Note: Descriptions are shown in the official language in which they were submitted.
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METHOD OF MEASURING TOTAL VASCULAR HEMOGLOBIN MASS
BACKGROUND AND SUMMARY
The National Trauma Data Bank Report for 2004 describes 576,247 hospital
admissions for trauma between 1999 and 2004. Of these cases, 109,080 patients
were
admitted to the intensive care unit (ICU), 100,050 were taken directly to the
operating room
(OR), and 7878 died. The remaining 332,928 were admitted for general care. For
many of
these patients (especially for the ICU and OR patients) it was necessary to
closely monitor
the hematocrit with multiple phlebotomy blood samples within the first few
hours. The key
to providing optimal care for these challenging patients is for the trauma
specialist to
provide rapid therapeutic interventions based upon informed decision-making.
The
clinician's ability to deliver such quality care is based primarily on
physical assessment
skills, training, and experience, and secondly upon the degree of patient
physiologic and
hemodynamic data available at the moment of decision-making. There is a clear
need for
the clinician to have quantitative data to base his or her treatment decision.
The process of frequent phlebotomy consumes valuable emergency staff time
and there can be substantial lag-time before results are available. Laboratory
techniques
have become more accurate and bedside devices have improved turn-around time
for in
vitro lab analysis, but these improvements have not alleviated the central
problem of lack of
real time information. The patient's condition may deteriorate within minutes,
and reasons
for the deterioration can be varied and not always obvious. Survival rates for
such patients
could be improved if needed data could be provided continuously, allowing
better
opportunity to act upon the vital information in a more timely manner. Patient
monitoring
methods have advanced over the decades with the development continuous
arterial blood
and oximetry pressure monitoring, but there remains no device that delivers
other necessary
physiologic data on a continuous basis.
In the current practice of critical care medicine, the only patient parameters
that are continuously monitored are the vital signs, pulse oximetry, and
temperature. Aside
from oximetry, the physiologic parameters are available only through
phlebotomy sampling
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and laboratory analysis. The hemodynamic parameters, other than vital signs,
are available
only with central vascular catheterization in the ICU or the cardiac catheter
laboratory. The
availability of these continuous physiologic and hemodynamic parameters during
patient
resuscitation would improve the delivery of appropriate, timely, and cost
effective patient
care and, thereby, improve outcomes. Such continuous monitoring would also
improve the
ability of the critical care team to effectively care for multiple patients
without the need for
numerous and laborious repeat lab tests.
The fundamental basis of successful shock resuscitation, no matter what the
cause, is the rapid assessment of vascular volume. Volume assessment, however,
can be as
difficult and ambiguous as it is essential to the resuscitation process.
Volume assessment
has been called the "cornerstone" of resuscitation. One cannot simply empty
the vascular
system, however, to see how much blood it contains. Therefore, throughout
history, a
number of techniques and methods have been utilized in the search for the best
way to
assess volume.
Nuclear medicine techniques for direct volume measurement have existed for
a few decades, but they involve the added risk of radiation and are slow and
therefore not
amenable for use during resuscitation. Central vascular pressure and pulmonary
artery
wedge pressure monitoring became the gold standard several decades ago but are
recently
falling out of favor due their invasive nature and lack of statistical proof
of reliability. New
hemodynamic monitoring methods aimed at predicting a patient's potential for
fluid
responsiveness have recently been put into practice. These include ultrasound
imaging for
measurement of the inferior vena cava size, arterial waveform analysis, and
noninvasive
cardiac output monitoring (NICOM). While these methods are a step forward in
terms of
hemodynamic monitoring, none are actually capable of determining circulating
blood
volume.
According to one aspect, the invention provides a method of measuring total
vascular hemoglobin mass. One step of the method is to store a volume of fluid
to be
injected using a processor. A hemoglobin value is measured in vivo on a
substantially
continuous basis. The method also measures a total circulating vascular volume
based on
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the volume of fluid to be injected by detecting a change in blood
concentration over a
period of time upon injection of a fluid into a body. The total hemoglobin
mass is
calculated based on the measured hemoglobin value and the measured total
circulating
vascular volume using the following formula:
TIIM = 11gb t¨ Bt CiL)
where THM is total hemoglobin mass in grams, Hgb is the measured hemoglobin
value in grams per deciliter, and BV is the measured total circulating
vascular volume in
deciliters. The calculated total hemoglobin mass is then displayed.
Additional features and advantages of the disclosure will become apparent to
those skilled in the art upon consideration of the following detailed
description of the
illustrated embodiment exemplifying the best mode of carrying out the
disclosure as
presently perceived.
BRIEF DESCRIPTION OF DRAWINGS
For the purpose of promoting an understanding of the principles of the
invention, reference will now be made to the embodiments illustrated in the
drawings and
specific language will be used to describe the same. It will nevertheless be
understood that
no limitation of the scope of the invention is thereby intended. Any
alterations or further
modifications of the described embodiments, and any further applications of
the principles
of the invention as described herein are contemplated as would normally occur
to one
skilled in the art to which the invention relates.
FIG. 1 is a diagrammatic view of a body part having the new and improved
noninvasive vital sign measurement device of the invention attached thereto
utilizing a
spaced apart sender and receiver;
FIG. 2 is a view like FIG. 1 of another version of the new and improved
noninvasive vital sign measurement device of the invention showing a single
sender and
receiver;
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FIG. 3 is a view of another version of the new and improved vital sign
measurement device using three or more transducers;
FIG. 4 us a view of still another version of the new and improved vital sign
measurement device similar to that shown in FIG. 1;
FIG. 5 is a view of still another version of the new and improved vital sign
measurement device in which the transducers are mounted on opposite sides of a
blood
vessel or vessels;
FIG. 6 is a diagrammatic view of the new and improved multiparameter
intravascular catheter vital sign measurement device of the invention for the
insertion into a
blood vessel utilizing separate spaced apart transducers connected to drivers
and recognition
analysis devices;
FIG. 7 illustrates another version of the new and improved multiparameter
intravascular catheter vital sign measurement device of the invention showing
a
transducer/receiver with a reflector held apart by struts;
FIG. 8 is a view like FIG. 7 showing still another version of the new and
improved multiparameter intravascular catheter vital sign measurement device
of the
invention showing a single sender/receiver transducer and a reflector mounted
at opposite
ends of a notch cut into the side of the catheter;
FIG. 9 is a view like FIGS. 6, 7 and 8 of still another version of the new and
improved multiparameter intravascular catheter vital sign measurement device
of the
invention showing a sender/receiver transducer and a reflector mounted in a
notch cut into
the side of the catheter with a membrane or diaphragm covering and sealing the
notch
forming a gas filled chamber pressure sensor; and
FIG. 10 is a graph showing a blood volume measurement made during
testing from which total vascular hemoglobin mass is determined according to
an
embodiment of this disclosure.
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DETAILED DESCRIPTION OF SELECTED EMBODIMENTS
The new and improved noninvasive vital sign measurement device 10 of the
invention is a medical device for supplying vital sign measurements for any
purpose and in
any setting where such information is useful to medical clinicians conducting
physical
examinations or monitoring patients (inpatient, outpatient, or ambulatory),
whether in well-
equipped hospitals, clinics, or on a battlefield. The invention would allow
the monitoring of
vital signs continuously. In the vascular application of the device, vital
signs that can be
measured would include arterial and venous blood pressure and pulse, blood
flow velocity,
and blood density. Peripheral vascular resistance could be calculated and
displayed using
data from the device. More conventional equipment could be mated with the
device in order
to continuously monitor such things as temperature and oxygen saturation.
Other potentially
measurable pressure parameters could include the extravascular space,
intracranial space,
intrathoracic (vascular, airway, and pleural) space, or any confined body
cavity, depending
upon the particular configuration of the device and where it is mounted upon
or applied to
the body. Examples of confined body cavities would include possibly the
urinary bladder,
gallbladder, intra-abdominal, ocular, and more probably extremity fascial
compartments.
Additional measurements that may be obtainable by the device could be other
vascular
parameters including possibly intracardiac chamber pressures and more possibly
central
venous pressures.
When arterial blood pressure is measured and monitored, both systolic and
diastolic blood pressure should be monitored beat-by-beat. This information
would be
useful in evaluating routine vital signs, hypertension, hypotension, and shock
from any
cause. The instantaneous monitoring by the application of the invention would
provide a
means by which the effectiveness of pharmaceutical intervention and surgical
intervention
could be immediately assessed. Venous and extravascular space monitoring can
be used to
determine tissue perfusion and lymphatic obstruction, as well as the general
state of
hydration of the patient. Vascular monitoring will provide information
regarding patient
shock from any cause, e.g., sepsis, blood loss, and autonomic malfunction.
Data from the
combined monitoring of arterial pressure and blood flow could be used to
calculate vascular
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resistance. For the clinician, knowing the level of vascular resistance and
continuously
monitoring blood pressure are key factors in determining not only the cause of
shock but
also the best course of treatment in each circumstance.
Intrathoracic measurements could include intrapulmonary and intracardiac,
as well as pleural and pericardial space pressures. Measurement of large,
medium, and small
airway and alveolar space pressures would give physicians both diagnostic and
treatment
monitoring tools for acute and chronic lung disease. The device could be used
to confirm
endotrachial tube placement. Intrapleural pressure measurements would provide
data for
rapid diagnosis or confirmation of hemothorax and pneumothorax, and could be
used in
both hospital and prehospital settings to help determine the urgency with
which these
conditions should be treated.
Intracardiac pressure measurements would allow diagnosis of valvular
failure, cardiomyopathy, congenital defects, myocardial ischemia/infarct, and
congestive
heart failure. Chamber pressure measurements together with echocardiogram data
and
pulmonary vascular readings would yield vital information regarding the
etiology of any of
the above maladies previously available only with cardiac catheterization.
More convenient and accurate ocular pressure measurement would allow
physicians improved means of diagnoses and treatment monitoring of ocular
diseases such
as glaucoma.
Intracranial pressure measurements would most likely be extremely difficult
because of signal attenuation through bone; however, if possible, it would
give physicians a
rapid estimate of tissue pressure, ventricular pressure, and vascular space
pressure when
dealing with patients suffering from head injury or stroke, and post-operative
neurosurgical
patients. It would also be useful in the diagnosis of such maladies as
pseudotumor cerebri
and hydrocephaly.
Currently, most all of the measurements above can be obtained accurately
only by the use of expensive and/or invasive procedures. Sphygmomanometer
blood
pressure cuff readings are accurate in normal and high ranges, but cumbersome
and slow, as
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well as painful for many patients. For automatic blood pressure cuff devices,
the
International Electrotechnical Commission has set international standards
regarding strict
limits on the pressure to which the cuff can be inflated. And, in order to
avoid tissue damage
and considerable discomfort to the patient, they have also set limits on the
period of rapid
inflation/deflation cycles. Blood pressure cuff readings are in fact
contraindicated for post-
mastectomy patients in the arm on the affected side. However, when vital signs
are unstable
or potent drugs are needed in order to maintain blood pressure, time-consuming
invasive
procedures are required for continuous monitoring. In the emergent setting,
clinical
decisions must often be made long before there is any x-ray or
echocardiography evidence
available and long before invasive vascular monitoring catheters can be
inserted and
calibrated.
Vital sign data which can be obtained by the device are useful in intensive
care units, operating rooms, all prehospital settings, emergency departments,
dialysis
centers, medical practice offices, medical research, pulmonary and veterinary
clinics, in
military installations or on a battlefield, and in aerospace installations for
monitoring pilots
and astronauts at work. On an ambulatory basis, such data would also be very
useful in
everyday life and in the sports world. We currently have no convenient way to
monitor the
businessman, the homemaker, or the athlete in action.
