U.S. patent application number 12/563887 was filed with the patent office on 2010-06-17 for apparatus and method for non-invasive and minimally-invasive sensing of parameters relating to blood.
Invention is credited to Xuefeng Cheng, Butrus T. Khuri-Yakub, Daniel Hwan Kim.
Application Number | 20100152559 12/563887 |
Document ID | / |
Family ID | 37889629 |
Filed Date | 2010-06-17 |
United States Patent
Application |
20100152559 |
Kind Code |
A1 |
Cheng; Xuefeng ; et
al. |
June 17, 2010 |
APPARATUS AND METHOD FOR NON-INVASIVE AND MINIMALLY-INVASIVE
SENSING OF PARAMETERS RELATING TO BLOOD
Abstract
Medical diagnostic system, apparatus and methods are disclosed.
Optical transmitters generate radiation-containing photons having a
specific interaction with at least one target chromophore in a
target structure, preferably a blood vessel such as the interior
jugular vein. The optical transmitters transmit the radiation into
at least a first area including a substantial portion of the target
structure and into a second area not including a substantial
portion of the target structure. Optical receivers detect a portion
radiation scattered from at least the first area and the second
area. A processor estimates oxygenation, pH or cardiac output based
on the scattered radiation detected from the first area, and the
scattered radiation from the second area.
Inventors: |
Cheng; Xuefeng; (Cupertino,
CA) ; Kim; Daniel Hwan; (Mountain View, CA) ;
Khuri-Yakub; Butrus T.; (Palo Alto, CA) |
Correspondence
Address: |
PATTERSON & SHERIDAN, L.L.P.
3040 POST OAK BOULEVARD, SUITE 1500
HOUSTON
TX
77056
US
|
Family ID: |
37889629 |
Appl. No.: |
12/563887 |
Filed: |
September 21, 2009 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11233308 |
Sep 22, 2005 |
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12563887 |
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11095091 |
Mar 30, 2005 |
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11233308 |
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Current U.S.
Class: |
600/322 |
Current CPC
Class: |
A61B 5/14551 20130101;
A61B 5/0261 20130101; A61B 5/029 20130101; A61B 5/14539 20130101;
A61B 5/1464 20130101; A61B 5/489 20130101; A61B 8/08 20130101; A61B
5/0059 20130101; A61B 5/0048 20130101; A61B 8/065 20130101; G01S
15/8968 20130101; A61B 5/413 20130101; A61B 5/14542 20130101; A61B
8/488 20130101 |
Class at
Publication: |
600/322 |
International
Class: |
A61B 5/1455 20060101
A61B005/1455 |
Claims
1. A system for monitoring one or more parameters relating to blood
of a patient comprising: one or more optical transmitters
configured to generate pulsed radiation containing photons having a
specific interaction with at least one target chromophore in a
target structure within the patient, said one or more optical
transmitters configured and positioned to transmit the pulsed
radiation from the optical transmitter into at least a first area
including a substantial portion of the target structure and into a
second area not including a substantial portion of the target
structure; one or more optical receivers configured and positioned
to detect a portion of the pulsed radiation scattered from at least
the first area and the second area; and a processor adapted to
estimate the one or more parameters relating to the patient's
blood, the estimation based in part on the scattered pulsed
radiation detected from the first area, and the scattered pulsed
radiation from the second area.
2. A system according to claim 1 wherein the target structure is a
blood vessel.
3. (canceled)
4. (canceled)
5. A system according to claim 1 wherein said one or more optical
transmitters is adapted to transmit the pulsed radiation into the
first and second areas simultaneously.
6. A system according to claim 1 wherein the processor comprises a
central processing unit and a memory device and said one or more
optical receivers are configured and adapted to measure optical
properties associated with the second area.
7. (canceled)
8. A system according to claim 6 wherein said processor is further
adapted to calculate a probability for the pulsed radiation to pass
through the first area and the second area.
9. (canceled)
10. A system according to claim 1 wherein said processor estimates
the one or more parameters in part by relating probabilities for
the pulsed radiation to pass through the first and second
areas.
11. (canceled)
12. A system according to claim 1 wherein the pulsed radiation
comprises photons having a first wavelength and photons having a
second wavelength, the first wavelength selected to have the
specific interaction with a first target chromophore, and the
second wavelength selected to have a specific interaction with a
second target chromophore.
13.-20. (canceled)
21. A system according to claim 1 wherein said one or more optical
transmitters comprise an optical source and one or more launch
optics coupled to the optical source via one or more optical
fibers.
22. A system according to claim 21 wherein said one or more optical
receivers comprise one or more optical detectors and one or more
collecting optics coupled to the one or more optical detectors via
one or more optical fibers.
23.-44. (canceled)
45. An apparatus for monitoring one or more parameters relating to
blood of a patient comprising: a pad of pliable material adapted to
be engaged with the skin of a patient; one or more launch optics in
communication with at least one optical source via optical fiber,
said one or more launch optics being mounted on the pad and
positioned so as to enable pulsed transmission of photons generated
by the at least one optical source through the patient's skin into
at least a first area including a substantial portion of a target
blood vessel and into a second area not including a substantial
portion of the target blood vessel; and one or more collecting
optics in communication with at least one optical detector via
optical fiber, said one or more collecting optics being mounted on
the pad and positioned so as to enable detection of the photons
having been scattered within the first area and the second area,
wherein the photons have a specific interaction with at least one
target chromophore in the target blood vessel.
