Note: Descriptions are shown in the official language in which they were submitted.
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SYSTEMS AND METHODS FOR MAKING
NONINVASIVE ASSESSMENTS OF CARDIAC TISSUE AND PARAMETERS
Reference to Priority Application
This application claims priority to U.S. Provisional Application No.
60/393,293 filed
July l, 2002 and U.S. Provisional Application No. 60/475,803 filed June 3,
2003. . , .
Technical Field of the Invention
This invention relates to systems and methods for assessing cardiac tissue and
cardiac
parameters noninvasively using ultrasound techniques.
Background of the Invention
Methods and systems for determining and characterizing various systems and
tissue
properties are known. Characterization of internal tissues using non-invasive
and non-
traumatic techniques is challenging in many areas. Non-invasive detection of
various cancers
remains problematic and unreliable. Similarly, non-invasive assessment and
monitoring of
important internal clinical parameters, such as intracranial pressure and
cardiac output, are
also practical challenges, despite the efforts devoted to developing such
techniques.
Ultrasound imaging is a non-invasive, diagnostic modality that is capable of
providing
information relating to tissue properties. In the field of medical imaging,
ultrasound may be
used in various modes to produce images of objects or structures within a
patient. In a
transmission mode, an ultrasound transmitter is placed on one side of an
object and the sound
is transmitted through the object to an ultrasound receiver. An image may be
produced in
which the brightness of each image pixel is a function of the amplitude of the
ultrasound that
reaches the receiver (attenuation mode), or the brightness of each pixel may
be a function of
the time required for the sound to reach the receiver (time-of flight mode).
Alternatively, if
the receiver is positioned on the same side of the object as the transmitter,
an image may be
produced in which the pixel brightness is a function of the amplitude of
reflected ultrasound
(reflection or backscatter or echo mode). In a Doppler mode of operation, the
tissue (or
object) is imaged by measuring the phase shift of the ultrasound reflected
from the tissue (or
object) back to the receiver.
Ultrasonic transducers for medical applications are constructed from one or
more
piezoelectric elements activated by electrodes. Such piezoelectric elements
may be
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constructed, for example, from lead zirconate titanate (PZT), polyvinylidene
diflouride
(PVDF), PZT ceramic/polymer composites, and the like. The electrodes are
connected to a
voltage source, a voltage waveform is applied, and the piezoelectric elements
change in size
at a frequency corresponding to that of the applied voltage. When a voltage
waveform is
applied, the piezoelectric elements emit an ultrasonic wave into the media to
which it is
coupled at the frequencies contained in the excitation waveform. Conversely,
when an
ultrasonic wave strikes the piezoelectric element, the element produces a
corresponding
voltage across its electrodes. Numerous ultrasonic transducer constructions
are known in the
art.
When used for imaging, ultrasonic transducers are provided with several
piezoelectric
elements arranged in an array and driven by different voltages. By controlling
the phase and
amplitude of the applied voltages, ultrasonic waves combine to produce a net
ultrasonic wave
that travels along a desired beam direction and is focused at a selected point
along the beam.
By controlling the phase and the amplitude of the applied voltages, the focal
point of the
beam can be moved in a plane to scan the subject. Many such ultrasonic imaging
systems are
well known in the art.
An acoustic radiation force is exerted by an acoustic wave on an object in its
path.
The use of acoustic radiation forces produced by an ultrasound transducer has
been proposed
in connection with tissue hardness measurements. See Sugimoto et al., "Tissue
Hardness
Measure Using the Radiation Force of Focused Ultrasound", IEEE Ultrasonics
Symposium,
pp. 1377-80, 1990. This publication describes an experiment in which a pulse
of focused
ultrasonic radiation is applied to deform the object at the focal point of the
transducer. The
deformation is measured using a separate pulse-echo ultrasonic system.
Measurements of
tissue hardness are made based on the amount or rate of object deformation as
the acoustic
force is continuously applied, or by the rate of relaxation of the deformation
after the force is
removed.
Another system is disclosed by T. Sato, et al., "Imaging of Acoustical
Nonlinear
Parameters and Its Medical and Industrial Applications: A Viewpoint as
Generalized
Percussion," Acoustical Imaging, Vo. 20, pg. 9-18, Plenum Press, 1993. In this
system, a
lower frequency wave (350 kHz) is used as a percussion force, and an
ultrasonic wave
(SMHz) is used in a pulse-echo mode to produce an image of the subject. The
percussion
force perturbs second order nonlinear interactions in tissues, which may
reveal more
structural information than conventional ultrasound pulse-echo systems.
2
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Fatemi and Greenleaf reported an imaging technique that uses acoustic emission
to
map the mechanical response of an object to local cyclic radiation forces
produced by
interfering ultrasound beams. The object is probed by arranging the
intersection of two
focused, continuous-wave ultrasound beams of different frequencies at a
selected point on the
object. Interference in the intersection region of the two beams produces
modulation of the
ultrasound energy density, which creates a vibration in the object at the
selected region. The
vibration produces an acoustic field that can be measured. The authors
speculate that
ultrasound-stimulated vibro-acoustic spectrography has potential applications
in the non-
destructive evaluation of materials, and for medical imaging and noninvasive
detection of
hard tissue inclusions, such as the imaging of arteries with calcification,
detection of breast
microcalcifications, visualization of hard tumors, and detection of foreign
objects.
U.S. Patents 5,903,516 and 5,921,928 (Greenleaf et al.) disclose a method and
system
for producing an acoustic radiation force at a target location by directing
multiple high
frequency sound beams to intersect at the desired location. A variable
amplitude radiation
force may be produced using variable, high frequency sound beams, or by
amplitude
modulating a high frequency sound beam at a lower, baseband frequency. The
mechanical
properties of an object, or the presence of an object, may be detected by
analyzing the
acoustic wave that is generated from the object by the applied acoustic
radiation force. An
image of the object may be produced by scanning the object with high frequency
sound
beams and analyzing the acoustic waves generated at each scanned location. The
mechanical
characteristics of an object may also be assessed by detecting the motion
produced at the
intersections of high frequency sound beams and analyzing the motion using
Doppler
ultrasound and nuclear magnetic resonance imaging techniques. Variations in
the
characteristics of fluids (e.g. blood), such as fluid temperature, density and
chemical
composition can also be detected by assessing changes in the amplitude of the
beat frequency
signal. Various applications are cited, including detection of
atherosclerosis, detection of gas
bubbles in fluids, measurement of contrast agent concentration in the blood
stream, object
position measurement, object motion and velocity measurement, and the like. An
imaging
system is also disclosed.
U.S. Patent 6,039,691 (Walker et al.) discloses methods and apparatus for soft
tissue
examination employing an ultrasonic transducer for generating an ultrasound
pulse that
induces physical displacement of viscous or gelatinous biological fluids and
analysis
techniques that determine the magnitude of the displacement. The transducer
receives
ultrasonic echo pulses and generates data signals indicative of the tissue
displacement. This
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apparatus and method is particularly useful for examining the properties of a
subject's
vitreous body, in connection with the evaluation and/or diagnosis of ocular
disorders, such as
vitreous traction.
U.S. Patent 5,086,775 (Parker et al.) describes a system in which a low
frequency
vibration source is used to generate oscillations in an object, and a coherent
or pulsed
ultrasound imaging system is used to detect the spatial distribution of the
vibration amplitude
or speed of the object in real-time. In particular, the reflected Doppler
shifted waveform
generated is used to compute the vibration amplitude and frequency of the
object on a
frequency domain estimator basis, or on a time domain estimator basis.
Applications of this
system include examination of passive structures such as aircraft, ships,
bridge trusses, as
well as soft tissue imaging, such as breast imaging.
Several U.S. Patents to Sarvazyan relate to methods and devices for ultrasonic
elasticity imaging for noninvasively identifying tissue elasticity. Tissue
having different
elasticity properties may be identified, for example, by simultaneously
measuring strain and
stress patterns in the tissue using an ultrasonic imaging system in
combination with a pressure
sensing array. The ultrasonic scanner probe with an attached pressure sensing
array may
exert pressure to deform the tissue and create stress and strain in the
tissue. This system may
be used, for example, to measure mechanical parameters of the prostate. U.S.
Patents to
Sarvazyan also describe shear wave elasticity imaging using a focused
ultrasound transducer
that remotely induces a propagating shear wave in tissue. Shear modulus and
dynamic shear
viscosity at a given site may be determined from the measured values of
velocity and
attenuation of propagating shear waves at that site.
Cardiac Performance
Cardiac output is important to the body for two reasons. The major limitation
in the
delivery of nutrients to the tissues of the body is the delivery of oxygen.
Delivery of
metabolic substrates ("food") and elimination of waste products require less
blood flow than
is necessary for adequate delivery of oxygen for the tissues' metabolic needs.
An inadequate
cardiac output translates into some tissues of the body receiving too little
oxygen and leads to
dysfunction of the affected organ or even tissue damage or cell death of the
deprived tissue.
The "gold standard" for measurement of cardiac output is the pulmonary artery
catheter. It measures cardiac output via the thermodilution technique. It is
effective, and not
difficult to use, but it requires placing the catheter into a vein and
threading the catheter
through the heart and into the lungs. The risks to the patient from using the
pulmonary artery
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catheter preclude routine use. Echocardiography can be used, either
transthoracically or
using esophageal echo. This technique is safer to the patient, but it is
technically more
difficult, less accurate, and impractical to use for longer than a few minutes
at a time. Other
techniques exist, but none have gained universal acceptance. A low risk method
for
measuring either cardiac output, or providing a good estimation of the
components of cardiac
output, would prove invaluable in critical care settings. Such a technique
would likely be
used in far more patients than is the number of patients who currently receive
a pulmonary
artery catheter.
Cardiac output is the product of heart rate and stroke volume (the amount of
blood the
heart pumps to the body in a single beat). Heart rate is easy to determine.
Stroke volume is
difficult to measure directly, so it is generally calculated by measuring or
estimating cardiac
output and then deriving stroke volume = cardiac output = heart rate. The
objective of
providing a non-invasive measurement of cardiac output thus becomes a problem
of how to
measure stroke volume in a non-invasive fashion. Heart rate is also usually
easy to
manipulate. Consequently, the difficult aspect in the clinical manipulation of
cardiac output
is generally reduced to a problem of how to manipulate stroke volume.
Stroke volume is a function of two basic properties of the heart: volume
status and
contractility. Each of these parameters is as important to blood pressure as
vascular
resistance and heart rate. Although the volume status of a patient is
manipulated by
increasing or decreasing the blood volume of the body, what is really
important is the volume
status of the right and left ventricles. The ventricles need to be "filled up"
prior to contraction
for two reasons. First, the ventricles cannot pump to the lungs or body (right
and left
ventricles, respectively) what the ventricles don't have in them at the start
of contraction. The
more blood in the chamber of the ventricle, the more blood could be
potentially pumped out.
Second, as more blood is put in the ventricle, the muscle cells of the heart
become more
stretched. The greater the stretch, the harder the heart muscle contracts at
the next heartbeat.
This phenomenon is known as the length-tension relationship, and is
illustrated in Figure 1.
Stronger contractions permit the heart to pump against a higher blood pressure
and/or pump
out a higher percentage of the blood in the ventricle. Expressed
mathematically, stroke
volume (SV) is equal to the product of end-diastolic volume (EDV, the amount
of blood in
the chamber of the ventricle just before contraction begins) and the ejection
fraction (EF, the
percent of the EDV that is pushed out of the ventricle during heart
contraction). SV = EDV x
EF.
s
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When treating a patient who is thought to have a low stroke volume, a common
clinical maneuver is to administer fluid. In a normal heart, the EF will not
decrease even if
blood pressure increases as a result of the improved stroke volume. However, a
heart with
poorly functioning muscle will have a low EF at baseline and will not
demonstrate much of
an improvement in its contraction when EDV is increased (See Figure 1). In
fact, more
volume may worsen the status of the patient if the heart does not improve its
performance in
response to the volume. If performance does not improve, the heart may become
distended,
which results in impaired function. Furthermore, even if over-distention does
not occur, the
increase in volume increases the filling pressures, that in turn must be
matched by increased
pressures in the atrium and veins. In the case of the right ventricle, high
venous pressures
cause congestion in the abdominal organs and legs that can lead to liver and
intestinal
dysfunction and to peripheral edema. In the case of the left ventricle, high
venous pressures
cause the pressure in the blood vessels in the lungs to increase. If these
pressures get too
high, fluid leaks out into the lungs and causes symptoms of heart failure
(shortness of breath,
inability to lay flat) or even pulmonary edema, a life-threatening event where
the air sacs in
the lung fill with fluid and limit the ability to get oxygen into the blood.
It is therefore important to know when giving a patient more fluid would
produce
these undesirable side effects. Current technology for this determination
largely rests with
the application of the pulmonary artery catheter. The catheter can measure the
pressure in the
atria and thus provide an estimate of the pressure in the ventricular chamber
during diastole
when the heart muscle is relaxed. If these pressures are already high, then
more fluid must be
administered with great care, if at all. Unfortunately, interpretation of
pressures provided by
the pulmonary artery catheter can be difficult, making optimal fluid
management
problematic. The difficulty, in part, is that the relationship between the
filling pressure (end-
diastolic pressure) and volume (end-diastolic volume) is not linear. Figure 2
illustrates this
relationship between end-diastolic pressure and volume for heart tissue that
is stiff and
compliant. A change in pressure of a few mmHg could represent a big or a small
change in
ventricular volume, depending on the character of the heart tissue.
Furthermore, as the
condition of the heart changes, the curve can shift around making it harder to
interpret the
pressure measurements as a measure of end-diastolic volume.