The function of the device depends upon subtle, but measurable, changes in
acoustic velocity that occur as a result of changes in density of the medium
through which
the sound wave is propagating. The noninvasive device would measure acoustic
transit
times, and thereby measure density within fluid or gas-filled body
organs/structures/vessels.
By monitoring transit times and minute shifts in transit time in rapid
sequence (10 to 100
times per second) during all phases of systole and diastole, such
measurements, if made
with precision, would result in accurate, reliable, and continuous vital sign
data.
In addition to arterial and venous pressure readings, this principle would in
like manner apply to the measurement of pressure in gas-filled structures such
as pulmonary
airways and possibly the bowel lumen. Similarly, ocular, intrauterine, and
possibly
extremity compartment pressures would be amenable to measurement. Intracranial
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pressures may be measurable with this technique as well. The acoustic
frequency
specifications and configuration of the device would be altered according to
the purpose at
hand; e.g., airway pressure measurement would require much lower frequencies
for better
intrathoracic sound penetration and since acoustic velocity is much slower in
gas than in
fluid.
The measurements are based upon the characteristics of acoustic waves as
they propagate through biologic tissue or fluids or gases. Since acoustic
velocity increases
with the density of the medium through which it is propagating, then there
must be a
measurable change in acoustic velocity through a fluid or gas-filled vessel,
cavity, or
compartment as the density within changes. Minute blood density fluctuations
will occur as
the blood pressure cycles between systole and diastole. Therefore, there must
be a
measurable change in the acoustic wave propagation velocity through the blood
as the
pressure changes. The common equation, V=D/T, indicates that changes in
velocity (V) are
inversely proportional to changes in transit time (T) over a fixed distance
(D). If the
measurements were done with precision, then the device output would consist of
highly
accurate, beat-by-beat digital pressure readings in the case of vascular
application of the
device.
The UNESCO equation describes the relationship between acoustic wave
velocity and pressure in water. The equation also takes into account other
factors that
contribute to the density of the fluid, such as the salinity and the
temperature. Although
fluids (blood included) are considered incompressible, the equation shows that
there should
be minute but measurable changes in velocity associated with changes in
pressure, even
within the human blood pressure range of 0 to 300 mmHg.
To form theoretical support for this method, the space between two
hypothetical transducers was assumed to be 10 cm. The UNESCO equation was then
used
to calculate acoustic velocity at pressure increments of 10 mmHg assuming
fluid
temperature is 37° Celsius, salinity is 9 psu (practical salinity
units) or ppt (parts per
thousand), and variable pressure is expressed in kPa. Calculations using the
formula V=D/T
indicate that in order for the device to have precision to within 1 mmHg, it
must be capable
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of detecting shifts in transit time of roughly 10 picoseconds. Trending of the
pressure could
be achieved by the detection of shifts of approximately 100 picoseconds.
Referring now to FIG. 1, the monitoring device 10 is shown attached to a
body part 12 from which the blood pressure and other vital signs are
monitored. Body part
12 can be any body part including the head, the neck, the chest, the abdomen,
the arms, and
the legs, to measure pressure in any blood vessel in good proximity to the
skin's surface.
Two transducers 14, 16 are spaced apart, longitudinally in line with a vessel,
a specific and
fixed distance, e.g., 10 cm, and applied to the skin using an acoustic
conductive medium. By
measuring the transit time of the acoustic signal between the two transducers
14, 16, the
velocity of the sound wave through the tissue can be calculated using the
equation V=D/T,
where V equals the velocity; D equals the space between the transducers 14,
16; and T
equals the time the signal takes to propagate (transit time) between the two
transducers 14,
16. One transducer 14 (the sender) generates the input signal and the other
transducer 16
(the receiver) generates the output signal.
Utilizing this pitch-catch method, with the two transducers 14, 16 both
serving the dual function of sender and receiver, measurements of both
upstream and
downstream transit times would be achieved. Blood flow velocity would be
calculated in a
conventional manner using the difference between downstream and upstream
transit times.
Since transit time oscillations resulting from blood flow are magnitudes
greater than transit
time oscillations associated with cyclical pressure changes, these flow
oscillations must be
effectively cancelled out of the calculation by the summation of downstream
and upstream
transit times. The data resulting from this summation would reflect the effect
of pressure
fluctuations on transit times. The summation would magnify the observed
systolic/diastolic
shift in transit times by a factor of two while canceling the effect of blood
flow. Also, this
technique would reduce artifact resulting from body movement and intravascular
turbulence. Mathematically this can be expressed as follows:
Ttotal=(Tdownstream+Tupstream)/2
Factors determining total transit time are (1) acoustic velocity due to blood
density, (2) acoustic velocity as it is influenced by blood flow and artifact
produced by body
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movement and vascular turbulence, and (3) the velocity of the acoustic wave as
it passes
through the surface conductive medium and the skin and subcutaneous tissues.
V. sub. total=V. sub.l+V. sub.2+V. sub.3
V3 (tissue and conductive medium contribution to velocity) will remain
constant. V2 (blood flow and artifact contribution to velocity) can be
readily measured
and canceled out of the equation by summing the velocity in both directions,
thereby
eliminating its contribution to the equation. Therefore, V1 (density
contribution to
velocity) remains as the only variable factor when transit times are measured,
and, within
blood vessels, pressure will be the only density determining factor that
fluctuates on a
moment by moment basis. Vtotal=Vdensity+Vconstant
Therefore: .DELTA.Vtotal=.DELTA.Vdensity
Since density--as it is determined by the momentary values of hematocrit,
salinity, and temperature--remains fixed, then it follows that any momentary
velocity
fluctuations are a result of fluctuations in pressure alone. Thus:
.DELTA.Vdensity=LDELTA.Vpres sure and therefore:
.DELTA.Vtotal=.DELTA.Vpressure
Therefore, any momentary fluctuation in transit times will also be a result of
fluctuations in pressure. These fluctuations in transit times can thus be
expressed
mathematically as: .DELTA.Ttotal=.DELTA.Tpressure and since:
.DELTA.Ttotal=(.DELTA.Tdownstream+.DELTA.Tupstream)/2 then
therefore:
.DELTA.Tpressure=(.DELTA.Tdownstream+.DELTA.Tupstream)/2
Signal processing and adequate sound conduction through the skin and
subcutaneous tissues to and from the structure of interest are critical steps
involved in
ensuring the accuracy and reliability of the device. Factors such as incorrect
device
placement, obesity, and edema will interfere with acoustic conduction and
possibly render
the device ineffective. In an aerospace application or during sports
participation, high G
forces may effectively dislodge the device from its proper position. The use
of bi-directional
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"pitch-catch" transducers will reduce the error resulting from imprecise
device placement
upon the body. This method will also likely reduce artifact from body
movement.
The output signal would appear as an amplitude spike (buried within noise)
that moves to and fro along the instrument's time scale indicating at its
limits a systolic and
diastolic transit time for each cycle. Shorter transit times are associated
with the systolic
pressure and longer transit times with the diastolic pressure. The point along
the larger
transit time scale where these minute pressure-related transit time shifts
will be observed
will drift as blood density drifts due to changing physiologic values such as
hematocrit,
salinity, and temperature. This drifting must be accounted for by the
continuous and precise
monitoring of blood density as it is affected by these varying physiologic
values. This can
be done using the precision acoustic transit time measurements described
earlier, and it is
essential for continuous calibration of the device as base-line drifting of
blood density
occurs.
Ideally, calibration measurements will be made by observing the density of
venous blood while it is under the influence of zero increased vascular
pressure, i.e., at
atmospheric or ambient pressure. As the device operates in its vascular mode,
continuous
calibration and ultra-precision is the core of its design and function. To
summarize, the core
of the device design and function depends upon ultra-precision and continuous
calibration
for changes in temperature and also for changes in transducer separation
distance (if not
fixed by use of a rigid housing). (See page 34.)
According to the UNESCO equation, there is a 98.4 picosecond change in
the transit time across a 10 centimeter distance for every 10 mmHg change in
pressure.
However, the variability of the above mentioned physiologic values may result
in as much
as several hundred nanoseconds of drift in transit times. According to the
UNESCO
equation, when salinity is fixed at 9 psu, a change in body temperature of 1
degree Celsius
would alter transit time by about 75 nanoseconds (a scale that is magnitudes
greater than
transit time shifts resulting from incremental changes in pressure).
Similarly, with
temperature fixed at 37° Celsius, alterations in salinity of only 0.1
psu would result
in a change in transit time of about 4.25 nanoseconds. These values were
calculated using
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velocity data obtained from the UNESCO equation at the National Physical
Laboratories
(NPL) interactive website and using the previously noted common equation,
V=D/T. Such
large scale changes in blood density would of course occur over a period of
hours and not
milliseconds and therefore should not affect momentary pressure readings.
However, if
blood density alterations are not monitored precisely and continuously, then
the minute
fluctuations in transit time related to pressure oscillations would have no
baseline or frame
of reference and would be useful only for pulse detection. Even trend
monitoring would be
difficult as such without a solid frame of reference.
Ideally, in order for the peak and trough (systolic and diastolic) density
values to be meaningful, the measurement of baseline density must be performed
within the
observed fluid or gas when it is under zero increased pressure. In vivo,
however, blood or
other physiologic fluids or gases are rarely without the influence of at least
minimal
pressure. This fact increases the challenge of device calibration. However, it
can be
predicted intuitively that there may be a measurable "zero" or baseline
density that could be
monitored by the device by selectively "capturing" venous system readings
during the
lowest point in the cycle (most likely during inspiration at end-diastole). It
can also be
predicted that there may be a mathematical relationship between peak and
trough arterial
and venous density and flow values and the baseline density value. This
prediction allows
for the potential determination of the baseline calibration density by means
of extrapolation.
Another means of device calibration, much less desirable because of its semi-
invasive
nature, would be the measurement of the density of an in vitro blood sample at
atmospheric
pressure.
The best method for ultrasound (US) or electromagnetic (EM) pulse delivery
and detection must be determined. Potential devices would include conventional
high
frequency ceramic piezoelectric US transducers, RF (radio frequency) US
transducers,
polymer piezoelectric US transducers, IR (infrared) receivers, and Fiber Bragg
Grating
(EBG) Laser receivers, not excluding other existing and/or future transducers
or sensors
which are found to be applicable. All of these devices are referred to herein
as "transducers"
and/or "sensors." The frequency and amplitude chosen for the input signal, as
well as the
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mechanism of its delivery, will depend upon requirements for patient safety
and
requirements for proper tissue penetration and conduction of the acoustic
wave.
The device must be capable of detecting transit time shifts as low as 9.8
picoseconds in order for it to have resolution of 1 mmHg pressure, which would
be ideal for
medical purposes. Medical ultrasound typically operates in the frequency range
of 1 to 10
MHz. This device will likely require a higher frequency acoustic input signal
for accuracy.
However, lower frequencies better penetrate tissues with less attenuation.
Input signal attenuation and penetration varies between tissue types and
according to the frequency. For example, according to Dowsett, Kenny and
Johnston: The
Physics of Diagnostic Imaging, chapters 17, 18; attenuation
coefficient/frequency (dBcm-1
Hz-1) are listed for the following tissue types: muscle: 1.8-3.3; fat: 0.6;
brain: 0.9; blood:
0.2; bone: 20. These variations in signal attenuation can be exploited in
order to enhance the
quality of the output signal, since blood is a better conductor of acoustic
energy and less
prone to signal attenuation when compared to biologic tissues. However, signal
attenuation
is much higher at high frequencies. Nevertheless, there exists enough of a
difference
between its value in blood and tissues that the principle remains the same.
Ideally, the
chosen frequency would attenuate within the skin and subcutaneous tissue
before directly
reaching the receiver yet conduct effectively along the vessel to the
receiver. This would
greatly enhance the signal-to-noise ratio.