46. An apparatus according to claim 45 further comprising two or
more transmitter-receiver pairs each pair comprising one launch
optic and one collecting optic, including a first pair having a
spacing and being positioned such that the launch optic and
collecting optic associated with the first pair transmits photons
into and detects scattered photons from the first area, and a
second pair having a spacing and being positioned such that the
launch optic and collecting optic associated with the second
47.-50. (canceled)
51. An apparatus according to claim 45 wherein some of the photons
have a first wavelength and some of the photons have a second
wavelength, the first wavelength selected to have the specific
interaction with a first target chromophore within the target blood
vessel, and the second wavelength selected to have a specific
interaction with a second target chromophore within the target
blood vessel.
52.-55. (canceled)
56. An apparatus according to claim 45 wherein said at least one
optical source and said at least one optical detector are housed in
a portable box being dimensioned and sized such the patient can
carry the portable box for extended periods, and said portable box
further housing a battery and a wireless transmitter for data
communication to a station box located at a hospital or clinic.
57.-64. (canceled)
65. A method for monitoring one or more parameters relating to
blood of a patient comprising the steps of: engaging one or more
optical transmitters and one or more optical receivers on a tissue
boundary of the patient; generating pulsed radiation containing
photons having a specific interaction with at least one target
chromophore in a target structure within the patient; transmitting
said pulsed radiation through the optical transmitters into a first
area including a substantial portion of a target blood vessel and
into a second area not including a substantial portion of the
target structure; detecting with the one or more optical receivers
a portion of the pulsed radiation scattered from at least the first
area and the second area; and estimating the one or more parameters
relating to a patient's blood based in part on the detected
scattered pulsed radiation from the first area, and the scattered
pulsed radiation from the second area.
66.-68. (canceled)
69. A method according to claim 65 further comprising the step of
calculating optical properties associated with the second area,
wherein said step of estimating is based in part on the measured
optical properties associated with the second area.
70. A method according to claim 69 wherein said optical properties
associated with the second area includes average tissue scattering
and absorption properties for the pulsed radiation associated with
the second area.
71. A method according to claim 69 wherein said step of estimating
the one or more parameters includes calculating a probability for
the pulsed radiation to pass through the first area and the second
area.
72. A method according to claim 71 wherein said probabilities are
calculated using a photon diffusion equation.
73.-81. (canceled)
82. A method according to claim 65 wherein the step of transmitting
said pulsed radiation includes transmitting the pulsed radiation
through the optical transmitters into a third area including a
substantial portion of a second target structure, and said step of
detecting comprises detecting pulsed radiation scattered from the
third area.
83.-84. (canceled)
85. A method according to claim 65 wherein said step of generating
pulsed radiation uses at least one optical source, said step of
transmitting uses one or more launch optics coupled to the at least
one optical source via one or more optical fibers, said one or more
optical receivers comprise one or more optical detectors and one or
more collecting optics coupled to the one or more optical detectors
via one or more optical fibers.
86.-96. (canceled)
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is related to co-pending U.S. patent
application Ser. No. 11/095,091, filed 30 Mar. 2005, in the name of
John F. Black, Daniel Hwan Kim, and Butrus T. Khuri-Yakub, entitled
"Apparatus and Method for Non-Invasive and Minimally-Invasive
Sensing of Venous Oxygen Saturation and pH Levels", which is hereby
incorporated by reference as if fully set forth herein.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] This invention is related to techniques for monitoring vital
bodily functions, including cardiac output. It relates in
particular to methods and apparatus for non-invasive and
minimally-invasive real-time monitoring of parameters such as
venous oxygenation saturation or pH in a vessel, an organ or tissue
containing blood.
[0004] 2. Description of the Related Art
[0005] Cardiac output is defined as the volume of blood circulated
per minute. It is equal to the heart rate multiplied by the stroke
volume (the amount ejected by the heart with each contraction).
Cardiac output averages approximately 5 liters per minute for an
average adult at rest, although it may reach up to 30 liters/minute
during extreme exercise.
[0006] Cardiac output is of central importance in the monitoring of
cardiovascular health, as discussed by Conway "Clinical assessment
of cardiac output", Eur. Heart J. 11, 148-150 (1990). Accurate
clinical assessment of the circulatory status is particularly
desirable in critically ill patients in the ICU and patients
undergoing cardiac, thoracic, or vascular interventions, and has
proven valuable in long term follow-up of outpatient therapies. As
the patient's hemodynamic status may change rapidly, continuous
monitoring of cardiac output will provide information allowing
rapid adjustment of therapy. Measurements of cardiac output and
blood pressure can also be used to calculate peripheral
resistance.
[0007] A recent review of the various techniques for measuring
cardiac output is given in Linton and Gilon, "Advances in
non-invasive cardiac output monitoring", Annals of Cardiac
Anaesthesia, 2002, volume 5, p 141-148. This article lists both
non/minimally invasive and invasive methods and compares the
advantages and disadvantages of each.
[0008] The pulmonary artery catheter (PAC) thermodilution method is
generally accepted as the clinical standard for monitoring cardiac
output, to which all other methods are compared as discussed by
Conway and Lund-Johansen ("Thermodilution method for measuring
cardiac output", Europ. Heart J. 11(Suppl 1), 17-20 (1990)). The
long history of use has defined the technology, suitable clinical
applications, and its inadequacies. Many new methods have attempted
to replace the thermodilution technique, but none have so far
gained acceptance.