Ideally, clinicians would like to have a direct measure of end-diastolic
volume. An
echocardiogram may provide a volume measurement, but this measurement does not
tell the
clinician whether that volume is too high, too low or just right. Measurement
of ventricular
wall stiffness, if it could be provided, would be helpful because wall
stiffness is directly
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affected by ventricular pressure. In fact, knowledge of a wall stiffness
parameter may be
more useful than knowledge of a pressure parameter because stiffness is also
affected by
ventricular size. Measurement of a ventricular wall stiffness parameter is
likely to be more
effective than measurement of a pressure parameter in determining when fluid
volume
administration will be ineffective or even harmful to a patient.
Ultrasound techniques, such as Doppler tissue imaging modes, have recently
been
proposed for use in the diagnosis of cardiac tissue and function. In general,
these techniques
involve tracking of tissue movement, or velocity. Tissue velocities are used
to derive an
estimate of strain rate, and from strain rate, an estimation of tissue strain
may be derived.
These techniques are dependent on accurate tissue motion estimates, when
tissues are moving
in different directions within a small spatial region.
U.S. Patent 6,527,717 discloses systems and methods for analyzing tissue
motion in
which motion estimates are corrected for transducer motion. Tissue motion may
be used to
determine a strain rate or strain, and motion estimates may be generated using
data acquired
by an intracardiac transducer array.
U.S. Patent 6,099,471 discloses ultrasound techniques for determining strain
velocity
from tissue velocity. Tissue velocity is determined based on measurements of
the pulse-to-
pulse Doppler shift at positions along an ultrasound beam.
U.S. Patent 6,517,485 discloses ultrasound systems and methods for calculating
and
displaying tissue deformation parameters, such as tissue Doppler and strain
rate imaging.
U.S. Patent 6,537,221 relates to strain rate analysis for ultrasound images in
which the spatial
gradient of velocity is calculated in the direction of tissue motion. U.S.
Patents 6,579,240,
discloses ultrasound display of a moving structure, such as a cardiac wall
tissue within a
region of interest, as a color representation.
The accuracy and clinical usefulness of tissue strain predictions based on the
estimation of strain rate from Doppler tissue velocities is problematic.
Existing methods of
measuring ventricular filling and cardiac contractility using infra-arterial
lines or
echocardiograms have limited application because of the risk to the patient,
high expense and
difficulty in interpretation of the information provided. Lack of direct, non-
invasive and
inexpensive methods to measure ventricular filling and cardiac contractility
means that
optimal management of stroke volume is missing from the care of many patients
who would
benefit from such optimization.
Arterial Blood Pressure
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Arterial blood pressure (ABP) is a fundamental objective measure of the state
of an
individual's health. Indeed, it is considered a "vital sign" and is of
critical importance in all
areas of medicine and healthcare. The accurate measure of ABP assists in
determination of
the state of cardiovascular and hemodynamic health in stable, urgent,
emergent, and operative
conditions, indicating appropriate interventions to maximize the health of the
patient.
Currently, ABP is most commonly measured noninvasively using a pneumatic cuff,
often described as pneumatic plethysmography or Korotkoff's method. While this
mode of
measurement is simple and inexpensive to perform, it does not provide the most
accurate
measure of ABP, and it is susceptible to artifacts resulting from the
condition of arterial wall,
the size of the patient, the hemodynamic status of the patient, and autonomic
tone of the
vascular smooth muscle. Additionally, repeated cuff measurements of ABP result
in falsely
elevated readings of ABP, due to vasoconstriction of the arterial wall. To
overcome these
problems, and to provide a continuous measure of ABP, invasive arterial
catheters are used.
While such catheters are very reliable and provide the most accurate measure
of ABP, they
require placement by trained medical personnel, usually physicians, and they
require bulky,
sophisticated, fragile, sterile instrumentation. Additionally, there is a risk
of permanent
arterial injury causing ischemic events when these catheters are placed. As a
result, these
invasive monitors are only used in hospital settings and for patients who are
critically ill or
are undergoing operative procedures. .
U.S. Patent 4,869,261 to Penaz discloses a method for automatic, non-invasive
determination of continuous arterial blood pressure in arteries compressible
from the surface
by first determining a set point with a pressure cuff equipped with a
plethysmographic gauge
of vascular volume and then maintaining the volume of the measured artery
constant to infer
arterial blood pressure. A generator producing pressure vibrations
superimposed on the basic
blood pressure wave, and the changes in the oscillations of the blood pressure
wave are
monitored by an active servo-system that constantly adjusts the cuff pressure
to maintain
constant arterial volume; thus, the frequency of vibration of the blood
pressure wave that is
higher than the highest harmonic component of the blood pressure wave is used
to determine
arterial blood pressure.
U.S. Patent 4,510,940 to Wesseling discloses a method for correcting the cuff
pressure in the indirect, non-invasive and continuous measurement of the blood
pressure in a
part of the body by first determining a set-point using a plethysmograph in a
fluid-filled
pressure cuff wrapped around an extremity and then adjusting a servo-reference
level as a
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function of the shape of the plethysmographic signal, influenced by the
magnitude of the
deviation of the cuff pressure adjusted in both open and closed systems.
U.S. Patent 5,241,964 to McQuilkin discloses a method for a non-invasive, non-
occlusive method and apparatus for continuous determination of arterial blood
pressure using
one or more Doppler sensors positioned over a major artery to determine the
time-varying
arterial resonant frequency and hence blood pressure. Alternative methods
including the
concurrent use of proximal and distal sensors, impedance plethysmography
techniques,
infrared percussion sensors, continuous oscillations in a partially or fully
inflated cuff,
pressure transducers or strain gauge devices applied to the arterial wall,
ultrasonic imaging
techniques which provide the time-varying arterial diameter or other arterial
geometry which
changes proportionately with infra-mural pressure, radio frequency sensors, or
magnetic field
sensors are also described.
U.S. Patent 5,830,131 to Caro et al. discloses a method for determining
physical
conditions of the human arterial system by inducing a well-defined
perturbation (exciter
waveform) of the blood vessel in question and measuring a hemo-parameter
containing a
component of the exciter waveform at a separate site. The exciter consists of
an inflatable
bag that can exert pressure on the blood vessel of interest, and is controlled
by a processor.
Physical properties such as cardiovascular disease, arterial elasticity,
arterial thickness,
arterial wall compliance, and physiological parameters such as blood pressure,
vascular wall
compliance, ventricular contractions, vascular resistance, fluid volume,
cardiac output,
myocardial contractility, etc. are described.
U.S. Patent 4,646,754 to Seale discloses a method for non-invasively inducing
vibrations in a selected element of the human body, including blood vessels,
pulmonary
vessels, and eye globe, and detecting the nature of the responses for
determining mechanical
characteristics of the element. Methods for inducing vibrations include
mechanical drivers,
while methods for measuring responses include ultrasound, optical means, and
visual
changes. Mechanical characteristics include arterial blood pressure, organ
impedance, intra-
ocular pressure, and pulmonary blood pressure.
U.S. Patent 5,485,848 to Jackson et al. discloses a method and apparatus for
non-
invasive, continuous arterial blood pressure determination using a separable,
diagnostically
accurate blood pressure measuring device, such as a conventional pressure
cuff, to initially
calibrate the system and then measuring arterial wall movement caused by blood
flow
through the artery to determine arterial blood pressure. Piezoelectric devices
are used in
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wristband device to convert wall motion signals to an electric form that can
be analyzed to
yield blood pressure.
U.S. Patent 5,749,364 to Sliwa, Jr. et al. discloses a method and apparatus
for the
determination of pressure and tissue properties by utilizing changes in
acoustic behavior of
micro-bubbles in a body fluid, such as blood, to present pressure information.
This invention
is directed at the method of mapping and presenting body fluid pressure
information in at
least two dimensions and to an enhanced method of detecting tumors.
PCT International Patent Publication WO 00/72750 to Yang et al. discloses a
method
and apparatus for the non-invasive, continuous monitoring of arterial blood
pressure using a
finger plethysmograph and an electrical impedance photoplethysmograph to
monitor dynamic
behavior of arterial blood flow. Measured signals from these sensors on an
arterial segment
are integrated to estimate the blood pressure in this segment based on a
hemodynamic model
that takes into account simplified upstream and downstream arterial flows
within this vessel.
A noninvasive, continuous ABP monitor would provide medical personnel with
valuable information on the hemodynamic and cardiovascular status of the
patient in any
setting, including the battlefield, emergency transport, clinic office, and
triage clinics.
Additionally, it would provide clinicians the ability to continuously monitor
the ABP of a
patient in situations where the risks of an invasive catheter are unwarranted
or unacceptable
(e.g., outpatient procedures, ambulance transports, etc.). Thus, the present
invention is
directed to methods and systems for the continuous assessment of ABP using non-
invasive
ultrasound techniques.
Summary of the Invention
The present invention provides methods and systems using the application of
ultrasound for noninvasively assessing, localizing and monitoring cardiac
properties and
parameters, and for diagnosing, localizing and monitoring various conditions,
responses and
disease states. Acoustic properties of tissues, including cardiac tissues, and
tissue
displacement, may be evaluated using the methods and systems described herein,
as well as
the techniques described in PCT International Publication WO 02/43564, which
is
incorporated herein by reference in its entirety.
Acoustic properties of cardiac tissue may be determined, for example, by
collecting
acoustic scatter data using an ultrasound transducer, or transducer array,
aimed at, or having a
focus on or in cardiac tissue. In a "passive" mode embodiment, measurements of
the
"intrinsic" properties of cardiac tissue, in situ, such as tissue stiffness or
tension or strain, etc.,
to
CA 02490999 2004-12-22
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are taken using ultrasound techniques. In another embodiment, focused
ultrasound beams)
are applied to cardiac tissue to deform localized cardiac tissue, and one or
more aspects) of
the deformation, or a biological response produced by the deformation(s), is
assessed and
related to cardiac tissue properties and parameters. The (intrinsic or
induced) acoustic
properties of cardiac tissue, such as (intrinsic or induced) displacements of
target tissue sites,
are related to physical and/or structural tissue properties, such as tissue
stiffness, distension,
tension, strain, strain rate, elasticity, compliance and the like, which are
related to clinically
important cardiac parameters and properties, such as cardiac output.
In another embodiment, an oscillatory radiation force is applied to localized
cardiac
tissue to induce localized tissue oscillations. Acoustic emissions produced by
the oscillating
cardiac tissue, and/or other properties of the oscillating tissue, are related
to the properties of
the cardiac tissue and may be related, according to the present invention, to
specified cardiac
parameters and properties. In yet another embodiment, focused ultrasound
beams) are used
to make local sonoelasticity measurements to assess the properties of cardiac
tissue. For
some applications, observations of changes and trends in the properties of
targeted cardiac
tissue over time are desired, rather than absolute measurements of targeted
cardiac tissue
properties at a given time.
The methods and systems of the present provide important information about the
health and condition of cardiac tissue, such as ventricular wall stiffness. By
the law of
LaPlace, wall stiffness is a function of ventricular chamber volume,
ventricular wall thickness
and the pressure in the ventricular chamber. If the heart muscle is
contracting, then wall
stiffness increases, if for no other reason than the ventricular chamber
pressure increases.
From these first principles a wide variety of useful information can be
extracted from the
measurement of myocardial tissue properties, such as wall stiffness, at
various times
throughout the cardiac cycle.
The cardiac cycle is divided into systole and diastole. During systole the
heart muscle
contracts and blood is ejected: During diastole the muscle relaxes and the
ventricular
chamber fills with blood from the atrium. Figure 3 illustrates the pressure
and volume
relationships of blood in the left and right ventricles during cardiac
cycling. The volume of
blood in the ventricle just before ejection begins is called the end-diastolic
volume (Point A,
Figure 3) and is associated with the end-diastolic pressure (Point B, Figure
3). Ventricular
end-diastolic volume affects both wall thickness (the wall thins as the heart
fills) and end-
diastolic pressure (pressure goes up as volume increases, but in a non-linear
fashion). At
end-diastole, the ventricular muscle should be maximally relaxed, and wall
stiffness is
CA 02490999 2004-12-22
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therefore determined by the intrinsic stiffness of the muscle, ventricular
chamber volume,
wall thickness and end-diastolic pressure. Consequently, ventricular wall
stiffness at the end
of diastole is heavily influenced by end-diastolic volume. Ventricular wall
stiffness is thus a
good parameter, measurable using methods and systems of the present invention,
for
determining end-diastolic volume and pressure.
Determinations of cardiac wall stiffness parameters provide useful information
throughout the cardiac cycle, and not just at end-diastole. Examination of the
pressure and
volume relationships during the cardiac cycle, as shown in Figure 3, reveals
that ventricular
chamber pressure changes continually (in this example, the left ventricle). Of
particular
interest are the periods when the ventricle begins to fill (Point C, when the
atrial pressure
exceeds the ventricular pressure); when the ventricle is rapidly relaxing
(Period D); and when
the ventricle is rapidly developing pressure (Period E). The changes in wall
stiffness during
Period D, along with the wall stiffness at Point B, are useful in the
assessment of ventricular
relaxation and in the diagnosis of diastolic dysfunction. The changes in wall
stiffness during
Period E are useful in the assessment of ventricular contraction
(contractility) and the
diagnosis of systolic dysfunction.
It is important to understand that at the end of ventricular contraction, the
ventricle
has squeezed down on itself and is similar to a compressed spring ready to
recoil open.
