anchored to the detector surface, mediated by
redox species in aqueous solution in contact with
monolayer relative to electric flow observed in the
absence of analyte-peptide conjugate. Other embodiments
employ a gravimetric biosensor, a surface plasmon
resonance biosensor, or an optical biosensor, as described
hereinbelow.
30p In a general embodiment of the invention, a second
charged coil-forming peptide 46 is anchored to the
biosensor surface. A first charged coil-forming peptide
44 is linked to analyte (or analog) as a conjugate 42.
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The peptides 44 and 46 form a heterodimer and are two non-
identical, preferably oppositely charged polypeptide
chains, typically each about 21 to about 70 residues in
length, having an amino acid sequence compatible with
their formation into stable two-stranded a-helical
heterodimeric coiled-coils. They are designated herein as
second coil-forming peptide (heterodimer-subunit peptide
1), and first coil-forming peptide (heterodimer-subunit
peptide 2). In the discussion below, second coil-forming
peptide will refer to the peptide attached to the
biosensor surface in the biosensor, and first coil-forming
peptide, to the peptide having an attached analyte. It
will be understood that these designations refer to the
functional role played by the subunit peptide, not the
actual peptide sequence.
Another embodiment of the invention is shown in FIG.
1B. In analogy with device 10, the device 700 includes a,
substrate 710, biosensor 712, sample flow path 714
comprising a bibulous material, a conjugate pad 716, a
plastic (e. g., polyurethane) spacer 718, and filter
element 720. The conjugate pad, comprising a bibulous
matrix, contains diffusively bound conjugate. The flow
path 714 comprises a bibulous matrix containing non-
diffusively bound analyte binding agent. Sample
introduced at a sample application area 724 permeates the
filter 720 that serves to remove particulate matter, such
as cells. Spacer 718 facilitates the distribution of the
fluid flowing from filter 720 into pad 716. The fluid
mixes with first coil-forming peptide/analyte conjugate in
pad 716 and enters matrix 714 that includes a mixing zone
726. The fluid emerging from the matrix 714 enters
biosensor 712 as described hereinabove. The device
preferably includes a background. control biosensor and a
CA 02420743 2003-02-26
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positive control biosensor in analogy with device 250
described herein.
Yet another embodiment of the device 600 is shown in
FIGS. 1C, 1D. The device 600 includes a housing 614,
sample application port 610, microchannel 611, biosensor
chamber 612, filter element 616, porous support element
618, conjugate pad 620, and matrix element 622. Sample
introduced at port 610 enters the top of the device
permeates the filter 616 which serves to remove
particulate matter, such as cells. The conjugate pad 620
contains diffusively bound conjugate. Matrix 622
comprises a bibulous matrix containing non-diffusively
bound analyte binding agent. Support element 618 supports
the filter 616 and has openings therethrough for downward
transfer, under the influence of gravity and also
capillary action, of fluid from filter 616 into conjugate
pad 620. The fluid mixes with first coil-forming
peptide/analyte conjugate in pad 620 and enters matrix
622. The fluid emerging from the matrix 622 enters
biosensor 612. A plurality of separate biosensors can be
arranged, preferably on a planar on a substrate, to
receive sample fluid delivered from microchannels
emanating from a central sample application area (FIG.
1C). The device preferably includes a background control
biosensor and a positive control biosensor in analogy with
device 250 described herein.
The devices of the present invention can include
filters which can include means for sample pre-treatment,
such as filtering red blood cells (U. S. Patent Nos.
5,658,444; 5.837,546; 5,747,274). The filters can
comprise separation material such as synthetic membranes,
fibrous depth filters such as glass fiber; plastic fiber,
metal fibex, cellulose fiber or any combination of filters
26
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and membranes. The separation material may be untreated
or can be coated with protein, dextran, sugars, or
carbohydrates for red cell stabilization, LDL ,
precipitation reagents such as magnesium chloride and
dextran sulfate, antibodies or red cell agglutination
agents to facilitate red cell removal. Sample
pretreatment within a filter can also adjust the pH to
within a specific range, reference salt concentration,
turbidity and or viscosity, and/or reduce or remove
interfering substances such as immunochemical cross-
reactants, redox substances and the like.
Fig. 4 shows a simplified schematic view of an
electrochemical biosensor 60 for detecting an analyte-
peptide quantitating conjugate in a liquid medium, in
Z5 accordance with the invention. The biosensor includes a
chamber 61 that is in liquid communication with the
sample-flow pathway of the device via opening 62.
Although not shown, the chamber may include a second port
or vent to facilitate liquid flow through the port. The
biosensor 60 includes a working electrode 64 having a
conductive detection surface 66, and, in a preferred
embodiment, a hydrocarbon-chain monolayer 68 formed on the
detection surface. In the embodiment shown, the detection
surface is the upper surface of a conductive film 70
deposited on substrate 71. Details of the monolayer
formed on the detection surface, and the method of forming
the monolayer on the surface, are discussed below.
A cover 73 in the apparatus has an upper wall 74, and
side walls, such as wall 76, which are joined to edge
regions of the substrate to form a closed chamber 61
therewith. The chamber serves to hold an aqueous solution
required for biosensor operation, as will be described.
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A reference electrode 78 and a counter electrode 80
in the apparatus are provided on the chamber-facing
surface of wall 74, as shown, and are thus both in
conductive contact with electrode 64 when the chamber is
filled with an electrolyte solution. In the device of the
invention, the sample liquid enters the chamber through an
opening such as shown at 62 downstream from the mixing
zone. In a preferred embodiment, electrolyte reagents are
provided within the chamber preferably in a dry form that
is readily dissolved within the sample fluid. For
example, reagents can be lyophilized and deposited, or
spotted and dried, in the reaction zone or in the sample
flow pathway as described in U.S. Patent No. 5,580,794.
The liquid entering the chamber mixes with ionic species
capable of undergoing a redox reaction, i.e., losing or
w gaining an electron, at a suitably charge electrode.
Exemplary redox species are Fe(CN)6~-~4-, as a negatively
charged species, and Ru (NH3) 6z+/s+ as a positively charged
species. Other probes which can be used include Mo(CN)63-
(ED = +800 mV) , W (CN) 63- (Ea=+580 mV) , Fe (CN) 4- (Eo=+580 mV) ,
Ce4+~3+, (Eo=+1.4V) , and Fe+3/z+ (ED= +666mV) . Typical redox
ion concentrations are between 0.01 and 10 mM. The
solution is contained in chamber 60 and is in contact with
reference and counter electrodes.
The Voltage potential placed on the electrode, i.e.,
between the electrode and reference electrode, is
typically at least 90 mV above the electrochemical
potential (ED) value of the redox species, for oxidation,
and at least 90 mV below the electrochemical potential,
for reduction of the species. Consider, for example,
Fe (CN) g3 /4-, with an Eo of 450 mV (vs . NHE) . Above about
550 mV electrode potential, any Fe2+ species is oxidized
to Fe3+, and at an electrode potential below about 350 mV,
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and Fe+3 is reduced to Fe+z. Similarly, Ru (NH3) 6a+/s+ has an
Eo of +50 mV (vs. NHE), so oxidation is achieved at an
electrode potential above about +150 mV, and reduction,
below about -50 mV.