Signal input from the sender must consist of brief pulses or "clicks"
generated at specific intervals (e.g., 10 to 100 times per second) in order to
detect all phases
of the pressure cycle. The brevity of the impulse will be important for
precision and will
guide the choice of acoustic energy to be considered for use in the device.
There will likely
be a need to focus the ultrasound beam in such a way as to effectively
maximize the
intravascular acoustic travel distance (the distance that the sound wave
actually travels
within the blood vessel on its path to the receiving transducer). Such
focusing will probably
take the form of a simple transducer array, possibly requiring the use of more
than one
frequency.
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Also for the sake of precision, the operational goal of device 10 is for the
sending transducer to create a focused shock wave "click," and to clock its
transit time
within the blood to the receiving transducer. Since sound waves travel in all
directions from
their point of origin, it would be difficult to know the exact length of the
intravascular sound
wave path. However, fixing the transducer separation by using rigid housing
and using a
technique such as Time-Reversal Minors should define the wave path well enough
to
accomplish the desired level of precision. Maximizing and closely defining the
length of the
sound wave path is a crucial step for the accurate determination of
intravascular sound
speed. Detection and timing of the first arrival wave would indicate the
transit time for the
most direct path between transducers. Since sound speed is higher in blood
than in the
surrounding tissues, then this first arrival wave would be considered to have
passed through
blood.
Referring to FIG. 2, there is shown the device 10 of the invention in a single
transducer variant comprising a transducer 18 applied to the skin of a body
part 12 using an
acoustic conductive medium such as above described. The single transducer
functions as the
sender, generating the acoustic input signal and as the receiver, generating
the output signal.
The single transducer 18 could yield the same data as the two-transducer
method, above
described. This may be accomplished by measurement of the transit time (the
echo) of the
acoustic wave to and from the far side of the specific vessel as the wave
reflects off the
vessel wall interface. Using this method, the angle of the transducer axis to
the vessel is
critical. The transducer must remain as close to perpendicular as possible to
the plane of the
vessel in order to eliminate errors caused by blood flow. As this method may
not be as
precise as the two-transducer method, it may be more useful for trending.
Phase-shift detection could be used as another signal processing technique in
both the single and two-transducer methods to detect transit time shifts in
vessels, chambers,
body cavities and compartments, or airways. Since the velocity of acoustic
transmission
changes with varying pressure, the phase of the reflected or transmitted wave
would shift
proportionately with changes in transit time and therefore would also shift
with changes in
pressure. When using this phase-shift detection technique during vascular
system or static
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fluid compartment measurement, very high frequencies (most likely within the
range of 50
MHz to 7.5 GHz, but not excluding higher or lower frequencies) would be
required in order
to ensure precision. When analyzing the pulmonary or pleural spaces, lower
frequencies
(probably ranging from 100 KHz to 1 MHz, but not excluding higher or lower
frequencies)
would be required.
Such a phase-shift detection technique would not truly utilize the Doppler-
effect in its detection of phase shift. Since there is no flow involved within
static
compartments, then there is no Doppler-effect possible. Within vessels,
however, the desire
is to cancel out any effect of flow and motion artifact. Therefore, while
phase-shift may still
be a measurable quantity, the Doppler-effect would not be applicable in either
the vascular
setting or the static compartment setting.
The single transducer transit time and phase-shift detection methods may be
more suitable for measurements of static compartments or hollow organs within
which flow
is not a significant factor. They may be less suitable for vascular pressure
measurements
where they cannot easily cancel out noise caused by flow.
IR, RF, and/or Laser technology may also be used in transducer design for
the single or two-transducer methods. The arrangement of the sender and
receiver
transducers would be as in FIGS. 1 and 2. Sensor function could be enhanced
with the use
of Laser technology with Fiber Bragg Gratings (FBG's) tuned to a specific US
frequency.
FBG Laser may be especially useful in sensor design due to its capability of
sensing high
frequencies and its resistance to RF interference.
A third type of arrangement for the transducers utilizes three or more
transducers, one sender and two receivers arranged in the order, receiver-
sender-receiver, as
they lay longitudinally over the vessel. Again, these transducers could be of
piezoelectric
design or could use any of the other advanced technology above described. The
two
receiving transducers would clock the US wave front as it passes upstream and
downstream
from the centrally located sending transducer. The velocity values would be
summed in
order to cancel out the effect of blood flow and to separate it from the
effect produced by
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pressure fluctuations. Like the two-transducer technique, this technique--
given very specific
placement of the transducers and chosen frequencies--would also take advantage
of the fact
that the attenuation coefficients of biologic tissues differ from that of
biologic fluids.
However, accuracy would likely not be as precise as with the two-transducer
method since
the wave paths upstream and downstream do not cross the same section of
vessel, and thus
cancellation of turbulence-induced signal variations may not be as effective.
See FIG. 3.
Each of the sensors 14, 16 and 18 and each of the monitoring devices 10 of
the invention illustrated in FIGS. 1-3 are connected to a computer that is
programmed with
recognition and analysis software. Depending upon the function of the sensor
14, 16 or 18,
i.e., whether the sensor is a sender, a receiver, or both, the computer
software will differ as
to each recognition and analysis computer 22 to receive the signal from its
individual sensor
14, 16, 18 and convert the same into a measurement of arterial and venous
blood densities
and blood flow velocities, blood pressure, pulse rate, vascular resistance,
cardiac output,
pressure pulse wave velocity, and the like. The display will include both
instantaneous
measurements and a plot of each measurement versus time.
Scanners 24 are provided to scan each of the computers 22 sequentially from
about 10 to about 100 times per second, depending upon the particular clinical
application.
Each of the scanners would be operatively connected to a display 26 that would
display the
data from each of the sensors 14, 16, 18 of each of the measurements, in the
form of both
instantaneous measurements and the historical trends of each measurement. The
display
would be combined with a selection switch by which each measurement and trend
could be
selectively displayed.
Attached to each display would be a printer 28 which would print out current
vital signs and a continuous record of highest, lowest, and trends of each
measurement, as
well as trends for each patient and each location of a sensor 14, 16, 18.
Each of the sensors 14, 16, 18, each of the recognition and analysis
computers 22, each of the scanners 24, each of the displays 26, and each of
the printers 28
are connected to a power source 30.
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In the single sensor device 10 illustrated in FIG. 2, the sensor 16 is
connected
to a single recognition and analysis computer 22 which is connected directly
to a display 26
and to a printer 28.
Precise measurement of biologic fluid density is the critical step in ensuring
accuracy by way of continuous calibration for the device using any of the
above methods.
Also, such continuous and precise monitoring of blood density would be
extremely useful in
the diagnosis and treatment of trauma patients and any other malady involving
rapid or
profound blood loss or physiologic fluid shifts. See device 100.
Morbid obesity would likely make this device unusable as it would increase
signal attenuation and would therefore make readings very difficult. There
would be a
marked decrease in the signal-to-noise ratio in patients with thick layers of
adipose tissue.
When working within normal physiologic blood pressure ranges, period
calibration of the
device using a conventional sphygmomanometer would solve the problem of
accuracy in
most situations where body habitus interferes with the normal function of the
device.
However, in the non-obese patient, even in cases where calibration is not
possible, such as a
profound hypotension or cardiac arrest, accurate readings may be attainable
with the device.
Another challenge would be the design of a stable transducer-to-skin
interface acoustic conductive medium. The interface must remain fixed in
position for a
number of hours. It must be comfortable to the patient, and provide reliable
ultrasound
conduction.
In summary, the vascular application of the device would be capable of
accurately and continuously measuring arterial and venous blood pressures,
pulse rate,
blood density, and blood flow velocity, and it would be capable of calculating
peripheral
vascular resistance. When the venous system and interstitial space are
monitored, the state
of hydration can be assessed. When applied to the chest, then pulmonary,
central venous,
pleural space and cardiac monitoring may also be possible. The device may have
many
other uses, including the measurement of compartment, ocular, intra-abdominal,
intracranial, and specific organ pressures. In addition, the device could be
mated to other
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more conventional equipment, e.g., measuring oximetry and temperature.
Device 100
Another version of the new and improved noninvasive blood density
measurement device 100 is a simpler form of device 10 of the invention for the
purpose of
supplying in vivo blood density information for medical monitoring and
research. The
function of device 100 is the same as that of device 10, except that ultra-
precision is not
required. As with device 10, the goal with device 100 is the noninvasive in
vivo
measurement of blood density. However, it will not have the necessary
precision to detect
the minute density fluctuations which represent pulse pressure. Therefore the
device
requires somewhat less sophistication.
The primary goal of monitoring blood density is to detect fluid shifts within
the body. Also, because hematocrit is the main contributor to the density of
whole blood,
then both device 10 and device 100 are continuous noninvasive hematocrit
monitors. They
could, therefore, both be used to monitor multiple parameters in critically
ill or injured
patients or be used to spot check patients for blood disorders.
Within the practice of nephrology, the monitoring of blood density during
dialysis is well known to be important as it is used to predict and preempt
sudden onset of
hypotension. However, the relevance of blood density values and trends as they
relate to the
status of critically ill or injured and potentially unstable "critical"
patients is not well
known. Under the current state of the art in blood density measurement,
comprehensive
research on the clinical relevance of blood density is not possible. Device
100 is needed so
that such clinical relevance, or lack thereof, can be discovered.
Currently the state of the art in blood density measurement is practiced using
only extracorporal methods. One method uses frequent blood sampling and
subsequent
laboratory analysis. Another less precise method uses a continuous optical
device during
hemodialysis which clamps onto the dialysis tubing and measures the
concentration of the
extracorporal blood. Continuous blood density monitoring is currently
unavailable for
patients who are not undergoing hemodialysis. Frequent blood sampling,
although precise,
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is labor intensive, expensive, and impractical. There is a need for a tool
such as device 100
which measures blood density conveniently, continuously, noninvasively, and in
vivo. In
addition, device 100 may be capable of providing continuous data relating to
arterial and
venous blood flow velocities, extravascular fluid stores, and analogs of
vascular resistance
and cardiac output.
This description of device 100 is also an addendum to the description of
device 10. Most of the technical aspects of device 10 are identical to that of
device 100, and
therefore, the text of this description applies fully to that of device 10.
FIGS. 4 and 5
illustrate one possible form of device 100 and device 10 of the invention.
Although the
basic function of the two devices is almost identical, the goals of precision
and clinical
application differ and would dictate certain technical variations.
Physiological Basis for Measurement of Blood Density in Critical Patients
During severe physiological stresses there are significant alterations in
blood
density as the blood becomes more concentrated or more dilute. These
alterations occur as a
result of transcompartmental fluid shifts that, in turn, are caused by
physiologic
compensations in for form of osmotic or hydrostatic effects. Significant fluid
shifts--and
thus dynamic changes in blood density--occur during shock from any cause
including
hemorrhage, sepsis, spinal cord injury, toxins, and cardiogenic causes. Less
significant, but
still noteworthy, are fluid shifts and blood density changes that occur during
more ordinary
clinical situations such as orthostasis (1), dehydration and rehydration,
various
pharmacological therapies, and weightlessness. Therefore, the measurement of
blood
density and its trends may thus become an important tool for ruling out
certain causes of
syncope and dizziness.
When the condition of low intravascular volume or pressure occurs,
physiologic compensations are triggered in an attempt to maintain blood volume
and
thereby blood flow to the vital organs. Under these conditions the vascular
system
osmotically draws fluid into the blood from the extravascular space. This
results in dilution
of the blood and a drop in blood density (hemodilution). For example,
hemorrhage results in
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rapid fluid movement from the extravascular to the intravascular space and
thereby causes
hemodilution. (2) The vascular system is, in effect, "borrowing" fluid from
the tissues in
order to preserve blood volume and flow.