"Springing" open is exactly what the ventricle will do if the muscular
contraction relaxes
quickly enough. If allowed to spring open, the ventricle will literally suck
blood into it from
the atrium. This phenomenon results in a rapid transfer of blood into the
ventricle, more so
than for the rest of diastole. Figure 4A illustrates the flow of blood into a
normal ventricle.
The E wave is the initial rapid filling as the ventricle draws in blood.
Thereafter there is
modest filling during the middle of diastole followed by another increase in
filling when the
atrium contracts (A wave) and forces more blood into the ventricle. Fig. 4B
shows the
pattern of filling in an abnormal circumstance. When the ventricular muscle
does not relax
rapidly (diastolic dysfunction), the residual muscle activity present at the
beginning of filling
does not permit the ventricle to spring open and limits the amount of blood
entering the
ventricle in early diastole. The rest of the diastolic period must now make up
for that limited
filling in early diastole - in this circumstance, the A wave is larger than
the E wave. To
accomplish this make-up filling, the atrium must increase in pressure, and
this pressure
increase is transmitted to both the ventricle and the organs upstream of the
atrium (such as the
lungs). If the pressure gets too high, then heart failure symptoms appear,
such as pulmonary
congestion.
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At one time, heart failure was thought to be due exclusively to poor
contractility.
Now it is understood that diastolic dysfunction alone can cause heart failure.
The problem is
that diastolic dysfunction cannot be diagnosed as easily as just described.
For example, old
age causes the abnormal filling pattern shown in Figure 4B to develop.
Furthermore, as atrial
pressure increases, the E wave becomes bigger, thereby preventing the
appearance of a
diminished E wave to diagnose diastolic dysfunction.
The clinical diagnosis of diastolic dysfunction is relatively easy if a
pulmonary artery
catheter is placed. If the patient has normal left ventricular function, yet
the pulmonary artery
catheter reveals a high left atrial pressure, then the diagnosis is confirmed.
However, most
clinicians are unwilling to place a pulmonary artery catheter, so a great deal
of effort has been
made to estimate left atrial pressure using echocardiography. At present, all
the techniques
that have been proposed have met with very limited success. Certainly there is
no consensus
on how to diagnose diastolic dysfunction with echocardiography alone. The
ability to make
determinations of cardiac properties and parameters noninvasively, such as
cardiac tissue
stiffness and contractility, however, makes the diagnosis of diastolic
dysfunction trivially
easy (and non-invasive) because wall stiffness in late diastole reflects left
atrial pressure.
Even more information may be obtained by examining wall stiffness during late
systole. The intrinsic defect in diastolic dysfunction is abnormally slow
relaxation of the
ventricular muscle at the end of the contraction. The portion of the cardiac
cycle between the
end of ejection and the opening of the mitral valve is known as isovolumic
relaxation.
During this time, wall stiffness is directly proportional to the magnitude of
the muscular
contraction. This is because the ventricular chamber pressure is being
generated by the
muscle activity, and the chamber size is not changing, as the ventricle is
neither emptying nor
filling. Therefore the rate at which wall stiffness decreases directly
reflects the rate at which
the muscle relaxes. The presence of slow relaxation would therefore provide
direct evidence
of diastolic dysfunction. Examination of wall stiffness during isovolumic
relaxation and at
the end of diastole should revolutionize the diagnosis of diastolic
dysfunction because the
procedure is simple, non-invasive and provides unambiguous results.
The methods and systems of the present invention provide high time resolution
information on myocardial tension (strain) throughout the cardiac cycle.
Strain
measurements can be further manipulated to yield strain rate, the rate of
change in strain over
time. This approach is fundamentally different from the technologies that use
measurement
of myocardial tissue velocities to predict strain rate and strain. The present
invention
provides a direct determination of tissue strain, so that strain no longer has
to be referenced to
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an arbitrary zero as a consequence of the use of integration to determine
strain from strain
rate. Specifically, methods and systems of the present invention provide
determinations of
myocardial contractility, myocardial strain and strain rate; detection of
myocardial ischemia
and infarction; determination of ventricular filling; and detection of
diastolic dysfunction.
Each of these particular applications is discussed below.
Myocardial Contractility
Classically, myocardial contractility has been defined as either dP/dt, the
rate of
change of intraventricular pressure, or as peak elastance as determined by the
highest value of
the intraventricular pressure - ventricular volume ratio during systole. dP/dt
peaks during
isovolumic contraction and therefore is relatively, but not completely,
uninfluenced by
loading conditions. The major drawback is that measurement of intraventricular
pressure
requires the invasive placement of a catheter into the ventricular chamber.
Peak elastance not
only requires ventricular pressure measurements, but ventricular volume
measurements as
well. Less accurate, but clinically useful estimates of peak elastance have
been achieved with
non-invasive brachial blood pressure measurements and echocardiographic
estimates of
ventricular volume or area.
Several methods have been applied in the laboratory in attempts to quantify
contractility. One method involves placing a catheter in the chamber of the
left ventricle and
measuring how rapidly pressure develops during ventricular contraction. Figure
5 shows a
sample intraventricular pressure tracing (bottom panel) and the rate of change
in pressure
tracing (dP/dT, top panel). The slope of the intraventricular trace equals the
rate of change in
pressure at any given instant. Usually, contractility is considered
proportional to the
maximum rate of change in pressure observed during the contraction (peak value
of the
dP/dT trace). Ventricles with high contractility contract more rapidly and
exhibit a higher
value for dP/dT. The peak rate of pressure development occurs before the
ventricle begins to
eject blood. This means that the volume of the ventricle is not changing and
that wall tension
is strictly proportional to the ventricular chamber pressure. Wall stiffness
directly reflects
both wall tension and chamber pressure. Contractility can therefore be
estimated as the
maximum rate of change in wall stiffness during the onset of contraction.
Peak systolic strain rate correlates with peak dP/dt and with peak elastance.
See, e.g.,
Pislaru C, Abraham TP, Belohlavek M: Strain and strain rate echocardiography.
Curr Opin
Cardiol 2002; 17:443-454; and Weidemann F, Jamal F, Sutherland GR et al.:
Myocardial
function defined by strain rate and strain during alterations in inotropic
states and heart rate,
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Am J Physiol Heart Circ Physiol 2002;283:H792-H799. The maximum rate of tissue
acceleration (rate of velocity increase) also correlates with dP/dt and
elastance when the heart
is subjected to positive or negative inotropic agents. See, e.g., Vogel M,
Cheung MMH, Li J
et al.: Noninvasive assessment of left ventricular force-frequency
relationships using tissue
doppler-derived isovolumic acceleration, Circulation 2003; 107:1647-1652.
Peak strain rate may thus provide the best clinical estimation of myocardial
contractility, particularly if strain rates can be measured in a completely
non-invasive fashion.
Currently, measurement of strain and strain rates require a high quality
echocardiogram
machine, or rely on predictions made from tissue velocity measurements.
Predictions of
strain based on tissue velocity measurements, though they can be made using
non-invasive
ultrasound techniques, are not consistently accurate. Myocardial velocity
measurements,
furthermore, conventionally relate to net, or bulk, tissue movement.
Continuous
measurement of contractility over a prolonged period of time using
echocardiogram
techniques is not practical or cost effective; especially in an intensive care
unit or operating
room, where it may be desirable to monitor many patients simultaneously.
Methods and systems of the present invention provide determinations of strain
rate as the rate
of change of strain, measured directly using non-invasive ultrasound
techniques over time.
Strain rate is measured, not as bulk movement of myocardial tissue but,
rather, as relative
movements of selected target sites within myocardial tissue. And, because the
time of peak
strain rate is evanescent, improvements in accuracy of peak strain
measurements provided
using methods and systems of the present invention, reduce the amount of
necessary time
averaging of the signal, and improve the cycling rate of the measurement. Both
passive and
active modes of the present invention may be implemented to determine strain
in myocardial
tissue. Moreover, the improved accuracy, non-invasiveness and cost-effective
attributes of
methods and systems of the present invention permit use of strain and strain
rate
measurements for monitoring myocardial contractility and tissue properties, as
well as
diagnosis of myocardial dysfunction.
Myocardial Ischemia and Infarction
Tissue Doppler ultrasound techniques have been used to detect myocardial
ischemia,
primarily in experimental situations that involve severe ischemia and
consequent impairment
of systolic dysfunction. Strain rate patterns change dramatically with the
onset of ischemia,
characterized by a delayed onset of (contraction) strain rate, decreased peak
systolic strain
is
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rate and strain, post-systolic shortening, and decreased peak strain rate
during early
ventricular filling. See, e.g., Pislaru C, Anagnostopoulos PD, Seward JB et
al.: Higher
myocardial strain rates during isovolumic relaxation phase than during
ejection characterize
acutely ischemic myocardium, J Am Coll Cardiol 2002; 40:1487-1494. The
magnitude of
infarction can be determined as well when the myocardium is exposed to
dobutamine.
Transmural infarction (total infarction) is identified by an almost complete
absence of strain
rate or integrated strain over the cardiac cycle, whereas incomplete
infarction demonstrates
reduced strain and strain rate at rest, and progressive post-systolic
increases in strain (post-
systolic shortening) in response to dobutamine. See, e.g., Weidemann F, Dommke
C, Bijnens
B et al.: Defining the transmurality of a chronic myocardial infarction by
ultrasonic strain-
rate imaging, Circulation 2003;107:883-888. The use of dobutamine stress, in
combination
with strain rate analysis, appears to be the best method to assess how much
myocardial tissue
remains alive after a myocardial infarction. See, e.g., Hoffinann R, Altiok E,
Nowak B et al:
Strain rate measurement by doppler echocardiography allows improved assessment
of
myocardial viability in patients with depressed left ventricular function. J
Am Coll Cardiol
2002; 39:443-449. Using PET scanning to define the degree of viable tissue in
areas of
myocardium that had suffered infarction, viable tissue demonstrated increases
in strain rate to
dobutamine whereas non-viable tissue did not. Prediction of viability from
strain rate was
better than standard 2-D echo analysis of wall motion and better than
examination of tissue
velocities alone.
Although determinations of strain rate may be made adequately using the
current
methods of tissue Doppler, the clinical detection of myocardial ischemia is
suboptimal at
present. In emergency rooms and intensive care units, ischemia is often not
detected unless
the patient complains of chest pain, or ECG changes happen to be noted. In the
operating
room, ECG detection of myocardial ischemia is really the only option, since
most patients are
asleep. Unfortunately, ECG changes are often a late event in ischemia. 2-D
echocardiography can detect ischemia via decreases in regional systolic wall
motion, but the
cost (of the machine and the operator) limits the extent to which this method
can be used in
routine clinical settings.
Methods and systems of the present invention that provide direct measurement
of
tissue properties, such as stiffness, tension, strain, etc., using non-
invasive ultrasound
techniques, are well suited for early detection of myocardial ischemia and
infarction.
Myocardial tissue properties determined using acoustic techniques may be used,
for example,
to monitor diastolic relaxation, which is often the first clinical indication
of cardiac ischemia.
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Ventricular Filling and Atrial Pressures
The degree of ventricular filling has important ramifications for management
of the
heart and heart function in virtually all critical care situations, including
the intensive care
unit and the operating room. There are two clinical issues that must be dealt
with: the
amount of blood in the ventricle (end-diastolic volume); and the pressure in
the left atrium.
These two are related, as ultimately the pressure in the atrium is responsible
for pushing
blood into the ventricle. However, the relationship is curvilinear and may
shift (for the same
ventricular blood volume) to higher or lower filling pressures, depending on
the stiffness of
the myocardium that is, in turn, influenced by many factors including tissue
injury and
diastolic function.
The blood volume of the ventricle is important, because the heart cannot pump
what it
does not have. Without blood entering the ventricle, there is no cardiac
output and no blood
pressure. Furthermore, the strength of the ventricular contraction is in part
dependent on the
stretch of the myocytes at the initiation of contraction. Greater stretch,
produced by greater
blood volume, generally increases the strength of the contraction. Volume can
be estimated
non-invasively by 2-D echocardiography, but the cost of the equipment makes it
difficult to
obtain multiple measurements over the course of a day, let alone to monitor
blood volume
continuously. Currently, central venous or pulmonary wedge pressure is often
used to
estimate ventricular end-diastolic pressure, but the interpretation of the
value is problematic
due to the curvilinear relationship between pressure and volume. This problem
is particular
true in the early part of the curve, where large changes in ventricular volume
may have only
small effects on end-diastolic ventricular and atrial pressure.
Using methods and systems of the present invention involving observation of
the
intrinsic and/or induced acoustic properties of myocardial tissue, such as
stiffness, tension,
etc., are measured to determine ventricular filling and/or volume. As the
ventricle fills,
radius increases and wall thickness decreases. Therefore, even if end-
diastolic pressure
changes minimally, increased volume results in increased tension. In fact, the
myocardial
tension changes more than pressure as the ventricle expands thereby making
tension a better
measure of volume status than pressure alone. Techniques that predict tissue
strain rate
and/or strain based on tissue velocity determinations are generally not
suitable for making
ventricular filling and/or volume predictions, because they don't directly
determine a zero
tension point at the beginning of diastole. The ability to measure absolute
myocardial strain,
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using methods and systems of the present invention, permits the utilization of
myocardial
strain as an index of ventricular volume.
Even if atrial pressure is not used to estimate ventricular volume, atrial
pressure is still
an important clinical parameter. Whatever the atrial pressure, it must be
exceeded by the
veins taking blood to the atrium. If the back-pressure gets too high, then
fluid leaks out of the
upstream veins and capillaries and can lead to clinical problems such as
anasarca, liver
dysfunction and pulmonary congestion or edema, all of which can be life-
threatening. Thus,
when a clinician attempts to optimize ventricular filling, the clinician must
also be cognizant
of the impact higher atrial pressures might have on the body. As direct
measurement of
central venous or pulmonary artery wedge pressures requires an invasive
catheter, attempts
have been made to estimate wedge pressure with non-invasive echocardiographic
techniques.