[0009] Jansen (J. R. C. Jansen, "Novel methods of
invasive/non-invasive cardiac output monitoring", Abstracts of the
7.sup.th annual meeting of the European Society for Intravenous
Anesthesia, Lisbon 2004) describes eight desirable characteristics
for cardiac output monitoring techniques; accuracy, reproducibility
or precision, fast response time, operator independency, ease of
use, continuous use, cost effectiveness, and no increased mortality
and morbidity. A brief description of some of these techniques
follows.
[0010] Indicator dilution techniques. There are several indicator
dilution techniques including transpulmonary thermodilution (also
known as PiCCO technology, from Pulsion Medical Technologies of
Munich, Germany), transpulmonary lithium dilution method (LiDCO
Group plc of London, UK), PAC based thermodilution and other
methods (Viligance, Baxter; Opti-Q, Abbott; and TruCCOMS, AorTech).
U.S. Pat. No. 6,757,544 to Rubinstein et al. teaches the technique
of optically monitoring indicator dilution in a non-invasive manner
for the purpose of computation of cardiac output, cardiac index,
and blood volume. Transpulmonary indicator dilution methods with
bolus injections are variations on the conventional bolus
thermodilution method. CO is calculated with use of the
Steward-Hamilton equation (Geddes, "Cardiac output using the saline
dilution impedance technique", IEEE Engineering in Medicine and
Biology magazine March 1989, 22 26). Application of this equation
assumes three major conditions; complete mixing of blood and
indicator, no loss of indicator between place of injection and
place of detection, and constant blood flow. The errors associated
with indicator dilution techniques are primarily related to the
violation of these conditions, as discussed by Lund-Johansen ("The
dye dilution method for measurement of cardiac output", Europ.
Heart J. 11 (Suppl 1), 6-12 (1990)) and de Leeuw and Birkenhager
("Some comments of the usefulness of measuring cardiac output by
dye dilution", Europ. Heart J. 11 (Suppl 1), 13-16 (1990)). Of the
mentioned methods the transpulmonary indicator dilution methods as
well as the so-called `continuous cardiac output` thermodilution
methods have been partially accepted in clinical practice as
described in, for example, Rodig et al. "Continuous cardiac output
measurement: pulse contour versus thermodilution technique in
cardiac surgical patients". Br J Anaesth 1999; 50: 525.
[0011] Fick principle. The direct oxygen Fick approach is currently
the standard reference technique for cardiac output measurement, as
discussed by Keinanen et al., "Continuous measurement of cardiac
output by the Fick principle: Clinical validation in intensive
care", Crit Care Med 20(3), 360-365 (1992), and Doi et al.,
"Frequently repeated Fick cardiac output measurements during
anesthesia", J. Clin. Monit. 6, 107-112 (1990). It is generally
considered the most accurate method currently available, although
there are many possibilities of introducing errors, and
considerable care is needed. However when using the Fick method to
trend cardiac output over a short time interval, i.e. during an
operation or in an intensive care unit stay, many of these sources
of errors are no longer pertinent. The NICO (Novametrix) system is
a non-invasive device that applies Fick's principle on CO.sub.2 and
relies solely on airway gas measurement as described by Botero et
al., "Measurement of cardiac output before and after
cardiopulmonary bypass: Comparison among aortic transit-time
ultrasound, thermodilution, and noninvasive partial CO.sub.2
rebreathing", J. Cardiothoracic. Vasc. Anesth. 18(5) 563-572
(2004). The method calculates effective lung perfusion, i.e. that
part of the pulmonary capillary blood flow that has passed through
the ventilated parts of the lung. The effects of unknown
ventilation/perfusion inequality in patients may explain why the
performance of this method shows a lack of agreement between
thermodilution and CO.sub.2-rebreathing cardiac output as described
in Nielsson et al. "Lack of agreement between thermodilution and
CO.sub.2-rebreathing cardiac output" Acta Anaesthesiol Scand 2001;
45:680.
[0012] Bio-impedance and conduction techniques. The bio-impedance
method was developed as a simple, low-cost method that gives
information about the cardiovascular system and/or (de)-hydration
status of the body in a non-invasive way. Over the years, a
diversity of thoracic impedance measurement systems have also
appeared. These systems determine CO on a beat-to-beat time base.
Studies have been reported with mostly poor results, but in
exceptional cases good correlations compared to a reference method.
Many of these studies refer to the poor physical principles of the
thoracic impedance method as described in Patterson "Fundamentals
of impedance cardiography", IEEE Engineering in Medicine and
Biology 1989; 35 to explain the discrepancies. The accuracy of this
technique is increased when the electrodes are placed directly in
the left ventricle, rather than on the chest, however this also
increases its invasiveness.
[0013] Echo-Doppler ultrasound. This technique uses ultrasound and
the Doppler effect to measure cardiac output. The blood velocity
through the aorta causes a `Doppler shift` in the frequency of the
returning ultrasound waves. Echo-Doppler probes positioned inside
the esophagus with their echo window on the thoracic aorta may be
used for measuring aortic flow velocity, as discussed by Schmidlin
et al., "Transoesophageal echocardiography in cardiac and vascular
surgery: implications and observer variability", Brit. J. Anaesth.