Similarly, an infusion of IV fluids will initially cause hemodilution and a
drop in blood density. However, the dilution from IV fluids will not persist
if the blood
volume and osmotic and hydrostatic pressures remain adequate, because the
fluid will
eventually migrate from the blood to the extravascular space (if it is not
first lost through
renal excretion or other insensible losses). The vascular system thus gives
back fluid that it
may have once "borrowed" from the tissues, and the blood density or
concentration will
drift back toward a more normal range.
Certain catastrophic vascular effects are triggered by prolonged or severe
shock, sepsis, burns, crush injuries, and toxins. The result of these effects
is capillary
damage and leak. This leaking causes fluid to shift from the blood to the
extravascular
space, and thus results in a blood volume decrease accompanied by a blood
density increase
(hemoconcentration). Since diuretics and blood transfusions effectively cause
hemoconcentration, and IV fluids cause hemodilution, then device 100 could
also become a
useful tool in the monitoring and management these types of therapeutic
interventions.
The clinical course for severely ill or injured patients is typically very
dynamic. As the disease process takes its course, the physician then responds
with
aggressive treatment using surgical techniques, vasoactive drugs, IV fluids,
and blood
products. Multiple events take place in rapid sequence, and each has its own
effect upon
blood density. In such dynamic situations, the interpretation of blood density
values and
trends would be complex. Clinical studies must be done in order to define the
parameters
for use of device 100.
Since blood density is already known to be an important indictor to observe
during hemodialysis, it is reasonable to assume that there is also a
relationship between the
clinical course experienced by the critical patient and the magnitude and
rapidity of the
changes in blood density. It is likewise reasonable to assume that there are
physiological
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limits to blood density and that high and low extremes are not compatible with
life.
Therefore, blood density monitoring may be as important in the management of
any critical
patient as it is in the management of a hemodialysis patient.
The density of blood is determined by multiple factors, the most influential
of which are the hematocrit and serum protein level. Other determining factors
are pressure,
temperatures, and dissolved sugars, salts, and gases. Since device 100 is not
designed for
ultra-precision as is device 10, the blood pressure contribution to density
values will be
negligible. For purposes of simplicity during earthbound research, pressure
would be
assumed to be fixed at one atmosphere. Also, since it has a significant effect
upon density, a
calibration or correction for temperature must be accomplished.
The device would be useful in any medical setting where care is provided for
critical patients. This would include such settings as emergency departments,
intensive care
units, surgical or post-operative areas, burn units, hemodialysis units,
military mash units,
and battlefield settings. It would also be useful in aerospace research
settings. There are
significant transcompartmental fluid shifts that occur in the zero gravity
environment. (3, 4)
In order to ultimately find its niche among the tools of clinical medicine,
the
device must initially serve the purpose of research tool. To date it appears
that there has
been very little research done and in the area of blood density and its
correlation to the
clinical status of critical patients. However, the few studies that have
surfaced do seem to
indicate that such monitoring would probably be useful, provided that
convenient method of
measurement is available. There is currently no existing device for monitoring
blood density
in vivo and noninvasively on a continuous basis.
Device Function--Another Version
The function of device 100 depends upon in vivo sound speed measurement
within blood. Its basic function is identical to that of device 10, but with
less required
precision. Acoustic transit time (time-of-flight) is measured using the "pitch-
catch" method
between proximal and distal transducers in both upstream and downstream
directions
simultaneously. Another option would to replace one transducer with a
reflector. The
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remaining transducer would act as both sender and receiver, and would clock
total time-of-
flight. Either device would be applied to the body wherever the best arterial
and venous
signal can be obtained. These time-of-flight measurements would be then be
used to
calculate acoustic velocities upstream and downstream. Readings would be taken
from both
arterial and venous blood and then the data from each or averaged data from
both would be
fed into a microprocessor where it would be converted to blood density and
flow values.
When using the time-of-flight method to obtain blood density readings, the
effect of blood flow velocity upon time-of-flight must be negated. Although
the condition of
"static" blood is not possible under most in vivo circumstances, a correction
can be made
for the effect of flow upon the measurements. BY using the upstream and
downstream
velocity values, a calculated "static equivalent" acoustic velocity can be
obtained. The effect
of flow is thus cancelled out mathematically by dividing the sum of the
upstream and
downstream acoustic velocities by two. The result is the acoustic velocity
equivalent as if it
were measured within static blood.
In practical terms it may be very difficult to separate the arterial and
venous
signals since, in the circulatory system; the arteries and veins are usually
paired and located
in close proximity to one another. The signal would thus be essentially
already averaged
along with the portion of the signal attributable to capillary blood.
Based upon the UNESCO equation for sound speed in water, this resulting
"static equivalent" acoustic velocity value has a direct mathematical
correlation with blood
density as discussed in the description of device 10. The UNESCO equation--
which was
developed for the study of sound speed in seawater--can be used to
substantiate this method,
since blood, like seawater, is simply water with certain substances in
solution or suspension.
Blood flow velocity would be calculated by subtracting the upstream from
the downstream acoustic velocity and dividing by 2. Both venous and arterial
flow
velocities can be calculated in this fashion. This time-of-flight method of
blood flow
velocity determination differs from that of the Doppler method, since its
operation depends
upon sound-speed measurements and not phase-shift data.
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Since the device senses blood flow in both directions, it would be able to
differentiate venous from arterial flow. In terms of practical function, the
fact that it is
sensing blood flow assures that the device is indeed reading the density
within intravascular
fluid and not within extravascular fluid.
Interestingly, device 100 may also be capable of providing a measurement of
acoustic velocity within the extravascular space. This would be useful in
monitoring the
body's stores of extravascular fluid, i.e., hydration status. Sound speed
should change
proportionately with the fluid "saturation" of the extravascular tissue.
Another method of assessing the dynamic status of the extravascular fluid
would be to monitor the difference between arterial and venous blood density
values. This
may present a challenge if the arterial venous blood density measurements can
not be
distinguished one from the other, but the data obtained from this method would
contain
relevant information relating to the movement of fluid into and out of the
extravascular
space. For example, when the arterial blood density is found to be higher than
the venous
blood density, the conclusion might be made that physiologic compensation is
underway
and that fluid is moving from the extravascular space into the blood, such as
might be seen
during blood loss. This differential arterial-venous blood density value would
probably be
detectable prior to any significant alteration in whole blood density and
would serve as a
very sensitive and contemporaneous real time gauge of intravascular-
extravascular fluid
movement.
By applying existing state of the art methods of measure, this device could
also sense pulse-wave velocity by clocking the pulse-wave as it passes by the
two
transducers. Pulse-wave velocity is the speed of the arterial pressure wave as
it propagates
from the heart to the peripheral tissues. It can be used to estimate cardiac
output when
adjusted for patient age. An analog of cardiac output could most likely be
calculated from
mean arterial blood flow and pulse-wave velocity. An analog of vascular
resistance could
also theoretically be calculated if mean arterial and venous pressures and
blood flow are
known. The actual analog values of cardiac output and vascular resistance
obtained in this
manner would be useful only for trending. (5)
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Another already existing method of measuring pulse-wave velocity involves
mating the device with an electrocardiogram (EKG) electrode and measuring the
time from
the QRS impulse to the arrival of the pulse-wave at the device. Also, contour
analysis of the
pulse pressure wave is a method currently being used to estimate cardiac
output. These
methods seem to be gaining some validity within the research literature as
being reliable for
trend monitoring of cardiac output. (6)
Like device 10, device 100 would digitally display the instantaneous arterial
and venous blood density and blood flow velocities, the analogs of cardiac
output and
vascular resistance, and their trend lines as well as rate of change.
The acoustic frequency range to be used would vary with the desired
separation distance between transducers. Both are yet to be determined. As
with device 10,
the electronics required would include a device driver 20 that controls each
transducer by
triggering impulses at a rate between 10 Hz and 100 Hz. Each transducer would
act as both
sender and receiver and would be connected to a signal processing computers 22
that would
note time-of-flight for each impulse and then calculate the dimensions and
location of the
vessels and arterial and venous blood density and flow velocities. The
receivers would also
send pulse-wave velocity data to the computers 22 for signal processing and
calculation of
vascular resistance and cardiac output analogs. Scanners 24 would collect data
from the
computers and transmit it to the display 26 and printer 28. A power source 50
would be
connected to all components.
The housing for the transducers (FIG. 4), which would most likely be
constructed from medical grade epoxy or silastic, would fix the distance
between the two
transducers and set the proper angle to the skin in order to create the most
optimal signal for
processing. The requirements for housing style might differ depending upon the
anatomic
location to be monitored. For use on the arm, for example, the housing must be
as flat as is
practical and secured in some fashion. Testing must yet be performed in order
to find the
most optimal transducer separation, angle, and frequencies, as well as the
best type of
acoustic conductive medium, arm band style, etc.
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There would possibly be a need to tune (power, transducer wavelength,
transducer triggering frequency, and beam spread or scatter) the ultrasound
beam in such a
way as to effectively maximize the intravascular acoustic travel distance (the
distance that
the sound wave actually travels within the blood vessel on its path to the
receiving
transducer). Such tuning might take the form of a transducer array possibly
requiring the use
of more than one frequency.
As with device 10, and for the sake of precision, the operational goal of
device 100 is for the sending transducer to create a shock wave "click," and
to clock its
transit time within the blood to the receiving transducer. Since sound waves
travel in all
directions from their point of origin, it would be difficult to know the exact
length of the
intravascular sound wave path. However, fixing the transducer separation by
using rigid
housing and using a technique such as Time-Reversal Minors should define the
wave path
well enough to accomplish the desired level of precision. Maximizing and
closely defining
the length of the sound wave path would be a crucial step for the accurate
determination of
intravascular sound speed. Detection and timing of the first arrival wave
would indicate the
transit time for the most direct path between transducers. Since sound speed
is higher in
blood than in the surrounding tissues, then this first arrival wave would be
considered to
have passed through blood.
Therefore, both sending and receiving transducers may need to be tuned
(focused or defocused) to maximize the path within the blood. Tuning may need
to be
individualized for each patient and for each anatomic location. The induced
intravascular
"click" would emanate in all directions from its point of origin within the
vessel. Much of
the resulting wave energy would then travel through the blood (arteries,
veins, and
capillaries) and be detected by the receiving transducer. The effect of this
tuned shock-wave
technique would be to maximize the intravascular travel distance of the
acoustic wave. It
would also improve signal-to-noise ratio for the time-of-flight impulse.
The remainder of the wave energy would reflect and/or scatter and would
ultimately also be detected by the receivers. The data collected from the
reflections would
reveal the status of the extravascular fluid balance, and would also be used
to determine the
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size and location and the selected vessel. This data would in turn be fed into
the signal
processors and could be used to guide the tuning of the ultrasound beam. By
applying signal
processing techniques, the true intravascular signal would be separated from
the signals
resulting from reflections and scatter.
For device 10 this technique of tuned input and detection would likewise be
applied for the purpose of signal enhancement and precision, even though input
frequencies
and transducer types might differ.
The type of transducer to be used in device 100 is also yet to be determined.
Common piezoceramic transducers may work well, but in order to assure the
proper brevity
of the input impulse "click", other types of transducers may be required,
including but not
limited to, polymer, piezo, laser, infrared, radio frequency, Fiber-Bragg
laser receivers, and
hybrid transducers.
Device Function--Still Another Version of Device 100
Although sound velocity measurements would ideally be made
longitudinally through a vessel, in practical terms this might be very
difficult to do non-
invasively. Sound waves prefer to travel along straight paths. Therefore the
most practical
body location chosen may involve the acoustic wave crossing a vessel
perpendicularly. If
high enough frequencies are used (1 to 20 MHz), then the desired precision
might still be
accomplished. The vessels adjacent to the external ear may be amendable to
such a
monitoring method. Other locations such as the perioral and post-auricular
arteries might
also be used.