The technique that utilizes tissue Doppler involves calculating the E/Ea
ratio, where E is the
peak blood inflow rate across the mitral valve in early diastole and Ea is the
peak tissue
velocity in early diastole as measured at the mural annulus (Sengupta et al,
2002). It is
believed that wall tension measurements and their rate of change may prove. as
useful as Ea
or even the E/Ea ratio.
Diastolic Dysfunction
Diastolic dysfunction involves slowed and even incomplete relaxation of the
ventricle
during diastole. The functional implication is that if the ventricle remains
stiff, particularly in
early diastole when it is supposed to be receiving rapid inflow of blood from
the atrium, then
either too little blood will enter the ventricle, or the pressure in the
atrium will have to
increase to force the blood into the ventricle. If the atrial pressures
increase to unacceptably
high values, then signs and symptoms of fluid overload develop. In patients
with known
diastolic dysfunction but normal systolic function, alterations are observed
in diastolic tissue
velocities. The ratio of myocardial velocity at the annulus in early diastole
(Em) to that in
late diastole (Am) has been applied in much the same manner that the E/A ratio
of blood flow
across the mitral valve has been used to detect diastolic dysfunction (Isaaz,
2002). There is
evidence that Em itself reflects diastolic relaxation and is not affected by
the atrial pressure,
making it more useful than the E/A ratio. When peak strain rates in diastole
are measured,
they are reduced in early diastole (Stoylen, 2001 ).
Although these velocity measurements show promise, they are still relatively
indirect
measurements of how fast the ventricle is relaxing during the isovolumic
relaxation phase of
the cardiac cycle and early diastolic filling. The "gold standard" for
assessment of diastolic
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relaxation is the rate of decay of ventricular pressure during isovolumic
relaxation (expressed
as the time constant, tau, of ventricular pressure relaxation) (Mandinov,
2000). Of course,
this measurement is highly invasive as it requires a catheter in the
ventricular chamber. The
rate of decrease of myocardial wall tension during isovolumic relaxation
should mimic the
decay of ventricular chamber pressure. Myocardial velocity in early diastolic
filling (Em)
correlates with tau and, like tau, appears to be little affected by loading
conditions (atrial
pressure, aortic pressure) (Waggoner, 2001). Therefore strain rate in early
diastolic filling
may not be affected by loading conditions, and so prove to be a useful measure
of diastolic
relaxation, too. Furthermore, the time trace of absolute tension near the
tension nadir may
reflect how quickly the myocardium relaxes.
The concept that the tension wave can be accurately determined over time,
especially
during isovolumic relaxation and early diastole, has the potential to
characterize uniquely a
wide range of disorders. Constrictive and restrictive pericariditis and
cardiomyopathies have
distinctive patterns of pressure in the ventricular chamber. If tension
follows the same
pattern as pressure, then tension measurements could easily accomplish what
currently
requires invasive measurement. Graded myocardial ischemia, as opposed to
abrupt total
occlusion of coronary artery blood flow, may first present as diastolic
dysfunction and
therefore precede ECG changes and even changes in systolic function. Thus, if
a system can
monitor both diastolic and systolic function, that system has the most chance
of detecting
ischemia early and provide the physician a greater opportunity for
intervention before the
condition worsens. As with all applications, assessment of diastolic function
may be
accomplished using either the passive or active ultrasound modes of the
present invention, or
both modes simultaneously or alternately.
Acoustic detection techniques that involve the application of acoustic
interrogation
signals to a target tissue site and acquisition of acoustic scatter data are
preferred, but
alternative detection techniques, including near-infrared spectroscopy (NIBS),
optical
coherence tomography (OCT), magnetic resonance techniques, positron-emission
tomography (PET), acoustic hydrophones and the like, may be used. A portable,
relatively
low-cost magnetic resonance scanner is described, for example, in the
California Institute of
Technology Engineering and Science publication, Vol. LXIV, No. 2, 2001. The
use of these
techniques to measure various spatial and temporal aspects of tissue
deformation and
associated biological responses is generally known.
Ultrasound sources and detectors may be employed in a transmission mode, or in
a
variety of reflection, palpation or scatter modes, including modes that
examine the
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transference of pressure waves into shear waves, and vice versa. Ultrasound
detection
techniques may also be used to monitor the acoustic emissions) from insonified
tissue.
Detection techniques involving measurement of changes in acoustic scatter,
particularly
backscatter, or changes in acoustic emission, are particularly preferred for
use in methods and
systems of the present invention operating in either the passive or active
modes, or in both
modes simultaneously or alternately. Exemplary acoustic scatter or emission
data that are
related to tissue properties include: changes in scatter or acoustic emission,
including changes
in the amplitude of acoustic signals, changes in phase of acoustic signals,
changes in
frequency of acoustic signals, changes in length of scattered or emitted
signals relative to the
interrogation signal, changes in the primary and/or other maxima and/or minima
amplitudes
of an acoustic signal within a cardiac and/or respiratory cycle; the ratio of
the maximum
and/or minimum amplitude to that of the mean or variance or distribution of
subsequent
oscillations within a cardiac cycle, changes in temporal or spatial variance
of scattered or
emitted signals at different times in the same location and/or at the same
time in different
locations, all possible rates of change of endogenous tissue displacement or
relaxation, such
as the velocity or acceleration of displacement, and the like. Multiple
acoustic interrogation
signals may be employed, at the same or different frequencies, pulse lengths,
pulse repetition
frequencies, intensities, and the multiple interrogation signals may be sent
from the same
location or multiple locations simultaneously and/or sequentially. Scatter or
emission from
single or multiple interrogation signals may be detected at single or at
multiple frequencies, at
single or multiple times, and at single or multiple locations.
Acoustic properties of scatter and/or emission data from selected target
tissue site(s),
or derivative determinations such as tissue displacement, tissue stiffness,
and the like, are
related, using empirical formulations and/or mathematical models, to tissue
properties and/or
clinical parameters. The relation of acoustic properties may be used in
combination with
other parameters, such as blood pressure, to assess tissue properties and/or
clinical
parameters. In one example, declining blood pressure during surgical
procedures may
indicate either diminished or elevated fluid volumes. Blood pressure may be
monitored
concomitantly with the acoustic properties of targeted cardiac tissue to
determine whether
declining blood pressure is a result of diminished or elevated fluid volumes.
In general,
increases in cardiac wall stiffness provide evidence of elevated fluid
volumes, while
reductions in cardiac tissue stiffness provide evidence of reduced fluid
volumes.
Single or multiple interrogation signals administered from different places
and/or at
different times may insonify single or multiple target tissue sites. Intrinsic
and/or induced
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acoustic properties of the insonated target tissue may be assessed, by
acquiring scatter or
emission data, simultaneously and/or sequentially. One of the advantages of
the methods and
systems of the present invention is that target tissue sites may be
volumetrically small, and
spatially resolved, to provide data from localized tissue sites with a high
degree of spatial
resolution. In this way, localized differences in tissue properties may be
identified and
associated with a spatial location within the interrogated tissue. According
to one
embodiment, tissue sites of varying size and/or location are assessed
simultaneously or
sequentially. For most applications, the use of acoustic sources) and/or
transducers)
capable of interrogating and detecting target tissue sites having a volume of
from 1 mm3 to
100 cm3 are suitable.
For assessment and/or monitoring of cardiac tissue properties based on the
acoustic
properties of tissue, the target tissue site is preferably at a selected site
within or on a surface
of cardiac tissue. For many applications, the ventricle or atrium walls are
targeted; for some
applications, for example, the right ventricular wall is targeted. Assessment
of cardiac tissue
properties based on their intrinsic and/or induced acoustic properties may be
supplemented
with data relating to mean and/or continuous arterial blood pressure, cardiac
cycle
information, heart rate, and the like.
Determinations of mean and/or continuous arterial blood pressure may be made,
using ultrasound according to methods and systems of the present invention, in
parallel with
determinations of cardiac tissue properties and parameters. Blood pressure
determinations
may be made, for example, by selecting a target tissue site within or on or in
proximity to a
blood vessel and, preferably, in proximity to cardiac tissue. In this way, a
single, integrated
acoustic system may be used for making determinations of mean and/or
continuous arterial
blood pressure in parallel with determinations of cardiac tissue properties
and parameters.
In yet another aspect, noninvasive systems and methods of the present
invention
provide a measure of arterial or venous blood pressure using acoustic
techniques to measure
alternating compression and dilation of the cross-section or other geometric
or material
properties of an artery or vein, using empirically established relationships
and/or
mathematical models. In another aspect, blood pressure is determined using
acoustic
techniques to measure alternating compression and dilation of tissue
surrounding blood
vessels that is displaced as the vessels are compressed and dilated with the
cardiac cycle.
Geometrical properties that may be determined using acoustic detection
techniques include
changes in diameter, cross-sectional area, aspect ratio, rates of changes in
diameter, velocity,
and the like. Material properties that may be determined using acoustic
detection techniques
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include the stiffness of vessel walls or tissue in proximity to vessel walls.
Blood pressure
may be assessed, for example, by acquiring acoustic data, in an active and/or
passive mode,
from target tissue sites at or in proximity to one or more blood vessels. The
acoustic data can
be related to the stiffness of vessel walls or supporting tissue, which can be
related to blood
pressure. Suitable target tissue sites for determination of arterial or venous
blood pressure
may comprise any blood vessel or surrounding tissue. Detection of ultrasound
scatter data
may be related, for example, with synchronous Doppler flow measurements within
the same
vessel.
A calibration step using a measure of blood pressure taken with a conventional
blood
pressure device, may be incorporated in the blood pressure determination.
Acoustic proxies
for the pulsatility of the blood vessel - such as oscillation rate of the
blood vessel wall - may
be substituted for direct measures of those quantities. In this method, the
spontaneous
changes in the diameter (or other geometric property) of the vessel being
monitored are
assessed using ultrasound, and this information is related (e.g., using
correlation techniques)
to synchronous Doppler flow measurements within the same vessel. Since the
diameter (or
other geometric property) of the vessel is a function of the pressure being
exerted against the
wall of the vessel by blood, and since the velocity of blood flow is dependent
on the diameter
(or radius) of the vessel through which the blood travels, blood pressure can
be calculated
from flow velocity measured by Doppler. By simultaneously measuring the
pulsatility of the
blood vessel of interest and the Doppler flow velocity proximal and distal to
this site,
continuous blood pressure can be determined.
In one embodiment, described in detail below, an acoustic detector, such as an
ultrasound transducer, detects ultrasound signals that are indicative of
tissue displacements,
or associated biological responses, in one or more of the following operating
modes:
transmission, reflection, scatter, emission, backscatter, echo, Doppler, color
Doppler,
harmonic, subharmonic or superharmonic imaging, a-mode, m-mode, or b-mode.
Ultrasonic
interrogation pulses having a known frequency, intensity and pulse repetition
rate are
administered to a desired target tissue site. The intensity, frequency and
pulse repetition rates
of the ultrasonic interrogation pulses are selected such that the
interrogation pulses do not
produce undesired side effects, and do not substantially interfere with
intrinsic tissue
displacements resulting, for example, from blood flow and respiration.
Transmitted signals,
signal reflections, acoustic emissions, scatter such as backscatter, and/or
echoes of the
interrogation pulses are detected and used to assess intrinsic tissue
displacements and/or
tissue properties at the target tissue site. In preferred embodiments of the
passive assessment
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mode, an acoustic detector is implemented to detect the backscatter of
administered
interrogation signals. An acoustic detector may additionally or alternatively
be operated in a
Doppler mode to measure the phase shift of ultrasound reflected back to the
detector.
A variety of techniques may be used to analyze the acquired acoustic data
relating to
intrinsic and/or induced cardiac tissue displacement or associated biological
responses. For
example, analytical techniques developed and employed in connection with
ultrasound
imaging, such as cross-correlation, auto-correlation, wavelet analysis,
Fourier analysis, CW
Doppler, sum absolute difference, and the like, may be employed to determine
various
properties of tissue deformation, and to relate tissue deformation to tissue
properties. Other
empirical techniques and systems, such as artificial neural networks (ANNs),
linear filters
(including those with both infinite impulse response IIR and finite impulse
response FIR
properties), Hidden Markov Models (HMMs), heuristics and fuzzy logic systems,
may be
used to relate one or more variables, such as tissue deformation,
displacement, ABP, etc., to
desired cardiac tissue properties and cardiac parameters. False peak
correction techniques
may be used to improve the accuracy of the assessment. Additionally,
properties of the major
and minor endogenous oscillations of cardiac tissue within a cardiac cycle, or
relationships
between major and minor endogenous oscillations within a cardiac cycle, or
across several
respiratory cycles, may be empirically related to cardiac tissue properties
and conditions.
These determinations may be made with, or without, additional information
relating to ABP
and/or respiration and/or exogenous tissue displacements. In one embodiment,
parameters
such as ABP are measured using other techniques, and one or more externally
measured
parameters are used for calibrating determinations made by systems of the
present invention.
Methods and systems of the present invention are preferably integrated with
control
and data storage and manipulation features similar to the control and data
storage and
manipulation features provided on other types of diagnostic and monitoring
systems. Various
types of control features, data storage features, data processing features,
data output features,
and the like, are well known in the art and may be adapted for use with the
present invention.
Various modes of operation of methods and systems of the present invention are
described below and in the description of preferred embodiments.