86(4), 497-505 (2001). Aortic cross sectional area is assumed in
devices such as the CardioQ, made by Deltex Medical PLC,
Chichester, UK) or measured simultaneously as for example in the
HemoSonic device made by Arrow International. With these minimally
invasive techniques what is measured is aortic blood flow, not
cardiac output. A fixed relationship between aortic blood flow and
cardiac output is assumed. CO can therefore be calculated using
this relationship. Abrupt changes in cardiac output are better
followed with Doppler systems than with the PAC based continuous
cardiac output systems as described in Roeck et al. "Change in
stroke volume in response to fluid challenge: assessment using
esophageal Doppler", Intensive Care Med 2003; 29:1729. This
measurement requires an above average level of skill on the part of
the operator of the ultrasound machine to get accurate reliable
results.
[0014] Arterial pulse contour analysis. The estimation of cardiac
output based on pulse contour analysis is an indirect method, since
cardiac output is not measured directly but is computed from a
pressure pulsation on basis of a criterion or model. The origin of
the pulse contour method for estimation of beat-to-beat stroke
volume goes back to the Windkessel model as described in, for
example, Manning et al. "Validity and reliability of diastolic
pulse contour analysis (Windkessel model) in humans", Hypertension.
2002 May; 39(5):963-8. Most pulse contour methods are based on this
model explicitly or implicitly as described in Rauch et al. "Pulse
contour analysis versus thermodilution in cardiac surgery", Acta
Anaesthesiol Scand 2002; 46:424, Linton et al. "Estimation of
changes in cardiac output from arterial blood pressure waveform in
the upper limb", Br J Anaesth 2001; 86:486 and Jansen et al. "A
comparison of cardiac output derived from the arterial pressure
wave against thermodilution in cardiac surgery patients" Br J
Anaesth 2001; 87:212.
[0015] Arterial pulse contour analysis techniques relate an
arterial pressure difference to a flow or volume change. Three
pulse contour methods are currently available; PiCCO (Pulsion),
PulseCO (LiDCO) and Modelflow (TNO/BMI). All three of these pulse
contour methods use an invasively measured arterial blood pressure
and they need to be calibrated. PiCCO is calibrated by
transpulmonary thermodilution, LiDCO by transpulmonary lithium
dilution and Modelflow by the means of 3 or 4 conventional
thermodilution measurements equally spread over the ventilatory
cycle. Output of these pulse contour systems is calculated on a
beat-to-beat basis, but presentation of the data is typically
within a 30-second window. A non-invasive pulse contour development
is the combination of non-invasively measured arterial finger blood
pressure with Modelflow as described in Hirschl et al. "Noninvasive
assessment of cardiac output in critically ill patients by analysis
of finger blood pressure waveform", Crit Care Med 1997;
25:1909.
[0016] None of the above-mentioned CO techniques combines all of
the eight "Jansen" criteria mentioned above. With respect to
accuracy and precision, a number of methods may approach the
thermodilution method with a precision of 15%. None of these new
techniques has displaced conventional thermodilution based on the
averaged result of 3 or 4 measurements done equally spread over the
ventilatory cycle as described in Jansen et al. "An adequate
strategy for the thermodilution technique in patients during
mechanical ventilation", Intensive Care Med 1990; 16:422. Under
research conditions the use of this conventional thermodilution
method remains the method of choice. However, in clinical settings,
the lower precision of the continuous cardiac output techniques may
be outweighed by their advantages of being automatic and
continuous.
[0017] In addition to measuring cardiac output, it is also
desirable in many critical care situations to continuously monitor
a patient's blood oxygen level. Currently, hospitals routinely
monitor blood oxygenation by pulse oximetry with a monitor attached
to the patient's finger or earlobe as described for example in
Silva et al., "Near-infrared transmittance pulse oximetry with
laser diodes", J. Biomed. Opt 8(3), 525-533 (2003). Typically the
oxygen monitor is a pair of light-emitting diodes (LED) and
photodiodes on a probe clipped to a part of the patient's body. Red
light from the LED reflects from the blood in a part of the
patient's body, such as an ear-lobe or finger-tip. As a patient's
oxygenation level drops, the blood becomes more blue, reflecting
less red light to the photodiode. Such blood-oxygen monitors
customarily measure percent of normal. Reassuring (normal) ranges
are from 95 to 100 percent. For a patient breathing room air, at
not far above sea level, an estimate of arterial oxygenation can be
made from the blood-oxygen monitor reading. Unfortunately,
measurements from such oxygen monitors cannot be reliably
correlated to oxygenation in the patient's venous blood. Venous
oxygen saturation is also a valuable parameter in the diagnosis of
septic and cardiogenic shock as described below.
[0018] Other methods of measuring oxygenation: Diffuse optical
tomography methods as described for example in Boas et al., Method
for monitoring venous oxygen saturation", US Patent application
20040122300 are conceptually appealing but are useful only where
the vessels in the vicinity of the diffusing photon field are
isolated veins. The presence of mixed arterial and venous blood
complicates the problem to as described by Wolf et al., "Continuous
noninvasive measurement of cerebral arterial and venous oxygen
saturation at the bedside in mechanically ventilated neonates",
Crit. Care. Med 25(9), 1579-1582 (1997).