A specific example is the superficial temporal vessels, located just anterior
to
the tragus of the external ear. In this case the vessel may be monitored by
placing the
sending transducer against the anterior wall of the external ear canal just
behind the tragus.
The receiving transducer would then be placed against the skin anterior and
superior to the
tragus, positioning the vessel between the two transducers. The sender would
emit its
impulse directly towards the receiver across the vessel. Continuous time-of-
flight
measurements would be continuously corrected for temperature changes and
converted into
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blood density values. If the separation between transducers can be maximized,
and if
frequencies used are high enough (500 KHz to 100 MHz), then desired precision
might be
accomplished. See FIG. 5. Temperature correction may be accomplished by
incorporating a
temperature probe into the device just adjacent to the transducer or with the
use of a dual-
mode oscillator crystal, which has the characteristic of self-temperature-
sensing.
In order to maintain accuracy of sound velocity readings and thus blood
density measurements, the device would incorporate a temperature probe and
also
electronics for continuously monitoring the mechanical separation between
transducers.
Since movements such as chewing and talking may change the separation of the
transducers, the device could instantaneously adjust for the change and
recalculate sound
speed based upon the new distance. In all practicality, the separation between
the
transducers or transducer and reflector should be fixed by using a rigid
housing or include a
monitoring mechanism for measuring changes in the separation distance between
the
transducers or reflector. This would apply to both device 10 and device 100.
Also, with
either device, it may be relatively easy to incorporate an oximeter. The
application of a
small array at the receiver may provide vessel diameter information and thus
an analog of
pulse pressure.
In summary, device 100 may provide the following continuous data from a
stable platform overlying a vessel near the ear: [0109] 1. Blood density
[0110] 2.
Temperature [0111] 3. Oximetry [0112] 4. Pulse rate [0113] 5. Pulse pressure
analog
Another device 120 of the invention is an intravascular catheter that contains
sensors for in situ measurements of multiple whole blood physiologic and
hemodynamic
parameters. It would be a unique and useful diagnostic tool that would
facilitate rapid
decision-making and optimization of medical management and resuscitation of
patients
suffering from the most challenging and life-threatening medical conditions.
The
instantaneous monitoring by the application of the invention would provide a
means by
which to immediately assess the effectiveness of fluid and blood
administration as well as
pharmaceutical or surgical interventions. Currently there exists a significant
lag-time
between the initial recognition of a problem, such as sudden hypotension, and
the obtaining
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of lab results that may or may not confirm clinical suspicions of the problem
cause(s).
The physiologic and hemodynamic data obtained with the use of device 120
would allow the clinician to more quickly differentiate the causes of
hypotension from any
cause or combination of causes. Intravascular monitoring would provide
information useful
in determining the cause of shock, e.g., sepsis, blood loss, and autonomic
malfunction such
as that caused by spinal cord injury. For the clinician, knowing the level of
vascular
resistance and vascular fullness would also improve the ability not only to
differentiate the
causes of abnormal vital signs, but also to determine the best course of
treatment in each
circumstance. Such treatment may include infusion of saline, blood, both
saline and blood,
or pressor medications; the performance of specific surgical inventions,
appropriate airway
management, or fracture stabilization; or, in some cases, simply the
administration of pain
medications. Continuous monitoring of this broad range of physiologic and
hemodynamic
parameters would also provide immediate feedback on the effectiveness of the
above
interventions.
Primarily, the device is intended for peripheral vascular (venous or arterial)
use in prehospital, emergency, surgical, post-surgical, burn unit, and ICU
environments.
The invention may be useful in any other medical patient or research
environment where it
might be desirable to measure and/or trend the above stated physiologic and
hemodynamic
values. These include but are not limited to military, aerospace, and
subsurface marine
environments. Other clinical or research applications may include the trending
of any
biologic fluid within any body space, whether it be peripheral or central
vascular or fluid
filled body cavity. Examples of confined body cavities would include the
urinary bladder,
gallbladder, intra-abdominal, brain ventricles or spinal fluid. The device
could possibly be
used to monitor the status of the above parameters within extremity fascial
compartments or
post-op plastic surgery skin flaps. It may be especially useful to also
monitor pH or lactic
acid level within a compartment where there is potential muscle necrosis to
determine the
need for and improve the timeliness of rapid surgical intervention. In
addition to vascular or
compartment pressure readings, the catheter could potentially be used to
measure the
pressure in gas-filled structures such as pulmonary airways and bowel lumen.
It would be
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useful in the clinician's office for routine on-the-spot laboratory analysis.
When the catheter device is placed intraarterially, systolic, diastolic and
mean blood pressures, pulse pressure, pulse rate, and the incorporated
physiologic
parameters would be displayed continuously. This type of information is
routinely used in
evaluating critical care patients. In addition, local blood flow velocity and
calculated local
vascular resistance (LVR) would also be displayed continuously. Ideally the
clinician would
want to know the level of total systemic vascular resistance (SVR) which
relates to total
sympathetic vascular tone. But SVR cannot be obtained from a peripheral
catheter. SVR,
which is commonly calculated in physiology research, is an important factor in
determining
the workload of the heart and how the vascular system is acclimating to
various insults,
such as trauma or infection. Although one cannot determine cardiac workload
from LVR, it
is theorized that LVR trends would parallel SVR and therefore be useful for
observing
changes in cardiac workload and sympathetic tone. LVR cannot be measured by
any blood
test and is not currently being measured for critical care resuscitation
purposes in any
clinical realm.
Since blood density is being precisely monitored, vascular volume can be
determined by injecting a small volume of IV fluid (saline) and observing the
change in
blood density. Vascular volume is then calculated by the common dilution
formula,
BV=Vi*.rho.1/.DELTA..rho., where BV is blood volume, Vi is the volume of
saline
injected, .rho. is blood density and .DELTA. .rho. is the change in blood
density observed as
a result of the given fluid injection. Cardiac output can therefore also be
calculated by the
formula CO=BV/T, where CO is cardiac output, BV is the above calculated blood
volume,
and T is the time that it takes for the change in blood density to occur after
administering a
given volume of saline (the time that it takes for the heart to circulate the
saline throughout
the vascular system one time). Previously, information about vascular tone,
vascular
volume, and cardiac output has been confined to the realm of medical research
and
unavailable for routine use in clinical medicine.
Venous monitoring with the present invention would yield the same valuable
blood density data as with arterial monitoring. Device 120 would also supply
venous
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physiologic and hemodynamic parameters. In some circumstances, it may be
valuable to
perform trending of venous physiologic data and hemodynamic data. Venous
placement of
the catheter would not require special expertise and could be initiated by
paramedics in the
prehospital environment. Like the arterial catheter, the venous catheter would
yield useful
data continuously.
Standard arterial pressure monitors, because of the hydraulic transducer, can
be clumsy due to the tubing and pressure bag. They require time consuming
calibration.
They do not provide information on blood concentration, blood flow velocity,
temperature,
blood gases, pH, oximetry or lactic acid level. Device 120's pressure sensor
could be
quickly calibrated to atmospheric pressure prior to insertion and thereafter
would
continuously self-calibrate for temperature changes. It would not require
calibration for
blood concentration or blood flow readings.
A. Blood Concentration and Hematocrit Measurement by Acoustic Methods:
It is known that sound speed in whole blood is determined by total protein
concentration11, the concentration of ions, and the temperature12.
The majority of
protein in whole blood resides within the hemoglobin of the red blood cells
(RBC's)10,
11. Therefore, not only can sound speed be used to very precisely measure
blood density, it
can also be used to measure the relative number and hemoglobin content of the
RBC's.
The function of the device 120's blood concentration monitor depends upon
the measurement of several properties of whole blood acoustic propagation.
These include
ultrasound velocity, attenuation, and backscatter measurements.
1) Ultrasound Velocity: By measuring high frequency sound wave time-of-
flight, velocity can be calculated very precisely9. Therefore, sound
velocity in whole
blood can be used to accurately measure and trend blood density.11, 12
The speed of sound in any liquid is also influenced by temperature13,
16. Therefore, if blood density is to be measured by the sound-speed
technique, then it is
imperative that temperature be taken into account. The primary determining
factors for
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blood density are the total blood protein content and salinity. Within the
human physiologic
range, however, changes in salinity do not significantly affect sound
speed11, 13.
Whole blood consists of particles (cells, platelets, and insoluble plasma
proteins) suspended within a water solution containing primarily salt ions,
sugar molecules,
and soluble proteins. The particles in suspension do not necessarily possess
the same
density as the solution itself. Therefore, the velocity of sound across a
sample of whole
blood is determined by the sum of the velocities of the blood components
within the path of
the sound wave.11
Most blood protein resides within the red blood cells in the form of the iron
containing protein hemoglobin. As a result, changes in sound velocity are
directly
proportional to changes in the concentration of hemoglobin within the sound
wave
path.10, 11 Velocity measurement is, therefore, an acceptable method of
total
hemoglobin estimation.7, 11, 12 Since the hemoglobin (Hgb) resides within
the red
blood cells (RBC's), then sound speed is also proportional to hematocrit
(Hct), the
percentage of whole blood consisting of RBC's.10 For the purposes of the
design of
device 120, Hgb and Hct are so closely linked in terms of their relationship
to sound speed
that they will be considered one and the same and will be referred to
heretofore as H/H.
When serum protein levels are unusually high or low, however, an error in
H/H estimation may be produced in the sound speed method.11 For example,
a high
serum protein level would result in overestimation of H/H, and a low serum
protein level
would result in underestimation of H/H. Nevertheless, methods exist for very
precise
measurement of sound speed9 and, within the normal human range of
H/H11,
sound speed has high overall correlation to H/H. In the very low H/H range,
however,
anomalies in serum protein level produce proportionately larger errors in H/H
measurement
by the sound speed method. Therefore, in order to create a device that is
accurate in all H/H
ranges, it would be desirable to primarily use sound speed for H/H measurement
and, to
correct the results obtained in the lower range of H/H.
Others have shown that sound speed11, sound attenuation15, and
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backscatter amplitude14, 15 each have their own individual relationship
to H/H within
certain ranges of H/H. Because each individual parameter seems to possess a
different level
of accuracy within different ranges of H/H, it can be theorized that each
parameter might be
used to correct the measurement results of the other parameters. In the work
for the above
referenced patent, averaging of the parameters was used to achieve modest
improvement in
the accuracy of H/H measurement. It is an object of the current invention to
enhance device
accuracy by developing specific mathematical algorithms that better describe
the influence
of each parameter upon the instrument's accuracy.
2) The Relationship Between the Accuracy of the Sound Speed Measurement
of H/H and the Serum Protein Content:
An in vitro study of random anonymous whole blood samples (n=30) to
determine the correlation of hematocrit with sound speed, correlation was
found to be high
(R=0.93) in the overall population. Best fit was performed in this population
of whole blood
samples, and it was found by regression analysis that the relationship between
sound speed
c and Hct can be described as follows: Hct=1.0135(c-1520)+8.8874 By further
regression
analysis (R=0.85), it was found that the percent of error in Hct measurement
produced by
variations in serum protein level (SP) can be described by the following
equation: % error=-
30.354Ln*(HCT/SP)+47.47
3) Continuous Precision Sound Speed Measurement:
The intravascular device would continuously measure acoustic transit times
in order to calculate sound speed in whole blood, and thereby measure whole
blood density.