First "Active" Acoustic Probin oar Palpation Mode
In a first "active" mode, methods and systems of the present invention
stimulate or
probe target cardiac tissue, or induce a response at a target cardiac tissue
site, by application
of focused ultrasound. The response of the targeted tissue to the application
of focused
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ultrasound may be deformation or displacement (a change in relative position),
a change in
temperature, a change in blood flow, or another detectable response. For
example,
application of an acoustic radiation force to "palpate" a target cardiac
tissue site may be
accomplished by administering one or more acoustic signals. Non-invasive
techniques, such
as ultrasound, optical techniques such as near infrared spectroscopy and
optical coherence
tomography, and other techniques, including magnetic resonance techniques,
external
electrophysiological stimulation, patient response, and the like are used to
assess at least one
response to the application of focused ultrasound. A visualization or imaging
technique, such
as ultrasound imaging or magnetic resonance imaging, may also be employed to
assist in
targeting the focused ultrasound pulses) and to assist in differentially
localizing responsive
tissues.
Acoustic techniques, such as ultrasound, may be used to induce biological
responses
in tissue and to deflect or deform biological materials. Biological materials
absorb some of
the ultrasound as it propagates into and through the material. See, e.g.,
Rudenko et al.
( 1996), "Acoustic radiation force and streaming induced by focused nonlinear
ultrasound in a
dissipative medium," J. Acoust. Soc. Am 99(5) 2791-2798. Also, at the
boundaries between
different tissue types, there is an 'impedance mismatch' (that is, differences
between the
product of density and speed of sound from one tissue to another) that allows
ultrasound to
push on the interface. See, e.g., Chu and Apfel (1982) "Acoustic radiation
pressure produced
by a beam of sound," J. Acoust. Soc. Am 72(6), 1673-1687.
For assessment of cardiac tissue and assessment of cardiac parameters, for
example,
one or more acoustic transducers) is placed in contact with or in proximity to
a subject's
chest. An initial environmental assessment, described below and preferably
employing
ultrasound techniques, may be made, if desired, to assess the characteristics
of the
environment between the acoustic source and the target tissue site, so that
the magnitude of
the acoustic force applied to the target tissue may be determined.
Environmental factors,
such as the distance between the acoustic transducer and various structural
landmarks, may
be determined. An initial environmental assessment may be determinative of
various method
and system parameters. Environmental assessments may additionally be updated
at intervals
throughout a diagnostic or monitoring procedure.
Following the (optional) environmental assessment, an acoustic force is
applied by an
acoustic transducer, at a predetermined frequency, to displace targeted
cardiac tissue at a
targeted location. The deformation may be produced at any desired location
within cardiac
tissue, depending on the focus (foci) of the ultrasonic transducers) producing
the acoustic
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radiation force. In some systems, variable foci ultrasonic transducers are
provided, and a
diagnostic procedure is carned out using a plurality of target tissue sites.
According to one
embodiment for assessment of cardiac output, the focus (foci) of the
ultrasonic transducers)
is preferably provided in proximity to the surface or a small distance below
the surface of a
ventricle wall, to maximize the tissue displacement induced by the radiation
pressure that
arises from the impedance mismatch between cardiac tissue and fluid.
The applied acoustic radiation force is sufficient to induce a detectable
displacement
in the cardiac tissue, or the applied ultrasound beam is sufficient to produce
a detectable
biological response, without producing any medically undesirable changes in
the examined
tissue. For example, the acoustic radiation force applied must not produce
shear in tissues in
proximity to the target tissue of a magnitude sufficient to tear or damage
tissue. The applied
ultrasound, moreover, must not appreciably increase the temperature of
examined tissue to
the point of causing unacceptable damage, and it must not induce extensive or
damaging
cavitation or other produce other deleterious mechanical effects in the
examined tissue.
Suitable ultrasound dosages may be determined using well known techniques. For
example,
Fry et al. studied the threshold ultrasonic dosages causing structural changes
in mammalian
brain tissue and illustrate, in their Fig. 1, the acoustic intensity v. single-
pulse time duration
producing threshold lesions in white matter of the mammalian (cat) brain. Fry
et al.,
Threshold Ultrasonic Dosages for Structural Changes in the Mammalian Brain,
The Journal
of the Acoustical Society of America, Vol. 48, No. 6 (Part 2), p. 1413-1417
(1970). One of
ordinary skill in the art may routinely determine safe ultrasonic dosages for
application to
cardiac tissue.
Additionally, the acoustic frequency must be low enough to penetrate the
tissues
between the skin surface and the cardiac tissue, and high enough to produce
measurable
deformation in the target tissue at the location of interest. Within the
parameters outlined
above, higher frequency acoustic waves are more easily focused and, therefore,
are preferred.
The intensity must be high enough to deform the tissue, but not be so great as
to induce
undesirable changes in the examined tissue. The pulse length is preferably
relatively short,
but long enough to create a measurable deformation or oscillation of the
target tissue, as
desired, while the pulse repetition frequency must be large enough to resolve
medically
interesting temporal features in the tissue, without inducing medically
unacceptable changes
in the tissue.
In general, at least one acoustic property related to tissue displacement, or
an
associated biological response, is determined and related to a tissue property
and, ultimately,
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to a clinically important parameter. For example, the magnitude, or amplitude,
of the
displacement induced by the known acoustic force is directly related to the
elasticity (or
stiffness or compliance, e.g., Young's modulus) of the cardiac tissue, and can
therefore be
empirically related to clinically relevant cardiac parameters, such as cardiac
output.
Additional properties of the target tissue displacement that may be determined
and related to
tissue properties include: various components of amplitude, such as maximum
amplitude in
the direction of the acoustic force or maximum amplitude perpendicular to the
direction of
acoustic force; all possible rates of change of the displacement or subsequent
relaxation of the
tissue, such as the velocity or acceleration of displacement or relaxation;
the amplitude or
rates of change of various components of the shape of the displacement;
changes in Fourier
or wavelett representations of the acoustic scatter signal associated with the
displacement;
properties of shear waves generated by the acoustic radiation force;
properties of induced
second harmonic deformation(s), and the like. Time displacements of pulse
echoes returning
from the target tissue are also indicative of the displacement amplitude and
may be
determined. These properties are all referred to as measures of
"displacement."
Second "Active" Acoustic Probin og~r Palpation Mode
In a second "active" mode of operation, application of focused ultrasound
produces
oscillation of targeted tissue, and data relating to the acoustic signals
emitted from the
targeted tissue are collected. These signals are referred to herein as
acoustic emissions. In
general, methods and systems of the present invention that relate to
application of focused
ultrasound may be used to produce oscillation of targeted tissue, and emitted
acoustic signals
are related to tissue properties and physiological conditions.
In one embodiment, methods and systems of the present invention employ a
confocal
acoustic system comprising at least two acoustic transducers, driven at
different frequencies,
or a focal acoustic system comprising a single acoustic transducer driven at a
given pulse
repetition frequency (PRF), to induce an oscillatory radiation force in the
target tissue, such
as cardiac tissue. The resulting oscillation is at a frequency that is the
difference of the
applied frequencies, at the target location that is marked by the overlap of
the two confocal
acoustic beams or, for the single transducer case, at the PRF. During and
after the application
of focused ultrasound, the targeted tissue emits acoustic signals related to
its intrinsic
properties. The second, active mode of operation may therefore be used to
characterize
tissue. Diagnostic ultrasound techniques may be used to measure the frequency
or other
properties of the emitted acoustic signal, which are empirically related to
tissue properties.
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"Passive" Acoustic Mode
In a "passive" acoustic mode, methods and systems of the present invention
employ
acoustic techniques, such as ultrasound, to acquire data relating to intrinsic
(endogenous)
tissue displacements. Ultrasound backscatter and/or emission data, for
example, are related
to intrinsic tissue displacements, which can be related to various tissue
properties.
Supplemental data, such as measures of mean and/or continuous arterial blood
pressure,
blood flow, and the like, may additionally be used in these determinations.
For example, the magnitude or amplitude or phase of acoustic scatter from
target
cardiac tissue sites undergoing intrinsic displacements during the course of
the cardiac cycle,
is directly related to the stiffness, e.g. Young's modulus, of the cardiac
tissue. Alternatively
or additionally, relationships between the major and minor intrinsic
oscillations of cardiac
tissue within a cardiac cycle, or within a cardiac cycle as modulated by one
or more
respiratory cycles, are empirically related to tissue properties. Properties
of the intrinsic
tissue displacement that may be assessed and related to tissue properties
include: various
components of amplitude, such as maximum amplitude within a cardiac cycle, the
ratio of the
maximum amplitude to that of the mean or variance of subsequent oscillations
within a
cardiac cycle, all possible rates of change of intrinsic cardiac tissue
displacement or
relaxation, such as the velocity or acceleration of displacement, and the
like. Additional data,
such as ABP measurements and/or respiration data, may be collected and used,
with the
acoustic data, to make various assessments and clinical determinations.
Relative trend determinations of the target cardiac tissue properties, such as
stiffness,
contractility, tension, strain and the like, at or near the relevant portions
of the heart (e.g.
ventricle walls and/or atrium walls) are made during certain portions of the
cardiac cycle, and
may be synchronized with EKG measurements. In general, the right ventricle is
relatively
easy to image with ultrasound. We discuss assessment of cardiac parameters
using the
physical properties (e.g., tension) in the right ventricle wall as exemplary,
though other
cardiac target sites may be used. For some embodiments, data may be collected
over many
cardiac cycles, in some embodiments starting when the patient's ventricle wall
tension is
known to be normal, such as before or early in the time course of surgery, and
continued until
the patient is stabilized. In one embodiment, a system of the present
invention comprises an
inexpensive transducer with its own power supply, controller and display unit,
designed to fit
onto standard cardiac diagnostic ultrasound scan heads and interface
electronically with
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standard diagnostic ultrasound machines. In another embodiment, one or more
transducer
arrays) are used for interrogation of and acquisition of acoustic data.
Brief Descriution of the Figures
Fig. 1 shows the relationship between stroke volume and end-diastolic volume
for
normal cardiac tissue, as well as cardiac tissue that has high and poor
contractility.
Fig. 2 shows the relationship between end-diastolic pressure and end-diastolic
volume
for stiff and compliant cardiac tissue.
Fig. 3 shows the pressure and volume relationships during the cardiac cycle.
Fig. 4A shows a normal ventricular filling profile, expressed in terms of
volume over
time, during a cardiac cycle.
Fig. 4B illustrates abnormal ventricular filling profile, expressed in terms
of volume
over time, during a cardiac cycle.
Fig. 5 shows a sample intraventricular pressure tracing (bottom panel) and the
rate of
change in pressure tracing (top panel).
Fig. 6 is a schematic diagram illustrating a system of the present invention
for
inducing and detecting tissue deformation for assessing cardiac tissue
properties.
Fig. 7 is a schematic diagram illustrating another system of the present
invention for
inducing and detecting tissue deformation for assessing cardiac tissue
properties.
Fig. 8 is a schematic cross-sectional diagram illustrating the use of confocal
acoustic
sources to produce tissue displacement and a diagnostic ultrasound probe to
measure the
amplitude of the displacement.
Fig. 9 shows a schematic illustration of a single cMCTT array transducer cell
structure.
Fig. l0A shows a plot demonstrating measured displacement of in vitro beef
brain as
a function of increasing simulated ICP and as a consequence to increasing
brain CSF volume.
Fig. lOB shows a backscatter trace of human brain, in vivo, while the subject
was
holding his breath.
Fig. lOC shows the displacement of human brain, in vivo, while the subject was
holding his breath.
Fig. l OD shows the displacement of human brain, in vivo, while the subject
first held
his breath and then inhaled.
Fig. 11 illustrates experimental results showing that the measured
displacement of
brain tissue, in vivo, is proportional to the acoustic radiation force
applied, as indicated by the
acoustic driving voltage.
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Detailed Descriution of the Invention
While the methods and systems of the present invention may be embodied in a
variety
of different forms, the specific embodiments shown in the figures and
described herein are
presented with the understanding that the present disclosure is to be
considered exemplary of
the principles of the invention, and is not intended to limit the invention to
the illustrations
and description provided herein. In particular, preferred embodiments of
methods and
systems of the present invention are described with reference to assessment of
cardiac tissue
properties and cardiac parameters, such as cardiac output. It will be
recognized by those
having skill in the art that the methods and systems of the present invention
may be applied to
other cardiac tissue targets and, more broadly, to other types of cardiac
tissue parameters.
Several exemplary systems of the present invention for acquiring data
indicative of
intrinsic and/or induced tissue displacements are described below. Although
such systems
may utilize commercially available components, the processing of the acquired
data and the
correlation of the acquired data to medically relevant physiological
properties provides new
modalities for noninvasively assessing numerous physiological parameters.
Exemplary data
processing techniques for detecting intrinsic and/or induced tissue
displacements using
acquired acoustic scatter data and correlating the acoustic scatter data or
the displacement
derivation with clinically important parameters, such as cardiac output, are
also disclosed
below. These techniques are exemplary and methods and systems of the present
invention
are not intended to be limited to the use of these exemplary techniques.