[0019] Ultrasound-tagged optical spectroscopy involves overlapping
an ultrasound wave and a diffusing optical field, and modulating
the frequency of the probe photons or their trajectories. A number
of different technologies have been developed that utilize some
interaction between ultrasound radiation and electromagnetic
radiation. U.S. Pat. No. 5,212,667 to Tomlinson et al. and U.S.
Pat. No. 5,174,298 to Dolfi et al. teach the technique of
ultrasound tagged frequency-modulated imaging. Other patents
teaching variations on the theme of frequency-modulated ultrasound
tagging techniques include U.S. Pat. No. 6,815,694 to Sfez et al,
U.S. Pat. No. 6,738,653 to Sfez et al., U.S. Pat. No. 6,041,248, to
Wang, U.S. Pat. No. 6,002,958 to Godik, U.S. Pat. No. 5,951,481 to
Evans, U.S. Pat. No. 5,293,873 to Fang. Trajectory modulation is
detected by monitoring the speckle pattern of the photons emerging
from the target. Image reconstruction techniques are then used to
recreate a map of the path the photons followed in the medium.
Imaging the speckle resulting from trajectory changes requires
significant computation power and post-processing to yield an
image. The technique has limited resolution, and is not yet capable
of yielding functional (oxygenation) information in a fast flowing
vessel.
[0020] Some variations of ultrasound-tagged frequency-modulated
imaging rely on observing the frequency shift induced by the
photoacoustic effect when an electromagnetic wave interacts in a
medium with a sound wave. The electromagnetic wave (having a
characteristic frequency .omega..sub.OPT) receives a frequency
shift at the ultrasound frequency .omega..sub.US to either the + or
- side of the carrier wave .omega..sub.OPT. Frequency modulation is
detected by measuring the frequency shifted photons by for example
using a Fabry-Perot etalon as described by Sakadzic and Wang, "High
resolution ultrasound modulated optical tomography in biological
tissues", Opt. Lett. 29(23) 2004, p 2770-2772. Since the Doppler
shifts induced by the ultrasound wave are very small compared to
the probe photon carrier wave frequency, the detection system must
be extremely sensitive to small frequency shifts. In addition, the
frequency shift can be to both larger and smaller frequency of the
initial carrier wave, and therefore some self-cancellation may
result.
SUMMARY OF THE INVENTION
[0021] There is a need in the art to be able to measure venous
oxygen saturation levels in various vascular structures in the
body, and from this be able to calculate cardiac output. There is a
need to make these measurements non-invasively or with minimal
invasiveness. There is a need to be able to make these measurements
in an MRI-/CT/X-Ray instrument compatible manner, thus preferably
not using ferromagnetic materials in construction, and using
designs such that the probe on/in the body may be remotely coupled
to the control system away from the magnetic field or ionizing
radiation sources generated by the MRI instrument or CT/X-Ray.
There is a need in the art to make these measurements in a manner
that does not depend on the melanin content of the skin. There is a
need to make these measurements in a manner such that the result
may be arrived at in a short time period, i.e. such that extensive
post-processing of the data is not required, so that the physician
may make accurate timely diagnostic and therapeutic decisions. Many
or all of the disadvantages associated with the prior art can be
significantly alleviated through embodiments of the present
invention.
[0022] According to certain embodiments of the present invention a
system for monitoring one or more parameters relating to blood,
such as oxygenation, pH or cardiac output, is provided including
optical transmitters configured to generate radiation containing
photons having a specific interaction with at least one target
chromophore in a target structure, preferably a blood vessel such
as the interior jugular vein. The optical transmitters are adapted
and positioned to transmit the radiation from the optical
transmitter into at least a first area including a substantial
portion of the target structure and into a second area not
including a substantial portion of the target structure. Optical
receivers are configured and positioned to detect a portion
radiation scattered from at least the first area and the second
area. A processor is adapted to estimate the one or more parameters
relating to the patient's blood, the estimation being based in part
on the scattered radiation detected from the first area, and the
scattered radiation from the second area. A processor preferably
estimates average tissue scattering and absorption properties for
the radiation associated with the second area, and calculates a
probability for the radiation to pass through the first area and
the second area using a photon diffusion equation. Estimates or
calculations are thereby made by relating probabilities for the
radiation to pass through the first and second areas. Radiation may
also be transmitted into a third area including substantial portion
of a second target structure, such as the carotid artery. The
launch optics and collecting optics are preferably mounted on a
sensor patch designed to be engaged to the patient's skin.
[0023] The system can also include a portable unit which houses the
optical source, detectors, processor and wireless communications.
The portable unit can be sized such the patient can carry the
portable unit for extended periods.
[0024] An ultrasound transducer can also be used in combination
with the optical methods to generate an image of tissues within the
first area including the target structure to enable placement of
the one or more optical transmitters and the one more optical
receivers on the patient so as to enhance the accuracy of the
monitoring of the system. The ultrasound transducer can also be
configured to provide an ultrasound radiation pressure field to
modulate the target structure at a modulation frequency. The
ultrasound pressure modulation can be preformed continuously, or it
can be operated temporarily, in order to calculate a calibration
adjustment to the optical measurement.
[0025] The present invention is also embodied in a pad of pliable
material adapted to be engaged on the skin of a patient. Launch
optics can be mounted on the pad and positioned so as to enable
transmission of photons generated by an optical source through the
patient's skin into at least a first area including a substantial
portion of a target blood vessel and into a second area not
including a substantial portion of the target blood vessel.