By continuously measuring transit times and minute shifts in transit times in
rapid sequence
(10 to 1000 Hz), such measurements, if made with precision, would result in
accurate,
reliable, and continuous blood density and H/H data. FIG. 1 illustrates the
sound speed
sensor as a single transducer-reflector pair with the reflector fixed at a
specific distance,
e.g., 5 mm. The transducer acts as both sender and receiver and is connected
to a power
source, recognition and analysis hardware and software, display, and printer.
The particular hardware that runs the sound speed sensor will depend upon
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available high fidelity equipment. An example of available equipment is that
developed by
Dr. Craig Hartley PhD, which is capable of detecting transit time of an
ultrasound wave by
detection of 1° of phase change9. Using a 20 MHz ultrasound
frequency, the
method has been shown to be capable of detecting a 139 psec increment change
in the time
of sound wave arrival. Assuming average sound velocity in the human body to be
1540 m/s,
the time increment of 139 psec time increment translates to a velocity
increment of 0.0329
m/s between fixed transducers. From the above referenced NPC in vitro study,
it was
discovered that a sound velocity increment of 0.0329 m/s translates into a
hematocrit
increment of 0.0333%. This level of precision would be excessive for the
purposes of blood
density measurement. However, such precision could potentially be used in
making pressure
measurements (see the below description of a novel method of pressure
measurement).
Sound speed changes that occur as a result of pressure changes are extremely
minute, and
would not alter measurements done for the purpose of hematocrit monitoring.
Temperature has a significant effect upon sound speed in liquid13.
Device 120 will include a probe which will allow for continuous calibration
for temperature
changes. A change in temperature of 1° Celsius produces an error of 7.5
nanoseconds
in time-of-flight across a sound wave path 10 millimeters in length. Therefore
the
conceptualized multiparameter catheter must have appropriate temperature
compensation.
This will most likely take the form of an off-the-shelf precision temperature
probe
embedded in the catheter tip, such as a thermistor or other miniature sensor.
An alternative
and unique method of precision temperature measurement utilizes an ultrasound
technique
called Dual-mode Oscillation. The actual method used will depend upon its ease
of
incorporation into the catheter and its accuracy and precision. Measurement of
the
temperature to within ±0.1° Celsius would probably be adequate for
purposes of
measuring H/H. It would provide a resolution of ±0.28% in hematocrit
determination.
4) Attenuation: Whole blood protein, whether contained within cells or
suspended in the plasma, absorbs sound. Therefore, as a sound wave travels
through a
sample of whole blood, the amplitude of the wave attenuates over a given
distance
depending upon the quantity of protein through which it passes. Most blood
protein resides
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within the hemoglobin molecule. Cell membranes also have significant effect
upon
attenuation because of their sound absorption and sound wave scattering
effect15.
Device accuracy for Hct estimation, therefore, may be enhanced by the
mathematical
incorporation of attenuation values. This can be accomplished by measuring the
amplitude
of the received wave at the sound speed measurement transducer.
5) Doppler signal backscatter measurement. A significant proportion of
ultrasound wave energy reflects off of cell membranes. The Doppler method
applies this
physical characteristic of sound to detect movement (flow) of whole blood
within vessels,
utilizing a change in pitch of the sound reflected off of the cells. When the
cells are struck
by the sound wave, they reflect and scatter the sound. The amplitude of the
reflected
Doppler signal is related to the number of cells within the sound field that
have reflected the
sound waves back to the transducer. Although the Doppler backscatter method
may not be
as amenable to precision measurement as sound speed, the backscatter
coefficient (BSC)
may be more representative of the actual number of cells present within the
sound path at
low Hct levels of <10%14. Interestingly, it is at such very low Hct
levels that sound
speed is least accurate. Therefore, measurement of backscatter level may
provide a way to
mathematically correct for inherent sound speed method errors in H/H
estimation that occur
secondary to anomalous serum protein levels when hematocrit is <10%. This can
be
accomplished using the same transducer as for sound speed measurement.
B. Arterial Blood Pressure Measurement
The parameters of systolic, diastolic, and mean blood pressure (and also
venous pressure) could be incorporated into device 120 using several methods.
Two existing
technologies that could easily be incorporated into the catheter are:
1) The most common method currently employed in clinical practice, the
traditional hydraulic pressure line with ex-vivo transducer, could easily be
used because the
catheter will be designed with a port. In this case no special adaptations
would have to be
made to the catheter.
2) A second type of intravascular pressure monitor involves a miniature solid
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state transducer within the tip of the catheter. This type of sensor is used
for cardiac
catheterizations and vascular research. The intravascular transducer would
have the
advantage of compactness with no need for the cumbersome hydraulic line and ex-
vivo
transducer. It could be calibrated prior to insertion. The disadvantage of the
intravascular
pressure transducer is the fact that position (if it is above or below the
level of the heart)
will affect the accuracy of the readings. But if the patient remains supine,
however, the
effect is minimal, producing only about a 2 mmHg change in pressure for every
inch of
elevation change. Position would not change pulse pressure readings.
3) Two novel methods for blood pressure measurement by sound speed
method are:
a) As discussed above in the section on blood density measurement, a
novel method for pulse pressure measurement involves precise continuous sound
speed
measurement. Theoretically, if changes in sound speed could be made with
extreme
precision (within incremental resolution of less than 2-4 picoseconds), then
pulse pressure
could be monitored. Due to calibration issues, systolic and diastolic
pressures could
probably not be measured. According to the UNESCO equation for sound speed in
sea
water13, there are minute changes in sound speed due to pressure changes.
The
advantage to using this method is that the data could be obtained using the
same sound
speed detection transducer(s) used for H/H measurement. A disadvantage to this
method is
the difficulty in obtaining the required picoseconds precision. Even if such
precision could
be obtained, it would be a very difficult matter to achieve accurate systolic
and diastolic
pressure measurements with this technique because of the fact that calibration
to zero would
be very difficult since the base-line would drift with changes in hematocrit,
temperature,
salinity, and changes in the elevation of the catheter-tip versus the level of
the heart.
The UNESCO equation describes the relationship between acoustic wave
velocity and pressure in water. The equation also takes into account other
factors that
contribute to alterations in sound speed, such as the salinity and the
temperature. Although
fluids (blood included) are considered incompressible, the equation shows that
there should
be minute but measurable changes in velocity associated with changes in
pressure, even
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within the human blood pressure range of 0 to 300 mmHg.
To form theoretical support for the sound speed method of pulse pressure
measurement, the space between two hypothetical transducers was assumed to be
2 cm (or 1
cm. between one transducer and a reflector). The UNESCO equation was then used
to
calculate acoustic velocity at pressure increments of 1 mmHg assuming constant
fluid
temperature at 37° Celsius, constant salinity at 9 ppm, and variable
pressure is
expressed in kPa. Calculations using the formula V=D/T indicate that, in order
for the
device to have precision to within 1 mmHg, it must be capable of detecting
shifts in transit
time of roughly 2-4 picoseconds. Trending of the pressure within ±5 mmHg
could be
achieved by the detection of shifts of approximately 20 picoseconds across a
sound path of
2 cm.
The method can be accomplished using either two transducers or one
transducer with a reflector. When using two transducers, one must account for
blood flow
velocity (which ranges from approximately 0.2-1.0 m/s) and its effect upon
time-of-flight
between the transducers. The preferable method, therefore, would be to utilize
a single
transducer-reflector pair. This method would have the advantage of
automatically
eliminating the effect of flow and of doubling the length of the flight path
(and thus
doubling precision) without increasing the length of catheter.
According to the UNESCO equation, there is a 19.68 picoseconds change in
the transit time across a 2 centimeter distance for every 10-mmHg change in
pressure. The
variability of the above mentioned physiologic values, however, may result in
a much larger
drift in transit times of several hundred nanoseconds of. According to the
UNESCO
equation, when salinity is fixed at 9 ppm, a change in body temperature of 1
degree Celsius
would alter transit time by about 15 nanoseconds (a scale that is roughly
three magnitudes
greater than transit time shifts resulting from incremental changes in
pressure). Similarly,
with temperature fixed at 37° Celsius, alterations in salinity of only
0.1 ppm would
result in a change in transit time of about 0.85 nanoseconds. These values
were calculated
using velocity data obtained from the UNESCO equation at the National Physical
Laboratories (NPL) interactive website16 and using the previously noted
common
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equation, V=D/T. Such large-scale changes in blood density would of course
occur over a
period of minutes to hours and not milliseconds and therefore would not affect
momentary
pulse-pressure readings.
The best method for ultrasound (US) pulse delivery and detection must be
determined. Potential devices would include conventional high frequency
ceramic
piezoelectric US transducers, RF (radio frequency) and Laser Optical-acoustic
US
transducers, polymer piezoelectric US transducers, IR (infrared) receivers,
and Fiber Bragg
Grating (FBG) Laser receivers, not excluding other existing and/or future
transducers or
sensors which are found to be applicable. All of these devices are referred to
herein as
"transducers" and/or "sensors." The frequency and amplitude chosen for the
input signal, as
well as the mechanism of its delivery, will depend upon requirements for
patient safety and
requirements to achieve proper blood penetration and conduction of the
acoustic wave.
This device will likely require a frequency greater than 20 MHz for
precision. In reality the attenuation of the signal at the higher frequencies
may prohibit
achievement of the desired level of resolution. Calculations using the UNESCO
equation
show that, even at a frequency of 100 MHz and a 2 cm transducer separation,
the best
resolution that can be achieved using current technology is only ±15 mmHg.
At
frequencies higher than 100 MHz, attenuation would most likely prevent
practical
transducer/reflector separation of more than a centimeter. However, at 20 Mhz
frequency,
resolution would be more than adequate for trending of blood density and blood
flow.
Future technologies could potentially arise that would make this method usable
for systolic
and diastolic blood pressure measurement, and not just for pulse pressure
measurement.
Time-of-flight measurement would be made by the most precise method
available. This could include such instrumentation as is available for
standard
sonomicrometry equipment. Phase-shift detection could be the most optimal
signal
processing technique because of its extreme precision. A sonomicrometer
technique
developed by Dr. Craig J. Hartley can detect changes in phase of as small as 1
degree of
arc9. At 20 MHz this sonomicrometer can detect movement in a mouse
carotid artery
as minute as 1 micron.ref This translates to a time resolution of 139
picoseconds.
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b) Another novel sound speed method that could be used involves the
incorporation into the catheter of a small chamber filled with an inert gas.
As with the above
H/H measurement by sound speed method, sound speed could be measured either by
using
the two-transducer (pitch catch) technique or utilizing the preferred
transducer-reflector
pair. Transit time measurements must be performed continuously at a rate of at
least 30
cycles per second in order to detect all phases of systole and diastole. The
zero pressure
could be easily calibrated to atmospheric pressure prior to vascular
insertion. Continuous
temperature calibration would also be required, but is already being done
continuously for
the H/H monitor. This method could be more accurate and precise than solid
state pressure
sensors.
An advantage to using this method may stem from the fact that the sound
speed analysis equipment is already incorporated into the system for H/H
measurement.
Applying this technique would require the addition of another transducer and
reflector pair
with the same electronics analyzing the signal. Potential disadvantages to
this acoustic
chamber approach may be that the close proximity of the chamber walls could
result in
anomalous sound speed readings. A lower frequency (50-500 KHz) transducer
would be
required for this technique. A potential pitfall could be the difficulty of
developing a low
frequency transducer that would be physically small enough to fit onto the
catheter. Also,
gas leakage could potentially limit the useful life of the monitor. If the gas
did leak, it would
not pose a danger to the patient because the volume of gas would so minute
that it would
immediately absorb into the blood. The type of gas utilized has yet to be
determined and,
except for safety reasons, would not be limited to any particular gas. Since
instrument
precision increases with frequency, the choice of gas used in this device
would be
dependent upon its ability to conduct higher acoustic frequencies. The
particular frequency
and gas characteristics are yet to be determined. The gas, of course, must not
possess any
properties that are toxic to the patient or that would deteriorate the plastic
of the catheter
itself.