In a simplified system (not illustrated), a single acoustic transducer may
provide the
interrogation signals) required for tissue assessment in passive modes, the
acoustic force
required for tissue displacement in active modes, and additionally may provide
for detection
of scattered interrogation signals) that are indicative of intrinsic (passive
mode) or induced
(active mode) tissue displacement. For example, commercially available
ultrasound
transducers have sufficient bandwidth, such that a single transducer may be
used to emit
interrogation signals) for measuring intrinsic tissue displacements when
operating at a first
frequency, a first pulse repetition rate and a first intensity; to induce
(exogenous)
displacement or oscillation of tissue when operating at a second frequency, a
second pulse
repetition rate and a second intensity, and to detect signals reflected or
backscattered or
echoed or emitted from the tissue, e.g. when operated at a third frequency, or
at additional
frequencies, to assess the intrinsic or induced tissue displacement or
emission, or to assess a
biological response to the intrinsic or induced tissue displacement. Multiple
acoustic
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transducers may also be used. In another embodiment, one or more diagnostic
ultrasound
probes and one or more displacement ultrasound probes may be embodied in a
single acoustic
element.
In general, acoustic interrogation pulses have larger peak positive pressure,
have a
higher frequency, and are shorter than acoustic palpation pulses. Acoustic
interrogation
pulses, for example, may have a typical frequency between 0.5 and 15 MHz, use
from 1-50
cycles per pulse, consist of 3-10,000 pulses per second, and have a time-
averaged intensity of
less than 0.5 W/cm2. Acoustic palpation signals may, for example, have a
frequency of from
0.5 to 10 MHz, consist of long tone bursts of from 0.1 - 100ms, consist of 1-
100 pulses per
second, and have a time averaged intensity of less than 100-1000W/cmz, where
longer pulses
have lower intensities, for example. Acoustic emissions from palpated or
oscillated tissue are
expected to be in the frequency range of SOOHz to IOKHz.
Fig. 6 is a schematic diagram illustrating a system of the present invention
for
inducing and/or detecting at least one aspect of intrinsic or induced tissue
displacement for
applications such as assessment of cardiac tissue properties. As shown in Fig.
6, systems of
the present invention comprise an acoustic source and receiver combination 10
for non-
invasively assessing tissue displacement or emission at a distance from the
source/receiver
combination. In one embodiment suitable for use in passive modes to assess
intrinsic tissue
displacement, acoustic source and receiver combination 10 comprises one or
more acoustic
sources) 12 for producing an interrogation signal. In another embodiment
suitable for use in
active modes to assess induced tissue displacement or emission, acoustic
source and receiver
combination 10 comprises one or more acoustic sources) 22 for generating an
acoustic
radiation force, or for generating an oscillatory radiation force, or inducing
an acoustic
emission. Acoustic sources) 12 are driven by and operably connected to an
amplifier or
power source 14, which is operably connected to one or more function
generators) 16, which
is operably connected to a controller 20. Controller 20 preferably has the
capability of data
acquisition, storage and analysis.
Controller 20, function generator 16 and amplifier 14 drive acoustic sources)
12 in an
interrogation (passive) or an acoustic radiation force (active) mode. In the
passive mode,
controller 30, function generator 28 and amplifier 26 drive acoustic sources)
22 through the
diplexer 24 at a desired frequency, intensity and pulse repetition rate to
produce an
interrogation signal for tissue target 32, such as cardiac tissue, without
producing undesired
side effects, and without producing a significant (exogenous) displacement.
The resulting
scattered signal is received at controller 30 via diplexer 24. In the active
mode, controller 20,
CA 02490999 2004-12-22
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function generator 16 and amplifier 14 drive acoustic sources) 12 at a desired
frequency,
intensity and pulse repetition rate to produce a displacement in tissue target
32, such as
cardiac tissue, without producing undesired side effects. In some embodiments,
the
controllers 20 and 30 communicate with one another to interleave their signals
in time, for
example. The system based on transducer 22 can monitor the displacements
and/or
emissions induced by transducer 12.
The operating acoustic parameters are related to one another and suitable
operating
parameters may be determined with routine experimentation. The focal point of
the acoustic
source(s), or transducer(s), may be fixed and non-adjustable as a consequence
of the
mechanical configuration of the transducer. Alternatively, multiple
transducers may be
provided and arranged to permit variation and adjustment of the focal point.
Acoustic
sources, or transducers, are preferably annular in configuration and, in
preferred embodiment,
acoustic source 12 comprises multiple annular transducers arranged in a
concentric
configuration. Acoustic sources and tranducers may be arranged axially or off
axis with
respect to one another.
A second acoustic source 13 driven by and operably connected to a diplexer 15,
which is operably connected to an amplifier or power source 17, which is
operably connected
to a function generator 19, which, in turn, communicates with controller 20
and/or controller
30 may also be provided, as shown in Fig. 6. Acoustic source 13 may be used
for assessing
the characteristics of the environment between the acoustic sources) and the
target tissue,
and may operate independently of transducer 12 and the related driver and
controller
components used for the assessment of the target tissue, or in coordination
with transducer
12.
Fig. 7 illustrates one embodiment of an acoustic source and probe combination
40 that
is especially suitable for use with the active mode of tissue assessment of
the present
invention. Source and probe combination 40 comprises confocal, annular
acoustic sources 42
and 44 and a diagnostic ultrasound probe 46. Phasing acoustic sources 42 and
44 at slightly
different frequencies produces a significant radiation force only at their
mutual focus,
indicated in the cardiac tissue, such as near the ventricular wall surface,
schematically
illustrated at location 48, and deforms the tissue. When a single acoustic
source is used, or
the sources are used such that there is no difference in frequency between the
sources, the
result is a unidirectional displacement of the target tissue that coincides
with their
overlapping foci, with negligible oscillatory component for the duration of
each acoustic
pulse. Under these circumstances, repeated single-frequency pulses will create
periodic
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" . a.~ ~...~ a.:.~. ...u, . .,~,.. ..... .. ,.... ..
pulsations of the tissue at the frequency of the PRF. In either embodiment,
acoustic
emissions may be generated from the transiently deformed tissue, with the
emissions
monitored by transducer 46 and related to tissue properties or physiological
conditions.
The acoustic source and probe combination 40 illustrated in Fig. 7 may also be
used,
in combination with an imaging system, to acoustically palpate tissue at
target sites to
localize tissue responses to the focused ultrasound. The imaging system may
employ
ultrasound or another tissue imaging modality, such as magnetic resonance
imaging,
computed tomagraphy, fluoroscopy, or the like. Using an acoustic source and
probe
combination having ultrasound imaging capability, for example, provides
visualization of the
target site and aids targeting of the acoustic radiation force and
localization of responses.
Fig. 8 illustrates another acoustic source and probe combination 50 comprising
a
plurality of ultrasonic transducers 51, 52, 53 and 54, arranged as concentric
annular elements.
Each annular acoustic source represents a single frequency source of
ultrasound that
cooperates, with the other acoustic sources, to interrogate and/or displace
tissue at a selected
location. The foci of the annular transducers is the focus of the
interrogation signal, or the
radiation force, and the location of assessment of intrinsic tissue
displacement and/or induced
tissue displacement and/or emissions. More or fewer ultrasonic transducers may
be used. A
larger number of annular transducers generally provide a greater degree of
control and
precision of where the interrogation signals, or the radiation force, is
focused. This
arrangement of annular transducers may also be used, in a variable frequency
mode, to
generate an oscillatory radiation force in target tissue. When multiple
acoustic sources are
used, each source is operated by a controller, amplifier and function
generator, but operation
of the separate acoustic sources is controllable using a centralized control
system. This
acoustic system may be further generalized or modified for specific
applications by using a
non-annular or non-axial distribution of transducers to allow for additional
ultrasound beam
forming or electronic steering.
Detection element 56 is provided in acoustic combination 50 to detect at least
one
aspect of intrinsic and/or induced tissue displacement. In one embodiment,
element 56
comprises a diagnostic ultrasonic probe that emits an ultrasonic pulse toward
the site of tissue
displacement and detects its echo to track the magnitude, or other aspects, of
tissue
displacement. In another embodiment, element 56 comprises an ultrasound probe,
such as a
transcranial Doppler, that detects the Doppler shift produced by the tissue
displacement. In
yet another embodiment, detection element 56 comprises a hydrophone that
detects the sound
waves emitted by tissue in which an acoustic radiation force is generated.
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Commercially available components may be used in systems of the present
invention.
The following description of specific components is exemplary, and the systems
of the
present invention are in no way limited to these components. High intensity
focused
ultrasound transducers are available from Sonic Concepts, Woodinville, WA.
Multi-element
transducers have been used by researchers and are described in the literature.
A multiple
focused probe approach for high intensity focused ultrasound-based surgery is
described, for
example, in Chauhan S, et al., Ultrasonics 2001 Jan, 39(1):33-44. Multi-
element transducers
having a plurality of annular elements arranged, for example, co-axially, are
suitable. Such
systems may be constructed by commercial providers, such as Sonic Concepts,
Woodinville,
WA, using technology that is commercially available. Amplifiers, such as the
ENI Model A-
150, are suitable and are commercially available. Diplexers, such as the Model
REX-6 from
Ritec, are suitable and are commercially available. Function generators, such
as the Model
33120A from HP, are suitable and are commercially available. Many types of
controllers are
suitable and are commercially available. In one configuration, a Dell
Dimension XPS PC
incorporates a Gage model CS8500 A/D converter for data acquisition, and
utilizes LabView
software from National Standards for data acquisition and equipment control.
In some
embodiments, an ATL transcranial Doppler probe, Model D2TC, is used for
detection.
One aspect of the present invention relates to acoustic source/detector
systems for use
in methods and systems of the present invention. In operation, an acoustic
source/detector
combination, such as a TCD transducer/detector, is stably mounted, or held, in
proximity to a
surface in proximity to an acoustic window, such that the focus of the
acoustic sources) is
adjustable to provide an acoustic focal point within, or on, or in proximity
to, myocardial
tissue. The acoustic source/detector combination is preferably provided as a
unitary
component, but separate components may also be used. The acoustic
source/detector
combination may be mounted on a stabilizer, or in a structure, on the chest.
An applicator
containing an acoustically transmissive material, such as a gel, may be placed
between the
surface of the acoustic source/detector combination and the chest. An acoustic
source/probe
combination may be provided in a holder that is steerable to facilitate
probing of various
targeted tissue sites within a general situs. Steering of the acoustic device
may be
accomplished manually or using automated mechanisms, such as electronic
steering
mechanisms. Such mechanisms are well known in the art.
In one embodiment, one or more transducer arrays) are used for acquisition of
acoustic data, and data is processed using accompanying processing, storage
and control
functions. In general, such transducer arrays may be referred to as "phased
arrays," since the
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individual acoustic elements within the array are coordinated with one
another. Transducer
arrays may be used in either or both passive and active modes of operation,
and may be used
in imaging modes to display data relating to cardiac tissue properties and
cardiac parameters.
Many imaging and display techniques are known in the art and may be used to
highlight
various types and aspects of acquired data.
In one embodiment, one or more transducer arrays may be operated
simultaneously,
or alternately, in active and passive modes of operation. Using a programmable
acoustic
transducer array, for example, multiple tissue sites may be acoustically
interrogated in an
active or passive mode simultaneously, or intermittently at pre-selected time
intervals.
Similarly, acoustic scatter data may be collected from multiple target cardiac
tissue sites
simultaneously, or intermittently. In one embodiment, for example, tissue
properties of target
myocardial tissue may be determined based on acquired acoustic data while mean
and/or
continuous ABP is determined simultaneously based on acoustic data acquired
from or in
proximity to blood vessel(s).
In one embodiment, acoustic arrays of the present invention comprise
capacitive
micromachined ultrasound transducers (cMUT). cMUT transducer arrays may be
used in
both active and passive modes of operation according to the present invention.
cMUT
ultrasonic transducers are manufactured using semiconductor processing
techniques and have
sufficient power and sensitivity to transmit and receive at diagnostic
ultrasound energy levels,
which is necessary and sufficient for our purposes. The transducers are made
by fabricating
very small capacitive diaphragm structures in a silicon substrate. Fig. 9
shows a single
cMUT array transducer cell structure. These diaphragm-structures convert
acoustic
vibrations into a modulated capacitance signal or vice versa. A DC bias
voltage is applied
and an AC signal is either imposed on the DC signal in transmission or
measured in
reception. A cMLTT array is composed of multiple individual cell structures
arrayed in rows
and/or columns.
In one embodiment, two cMUT acoustic arrays are aligned in a sparse two-
dimensional (2D) array known as a "Mills Cross" configuration, which allows
one array to
sweep vertically in send and receive modes and the other to sweep horizontally
in receive and
send modes. In this implementation, two crossed linear cMUT arrays
alternatively transmit
and receive ultrasound while electronically steering the sending and listening
beams, to
identify and focus on the acoustic signal that has the largest Doppler shift
using, for example,
range-dependent Doppler methodologies described below. In alternative
embodiments, the
send and receive modes of the acoustic arrays may be reversed, or a single
array may be used
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to both send and receive acoustic signals. Full 2D transducer arrays having
acoustic elements
arranged in any two-dimensional configuration may also be used. Three
dimensional
transducer arrays may also be used with appropriate control and processing
systems. In yet
another embodiment, a cMUT array may be used in combination with a PZT
transducer, with
the PZT transducer serving as the acoustic source and transmitting around the
cMUT array,
and the cMUT array serving as the acoustic detector.
cMLTT transducer arrays have the potential of being produced very
inexpensively, and
may also have the support electronics integrated onto the same chip. In one
embodiment,
acoustic arrays of the present invention are provided as a disposable
component of an ICP
monitoring device comprising one or more transducer arrays in operative
communication
with a data processing, storage and display device. The one or more transducer
arrays may
communicate with a data processing, storage and display device by means of one
or more
cables, or using a radio frequency or other wireless technology. The
transducer arrays) may
be steerable and may be programmed to scan, identify one or more desired
target site(s), and
maintain focus on that target site in an automated fashion. Transducer arrays
of the present
invention may also be programmed to collect acoustic data from multiple target
sites
simultaneously, or at different times. In one embodiment, a transducer array,
or a plurality of
arrays, may be programmed to operate alternatively as acoustic sources and
detectors. In one
embodiment, multiple transducer arrays used for monitoring multiple patients
provide data to
and communicate with a single data processing, storage and display device.