Collecting optics can be mounted on the pad and positioned so as to
enable detection of the photons having been scattered within the
first area and the second area.
[0026] The present invention is also embodied in a method for
monitoring one or more parameters relating to blood of a patient.
Optical transmitters and receivers are engaged on a tissue
boundary, such as skin, of the patient. Radiation is generated
containing photons having a specific interaction with a target
chromophore in a target structure within the patient. The radiation
is transmitted through the optical transmitters into a first area
including a substantial portion of the target blood vessel and into
a second area not including a substantial portion of the target
structure. Optical receivers detect a portion of the radiation
scattered from at least the first area and the second area. The
parameter(s) relating to a patient's blood can be estimated based
in part on the detected scattered radiation from the first area,
and the scattered radiation from the second area.
BRIEF DESCRIPTION OF THE DRAWINGS
[0027] The teachings of the present invention can be readily
understood by considering the following detailed description in
conjunction with the accompanying drawings, in which:
[0028] FIG. 1 is a schematic view of an embedded vascular structure
that is an example of a suitable target for measurement with
embodiments of the present invention.
[0029] FIG. 2A is a schematic diagram of an apparatus according to
an embodiment of the present invention.
[0030] FIG. 2B is a close-up cross-sectional schematic diagram
illustrating an example of use of the apparatus of FIG. 2A.
[0031] FIG. 3 is a schematic diagram of a three-wavelength pulsed
optical source for use in embodiments of the present invention.
[0032] FIG. 4 is a schematic diagram of an all-electronic optical
source for use in embodiments of the present invention.
[0033] FIG. 5 is an example of a source of three wavelengths using
an Optical Parametric Oscillator for use with embodiments of the
present invention.
[0034] FIG. 6 is a schematic diagram illustrating an example of
signal broadening expected at a tissue boundary.
[0035] FIG. 7 is a schematic diagram of an apparatus using the
principle of time gated upconversion according an alternative
embodiment of the present invention.
[0036] FIG. 8 is a schematic diagram of an apparatus having two
pulsed optical sources according another alternative embodiment of
the present invention proposed implementation of the present
invention.
[0037] FIG. 9A is a schematic diagram depicting time-gated
upconversion detector that can be used in the apparatus of FIG.
8.
[0038] FIG. 9B is a schematic diagram depicting an alternative
time-gated upconversion detector that can be used in the apparatus
of FIG. 8.
[0039] FIG. 10 is a schematic diagram depicting a second apparatus
having a background-free time-gated upconversion detector according
to another embodiment of the present invention.
[0040] FIG. 11 is a graph of the absorption of oxy-hemoglobin and
water in the range 700-1200 nm, an expected variation of the
scattering coefficient as a function of wavelength, and an expected
difference between an artery with fully oxygen-saturated blood and
a vein where the oxygen saturation is 55%.
[0041] FIGS. 12A-12B are schematic diagrams of sensors that can be
used with embodiments of the present invention.
[0042] FIG. 12C is a three-dimensional diagram of an alternative
sensor according to an embodiment of the present invention.
[0043] FIG. 12D is a cross-sectional diagram taken along line D-D
of FIG. 12C.
[0044] FIG. 13 is a schematic diagram illustrating an example of
trans-dermal measurement of oxygenation of blood in the internal or
external jugular veins.
[0045] FIG. 14 is a schematic diagram of a portion of the
circulatory system showing examples of locations that may be probed
for blood oxygenation using embodiments of the present
invention.
[0046] FIG. 15 is a horizontal cross-section through the chest
showing examples of locations that may be probed for
blood-oxygenation using embodiments of the present invention.
[0047] FIG. 16 is a close-up vertical thoracic cross-section
illustrating a sensor placed in the left bronchus to probe
oxygenation of the left pulmonary artery and descending thoracic
aorta.
[0048] FIG. 17 is a schematic thoracic diagram illustrating an
example of trans-tracheal placement of a sensor according to an
embodiment of the present invention.
[0049] FIG. 18A is sagittal cross-sectional schematic diagram
illustrating a normal heart.
[0050] FIG. 18B is a sagittal cross-sectional schematic diagram
illustrating a heart exhibiting Patent Ductus Arteriosus (PDA).
[0051] FIG. 18C is a sagittal cross-sectional schematic diagram
illustrating a heart exhibiting Patent Foramen Ovale (PFO).
[0052] FIG. 19 is a thoracic axial cross-sectional schematic
diagram illustrating examples of sensor placement for cardiac
mapping in newborn infants according to an embodiment of the
invention.
[0053] FIG. 20 is a sagittal cross-sectional schematic diagram
illustrating examples of sensor placement for monitoring of fetal
blood oxygenation.
[0054] FIG. 21 is a system for making relative measurements
relating to blood oxygenation and/or pH according to an embodiment
of the invention.
[0055] FIG. 22 is a schematic block diagram of a monitoring system
according to an embodiment of the invention.
[0056] FIG. 23 is a side view of a monitoring patch placed in close
proximity to tissues containing two blood vessels of interest,
according to an embodiment of the invention.
[0057] FIG. 24 illustrates amplitude modulation of two different
wavelengths of electromagnetic radiation, according to an
embodiment of the invention.
[0058] FIGS. 25a and 25b are top views of sensor patches according
to embodiments of the invention.
[0059] FIG. 26 is a cross section along I-I' of the sensor patch of
FIG. 25a placed on tissues containing a blood vessel of interest,
according to an embodiment of the invention.