C. Blood Flow Velocity Measurement:
The arterial blood flow velocity parameter would be measured by use of a
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standard Doppler probe mounted at or near the tip of the catheter. The best
mount position
would be chosen in order to detect flow in a location where flow is unaffected
by the
presence of the catheter itself. It may be possible to obtain Doppler flow
values using the
same transducer as is used for the blood concentration measurement. The blood
flow
velocity measurement could be displayed continuously and it would be used to
continuously
calculate, display, and trend local vascular resistance or local vascular
tone.
D. Calculation of Local Vascular Resistance (LVR):
It must be understood that by placing the device 120 multiparameter catheter
into a peripheral vessel, total systemic vascular resistance (SVR) can not be
obtained. SVR
can only be accomplished by monitoring a central vessel. However, for purposes
of trending
SVR, the monitoring of a peripheral artery could be useful. Trends of LVR may
parallel
trends for SVR. It is possible to calculate vascular resistance or its
inverse, vascular
conductance, by using the measured values of blood pressure, flow velocity,
and vessel
cross-sectional area CSA. The mathematical relationships are as follows:
[0153]
F=V*CSAexpressed in ml/min [0154] LVR=MAP/Fexpressed in
mmHg/(ml/min)
[0155] VC=F/MAPexpressed in (ml/min)/mmHg where F is local blood flow
volume
in the peripheral artery, V is local blood flow velocity, and CSA is the cross-
sectional area
of the cannulated vessel, PVR is the local vascular resistance within the
cannulated vessel,
MAP is mean arterial pressure, and VC is vascular conductance.
The monitoring of these particular vascular system parameters would give
the clinician valuable information on the general state of vascular tone. For
example, when
blood pressure plummets, generally there is compensation by sympathetic and
adrenergic
mechanisms to increase PVR in order to maintain adequate blood pressure. This
is certainly
true in the case of hemorrhagic shock, in which case device 120 would show an
increasing
PVR coupled with a decreasing blood pressure and possibly a decrease in vessel
CSA.
Conversely, when the vascular volume is restored via blood product and/or IV
fluid
administration, device 120 would show a returning to normal of the PVR from
high levels.
In the case of spinal shock, however, device 120 would show that normal
compensatory
mechanisms are not occurring and that PVR is inappropriately low given the
current state of
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hypotension.
In reality, accurate intravascular measurement of vessel CSA may be
cumbersome. It may be better to employ other techniques for vascular tone
calculation.
Device 120 could employ the use of pulse wave velocity (PWV) calculation to
determine
vascular tone. PWV in a region of an artery is mainly related to the elastic
properties of the
arterial wall of that region. Two techniques have been described by Harada et
al that utilize
one-point measurement of (PWV) and wave intensity (WI)8. These two
methods
involve calculating 1) the characteristic impedance of an artery and 2)
calculating the
stiffness parameter. Interestingly, device 120 is already designed to measure
the required
input data for these equations including blood density, flow velocity, and
pressure. Please
see the reference article for the applicable equations.
E. Continuous Arterial Blood Gas Measurement:
An example of an intravascular blood gas monitor that could potentially be
incorporated into device 120 is NeoTrend made by Diametrics Medical, Inc.
NeoTrend was
evaluated in a journal article entitled "Continuous Neonatal Blood Gas
Monitoring Using a
Multiparameter Intra-arterial Sensor;" by C. Morgan, S. J. Newell, D. A.
Ducker, J.
Hodgkinson, D. K. White, C. J. Morley, J. M. Church; "Arch Dis Child Fetal
Neonatal Ed,
March 1999;80:F93-F98
FIG. 6 is a view of the new and improved multiparameter intravascular
catheter 110 vital sign measurement device of the invention showing a single
sender/receiver transducer 112 paired with a reflector 114. Another potential
configuration
with two transducers, one sender and one receiver, is not illustrated since
the receiver
transducer would simply occupy the location of the reflector in FIG. 6.
Included are
diagrammatic representations of the incorporated probes for measuring sound
speed and
blood flow 112-114, temperature 116, blood pressure (BP) 118, arterial blood
gases
(AB G's) 20.
Each of the sensors 12, 14, 16, 18, and 20 illustrated in FIG. 6 is connected
to hardware drivers 22, and a computer 24 that is programmed with recognition
and analysis
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software. Depending upon the function of the sensor 112, 116, 118, or 120, the
computer
software will differ as to each recognition and analysis computer 124 to
receive the signal
from its individual sensor 112, 116, 118, and 120, and convert the same into a
measurement
of blood concentration, hematocrit and/or hemoglobin, blood flow velocity,
blood pressure,
pulse rate, local vascular resistance, pH, p02, pCO2 and any other data
gathered from any
other technology incorporated into the device. The display 128 will include
both
instantaneous measurements and a plot of each measurement versus time as it
seems
prudent to display for clinical use.
Scanners 126 are provided to scan each of the computers 124 sequentially
from about 10 to 1000 times per second, depending upon the particular clinical
application.
Each of the scanners would be operatively connected to a display 128 that
would display the
measured data from each of the sensors 112, 116, 118, and 120 in the form of
both
instantaneous measurements and the historical trends of each parameter. Data
derived
mathematically, such as local vascular resistance and pulse pressure would
likewise be
displayed and trended. The display would be combined with selection switches
by which
each parameter and trend could be selectively displayed.
Attached to each display would be a printer 130 which would be capable of
printing current vital parameters and a continuous record of trends and rate
of change of
each measurement or calculated parameter.
Each of the sensors 112, 116, 118, and 120, the hardware drivers 122, each
of the recognition and analysis computers 124, each of the scanners 126, each
of the
displays 128, and each of the printers 130 are connected to a power source
132.
FIG. 7 illustrates a potential configuration for the sound speed sensor with
transducer 112 and reflector 114 held apart by stainless steel struts 140. The
scale image
provides an example of a catheter 110 that is 1 mm diameter and the transducer-
reflector
separation is 5 mm. The actual dimensions of the catheter may vary according
to the
particular logistics of construction and function, and are not limited to
these dimensions.
FIG. 8 illustrates another potential configuration for the sound speed sensor
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with transducer 112 and reflector 114 mounted at opposite ends of a notch cut
into the side
of the catheter 110.
The particular form(s) that these sound speed measurement devices might
assume are not limited and would include any configuration that would allow
intravascular
measurement of speed of sound in whole blood.
FIG. 9 illustrates a potential configuration of the gas-filled chamber
pressure
sensor. This particular example shows transducer 134 and reflector 136 mounted
in a notch
cut into the side of the catheter 110, with a membrane or diaphragm 138
covering and
sealing the gas-filled chamber 142. The transducer-reflector separation is 5
mm in this
example. The actual dimensions of the chamber, the particular transducer
frequency, and
type of gas used may vary according to the particular logistics of
construction and function.
The particular forms that the this sound speed measurement device might assume
are not
limited and would include any configuration that would allow measurement of
the speed of
sound in any gas contained within any intravascular device.
The new and improved multiparameter intravascular catheter vital sign
measurement device 120 of the invention is a medical device for supplying
multiple
physiologic and hemodynamic parameter measurements for any clinical medicine,
veterinary, military, or research setting where such information is useful to
clinicians
conducting physical examinations, or monitoring or treating patients or
personnel. It would
be especially useful in the care of critically ill or injured patients.
When the device is placed intraarterially, it would be capable of accurately
and continuously measuring arterial blood pressure, pulse rate, blood density,
H/H,
temperature, and blood flow velocity without the usual hydraulic arterial-line
system, and it
would also be capable of trending a calculated local vascular resistance.
Vascular volume
could be obtained by a common dilution method. When the device is placed
intravenously,
blood concentration, H/H, and temperature would be continuously monitored.
In addition, device 120 could incorporate any existing or future available
intravascular technology for physiologic and hemodynamic measurements which
might be
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desirable to monitor in the various environments discussed above. Examples of
desirable
venous or arterial physiologic parameters include, but are not limited to,
p02, pCO2, pH,
lactic acid levels, electrolytes, PT, PTT. Future advances in technology could
bring the
eventual incorporation into the catheter of such entities as genetic testing
or blood-typing
for cross-match for anticipated transfusion. In addition to its value in
critical care medicine,
the device could have many other uses, including but not limited to use in
many types of
medical or veterinary research, in clinicians offices, in military, aerospace,
and subsurface
marine applications, and in cardiovascular and pharmacological research.
Potentially, any other conventional equipment that can be miniaturized could
be mated with the device in order to continuously monitor a multitude of
desirable
physiologic parameters. The particular existing and available technology that
may be
incorporated into device 120 will depend upon the amenability of each
individual parameter
to miniaturization and to its accuracy, precision, licensing and safety
restrictions, and
compliance to specific regulatory guidelines.
F. Calculation of Total Vascular Hemoglobin Mass:
According to another embodiment, the device 120 can be programmed to
determine total vascular hemoglobin mass on an intermittent basis via
injection of a small
bolus of IV fluid. The usage environment for the device 120 will primarily be
the operating
room, ICU, emergency department, and military MASH units. Its use is mainly
intended
for critically ill or injured patients who are suffering from shock or who may
have the
potential to develop shock from any cause. Additionally, since the instrument
has the
capability to rapidly measure absolute circulating blood volume, it will also
be useful in
guiding the management of CHF patients and dehydrated or fluid overloaded
patients who
are especially fragile due to preexisting renal, cardiac, and pulmonary
disease, especially
when the need to undergo surgery arises. The device 120 is designed to take
the guesswork
out of decisions about when to transfuse, how much blood to transfuse, and
when to give IV
fluids and/or blood vs. pressor agents.
In some embodiments, the device 120 is designed for those particular
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patients in which rapid diagnosis is required and those in which it is crucial
to maintain
hemoglobin and vascular volume within certain limits in order to avoid further
stressing
already fragile cardiovascular, renal, pulmonary, and/or central nervous
systems. These
include trauma, sepsis, neurogenic shock, acute renal failure, severe
dehydration, major
burns, acute respiratory failure with possible volume overload, GI bleeding,
and major non-
trauma related surgical cases (scheduled or unscheduled) such as open heart
surgery, aortic
aneurysm, bowel resection, gynecologic, urologic, intra-thoracic or intra-
abdominal cancer
surgeries, large joint replacement, major open spine surgeries, and major
peripheral vascular
procedures. The percentage of surgical patients that will require monitoring
of the device
120 will increase with the age of the patient and degree of preexisting
disease. It will also
be an important tool in the diagnosis of unexplained hypotension and syncope.
The point
and purpose of the device 120 is to improve speed of diagnosis and to give
guidance to goal
directed therapy for these difficult patients, ultimately optimizing care,
improving
outcomes, and shortening hospital stays, thus limiting iatrogenic
complications as well as
reducing cost of care.
It is well known to critical care providers that hemoglobin level in an
actively hemorrhaging patient is rather meaningless because it does not
accurately reflect
the overall (whole body) oxygen carrying capacity of the blood. Decisions to
transfuse
must, therefore, be based on other indicators such as signs and symptoms,
hemodynamic
parameters, clinical course over time, and test results (including hemoglobin
level). Since
the device 120 delivers both hemoglobin level and circulating blood volume, as
described
above, the actual quantity of hemoglobin contained within the circulatory
system that is
available for oxygen transport can be determined. As a result, a new, more
relevant and
beneficial hemodynamic parameter, Total Hemoglobin Mass ("THM"), can be
introduced.