The reference electrode 78, which is held at ground,
serves as the voltage potential reference of the working
electrode 64 when a~selected potential is placed on the
working electrode by a voltage source 82. This potential
is measured by a voltage measuring device 84_which can
additionally include conventional circuitry for
maintaining the potential at a selected voltage, typically
between about -500 to +800 mV.
Voltage source 82 is connected to counter electrode
80 through a current measuring device 86 as shown, for
l5 measuring current flow between the two electrodes during
biosensor operation. The reference and counter electrodes
are Pt, Ag, Ag/AgCl, or other suitable electrodes. The
reference and working electrodes, the~circuitry connecting
them to the working electrode, and voltage source, are
referred to herein, collectively, as means for measuring
ion-mediated electron flow across the working-electrode
monolayer, in response to heteroduplex formation between
first coil-forming peptide-analyte conjugate and a
charged, coil-forming peptide second coil-forming peptide
46 anchored to the surface 66 of the working electrode 64.
Figs. 5A-5B are an enlarged view of a portion of the
working electrode 64 having a conductive detection surface
66 and a hydrocarbon-chain monolayer 68 formed on the
detection surface. The chains forming the monolayer are
typically 8-22 carbon, saturated hydrocarbon chains,
although longer chains, chains with some unsaturation,
chains with non-carbon chain atoms, such as lipid ethers,
and/or chains with minor branching, such as by non-chain
29
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methyl groups, may be employed, within the constraint that
the chains, at a sufficient packing density, form a
sufficiently close packed and ordered monolayer to be
effective as a barrier to electron flow, under biosensor
operating conditions, as discussed below. This density is
calculated to be between 3-5 chainsJnm2.
In the embodiment shown in Fig. 5A-5B, the chains 90
are coupled to the electrode detecting surface through
sulfhydryl linkages 92, although other suitable coupling
groups may be employed. One method for producing
monolayers having suitable hydrocarbon chain densities is
passive diffusion of chains onto the surface of an
electrode. A preferred method consists of actively
driving the chains onto the surface by applying a positive
voltage potential to the conductor surface. The latter
method achieves rapid monolayer formation and highly
reproducible electrode characteristics.
In a preferred embodiment of the invention, the
hydrocarbon-chain mixture which is actively driven onto
the conductor surface includes peptide second coil-forming
peptide peptide 46 that is capable of forming a
stabilized, a,-helical peptide heterodimer with an
oppositely charged, complementary subunit, first coil-
forming peptide. Such heterodimer subunits are described
in PCT patent application WO 95/31480 "Heterodimer
Polypeptide Immunogen Carrier Composition and Method",
publication date 23 November 1995, which is incorporated
herein by reference. Exemplary subunits are referred to
herein as K-coils, referring to positively charged
subunits whose charge~is provided dominantly by lysine
residues, and E coils, referring to negatively subunits
whose charge is provided dominantly by glutamic acid
residues. Preferred examples from the above-mentioned PCT
CA 02420743 2003-02-26
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patent Application 95/31480 include the following
sequences and their original sequence numbers in brackets
as found in said PCT application.
Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala Leu Glu
Lys Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala Leu
Glu Lys Glu Val Ser Ala Leu Glu Lys {PCT Sequence No:
19}(SEQ ID NO: 1)
Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala Leu
Glu
Lys Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala
Leu
Glu Lys Glu Val Ser Ala Leu Glu Lys {PCT Sequence
No:
20} (SEQ N0: 2)
ID
Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala Leu Lys
Glu Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala Leu
Lys Glu Lys Val Ser Ala Leu Lys Glu {PCT Sequence No:
2l} (SEQ ID N0: 3)
Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala Leu
Lys
Glu Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala
Leu
Lys Glu Lys Val Ser Ala Leu Lys Glu {PCT Sequence
No:
22} (SEQ N0: 4)
ID
Second coil-forming peptide peptide 46 can be
attached to the distal end of a short hydrocarbon chain
(end opposite the chain's thiol group) by suitable lipid-
to-peptide conjugation, e.g., by ester linkage to a
hydrocarbon fatty acid. Alternatively, the peptide may be
linked to the electrode surface through a peptide spacer,
e.g., a tripeptide spacer that extends from one end of the
subunit and includes cysteine as a terminal residue, for
sulfhydryl attachment to the electrode surface. In both
cases, the modified peptide is mixed with the hydrocarbon
chains, at a selected mole ratio, then driven into a
monolayer formation by applying a positive voltage to the
electrode, resulting in.a densely packed hydrocarbon-chain
monolayer 68 which includes charged, coil-forming peptide
46 embedded in the planar chain matrix, while still
31
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retaining a low dielectric barrier to ion flow through the
monolayer. The second coil-forming peptide peptide 46 is
included in the monolayer in a mole ratio
peptide/hydrocarbon chains of preferably between 1:100 to
1:5.
In a preferred method for forming the monolayer, a
mixture of thiol-containing chains and thiol-terminated
second coil-forming peptide peptide, at a selected mole
ratio, is actively driven to the surface by applying a
positive voltage potential to the. substrate surface, e.g.,
gold film. In practice, the hydrocarbon chain mixture
(about 1 mM hydrocarbon chalnsj in an ethanolic solution
of 100 mM T~i perchlorate,. neutral pH, is placed over the
electrode, and a selected potential is applied to the
electrode. The buildup of the monolayer can be monitored
by increase in layer thickness. Alternatively, monolayer
formation is monitored by measuring current across the
monolayer, as described below. In this case, formation of
the monolayer will be characterized by a steady drop in
electrode current, until minimum current is reached, at
which point maximum chain paoking has been achieved.
Active deposition of the C16 and peptide subunit can
be carried out sequentially in addition to the "mixed
mode", or simultaneous deposition, described hereinabove.
The conditions for the sequential deposition are
essentially the same except that the peptide subunit is
deposited first and the C16 subsequently.
The time required to achieve saturation. packing
density will vary with applied voltage, and can be as
short as 10 seconds--about 4 orders of magnitude faster
than monolayer formation by diffusion. Complete or nearly
complete monolayer formation (30 A thickness) occurs
within 20 minutes at about 1V potential and above. At
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lower positive voltages, additional reaction time is
required. Preferably the voltage applied to the electrode
is at least between about -~-250 mV relative to a normal
hydrogen electxode (+250 vs. NHEj and 1.2V (vs. NHEj.
Not only are rapid monolayer formation times
achieved, but the percentages of peptide and hydrocarbon
chains present in the reaction mixture are precisely
represented in the monolayers, giving highly reproducible
electrode characteristics.