[0060] FIGS. 27a and 27b are plan views of sensor patches according
to further embodiments of the invention.
[0061] FIG. 28 is a cross section along II-II' of the sensor patch
of FIG. 27a placed on tissues containing blood vessel of interest,
according to an embodiment of the invention.
[0062] FIG. 29 is a plan view of a sensor patch having an array of
transmitters and receivers, according to an embodiment of the
invention.
[0063] FIG. 30 is a flowchart illustrating several steps relating
to measuring cardiac output according to embodiments of the
invention.
[0064] FIGS. 31a and 31b is a sensor patch according to a further
embodiment of the invention.
[0065] FIG. 32 shows steps involved in monitoring oxygenation and
cardiac output according to a further embodiment of the
invention.
[0066] FIG. 33 shows steps involved in monitoring oxygenation and
cardiac output according to a further embodiment of the
invention.
[0067] FIG. 34 shows steps involved in monitoring oxygenation and
cardiac output according to a further embodiment of the
invention.
DESCRIPTION OF THE SPECIFIC EMBODIMENT
[0068] Although the following detailed description contains many
specific details for the purposes of illustration, anyone of
ordinary skill in the art will appreciate that many variations and
alterations to the following details are within the scope of the
invention. Accordingly, the exemplary embodiments of the invention
described below are set forth without any loss of generality to,
and without imposing limitations upon, the claimed invention.
Glossary:
[0069] As used herein, the following terms have the following
meanings:
[0070] Continuous wave (CW) laser: A laser that emits radiation
continuously rather than in short bursts, as in a pulsed laser.
[0071] Diode Laser: Refers to a light-emitting diode designed to
use stimulated emission to generate a coherent light output. Diode
lasers are also known as laser diodes or semiconductor lasers. A
diode-pumped laser refers to a laser having a gain medium that is
pumped by a diode laser.
[0072] Mode locked laser: A laser that emits radiation in short
bursts, as in a pulsed laser. Typically these pulses are on the
order of 0.1-100 picoseconds in temporal length and preferably 1-50
picoseconds.
[0073] Highly Non-linear Fiber: A fiber characterized by having a
guiding core with properties that can be used to convert
electromagnetic radiation at one frequency to another provided
there is sufficient intensity at the originating frequency and the
fiber has sufficient length.
[0074] Upconversion Process: A process by which photons of a given
frequency are converted to photons of shorter wavelength (higher
frequency). This technique may be used, e.g., to bring infra-red
photons into the detection range of silicon detectors for example,
or may be used in a pulsed configuration to give temporal
selectivity in which photons are upconverted and hence
detected.
[0075] Non-Linear Crystal: A crystal made of a material having
special optical properties allowing the frequency of an incoming
electromagnetic wave to be shifted according to predictable rules
and conditions.
[0076] Optical Parametric Oscillator: A process by which a photon
at a pump frequency .omega..sub.P is converted in a material inside
a resonator to two photons of lower frequency, typically called the
signal and idler photons with the relationship:
.omega..sub.P=.omega..sub.SIG+.omega..sub.IDL
[0077] Optical Parametric Amplifier: A process by which a photon at
a pump frequency .omega..sub.P is converted in a material (but
without the need for an external resonator) to two photons of lower
frequency, typically called the signal and idler photons with the
relationship:
.omega..sub.P=.omega..sub.sig+.omega..sub.idl
[0078] As stated above, there are eight desirable characteristics
for cardiac output (CO) monitoring techniques: accuracy,
reproducibility or precision, fast response time, operator
independency, ease of use, continuous use, cost effectiveness, and
no increased mortality and morbidity associated with its use. None
of the present CO monitoring techniques satisfactorily combines all
eight criteria mentioned above.
[0079] The Fick principle involves measuring the oxygen consumption
(VO.sub.2) per minute (e.g., using a spirometer), measuring the
oxygen saturation of arterial blood using for example standard
pulse oximetry on the finger, and measuring venous oxygen
saturation in the pulmonary artery or superior vena cava.
[0080] From these values, one can calculate:
CardiacOutput = Oxygen_Consumption ( ArterialSa O 2 - VenousSa O 2
) .times. [ Hb ] .times. 1.36 ##EQU00001##
where Arterial SaO.sub.2 and Venous SaO.sub.2 are respectively the
arterial and venous oxygen saturation, [Hb] is the blood hemoglobin
concentration and 1.36 is a factor subsuming the oxygen carrying
capacity of the hemoglobin. [Hb] can be related simply to the
hematocrit (Hct), a routinely measured parameter defined as the
percent of whole blood that is composed of red blood cells
(erythrocyte volume to total volume expressed as a percentage). The
range for Hct is 32-50% in "normal" "healthy" people. Hct does not
tend to change dramatically and quickly (unless the patient is
bleeding severely), so it is sufficient to take a sample every
4-6-8 hours for example and update the Fick calculation
periodically. Hematocrit (hct) can be measured, e.g., by taking a
sample of blood and spinning it down in a centrifuge and
calculating the volumes.