THM is defined as the total amount of hemoglobin contained within the vascular
system
(typically expressed in grams). Another way to state this is to say that THM
represents total
blood oxygen carrying capacity. THM is a much more accurate indicator of the
entire
body's total oxygen carrying capacity than is the hemoglobin level, which is
essentially a
simple "spot check" of the hemoglobin concentration. According to some
embodiments,
the THM could be calculated by the device 120 using the following equation:
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THM= H91). IL)* BV(dL)
where THM is Total Hemoglobin Mass in grams, Hgb is hemoglobin in grams per
deciliter,
and BV is the measured total circulating vascular volume in deciliters. A
typical 70kg male
patient with normal hemoglobin between 13 and 16 g/dL would have a THM of
roughly 600
to 800 grams.
Accordingly, a Total Hemoglobin Mass index (THMi) can be introduced.
THMi is defined as THM per kilogram body weight. THM and THMi may be "new" but
they are not a foreign concept to physicians who frequently order transfusions
and think in
terms of how much blood a patient will need. THM and THMi are new tools that
will be
helpful in determining exactly when and how much blood to transfuse. According
to some
embodiments, the THMi could be calculated by the device 120 using one of the
following
equations:
THM g = WO)
THMi. = ____________________________________________ THAR = Bgb
141 ti_kg) or ',c/V WCkg)
where THMi is the Total Hemoglobin Mass Index in grams per kilogram body
weight.
Conversely, it follows that when Wt(kg) and BV(dL) are known, then a specific
THMi can
be applied to help maintain a specific hemoglobin goal:
Wt(kg)
H gb ¨L. )= THMi ___________________________________
\-c1 ifiVW t)
As implied by the latter equation, the caregiver will be able to use specific
goal directed infusions of packed RBCs and IV fluids in order to achieve
specific
hemoglobin and volume goals during resuscitation. If it is a goal to prevent
the hemoglobin
from dropping below a certain preset level, e.g., 10g/dL, as blood volume is
being restored
during resuscitation, then THMi can be used to determine the minimum volume of
packed
RBC's and IV fluids to give at any particular point in time. Since THMi is
based upon
body weight, and the goal is to return the blood volume back to normal
(70mL/kg), the math
computes in such a way that a goal of 10g/di hemoglobin will always equate to
a THMi goal
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of 7g/dL no matter what the patient weighs. In any patient, therefore, if the
THMi is below
7g/kg, then the total body hemoglobin deficiency equals (7 ¨ nrimi).k Wt(kg)
For
example, if the patient weighs 70kg and the THMi is calculated to be 6.3, then
the
hemoglobin deficiency equals (7-6.3)x70, or approximately 50 grams. Since a
unit of
packed RBCs has a hemoglobin level of between 25 and 30g/dL, then a
transfusion of 2dL
(200mL) of packed RBCs will deliver at minimum of 50 grams and will therefore
replenish
THMi back to a level of 7g/kg and THM back to 490 grams. Once the THM has been
replenished, even if the blood volume were to be restored back to normal with
an infusion
of normal saline or LR, then the hemoglobin would not be diluted to less than
10g/dL. In
most cases, the IV fluids and packed RBCs would be infused simultaneously.
Due to the fact that THM and THMi better characterize total blood oxygen
carrying capacity, it can be anticipated that their introduction as new
hemodynamic
parameters may give rise to the eventual updating of transfusion guidelines
which are
currently based predominantly upon hemoglobin level.
According to one embodiment, THM could be determined through a rapid
injection of a small volume (e.g., 50 to 100mL) of normal saline or Lactated
Ringers
solution ("LR") in a separate IV site, which allows determination of absolute
circulating
blood volume. Figure 10 is a graph representing measurement of absolute blood
volume
from the animal study that was done by Applicant on 9/18/2012. Absolute
Circulating
Blood Volume is the volume of circulating blood immediately available to the
heart for
distribution to the central circulation and major organs. It is calculated as
the volume of
blood into which the indicator (normal saline or lactated ringers) mixes after
rapid injection.
Data points shown in Figure 10 are displayed every 5 seconds following a rapid
bolus
injection of Ringers lactate IV solution, in this case, administered via the
femoral vein. The
graph demonstrates the timing and degree of decrease in blood concentration
(dilution) of
the circulating intravascular blood volume. The line with diamond indicated
data points is
the instant data and other line is a 6 point (30 seconds) moving average. It
was found that
absolute blood volume (ABV) is best calculated using the change in
concentration from the
initial concentration (Cl) to the lowest point of the moving average (C2) by
the common
equation:
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ABV=C1*Vi/(C1-C2)
Where Vi is the amount of IV fluid injected.
It was also found that cardiac output (CO) can be calculated using the time
that it takes for
the dilute to its first bottom, in this case at 45 seconds after the bolus
injection. We will call
this point TC3 (time to C3) for reasons of clarity. The equation for this
calculation is a
simple ratio. Since the definition of CO is the amount of blood circulated in
1 minute (60
seconds), the equation is:
CO=ABV*60/TC3
Prior to the injection, the user will enter the amount of fluid being injected
and the patient's weight into the user interface of the device 120, which
would be stored in
the memory of the device 120. If desired, the hemoglobin goal could be entered
into the
user interface of the device 120, which would also be stored in memory. After
pressing a
START button, the user will inject the predetermined volume of fluid over 10-
20 seconds.
The device 120 will then automatically calculate and display blood volume,
cardiac output,
and total hemoglobin mass within 60-90 seconds. If the user has entered a
predefined
hemoglobin goal, the device will also display the minimum volume of packed
RBC's that
must be transfused in order to meet that hemoglobin goal at that particular
point in time.
The following is an example of how the device 120 could be used and how
THM and THMi might be applied in a trauma resuscitation scenario. Consider an
example
in which a 70kg male patient with head injury and possible internal injuries
arrives at the
trauma center having had a single episode of hypotension (BP<90 systolic)
which the
paramedics have treated with a fluid bolus while en route. The trauma
literature indicates
that head injured patients do best when hemoglobin level does not drop below
10g/dL. In
this case the goal would therefore be to prevent the THMi from dropping below
7.0g/kg and
THM from dropping below 490 grams. The caregiver, however, may at first see no
indication to order a blood transfusion if the initial hemoglobin level is
12g/dL, even though
the patient might have lost a large amount of blood volume from hemorrhage due
to yet
undetected internal injuries.
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In this example, the patient (whose normal blood volume should be
approximately 4900mL) has lost 25% of his blood volume due to hidden internal
hemorrhage. Using the device 120, the blood volume in this 70kg male can be
rapidly
measured and found to be 25% below normal, or 3675mL. Knowing the hemoglobin
to be
12g/dL and circulating volume to be 3675mL (36.75dL), then THM can be
calculated to be
441 grams and the THMi to be 6.3g/kg, both below the goals of 490 grams and
7.0g/dL
respectively. In order to prevent the hemoglobin level from dropping below
10g/dL during
resuscitation, the patient must therefore be transfused when his hemoglobin
level is actually
near normal at 12g/dL.
Although the decision to transfuse is always left to the attending physician's
best judgment, standard trauma shock resuscitation protocol is to initially
refill the vascular
space by giving fluid boluses of as much as 30mL/kg. While the fluid is
infusing, the
patient's response (or lack of response) to the treatment is observed and
transfusion is
typically ordered based upon the particular physician's intuition and
experience. Using the
device 120, however, it can easily be predicted that, if the vascular volume
is restored by the
necessary fluid bolus, then the hemoglobin level will be diluted to the
undesirable level of
9g/dL. Furthermore, this scenario assumes that the patient was healthy to
begin with and
that the bleeding has been controlled or has ceased. The need for immediate
transfusion is
even more urgent if bleeding is not yet controlled or if the patient has other
comorbidities.
If indeed all bleeding has been controlled, then a THM of 441 grams and THMi
of 6.3g/kg
indicate that the patient is now nearly 50 grams short of the goal of 490
grams in terms of
total hemoglobin mass (THM). The patient will thus require a transfusion of at
least 50
grams of hemoglobin in order to prevent the hemoglobin from dropping below the
goal of
10g/dL. Furthermore, since the device 120 provides continuous hemoglobin data,
the
response to treatment can be assessed in real-time and the circulating volume
can be
reassessed periodically as indicated by the clinical course of the patient.
The head injured patient may arrive at the trauma center in a hypotensive
state with or without internal injuries. Internal bleeding may be suspected
but is not yet
confirmed or ruled out. Considering the fact that trauma itself and trauma
resuscitation is a
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very dynamic state, aiming for a hemoglobin level of 10 is rather meaningless
and
somewhat futile if the circulating blood volume and THM are unknown. To
complicate the
situation further, the hypotension may be secondary to spinal shock with or
without other
internal injuries. In the spinal shock patient without internal injuries,
rapid fluid infusion
may not restore adequate vital signs and may in fact be harmful. In such a
case, the device
120 would immediately reveal normal hemoglobin, normal blood volume, normal
THM,
and normal THMi. Having this information immediately available would allow the
caregiver to avoid initial treatment that might be potentially detrimental
(albeit still proper
according to current guidelines). Having access to such data early in the
course of treatment
would clarify a choice in favor of more focused therapy such as pressor
agents, which
would not ordinarily be considered this early in the course of treatment when
shock from
internal bleeding has not yet been ruled out. The normal readings of the
device 120 would
also be a clear indication to proceed immediately to the CT scanner as opposed
to spending
more valuable time contemplating operative intervention or further fluid
therapy and/or
transfusions for wrongly suspected hemorrhagic shock. The ultimate result
would be that
the patient's care would be optimized, offering the best chances for survival
and improved
outcome as well as shortened hospital stay and reduced cost of care.
Ultimately,
resuscitation guidelines would need to change according to the new information
provided
by the device 120.
Without a way to accurately measure blood volume, critical treatment
decisions made in the early stages of resuscitation are, out of necessity,
based upon intuition
and experience. The above "multiple-trauma with brain injury" example was
chosen
because it represents one of most challenging types of cases involving
hemodynamic
instability and the need for balanced goal directed therapy during
resuscitation. Short of
giving multiple parallel detailed examples of each type of potential patient,
it can be said
that the device 120 could similarly be used to optimize the diagnosis and
treatment of
critically ill patients who are hemodynamically unstable or who are likely to
become
unstable. In the same way, the device 120 could be used to monitor the above
referenced
surgical patients who are at risk for sudden changes in blood volume as well
as
cardiovascular stress from anesthesia and preexisting disease.
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Use of the device 120 will reduce the need for frequent blood draws to
recheck hemoglobin levels in unstable patients. It is believed to eliminate
the need for:
- CVP and pulmonary capillary wedge pressure monitoring.
- Arterial wave form analysis equipment (note* arterial waveform
analysis is reliable only in patients who are unconscious and fully
ventilator dependent).
- Ultrasound imaging for measurement of the inferior vena cava size
(note* this procedure has been found to be operator dependent and
reliable only when the vena cava is either very distended or very
collapsed).
The device 120 will not eliminate the need for arterial pressure monitoring
and in fact may work most optimally when incorporated with an arterial
pressure catheter.
Noninvasive Cardiac Output Monitoring (NICOM) would be an adjunct to the
device 120.
The two devices would likely complement each other since NICOM is continuous
but not
capable of delivering hemoglobin values or blood volume. Continuous
noninvasive
hemoglobin monitoring is a relatively recent addition to options for
hemodynamic
monitoring. This oximeter-based technology, however, lacks the precision
required for
accurate blood volume analysis. It is also known to fail during severe shock
when its
information would potentially be most essential.
While the invention has been illustrated and described in detail in the
drawings and foregoing description, the same is to be considered illustrative
and not
restrictive in character, it being understood that only selected embodiments
have been
shown and described and that all changes, equivalents, and modifications that
come within
the scope of the inventions described herein or defined by the following
claims are desired
to be protected.
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