In aqueous medium, the isolated heterodimer-subunit
peptides are typically random coils. When second coil-
forming peptide and first coil-forming peptide are mixed
together under conditions favoring the formation of a,-
helical coiled-coil heterodimers, they interact to form a
two-subunit a,-helical coiled-coil heterodimeric complex.
Peptides in an a-helical coiled-coil conformation interact
with one another in a characteristic manner that is
determined by the primary sequence of each peptide. The
tertiary structure of an oc-helix is such that 7 amino acid
residues in the primary sequence correspond to
approximately 2 turns of the a,-helix. Accordingly, a
primary amino acid sequence giving rise to an cc-helical
conformation may be broken down into units of 7 residues
each, termed heptads. The heterodimer-subunit peptides
are composed of a series of heptads in tandem. When the
sequence of a heptad is repeated in a particular
hetexodimer-subunit peptide, the heptad may be referred to
as a "heptad repeat", or simply "repeat".
The dimerization of second coil-forming peptide and
first coil-forming peptide is due to the presence of a
repeated heptad motif of conserved amino acid residues in
each peptide°s primary amino acid sequence. Repeating
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heptad motifs having appropriate amino acid sequences
direct the second coil-forming peptide and first coil-
forming peptide polypeptides to assemble into a
heterodimeric a-helical coiled-coil structure under
permissible conditions. The individual a-helical peptides
contact one another along their respective hydrophobic
faces.
Second coil-forming peptide and first coil-forming
peptide may assemble into a heterodimer coiled-coil helix
l0 (coiled-coil heterodimer) in either parallel or
antiparallel configurations. Tn a parallel configuration,
the two heterodimer-subunit peptide helixes are aligned
such that they have the same orientation (amino-terminal.
to carboxyl-terminal). In an antiparallel configuration,
the helixes are arranged such that the amino-terminal end
of one helix is aligned with the carboxyl-terminal end of
the other helix, and vice versa.
Heterodimer-subunit peptides designed in accord with
the guidance presented in the above-referenced PCT
30 application typically show a preference for assembling in
a parallel orientation vs. an antiparallel orientation.
Fog example, the exemplary peptides identified by Glu Val
Glu Ala Leu Gln Lys Glu Val Ser Ala Leu Glu Lys Glu Val
Ser Ala Leu Glu Cys Glu Val Ser Ala Leu Glu Lys Glu Val
Glu Ala Leu Gln Lys (SEQ ID N0:5) and Lys Val Glu Ala Leu
Lys Lys Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala Leu
Lys Cys Lys Val Ser Ala Leu Lys Glu Lys Val Glu Ala Leu
Lys Lys (SEQ ID N0:6), form parallel-configuration
heterodimers as do other peptide sequences (as discussed
in the PCT application). When attaching an analyte to
first coil-forming peptide, it is generally desirable to
attach the analyte at or near the end of the peptide that
will form the distal end of the heterodimer. In
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particular, where the heterodimer forms a parallel
configuration, the second coil-forming peptide peptide is
preferably anchored to the biosensor surface at its C
terminus, and the analyte conjugated to the first coil-
s forming peptide peptide at its N terminus.
As just noted, one of the two subunit peptides
(second coil-forming peptide) in the heterodimer is
anchored to the biosensor surface, and the other peptide
(first coil-forming peptide) contains an analyte intended
l0 to participate in a binding reaction in the reaction zone
of the device. In both cases, the peptide is synthesized,
or derivatized after synthesis, to provide the-requisite
attachment function and analyte, respectively.
The first coil-forming peptide or second coil-forming
15 peptide may further comprise a detectable feature. The
detectable feature may be connected or conjugated to the
either coil-forming peptide in a variety of different ways
readily apparent to one of ordinary skill in the art.
Exemplary connecting or conjugating methods are disclosed
~0 in the Examples section, below. In general, most
conjugating methods do not disrupt the coil-forming
activity of the either coil-forming peptide, nor do such
conjugations disrupt the activity of conjugated analyte,
if present.
25 Considering the modification of second coil-forming
peptide, the peptide may be synthesized, at either its N
or C terminus, to carry additional terminal peptides that
can function as a spacer between the biosensor surface and
the helical-forming part of the peptide. Alternatively,
30 the second coil-forming peptide peptide can be attached to
the biosensor surface thorough a high-affinity binding
reaction, such as between a biotin moiety carried on the
CA 02420743 2003-02-26
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peptide and an avidin molecule covalently attached to the
surface.
Where second coil-forming peptide is embedded in a
hydrocarbon-chain monolayer (Figs. 5A-5B) the spacer
anchoring the second coil-forming peptide peptide to the
biosensor surface may be a hydrocarbon chain. The chain
is preferably a fractional length.of the chains making up
the bilayer, such that the distal ends of the heterodimer
which forms upon binding of the two peptides in the
assembled monolayer are at or near the exposed surface of
the monolayer. Thus, for example, if the monolayer is
made up of 18-carbon chains, the spacer is preferably 2-10
carbons in length, depending on the length of the
heterodimer.
The hydrocarbon-chain spacer, in the form of a omega-
thio fatty acid, may be coupled to a terminal hydroxyl or
amine coupling during solid-phase synthesis, as outlined
above. The derivatized peptide, in turn, can be attached
to a metal surface by standard thiolate coupling
(Dakkouri, et al., Zangmuir (1996) 12:2849-2852).
Where the analyte is a polypeptide, the analyte can
be'synthesized by either solid-state or recombinant
methods, to include the peptide analyte at the end of the
first coil-forming peptide peptide that will orient
distally in the assembled heterodimer. Where the analyte
is a non-peptide moiety, e.g., a non-peptide hormone,
drug, or nucleic acid, the first coil-forming peptide
peptide can be synthesized to include one or more residues
that can be specifically derivatized with the analyte. In
forming the conjugate, such as 42, the analyte is
preferably covalently attached to the N-terminal amino
acid residue, or to one of the residues facing the exposed
face of the heterodimer. Preferred coupling groups are
36
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the thiol groups of cysteine residues, which are easily
modified by standard methods. Other useful coupling
groups include the thioester of methionine, the imidazolyl
group of histidine, the guanidinyl group of arginine, the
phenolic group of tyrosine and the indolyl group of
tryptophan. These coupling groups can be derivatized
using reaction conditions known to those skilled in the
art.
To bind the analyte-first coil-forming peptide
conjugate 42 to the surface-immobilized second coil-
forming peptide peptide 46, the two peptides are contacted
under conditions that favor heterodimer formation.. A
medium favoring coiled-coil heterodimer formation is a
physiologically-compatible aqueous solution typically
having a pH of between about 6 and about 8 and a salt
concentration of between about 50 mM and about 500 mM.