In another embodiment, an acoustic array comprising PVDF (polyvinylidene
fluoride)
film transducers is used as an acoustic detector array, in combination with a
cMCTT array or a
single element PZT transducer employed as the source. In this embodiment, the
source
transducer or array transmits sound through the PVDF array, sweeping the sound
in a single
dimension generally perpendicular to the arrangement of the PVDF array. The
PVDF array
serves as the acoustic detector, receiving and processing acoustic signals. An
acoustic array
of the present invention may comprise a combination of PVDF and cMLTT arrays.
The
combined depth of the arrays may be on the order of 1 cm. The cMUT array is
arranged
below the PVDF array and transmits sound through the PVDF array. The PVDF
array may
be made in two dimensions, so that it can detect acoustic signals in two
directions, rather than
the single direction illustrated.
Alternatively, an acoustic array of the present invention may comprise a
combination
of a PVDF array and one or more PZT transducer(s). The PVT transducer may be
mounted
below the PVDF array and transmit through the PVDF array in a single, broad
beam. The
CA 02490999 2004-12-22
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PVDF array may be constructed as a single dimension array, or as a two
dimensional array.
An acoustic array having a two dimensional PVDF array has the capability of
receiving
acoustic signals in two dimensions and an underlying PZT transducer. This
system may
alternatively employ a cMCJT array in the place of the PZT transducer.
Systems of the present invention may comprise both non-disposable and
disposable or
reusable components. Costly elements of the acoustic system are provided as
non-disposable
components, while less costly components, which require close interaction with
a patient and,
perhaps, sterilization, are provided as disposable components.
In one embodiment, an acoustic array is provided as part of a disposable
system
element, in combination with a patient interface component. The acoustic array
is preferably
in contact with acoustically transmissive material, such as an acoustic gel,
that provides high
fidelity acoustic transmission into and from the target area. The acoustically
transmissive
material is preferably interfaced with a contact material, such as an adhesive
material, that
facilitates temporary positioning and affixation of the disposable system
element to a
patient's skin. The patient contact material may be protected by a removable
cover, which is
removable at the time of use. The disposable system element, including the
acoustic array,
may be provided as a unitary element that may be sterilized and packaged for
one-time use.
Alternative disposable systems and elements may also be employed. In one such
alternative system, acoustically transmissive material layers may be provided
as a separately
sterilized, packaged component that is designed to interface with a non-
disposable component
including the acoustic array(s). Such layers may be provided with an adhesive
layer on one
side for contact with the patient's skin. Or, a recess may be provided for
manual application
of acoustically transmissive material. It will be evident that many different
embodiments and
arrangements of disposable and non-disposable elements may be employed.
This compact, disposable array element may be placed in contact with the skin
of a
patient at an acoustic window and, when activated, electronically focuses the
acoustic
sources) and detectors) on the target site of interest, such as a target
myocardial tissue site.
The acoustic array monitors and stays focused on the target area of interest
during operation.
In this embodiment, the acoustic array forms part of a disposable assembly
including an
acoustic gel, or another acoustic material that facilitates transmission of
acoustic signals at
the interface with the patient's skin during operation. The exposed surface of
the acoustic gel
is preferably interfaced with one or more adhesive elements that facilitate
temporary
placement on and consistent contact with a desired patient surface. A
removable cover may
be provided over the acoustic gel to preserve the acoustic array and other
components. These
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elements may be provided as a disposable unit that is mountable on non-
disposable elements
of the system. Non-disposable elements of the system may include mounting
hardware, one
or more cables or wireless transmission interfaces, and a data processing,
storage and display
device.
Placement of the acoustic sources) and detectors) on a subject for assessment
of
acoustic properties of myocardial tissue (including blood and blood vessels)
may be at known
"acoustic windows." The placement of the sources) with respect to the
detectors) will
depend on the acoustic data desired - e.g., for collection of back scatter
acoustic data, the
sources) and detectors) are in proximity to one another, while the sources)
and detectors)
are positioned generally opposite one another for collection of forward
scatter acoustic data.
Acoustic scatter or reflection data may be collected at various angles by
placing the sources)
and detectors) at various locations on the patient.
To ensure that representative target tissue is sampled, the target tissue
location must
be volumetrically large enough to provide a representative sample. The
volumetric sampling
requirements will vary, of course, according to tissue type and location. In
general, target
sites having tissue volumes of from 1 mm3 to about 100cm3 are suitable, and
target tissue
sites having tissue volumes of less than about 5 cm3 are preferred. Acoustic
data acquisition
techniques of the present invention may be used in combination with known
ultrasound
imaging techniques to provide visualization of the target tissue sites.
Data, such as acoustic scatter data, relating to intrinsic and/or induced
tissue
displacements is processed according to methods and systems of the present
invention and
related to medically relevant physiological properties, such as cardiac output
and other
cardiac parameters. Exemplary data processing techniques for making various
correlations
based on various types of acquired data are well known. Although these data
processing
techniques are based on the acquisition of acoustic scatter data, they may be
applied, as well,
with modifications that would be well known in the art, in other modalities,
such as near
infrared spectroscopic (LAIRS) modalities and magnetic resonance modalities.
For some applications, as mentioned previously, relative trend determinations
of
cardiac tissue properties over time, or at different points within or over
multiple cardiac
cycles, are useful. In other applications, it is useful to compare measured,
or determined
values for cardiac tissue properties to standard values based, for example, on
empirical data.
In this way, abnormal, or dysfunctional tissue may be identified by comparison
to "normal"
or "functional" tissue values.
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In one embodiment, a small specialized ultrasonic palpation device is placed
on the
subject's chest and aimed, through the ribs, to a target cardiac tissue site
at or near the right
ventricular wall. This may be achieved using a diagnostic ultrasound scan head
placed
confocally with the palpation device, so that the focus of the palpation
device is registered on
the screen and visible to the person implementing this procedure. In another
embodiment, a
simple A-mode transducer/hydrophone is used to aim, palpate and display data,
and provides
a stand-alone device. The right ventricle is exemplary, but, in practice, this
technique may be
used to with focus ultrasound beams to targeted site at or near cardiac
tissue. This combined
palpation and aiming scan head is preferably secured to the outside of the
chest for the
duration of the medical procedure, with the assessment being initiated when
the patient's
blood volume and cardiac volume are normal.
With the specialized palpation device properly aimed, one can, in one
embodiment,
apply a constant-amplitude oscillatory radiation force to the right
ventricular wall, which
causes that focal portion of tissue and a rim of adjacent tissue to oscillate.
This may be done
by the application of focused ultrasound with a dual-annular array, with each
annulus
operating at slightly different frequencies from one another. The frequency of
the oscillatory
radiation force will be that of the difference frequency of the two annuli.
For a given tension
in the right ventricular wall, i.e., at a given part of the cardiac cycle, and
for a given constant
amplitude forcing, there will be a difference frequency, hence oscillation
rate, in the radiation
force that maximizes the acoustic emissions from the point of application of
the radiation
force. This frequency may be referred to as the resonant frequency of the
ventricle wall. As
the wall tension changes for a variety of reasons, this resonant frequency
will change: the
greater the tension the higher the resonant frequency, while the lower the
tension the lower
the resonant frequency.
Commercially available hydrophones may easily be integrated into the
ultrasonic
palpation device for tracking the acoustic signals emitted from the target
cardiac tissue. By
tracking this resonant frequency, starting with a baseline determined while
the patient is
awake" preoperatively or newly anesthetized but before a change in blood
volume, one can,
with concomitant blood pressure measurements, assay where the patient is on
the Starling
curve. For example, if the resonant frequency dips significantly lower than
the patient-
specific average normal value, this would be consistent with the ventricle
walls becoming
more flaccid. If this were to occur while blood pressure drops, then these
observations would
be strong evidence of hypovolemia. If the resonant frequency becomes
significantly higher
than the patient-specific average normal value, this would be consistent with
the ventricle
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walls becoming stiffer. If this were to occur while blood pressure drops, then
these
observations would be strong evidence of hypervolemia. One could continue
tracking
cardiac wall stiffness in this fashion throughout the medical procedure of
interest, until the
patient is safely stabilized.
In an alternative embodiment that is otherwise similar to the embodiment
described
above, one could use more than two annuli in the array. And, in another
alternative
embodiment that is similar to that described in the previous paragraph, one
could use a single
or multi-element array operated in a continuous wave (CW) mode, and vary the
amplitude of
the applied signal at a frequency that would induce the desired oscillations
in the tissue. And,
in yet another alternative embodiment that is similar to that described in the
previous
paragraph, one could use a single or multi-element array operated in a pulsed
mode, and vary
the pulse repetition frequency of the applied signal until the resulting
temporal series of
pulses induces the desired oscillations in the tissue. There are costs and
benefits associated
with each choice of palpation device.
In another embodiment, we apply a constant-amplitude oscillatory radiation
force,
using one of the several methods described above. Rather than search for a
resonant
frequency, however, we work with a given frequency that from experience is
known to be
above or below the resonant frequency of the heart's right ventricle wall. We
then track the
amplitude of the palpation-induced acoustic emission from the heart, both
within a cardiac
cycle and over many cardiac cycles, starting while the patient's cardiac
volume is normal,
and then proceeding throughout the medical procedure of interest until the
patient is safely
stabilized. For example, consider the case where one was driving the local
heart tissue into
an oscillation whose frequency was always below the resonant frequency of the
local heart
tissue. If the average amplitude of the ultrasound-induced acoustic emission
were observed
to increase over time, this would be consistent with a reduction in the
resonant frequency of
the heart tissue, approaching the driving frequency of the acoustic radiation
force from above.
This would suggest that the ventricle walls were becoming more flaccid than on
average.
This observation, in conjunction with an observed drop in blood pressure,
would give strong
evidence of hypovolemia. If, under the same assumptions, the average amplitude
of the
ultrasound-induced acoustic emission were observed to decrease over time, this
would be
consistent with an increase in the resonant frequency of the heart tissue,
moving up and away
from the driving frequency of the acoustic radiation force. This would suggest
that the
ventricle walls were becoming stiffer than on average. This observation, in
conjunction with
an observed drop in blood pressure, would give strong evidence of
hypervolemia.
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In another embodiment, any one of several aspects of safe, ultrasound-induced
deformation of the right ventricular wall of the heart is assayed, using, for
example, an A-
mode transducer placed confocally with the ultrasound palpation device. As
described above,
the palpation device may have one of several manifestations. Also, one would
likely not
need the absolute value of the deformations, just the trend in those
deformations over time, as
well as concomitant measurements of blood pressure, starting when the
patient's cardiac
volume is normal, and ending when the patient is safely stabilized.
According to yet another embodiment, cardiac tissue is not "palpated" at all.
Instead,
the local strain within a small portion of the cardiac ventricle wall tissue
is tracked using, for
example, an A-mode ultrasound system, optionally in conjunction with standard
diagnostic
ultrasound image. The local strain is assayed using sonoelasticity analysis on
the resulting
acoustic backscatter signal, a well-known technique developed over the last 15
years and
often applied for assaying the presence of breast cancer. Sonoelasticity
analysis would give a
measure of the scale of the intrinsic deformations of tissue, essentially the
average change in
spacing of two close points within the tissue (distances of millimeters or
less) divided by their
average spacing at systole or diastole, for example. By tracking such
intrinsic deformations
within or near the same place in the ventricle, the stiffness of the ventricle
walls is monitored,
which relates to cardiac output, as discussed above. For example, for a fall
in blood pressure
and cardiac output, as the stiffness of the heart tissue decreased, the
intrinsic displacements of
portions of the ventricle wall would increase, thereby suggesting hypovolemia.
As the
ventricle walls increased in stiffness, the intrinsic displacements of
portions of the ventricle
wall would decrease, thereby suggesting hypervolemia. In an alternative
embodiment, or one
that might be of use as a complement to the sonoelastic measurement scheme
described
above, one could gain useful information by tracking through time the
macroscopic
displacement of the heart tissue in one place in the heart, on the scale of a
centimeter or so.
Large macroscopic displacements of a fixed portion of heart tissue with low
blood pressure
would suggest low cardiac output due to hypovolemia, while small macroscopic
displacements of heart tissue - likely after a large net displacement away
from the center of
mass of the heart, towards the transducer - coupled with low blood pressure,
would be
consistent with low cardiac output due to hypervolemia.
Arterial blood pressure using"passive" or "active" mode
In another aspect of methods and systems of the present invention, intrinsic
and/or
induced changes in the diameter or other geometric properties of a blood
vessel, or changes in
CA 02490999 2004-12-22
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" ... .,...,. ....r. ~;e." .",~ . ..~... .._. ..
the intrinsic or induced displacement in tissue surrounding blood vessels, are
monitored and
assessed using ultrasound, and this information is related to synchronous
Doppler flow
measurements within the same vessel. In an active mode, tissue displacement
may be
induced in a blood vessel or in tissue surrounding a blood vessel by
application of an acoustic
radiation force, as described above. Similarly, in a passive mode, intrinsic
tissue
displacements at or near a blood vessel may be detected using a variety of
techniques, with
the use of ultrasound techniques being preferred. In some embodiments, an
initial assessment
is performed, using Doppler flow measurements or ultrasound detection
techniques, to locate
a desired blood vessel and thereby provide a focus for identifying intrinsic
and/or induced
displacements at or near the vessel.