Preferably, the salt concentration is between about 100 mM
and about 200 mM. An exemplary benign medium has the
following composition: 50 mM potassium phosphate, 100 mM
KCl, pH 7. Equally effective media may be made by
substituting, for example, sodium phosphate for potassium
phosphate and/or NaCI for KCl. Heterodimers may form
under conditions outside the above pH and salt range,
medium, but some of the molecular interactions and
relative stability of heterodimers vs. homodimers may
cutter trom characteristics detailed above. For example,
ionic interactions between the ionic groups that tend to
stabilize heterodimers may break down at low or high pH
values due to the protonation of, for example, Glu side
chains at acidic pH, or the deprotonation of, for example,
Zys side chains at basic pH. Such effects of low and high
pH values on coiled-coil heterodimer formation may be
overcome, however; by increasing salt concentration.
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Increasing the salt concentration can neutralize the
stabilizing ionic attractions or suppress the
destabilizing ionic repulsions. Certain salts have
greater efficacy at neutralizing the ionic interactions.
For example, in the case of the K-coil peptide 44 in Fig.
5B, a 1M or greater concentration of C104- anions is
required to induce maximal a-helical structure, whereas a
3M or greater concentration of C1- ions is required for the
same effect. The effects of high salt on coiled-coil
formation at low and high pH also show that interhelical
ionic attractions are not essential for helix formation,
but rather, control whether a coiled-coil tends to form as
a heterodimer vs. a homodimer.
Figs. 5A-5B show a biosensor electrode 64 in which
the hydrocarbon chain monolayer 68 includes an E-coil
peptide subunit, such as subunit 46, as described above.
In the embodiment shown, each peptide subunit is coupled
to the electrode surface via a tripeptide spacer, such as
spacer 94 in subunit 46, which is itself attached to the
electrode surface through a sulfhydryl linkage, as shown.
The peptide, including the peptide spacer, is formed
conventionally, e.g., by solid phase synthesis. The
monolayer was formed according to the method described
above.
Because of the negative charge imparted to the
monolayer by the E coil subunits 46, the monolayer shows
relatively low conductance to negatively charged redox
species, such as Fe(CN)63~, as evidenced by a relatively
low oxidation or reduction current with the redox species.
Fig. 5B shows the same monolayer, but after addition
of complementary, positively charged K-coil subunits 44
conjugated to analyte, such as indicated at 42. As shown,
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oppos~.tely charged subunits pair to form charge-neutral
heterodimers in the monolayer.
Without wishing to be bound by theory, in the absence
of heteroduplex formation between the two charged coiled
peptides, the monolayer retains its net negative charge,
forming an effective barrier to electron flow across the
monolayer mediated by a redox ion species of the same
charge, when a suitable oxidizing or reducing potential is
placed across the monolayer. This is reflected by a low
measured current across the membrane. With binding of an
analyte-first coil-forming peptide conjugate to an
anchored second coil-forming peptide in the monolayer, the
repulsive negative of the monolayer is reduced
sufficiently to allow the movement of redox species
through the monolayer, producing electron flow through the
electrode. The biosensor records this binding event as an
increase in current across the electrode, i.e., between
the working and counter electrodes. It will be recognized
that the peptide used for second coil-forming peptide can
have a negative or positive charge, but that the preferred
redox ion has the same charge as second coil-forming
peptide. Thus, in an alternative embodiment of this
aspect of the invention, second coil-forming peptide can
be a K-coil, and a redox species of the same charge can be
used, e.g., Ru(NH3)s+, with a negatively charged first
coil -forming peptide used in the conjugate.
By analogy to a transistor, the redox solution serves
as the "source", the monolayer as-the "gate", and the
underlying electrode as the "drain". Current flow in a
transistor is initiated by applying a threshold voltage to
the gate. In the biosensor of the invention, current flow
is initiated by a stimulus --in this case, heteroduplex
formation-- to the monolayer "gate".
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Figs. 6 and 7A-7B show a diagnostic device 110 for
detection and quantitation of Pseudomonas PAK pilin
peptide 112 constructed in accordance with the invention.
The sample flow pathway 224 of the device contains a
polystyrene-divinylbenzene matrix and has a mixing zone
116 containing a conjugate of first coil-forming peptide-
PAK peptide bound in a releasable form. Binding zone 118
contains (16-thiohydroxy)hexadecanyl-O-(2-acetamido-2-
deoxy-(3-D-galactopyranosyl-(1,4)-(3-D-galactopyranoside
immobilized to the bibulous material through the 16 thiol
moiety (not shown), This disaccharide is specifically
reactive with PAK peptide, forming a ligand-receptor pair
with the peptide.
Figs. 7A-7B show a magnified view of the biosensor
electrode 119 of device 110 which includes a hydrocarbon
monolayer 120 with embedded second coil-forming peptide
122 covalently attached to the electrode surface 124. The
biosensor electrode was prepared as described with
reference to Fig. 5B; employing a ratio of non-second
coil-forming peptide to second coil-forming peptide-chains
of about 4 to 1. Figs. 7A and 7B show second coil-forming
peptide before and after binding of an first coil-forming
peptide-PAK conjugate 126, respectively.
The operating response of the biosensor is
illustrated in Figs. 8A-8B. An increase in Pseudomonas
PAK protein receptor in the test sample produces an
increase in biosensor signal. The signal increases by
about 2-fold, from 225 nA to 400 nA, over a concentration
range of 0 to 250 (M PAK in the test sample.
FIG. 15 shows a biosensor electrode 524 in which the
hydrocarbon chain monolayer, indicated at 526 includes K-
coil peptide subunits, such as subunit 528, as described
above. In the embodiment shown, each peptide subunit is
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coupled to the electrode surface via a tripeptide spacer,
such as spacer 530 in subunit 528, which is itself
attached to the electrode surface through a sulfhydryl
linkage, as shown. The peptide, including the peptide
spacer, is formed conventionally, e.g., by solid phase
synthesis. The amount of peptide subunit in the monolayer
is about 20 mole percent. The nionolayer was formed
according to the method described above with respect to
FIGS. 5A, 7A.
Presumably because of the positive charge imparted to
the monolayer by the K-coil subunits, the monolayer shows
relatively high conductance to negatively charged redox
species, such as Fe(CN)63-, as evidenced by a relatively
high oxidation or reduction current with the redox
species.
FIG. l6 shows the same monolayer, but after addition
of complementary, negatively charged E-coil subunits/PAK
conjugate, such as indicated at 530. As shown, oppositely
charged subunits pair to form charge-neutral heterodimers
in the monolayer. This pairing is effective to reduce
monolayer conductance substantially, as evidenced by the
time-dependent fall in measured oxidation or reduction
current in the presence of Fe(CN)63 ions (FIG. 16). A
biosensor for determining PAK levels was constructed as
described in Example 1 and demonstrates lower current at
increased levels of PAK (FIG. 14). In Example l,
microtiter plates were coated with antibody to PAK protein
receptor (PAK). Various levels of PAK were incubated in
the wells with E-coil/PAK conjugate. At higher PAK
concentration, more of the E-coil/PAK conjugate is free in
solution due to competitive binding with the antibody.
Originally, the surface of the K/C16 chip is positively
charged allowing a high flow of negatively charged redox
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probe. At higher amounts of conjugate, the E-coil binds
with the K-coil and "neutralizes" the positive charge by
forming charge-neutral heterodimers that led to a decrease
in conductance.
Fig. 9 shows basic elements of a gravimetric
biosensor 150 incorporating the novel biosensor surface of
the invention. The~biosensor has a piezoelectric crystal
90 whose biosensor surface is indicated at 154. Second
coil-forming peptide peptide 158 is anchored to the
biosensor surface. A preferred embodiment includes a
hydrocarbon monolayer 156 with second coil-forming peptide .
peptide 158 embedded therein.
Surface acoustic waves (SAW) are generated in the
crystal by an oscillator 160. According to known
piezoelectric biosensor principles, the change in mass in
the biosensor surface resulting from heterodimer formation
alters the frequency, resonance frequency, and wavelength
of the SAW, and at least one of these wave characteristios
is measured by a detector 162. The oscillator and
detector collectively form detector means for detecting
heterodimer formation. Details of crystal construction
and associated detector means in gravimetric biosensors
are given, for example, in US Patent Nos. 5,478,756 and
4,789,804, and in PCT application WO 96/02830.
Fig. 10 shows basic elements of a surface plasrnon
resonance (SPR) biosensor 170 incorporating the novel
biosensor surface of the invention. A chamber 172 in the
biosensor contains a waveguide composed of a dielectric
film 174 (e.g., glass), a thin evaporated metal film 176
(e.g., chromium or titanium), and a thin film 178
(preferably of gold) constructed to support surface
plasmon waves at the dielectric/metal film interface. The
waveguide surface forms a biosensor surface having second
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coil-forming peptide, such as E coil peptide 182, anchored '
thereon. A preferred embodiment includes a hydrocarbon
monolayer 180 with second coil-forming peptide peptide 182
embedded therein,
A light source 184 directs a divergent light beam
onto the biosensor surface through a lens 186. At some
region along the length of the biosensor surface, the beam
angle strikes the surface at an absorption angle at which
absorption from the evanescent wave by surface plasmons
occurs. The absorption angle will shift with changes in
the composition of the material near the interface, that
is, in response to binding events occurring on the
monolayer surface.
The intensity of reflected light from each region
along the biosensor surface is monitored by a photosensor
188 whose photosensing grid is matched to specific
detector surface regions, and which is operatively
connected to an analyzer 190. .The light source and
photosensor are also referred to herein as biosensor
means.
In operation, the SPR absorption angle on the
biosensor surface is measured before and after application
of test sample, with the measured shift in angle being
proportional to the extent of surface heterodimer
formation.
A variety of biosensor devices which rely on changes
in the optical properties of a biosensor surface, in
response to ligand/anti-ligand binding events, have been
proposed. Fig. 11 shows basic elements of an optical
biosensor apparatus 200 having an open chamber 202 and a
biosensor surface 204 with second coil-forming peptide
anchored, such as shown at 207. A preferred embodiment
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includes a hydrocarbon monolayer 206 with second coil-
forming peptide peptide 207 embedded therein.
The detector means in the apparatus for detecting
binding events on the biosensor surface includes a source
208 of polarized light and a lens system 210 for directing
the light in a beam through the region of the monolayer.
A photodetector 212 at the opposite side of the biosensor
surface functions to measure intensity of light at a given
polarization angle, through a polarization filter 214.
Detection of heterodimer formation is based on the change
of polarization, angle and intensity of light transmitted
by the monolayer in response to perturbation of the
regular order of the monolayer by surface binding events.
These changes are recorded by an analyzer 216 operatively
connected to the photosensor.
The detection devices described hereinabove are used
in the detection of a single analyte. A multi-analyte
detection device is readily constructed according to Figs.
12 and 13 which show a multi-analyte detection device 250,
and employing a biosenso.r selected form those as described
hereinabove. The preferred device has a sample
introduction area 252, such as a port, in liquid
communication with each of a plurality of sample-flow
pathways, such as flow paths 254, 256. Each sample flow
path preferably is filled with a bibulous material and has
a separate mixing zone, reaction zone, and biosensor for
measuring a different analyte present in the test sample.
Each mixing zone, such as 258, 260, contains a separate,
releasably bound first coil-forming peptide-analyte (or
analyte analog) conjugate for each analyte being detected.
Each reaction zone, such as 262, 264, contains an
immobilized analyte-binding agent corresponding to the
first coil-forming peptide-analyte conjugate in the
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associated mixing gone. In one embodiment, all of the
biosensors, such as 266, 268, contain second coil-forming
peptide anchored to an electrochemical detection surface
within a hydrocarbon monolayer as described hereinabove.
In a preferred multi-analyte device, at least one of the
sample flow paths, such as 270, is a control sample-flow
pathway which lacks analyte first coil-forming peptide
conjugate. The control sample-flow pathway is in liquid
communication with background control biosensor 272
containing second coil-forming peptide anchored to a
detection surface within a hydrocarbon monolayer, as
described hereinabove. In the operation of this
embodiment of the invention, test sample applied to the
sample application region 252 flows into each of the
separate sample-flow pathways and interacts with the
respective reagents within each sample flow pathway in
analogy to the dual lane electrochemical device
illustrated in Fig. 1A. Each biosensor is connected to a
detector for measuring the change in signal generated by
each biosensor, in response to heteroduplex formation.
The device preferably is provided with a lid 274.
The device preferably includes at least one biosensor
which will be used as a positive control biosensor. This
biosensor is not connected to the sample application port.
It includes a fixed amount of first coil-forming peptide-
analyte conjugate pre-added, during manufacture of the
positive control biosensor chamber, in an amount to give
the limit of maximum expected response. Preferably, the
positive control biosensor chamber contains the first
coil-forming peptide/analyte conjugate, redox species
probe, and buffer-salts in a dried form which are
rehydrated at the time of use of the device. For example,
an aliquot of aqueous solution (e.g., .albumin) can be
CA 02420743 2003-02-26
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contained in a reservoir (not shown). The control
solution can be injected into the positive control
biosensor chamber through a passage between the chamber
and reservoir. The injection can be performed
electronically, e.g., by a minipump or electroosmotic
movement of solution through the passage, or manually,
. e.g., by depressing an injection means such as a flexible
bulb associated with the reservoir (not shown).
More generally, the diagnostic device of the
invention includes a reaction reagent effective to react
with analyte to generate a solution form of a first~coil-
forming peptide of the type described above. The device
further includes a biosensor, e.g., of the type described
above, having a detection surface with surface-bound
molecules of a second charged, coil-forming peptide
'capable of interacting with the first oppositely charged
coil-forming peptide to form a stable cc-helical coiled-
coil heterodimer, where the binding of the first peptide
to the second peptide, to form such heterodimer, is
effective to measurably alter a signal generated by the
biosensor, and a detector for measuring the change in a
signal generated by the biosensor, in response to
conjugate binding to the first charged, coil-forming
peptide, as detailed above.