Since the diameter (or other geometric properties) of the vessel is a function
of the
pressure being exerted against the wall of the vessel by blood, and since the
velocity of blood
flow is dependent on the diameter (or radius) of the vessel through which the
blood travels,
blood pressure can be calculated from flow velocity measured by Doppler.
Geometric
properties of vessels that may be evaluated using methods and systems of the
present
invention include changes in diameter, cross-sectional area, aspect ratio,
rate of change of
diameter, velocity, and related parameters. By simultaneously measuring the
pulsatility of
the blood vessel of interest and the Doppler flow velocity proximal and distal
to this site,
continuous blood pressure is determined. Specific methods for assessing ABP
are described
below.
Blood pressure may also be assessed, in an active or passive mode, by
examining
acoustic properties of target tissue sites at or in proximity to blood
vessels. The acoustic
properties of target tissue at or in proximity to blood vessels can be related
to tissue stiffness
or compliance, which can be related to blood pressure.
Blood pressure measurements made using the passive or active acoustic modes
described herein may also be used for calibration of existing invasive or non-
invasive blood
pressure monitoring devices. Thus, the methodology described below,
particularly with
reference to blood pressure determinations using the active acoustic mode, may
used in
combination with existing blood pressure monitoring devices, which are
available, for
example, from Medwave Corporation, St. Paul, Minnesota.
Correlation of non-invasively measured spontaneous vessel wall displacement
with Doppler
flow and ABP
This method uses a derived relationship between spontaneous vessel wall
displacement (due to blood pressure and smooth muscle tonal responses to the
hemodynamic
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CA 02490999 2004-12-22
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state), synchronous velocity of blood flow within the vessel of interest, and
invasively
monitored ABP to estimate ABP from non-invasively measured vessel wall
displacement and
Doppler flow velocity. Using an ultrasound probe, the given vessel of interest
is insonated
with a waveform of specific frequency and amplitude, and the time or phase
shift of a
particular reflected or backscattered or echo signal is used to calculate
spontaneous tissue
displacement.
The equation that relates time or phase shift to tissue displacement is d = t*
1500
m/sec, where d = tissue displacement, t = the time or phase shift of the
reflected signal, and
1500 m/sec is the estimated speed of sound through tissue. The relationship
between d,
synchronously measured Doppler flow velocity within the vessel of interest
(i), and
invasively measured ABP is then determined by taking simultaneous measurements
of
spontaneous vessel wall displacement, flow velocity, and ABP and solving for
the equation:
ABP = F(d, i), where F can be any function, such as an exponential, vector,
matrix, integral,
etc., or a simply an empirical relationship. Once F is established (by means
of multiple
empirical measurements from a variety of patients under various
circumstances), the non-
invasive determination of vessel wall displacement and flow velocity is used
to calculate
ABP. A calibration step using, for example, a cuff plethysmograph to measure
ABF, may be
implemented before continuous, noninvasive ABP measurements are made.
Correlation of ABP with amplitude of vessel wall signal and Doppler flow
velocity
This method uses a derived relationship between the amplitude of the reflected
vessel
wall signal, Doppler flow velocity, and invasively monitored ABP to estimate
ABP from
non-invasively measured vessel wall signal and Doppler flow velocity (i).
Using an
ultrasound probe, a particular vessel of interest is insonated with a waveform
of specific
frequency and amplitude, and the amplitude of the backscatter is used to
create a waveform
of vessel wall reflection/absorption. This new waveform, a, is generated by
integrating the
amplitude of the backscatter over a finite epoch (such as the cardiac cycle,
measured with
ECG tracing) and normalizing this by the time period of the epoch. The
relationship between
this derived waveform, a, and invasively measured ABP is then determined by
taking
simultaneous measurements of the backscatter signal, Doppler flow velocity,
and ABP and
solving for the equation: ICP = F(a,i), where F can be any mathematical
function, or simply
an empirical relationship. Once F is established (by means of multiple
empirical
measurements from a variety of patients under various circumstances), the non-
invasive
determination of a can be used to calculate ABP. A calibration step using a
cuff
plethysmograph to measure ABP may be implemented before continuous,
noninvasive ABP
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measurements are made.
Correlation between peak backscatter amplitude and ABP
In a manner similar to that described above, the peak amplitude of the
backscatter
signal over a given epoch (e.g., cardiac cycle) is normalized by the baseline
value of the
backscatter signal over the same epoch, and this, along with Doppler flow
velocity, is related
to the simultaneous invasive measurements of ABP. A calibration step using a
cuff
plethysmograph to measure ABP may be implemented before continuous,
noninvasive ABP
measurements can be made.
Methods and systems of the present invention may be used in a variety of
settings,
including emergency medicine settings such as ambulances, emergency rooms,
intensive care
units, and the like, surgical settings, in-patient and out-patient care
settings, residences,
airplanes, trains, ships, public places, and the like. The techniques used are
non-invasive and
do not irreversibly damage the target tissue. They may thus be used as
frequently as required
without producing undesired side effects. The methods and systems of the
present invention
do not require patient participation, and patients that are incapacitated may
also take
advantage of these systems. The methods and systems of the present invention
for assessing
cardiac tissue may be used on a continuous or intermittent basis.
EXAMPLE 1
Brain tissue was used as a model experimental system. We have shown in vitro
(Figure l0A) and in vivo (Figure lOB-D) and describe in detail below, that
intrinsic
displacements of brain tissue (e.g. compressions and distensions), and their
various acoustic
scatter properties, can be directly measured using a standard transcranial
Doppler (TCD)
transducer, off the-shelf data acquisition systems, and novel analysis of the
acoustic
backscatter signal from brain. Myocardial tissue displacement may be measured
in the same
fashion and related to tissue strain, tension, and the like, as described
above, to make
noninvasive assessment of cardiac tissue and parameters.
An in vitro model for examining changes in ICP using acoustic techniques was
constructed using fresh bovine brain immersed in fluid in a water-tight,
visually and
acoustically transparent bottle attached to a hand-pump for changing the
pressure on the
brain. An acoustic transducer (ATL/Philips Medical Systems, Bothell, WA), and
the bottle,
were placed in water so that the focus of the interrogation transducer was
near the edge of the
brain, but within the brain. Using a transducer whose amplifier was driven at
200 mV and a
LeCroy Waverunner oscilliscope, we collected acoustic waveforms backscattered
from the
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tt'°" tf...,. f! ..~ ~i...P '.",;:Ic !f",[i ..:;.G ~'- ti:::: Pt:..P
.n'' V::d~ "..t~..
brain generated by the interrogator that showed, measured by changes in
arrival times, that
increases in displacement of beef brain as a function of increased pressure on
the in vitro beef
brain, as determined by a gauge on the hand pump, were linearly related (See
Fig. l0A).
This was the expected result: as the pressure on the brain (ICP) increases as
a consequence
of increasing liquid (CSF) volume in a confined space, we would expect to see
the brain
move away from the container.
The displacement (compression and distension) waveforms shown in Figs. lOB-D
were produced using ultrasound techniques to measure acoustic scatter signals
associated
with intrinsic displacements of human brain tissue in situ. An acoustic
transducer
(ATL/Philips Medical System, Bothell,WA) was used to insonate target CNS
tissue with
acoustic interrogation signals having 10-103 acoustic pulses per second at
2.25 MHz
containing 3-15 cycles of ultrasound with peak negative pressures less than
2MPa or 20 bar.
Using a LeCroy Waverunner oscilliscope, we collected acoustic waveforms
backscattered
from the brain generated by the interrogator and calculated the tissue
displacement.
This calculation was made using a normalized correlation of paired received
signals.
Given an estimate of the speed of sound in brain and the calculated temporal
displacement,
the spatial displacement of the tissue at a given moment may be calculated.
Tracking the
spatial displacement over time provides a direct measure of the displacement
of the brain
tissue that is being noninvasively interrogated by the diagnostic ultrasound.
This calculation
can also be made by correlating the backscattered signal with a reference
interrogation signal,
noting when the interrogation signal is sent and when the backscattered signal
is received.
Changes in the amplitude of the backscatter from the region of interest may
also be
monitored to determine the ICP waveform. For example, we have found that by
integrating
the acoustic backscatter signal over a short time interval of about 5 to 10 ms
at the region of
interest, and normalizing that integral by the length of that time interval,
we developed a time
series that has the salient features of a typical ICP waveform. In particular,
for small volumes
of measured brain displacement, the signal derived from following
displacements or from
following the normalized integral of the backscatter looks identical to the
time course of the
mean velocity of blood in the middle cerebral artery of the test subject.
Figs. lOB-D show changes in properties of a human brain over time, measured in
situ,
using ultrasound techniques according to the present invention, as described
above. Certain
physiological behaviors, such as holding breath, sneezing, etc., are known to
transiently
increase or decrease ICP.
Fig. l OB shows changes in the normalized amplitude of the acoustic
backscatter as the
44
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II". t6..,. Lf ..' H..,V ~....rv ::...t: .....: . ~..,... ...,. .. ....
human subject held his breath. Fig. lOC shows the displacement of human brain
as the
human, based on correlation techniques, while the subject was holding his
breath, using
pulses with 1 S cycles of ultrasound. In particular, Fig. l OC shows the net
increased
displacement of brain towards the transducer as the pressure on the brain
increased due to an
accumulation of blood volume in the brain, along with the cardiac-induced
brain
displacement signals.
Fig. l OB shows the same kind of received signal characteristics as Fig. l OC,
where we
used pulses with 5 cycles, but analyzed the data by integrating over the
acoustic backscatter
signal as described above. As in Fig. IOC, both waveforms changed over the 10
seconds
while the subject held his breath, consistent with known transient changes in
ICP when
subjects hold their breath. The vascular pulse and autoregulation waveforms
are present, in
modified form, in Fig. l OC. 'The time series of Figs. l OB and l OC look
similar to the velocity
pattern found in the patient's middle cerebral artery (data not shown). This
measurement is
therefore an accurate representation of the compression and distension of
brain parenchyma
in response to the major cerebral arteries, supplemented by contributions from
the rest of the
cerebral vasculature.
Fig. lOD shows an example of changes in near-surface brain displacement as the
subject first held his breath for 2-3 seconds, then inhaled. Changes in
respiration and the
respiratory cycle are known to transiently change ICP. At first, the brain
surface's net
displacement toward the transducer increased. Upon inhalation, the brain
tissue moved, over
several cardiac cycles, away from the transducer. The observed displacement is
consistent
with the transient changes in ICP expected when a subject holds his breath
(transient blood
volume and ICP increase) and then inhales (transient blood volume and ICP
decrease).
Our measurements were made over a small volume of brain tissue (of order 1.0
cm3).
We anticipate that measurements of brain tissue displacement (e.g. compression
and
distension) of a relatively large volume of brain tissue (on the of order 10
cm3) will produce a
signal that looks identical to a typical ICP trace. This signal is used
directly, or with ABP
data, to assess ICP and/or autoregulation status, as discussed above.
Contributions to the
acoustic backscatter signal over a large volume of brain tissue are the result
of the average
displacements (distension and compression) of brain tissue produced by a
plurality of
cerebral blood vessels, whose particular intrinsic oscillations will cancel,
except for the major
ones (dicrotic notch, etc), which will reinforce one another, as observed
invasively.
CA 02490999 2004-12-22
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EXAMPLE 2
We have shown, in vitro, using a beef brain model similar to that described
above,
that a palpation pulse of ultrasound across a range of acoustic intensities
can cause increasing
displacements of brain without causing gross tissue damage. Palpation of
myocardial tissue
using ultrasound pulses may be achieved in a similar fashion.
Fresh bovine brain was immersed in fluid in a water-tight, visually and
acoustically
transparent bottle attached to a hand-pump for changing the pressure on the
brain. ATL
acoustic transducers (ATL-Philips Medical Systems, Bothell, WA), and the
bottle, were
placed in water so that the focus of the acoustic palpation and interrogation
transducers were
near the edge of the brain, but within the brain. Using LeCroy Waverunner
oscilloscope, we
collected acoustic interrogation waveforms backscattered from brain. For
palpating and
interrogating beef brain, in vitro, the interrogation pulses were administered
as described with
respect to Fig. 10A, while the palpation pulses had a pulse repetition
frequency of lHz,
contained 30,000-50,000 cycles, and had a time-averaged intensity of less than
SOOW/cmz.
As shown in Fig. 1 l, as the acoustic force of the ultrasound increases
(proportional to
the driving voltage given in mV) at ambient (0 mmHg) pressure, so does the
measured
displacement of the beef brain, given in microns. We have also shown in the
experimental
beef brain model described above, in vitro, that brain displacement due to
identical ultrasonic
palpation pulses decreases from 300 ~,m to 210 ~m as the pressure on the brain
increases
from 0 to 55 mm Hg. Therefore, when the same acoustic force is applied with
ultrasound,
brain-tissue displacement in vitro is inversely proportional to ICP, as
expected. Noninvasive,
ultrasound-based measurements of ultrasonic palpation of brain tissue can be
safely used to
directly measure ICP in humans, without the need for blood pressure
measurements, because
by this method the brain will be subjected to a known (ultrasonic) force.
Alternatively, using
a focused ultrasound beam with an intensity less than a value easily
determined to be safe,
probing or palpation of brain tissue with a known force will also yield data
ancillary to the
passive method of ICP determination, by calibrating the amount of deformation
brain tissue
undergoes when subjected to a known compressive force.
All of the publications described herein, including patents and non-patent
publications, are hereby incorporated herein by reference in their entireties.
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