The reaction reagent in the device may include, for
example, an immobilized or aggregated anti-analyte binding
agent having a conjugate of the first coil peptide and
analyte bound thereto, with the presence of analyte acting
to release conjugate from the binding agent in proportion
to the amount of analyte. The analyte, might be, for
example, one of a large number of combinatorial species,
where the assay is used for high-throughput screening for
compounds effective to bind the binding agent.
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Alternatively, the reaction reagent could be an
immobilized enzyme-cleavable conjugate designed to cleave
off and release coil peptide in the presence of an enzyme
analyte. In still another embodiment, the first coil is
produced recombinantly, either alone or as a fusion
protein, in a recombinant system, and the presence of
_ protein expression is detected by reaction of the
recombinantly made coil protein with the biosensor. Other
reagents and reagent formats in which the presence of
10~ analyte is effective to generate a first coil peptide in
soluble form, such as are known to those skilled in the
art, are also contemplated herein.
In the embodiment disclosed above, having a sample-
introduction region and a sample-flow pathway between a
l5 sample-introduction region and the biosensor, the reaction
reagent is disposed in the sample-flow pathway and
includes a conjugate of the first coil-forming peptide and
the analyte or an analyte analog, in a form releasable
into the sample liquid, and an analyte-binding agent.
20 Also forming part of the invention, is a general
assay method for detecting or quantitating an analyte
present in a liquid sample. The method includes reacting
the liquid sample with an analyte-reaction reagent,
thereby to generate a solution form of a first coil-
25 forming peptide having a selected charge and being capable
of interacting with a second, oppositely charged coil
forming peptide to form a stable a,-helical coiled-coil
heterodimer. The first peptide so generated is contacted
with a biosensor having a detection surface with surface-
30 bound molecules of such second, oppositely charged coil
forming peptide, under conditions effective to form a
stable oc-helical coiled-coil heterodimer on said detection
surface, where the binding of the solution form of the
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coil-forming peptide to the immobilized coil-forming
peptide is effective to measurably alter a signal
generated by the biosensor. The signal generated by the
biosensor is then measured, to determine whether such
coiled-coil heterodimer formation on said detector surface
has occurred.
From the foregoing, it can be seen how various
objects and advantages of the invention are met. The
device of the invention can be formed under controlled
manufacturing conditions consistent with microchip scale
and photomask processes, to produce highly uniform and/or
miniaturized and/or high-density array biosensor devices
with sample introduction region, biosensor, and sample-
flow pathway microfabricated on the substrate. The
invention can be used to create mufti-analyte assay
surfaces by photomasking techniques that are capable of
producing diagnostic devices having a plurality of unique
sample-flow pathways, in fluid communication with highly
reproducible biosensor elements.
After manufacture of a device with a plurality of
identical biosensors, the plurality of sample-flow
pathways can be readily adapted to a wide variety of
analytes(s), by binding an first coil-forming peptide
peptide conjugate in a releasable form within the mixing
zone and by binding analyte-binding agent in an
immobilized form within the reaction zone, carried out
under relatively simple production conditions, thus
combining both manufacturing precision at the initial
production stage, and assay flexibility at the analyte and
analyte-binding agent addition stage.
The invention is easily adapted to any of a variety
of biosensor devices, such as those illustrated above.
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Figure 17A depicts the structures of first coil-
forming peptide 1701 of the electro-active group
embodiment of the present invention. Figure 17B depicts
second coil-forming peptide 1703 with two DNP groups 1705a
and 1705b attached thereto.
In a particularly preferred embodiment, an electro-
active group is attached to the second coil-forming
peptide enabling the peptide to become electro-active and
therefore detectable by electrochemical means. One
skilled in the art would recognize that several different
species of electro-active groups can be used, for example,
dinitrophenyl (DNP) groups preferably attached to a
cysteine residue of the second coil-forming peptide
preferably through two histidine residues modified with
DNP as a side chain.
Figure 18 depicts what is believed to be the
electrochemical reaction of DNP group 1800 that results in
a detectable signal. DNP group 1800 is reduced in three
steps. Nitro-groups 1801 and 1803 are electrochemically
reduced to two nitroso groups (-NO) 1807 and 1809 during
first step 1805. Then they are further reduced in step
1811 to -NHOH 1813 and 1815, and finally in step 1817 to
amino groups 1819 and 1821. Such reduction steps produce
multiple peaks on a current (signal) to voltage graph,
however, in certain circumstances such peaks may overlap.
Variation of the reaction pH may also be manipulated to
elucidate individual peaks from overlapping peaks or
unusual peak shapes.
Figure 19 depicts exemplary data produced by the
practice of the electro-active group embodiment of the
present invention. Graph 1900 compares the effect of the
presence of dissolved oxygen, line A, with the effect of
purging the biosensor system with argon for ten minutes to
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remove dissolved oxygen as shown in line B. Shown are
square-wave voltammograms for both A and B. The reaction
solution contained lOmM phosphate buffer at pH 7Ø The
sensors were made from gold and prepared by first cleaning
with a piranha solution, then washing with water, followed
by incubation in a 20 uM second coil-forming peptide
solution for 3 hours, then rinsing again with water,
followed by incubation in 1.6 ~M DNP tagged first coil-
forming peptide solution for 40 minutes.
Figures 20A and 20B depict model coiled-coil
interaction 2001 resulting in the close association of
second coil-forming peptide 2003a with electro-active
group 2003b with first coil-forming peptide 2005 attached
to surface 2007 of gold electrode 2009 resulting in
coiled-coil heterodimer 2011 attached to surface 2007 with
electro-active group 2003b adjacent surface 2007. Zn this
embodiment, first coil-forming peptide is attached to
surface 2007 of gold electrode 2009 via a covalent bond
between a sulfur group and gold atoms, not shown, on
surface 2007. Tn the absence of first coil-forming
peptide 2005, attached first coil-forming peptide 2005
assumes a random, non-coiled structure as shown in frame
A. Zikewise, second coil-forming peptide 2003a with
electro-active group 2003b assumes, in the absence of
first coil-forming peptide 2005, a random, non-coiled,
bent shape. Fig, 20A depicts the before interaction
states of the first and second coil-forming peptides.
Fig. 21B depicts the effect of coiled-coil interaction
2001 resulting in coiled-coil heterodimer structures 2011.
After the coiled-coil interaction, both coils stretch up
with the electro-active groups 2003b facing down towards
gold electrode surface 2007.
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Figures 21A - 21C depicts the detection of the
electo-active group embodiments of the invention upon
their contact to an electrode surface occurs. Upon
formation of coiled-coil heterodimer complex 2101, Fig.
21A, electro--active group 2103 is brought into close
proximity to gold electode surface 2105, Application of
an electrical potential (voltage) causes a reduction of
electro-active group 2103 to 2103(b) as shown in Fig. 21B.
A signal is directly proportional to the number of
electro-active groups 2103 in proximity of gold electrode
surface 2105. In particular, electrochemical signals
result from the redox reaction involving gold electrode
surface 2105 and electro-active group 2103 from a square-
wave potential scan as shown in Fig. 21C. During the
scan, the electro-active groups shown in Fig. 21A are
reduced via multiple steps to amine groups as represented
by the darkened 2103(b) (Fig. 21B) producing voltammetric
peaks. Since the two nitro groups have slightly different
chemical environments, they should produce multiple peaks.
Peak 2110 in Fig. 22C is the overall signal from the
combined signals caused by overlapped peaks resulting from
each stage of the redox reaction.
Other possible electro-active groups exist. A good
electro-active group suitable for practicing the electro-
active group embodiments of the invention are generally,
but not limited to being, small in size, soluble in water,
easily conjugated to the second coil-forming peptide,
having a redox potential away from interfering substances
such as oxygen, For example, Bromo-dinitrophenol with or
without histidine, bromo-phenol, bromo-(3,5
dihydroxy)benzene, and bromo-methyl hydrazone. These
candidates represent molecules from the phenol and
hydrazine groups of chemicals. Other compounds such as
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from the catechol- family are also useful as electro-
active groups.
Example 1
Immunoassay for PAK protein
In a competitive ELTSA format, anti-PAK mouse IgG
(0.5 ug/well) was coated on a 96-well microtitre plate via
an overnight incubation in 10 mM sodium carbonate buffer,
pH, 9.5. The wells were blocked with 3% BSA in PBS, pH
7.4. A mixture of E-conjugate (E-coil/PAK protein
conjugate) and various amount of PAK was prepared and
incubated with the immobilized antibody at 37oC for 2
hours. Aliquots removed from the plate after the
incubation and applied to the K/C16 chip for analysis.
A biosensor chip (K/C16 chip) comprising C16
hydrocarbons and K-coil was inserted into a modified
Hewlett Packard HPLC electrochemical system. The sensor
chip was equilibrated with a continuous flow of PBS and
baseline response was taken by repeat injections (20 ~,L).
of a probe solution (ferricyanide probe 1mM in PBS).
Measurement was taken at 180 mV versus a Ag/AgCl reference
electrode.
Samples from the competitive ELISA were injected
using the autosampler. Each sample injection was followed
by injections of the probe solution. The signal from the
probe injections were integrated, averaged and reported.
The data are presented in FIG. 14 which shows the
area under the peak (~.A * sec) vs. concentration of PAK in
the incubation.
Although the invention has been described with
respect to particular devices and methods, it will be
understood that various changes and modifications can be
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made without departing from the invention, as encompassed
by the accompanying claims.
All publications and patent applications cited in
this specification are herein incorporated by reference as
if each individual publication or patent application were
specifically and individually indicated to be incorporated
by reference.
Although the foregoing invention has been described
in some detail by way of illustration and example for
purposes of clarity of understanding, it will be readily
apparent to those of ordinary skill in the art in light of
the teachings of this invention that certain changes and
modifications may be made thereto without departing from
the spirit or scope of the appended claims.
l5
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Example 2
Preparation of E- and K-coil peptide conjugates
The E- and K-coil peptides can be conjugated to other
proteins or biomolecules by two approaches:
A. Addition of a reactive chemical group to the peptide
during the synthesis process. Examples are:
l0
1. Cysteine
Coupling of a cysteine residue at either
the C- or N-terminus of the peptides so as
to introduce a free sulfhydryl (-SH) group.
l5 Conjugation of the peptide to other
biomolecules can be done via Reactions 2,
3, 5, or 8.
2. Bromoacetylation of the N-terminal end of the
20 peptides
The introduced halogen is reactive towards
proteins containing cysteine residues and
sulfhydryl-modified biomolecules,
preferably at pH higher than 7.5 (Reaction
25 3) .
3. Benzoylbenzoic acid (BBA)
A benzophenone is formed by coupling the BBA to the
N-terminus of the peptide using standard solid phase
30 peptide synthesis procedures. The photoreactive
benzophenone enables the peptide to react with
primary, secondary, and tertiary carbons along with
various sidechain of a protein (Reaction 6).
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B. Use of cross-linking reagents (available from Pierce
Chemical Company and Sigma-Aldrich Fine Chemicals):
1. Reactive towards amine (-NH~)and sulfhydryl (-
SH) groups
a) Employing Reactions 1 and 2
AMAS (N-[a-Maleimidoacetoxy]-succinimide ester)
BMPS (N-[(3-Maleimidopropyloxy]-succinimide ester)
EMCS (N-[s-Maleimidocaproyloxy]-succinimide ester)
GMBS (N-[y-Maleimidobutyryloxy]-succinimide ester)
ZC-SMCC (Succinimidyl-4-[N-maleimidomethyl]-
cyclohexane-1-carboxy-[6-amidocaproate])
MBS (m-Maleimidobenzoyl-N-hydroxysuccinimide ester)
I5 SMCC (Succinimidyl-4-[N-maleimidomethyl]-
cyclohexane-1-carboxylate)
SMPB (Succinimidyl-4-[p-maleimidophenyl]-butyrate)
SMPH (Succinimidyl-6-[((3-maleimidopropionamido)-
hexanoate])
~0 Sulfo-EMCS (N-[E-Maleimidocaproyloxy]-
sulfosuccinimide ester)
Sulfo-GMBS (N-[y-Maleimidobutyryloxy]-
sulfosuccinimide ester)
Sulfo-KMUS (N-[tc-Maleimidoundecanoyloxy]-
25 sulfosuccinimide ester)
Sulfo-MBS (m-Maleimidobenzoyl-N-
hydroxysulfosuccinimide ester)
Sulfo-SMCC (Sulfosuccinimidyl-4-[N-
maleimidomethyl]-cyclohexane-1-
30 carboxylate)
Sulfo-SMPB (Sulfosuccinimidyl-4-[p-
maleimidophenyl]-butyrate)
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b) Employing Reactions 1 and'3
SBAP (Succinimidyl-3-[bromoacetamido]-propionate)
SIA (N-Succinimidyl-iodoacetate)
SIAB (N-Succinimidyl-[4-iodoacetyl]-aminobenzoate)
Sulfo-SIAB (Sulfosuccinimidyl-[4-iodoacetyl]-
aminobenzoate)
c) Employing Reactions 2 and 4
BMPA (N-~3-Maleimidopropionic acid)
EMCA (N-s-Maleimidocaproic acid)
KMUA (N-7c-Maleimidoundecanoic acid)
d) Employing Reactions 1 and 8
SVSB (N-Succii~.imidyl-[4-vinylsulfonyl]-benzoate)
e) Employing Reactions 1 and (2 or 3 or 8)
SATA (N-Succinimidyl-S-acetylthioacetate)
SATP (N-Succinimidyl-S-acetylthiopropionate)
2. Reactive towards sulfhydryl (-SH) group and
carbohydrate
~ Employing Reactions 2 and 9
BMPH (N-[[3-Maleimidopropionic acid]-hydrazide~TFA)
EMCH (N-[~-Maleimidocaproic acid]-hydrazide)
KMUH (N-[~c-Maleimidoundecanoic acid]-hydrazide)
MZCzH (4-[N-Maleimidomethyl]-cyclohexane-1-
carboxylhydrazide~HCl~'~ dioxane)
MPBH (4-[4-N-Maleimidophenyl]-butyric aCid-
hydrazide~HC1)
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3. Reactive towards amine (-NHS) and carboxyl (-
COOH) groups
a) Employing Reaction 4
EDC (1-Ethyl-3-[3-dimethylaminopropyl]-
carbodiimide~HCl)
b) Employing Reactions 1 and 4
TFCS (N-[s-Trifluoroacetylcaproyloxy]-succinimide
ester)
4. Reactive towards sulfhydryl (-SH) and hydroxyl
(-OH) groups
~ Employing Reactions 2 and 10
PMPI (N-[p-Maleimidophenyl]-isocyanate)
5. Reactions involve nonselective photoreaction
and specific functional groups
a) Employing Reactions 2 and 7
ANB-NOS (N-5-Azido-2-nitrobenzoyloxysuccinimide)
NHS-ASA (N-Hydroxysuccinimidyl-4-azidosalicylic
acid)
SANPAH (N-Succinimidyl-6-[4'-azido-2'-
nitrophenylamino]-hexanoate)
Sulfo-HSAB (N-Hydroxysulfosuccinimidyl-4-
azidobenzoate)
Sulfo-SANPAH (Sulfosuccinimidyl-6-[4'-azido-2°-
nitrophenylamino]-hexanoate)
b) Employing Reactions 4 and 7
ASBA (4-[p-Azidosalicylamido]-butylamine)
c) Employing Reactions 5 and 7
57
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APDP (N-[4-(p-A~idosalicylamido)-butyl]-3'-(2'-
pyridyldithio)-propionamide)
d) Employing Reactions 9 and 7
ABH (p-Azidobenzoyl Hydrazide)
Reaction Schemes
For the purpose of this example, the terms protein
and peptide are interchangeable.
Reaction 1: NHS-ester
0 0
O R~.~NHa PH > 7 ~ R-CI-N-R + OH-N
H ,
O O
l5
Reaction 2: Maleimide
H
pH > 6.5-7.5
R-N I + R'-SH ~ R-
S-R'
O O
Reaction 3: Active Halogen
O H O H
pH > 7.5
-C- ~ -X + R'-SH ---~ R-C' ~ -S-R' + HX
H H
2 0 where X = I or Br
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Reaction 4: EDC coupling
H
I
O Ri-N-C=N-Rz
N R I H R II
O
H
R,-N- i ~N-R2 II I II
R4NH2
R3 C-O ---~- R3 C-NRQ + /C
II R~HN NHRZ
EDC activates tke carboxyl group of R3COOH and allows it to be coupled to the
amino group of RqNF~.
Reaction 5: Pyridyl disulfide
R-S-S + R~-SH p~ R-g-s-R' +
N N
H~ s
Reaction 6: Ben~oylbenzoyl photolysis
_ o _ o
Protein + ~ ~ CI ~ ~ CI - Peptide
' ( 350 rnn
1
_ OH _ O
CI Pe tide
Protein
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Reaction 7: Azidophenyl photolysis
2G5-275 nm
N3 + Protein ~ - Protein
(300-400 nm
with vitro group)
Reaction 8: Vinyl-sul~one
0
II -R + R~-SH ~ R. S O
O
III R
O
Reaction 9: Hydrazide
I H I H NaI04 II
Protein ~ - ~ -R t Protein C -H
H H
Oxidation of a protein carbohydrate to an aldehyde.
II II I
Protein C-H + N3 ~ C-N-NHz
O H H
N3 C-N-N= ~ Protein
ABH reacts with the aldehyde on the protein to form an arylazide activated
protein.
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SEQUENCE LISTING
<110> Helix BioPharma Corporation
<120> Biosensor Assay Device and Method
<130> 08892610W0
<140> Not Yet Assigned
<141> Filed Herewith
<150> US 09/654,192
<151> 2000-09-01
<160> 6
<170> FastSEQ for Windows Version 4.0
<210> 1
<211> 35
<212> PRT
<213> Artificial Sequence
<220>
<223> coil-forming peptide
<400> 2
Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala Leu Glu Lys Glu Val
1 5 10 15
Ser Ala Leu Glu Lys Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala
20 25 30
Leu Glu Lys
<210> 2
<211> 35
<212> PRT
<213> Artificial Sequence
<220>
<223> coil-forming peptide
<400> 2
Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala Leu Glu Lys Glu Val
Z 5 10 15
Ser Ala Leu Glu Lys Glu Val Ser Ala Leu Glu Lys Glu Val Ser Ala
20 25 30
Leu Glu Lys
1/3
CA 02420743 2003-02-26
WO 02/18952 PCT/CA01/01246
<210> 3
<211> 35
<212> PRT
<213> Artificial Sequence
<220>
<223> coil-forming peptide
<400> 3
Lys Val Ser A1a Leu Lys Glu Lys Val Ser Ala Leu Lys Glu Lys Val
Z 5 10 15
Ser A1a Leu Lys Glu Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala
20 25 30
Leu Lys Glu
<210> 4
<211> 35
<212> PRT
<213> Artificial Sequence
<220>
<223> coil-forming peptide
<400> 4
Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala Leu Lys Glu Lys Val
l 5 10 15
Ser Ala Leu Lys Glu Lys Val Ser Ala Leu Lys Glu Lys Val Ser Ala
20 25 30
Leu Lys Glu
<210> 5
<21l> 35
<212> PRT
<213> Artificial Sequence
<220>
<223> coil-forming peptide
<400> 5
Glu Val Glu Ala Leu GIn Lys Glu Val Ser Ala Leu Glu Lys Glu Val
1 5 10 15
Ser Ala Leu Glu Cys Glu Val Ser Ala Leu Glu Lys Glu Val Glu Ala
20 25 30
Leu Gln Lys
<210> 6
<211> 35
<212> PRT
<213> Artificial Sequence
2/3
CA 02420743 2003-02-26
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<220>
<223> coil-forming peptide
<400> 6
Lys Val Glu Ala Leu Lys Lys Lys Val Ser Ala Leu Lys Glu Lys Val
io i5
Ser Ala Leu Lys Cys Lys Val Ser Ala Leu Lys Glu Lys Val G1u Ala
20 25 30
Leu Lys Lys
3/3