[0081] The Fick principle relies on the observation that the total
uptake of (or release of) a substance by the peripheral tissues is
equal to the product of the blood flow to the peripheral tissues
and the arterial-venous concentration difference (gradient) of the
substance. In the determination of cardiac output, the substance
most commonly measured is the oxygen content of blood, and the
venous saturation is measured in the pulmonary artery using a
catheter as for example described by Powelson et al., "Continuous
monitoring of mixed venous oxygen saturation during aortic
operations", Crit. Care Med. 20(3), 332-336 (1992). This gives a
simple way to calculate the cardiac output. The drawback of drift
associated with this type of catheter has been discussed by Souter
et al., "Jugular venous desaturation following cardiac surgery",
Brit. J. Anaesth. 81, 239-241 (1998). It is also highly invasive,
incompatible with ambulatory measurement, and poses risks of
infection due to vascular system breach (femoral or jugular vessel
insertion). The nature of the challenge is illustrated
schematically in FIG. 1. An embedded vascular structure of a body
100 includes an artery 102 and vein 104, for example the internal
jugular vein and artery in the neck. The vein 102 and artery 104
are located beneath the epidermis 106 and dermis 108 of the body
100. The vein and artery are embedded in and around subcutaneous
structures 110, e.g., fat, muscle, tendon, etc.
[0082] Assuming there are no shunts across the cardiac or pulmonary
system, the pulmonary blood flow equals the systemic blood flow.
Measurement of the arterial and venous oxygen content of blood
involves the sampling of blood from the pulmonary artery (low
oxygen content) and from the pulmonary vein (high oxygen content).
In practice, sampling of peripheral arterial blood is a surrogate
for pulmonary venous blood.
[0083] Embodiments of the present invention allow non-invasive or
minimally invasive measurement of venous oxygen saturation at a
point where the value trends correctly with a direct pulmonary
artery catheter measurement. One can apply the above-described Fick
principle to such a measurement thereby enabling measurement of
cardiac output in a non- or minimally invasive manner. Embodiments
of the present invention for measuring venous oxygen saturation can
also be made insensitive to the presence of shunts in the heart,
such as for example acquired ventricular septal defects, and as
such offer valuable adjunct information if PAC thermodilution or
Fick data are already available. This is the case when the sensor
is placed on the internal jugular vein.
[0084] The value of the venous oxygen saturation is also a useful
adjunct diagnostic parameter in its own right. Patients with low
cardiac output tend to have low venous oxygen saturation, for
example around 50. This low value results from the increased
extraction of oxygen in the body tissues due to the poor perfusion
resulting from low flow. However high mixed venous oxygen
saturation with low cardiac output can indicate a significant
left-to-right shunt across the heart, such as an acquired
ventricular septal defect. Embodiments of the present invention
where the sensor is placed on the internal jugular will allow a
measurement of venous oxygen saturation before the heart and
pulmonary system, and thus will be insensitive to the presence of
these shunts.
[0085] Also by way of example a presentation of high cardiac
output, high venous oxygen saturation, narrow arterio-venous
difference and low peripheral resistance might suggest to the
physician to test for septic shock. On the other hand cardiogenic
shock is associated with high peripheral resistance. Thus
measurement of cardiac output can help guide and monitor the
administration of drugs such as vasodilators/vasoconstrictors and
inotropes.
[0086] A number of different technologies have been developed that
utilize some interaction between ultrasound radiation and
electromagnetic radiation. However, these prior art technologies
are all distinguishable from the techniques described herein. For
example, embodiments of the present invention are superior to
standard ultrasound-tagged photon techniques in that they are not
limited by the ability of the apparatus to detect very small
frequency shifts on the detected photons. U.S. Pat. No. 5,212,667
to Tomlinson et al. and U.S. Pat. No. 5,174,298 to Dolfi et al.
teach the technique of ultrasound tagged frequency-modulated
imaging. Other patents teaching variations on the theme of
frequency-modulated ultrasound tagging techniques include U.S. Pat.
No. 6,815,694 to Sfez et al., U.S. Pat. No. 6,738,653 to Sfez et
al., U.S. Pat. No. 6,041,248, to Wang, U.S. Pat. No. 6,002,958 to
Godik, U.S. Pat. No. 5,951,481 to Evans, U.S. Pat. No. 5,293,873 to
Fang.
[0087] Ultrasound-tagged frequency modulated imaging relies on
observing the frequency shift induced by the photoacoustic effect
when an electromagnetic wave interacts in a medium with a sound
wave. The electromagnetic wave (having a characteristic frequency
.omega..sub.OPT) receives a frequency shift at the ultrasound
frequency .omega..sub.US to either the + or - side of the carrier
wave .omega..sub.OPT. Heterodyne or interferometric techniques are
then used to decouple the frequency shifted wave from the carrier
wave. Implementation of the technique requires sophisticated lasers
with narrow linewidths and concomitantly long coherence lengths in
order to resolve the two frequencies. U.S. Pat. No. 6,002,958 to
Godik teaches the study of the amplitude modulation induced on an
electromagnetic wave by the ultrasound beam and scanning the
ultrasound beam in order to form an image of the absorber.
[0088] U.S. Pat. No. 6,264,610 to Zhu teaches the use of ultrasound
and near-IR imaging as adjunctive imaging techniques, but does not
attempt a physical link between the two techniques.
[0089] U.S. Pat. No 5,452,716 to Clift teaches the use of
two-wavelength probing using one wavelength specific to the
substance being probed and a reference field characterized by
another wavelength. This patent does not teach any form of temporal
gating, any form of targeting a structure, or any form of depth
control using co-located optical and ultrasound fields.
[0090] U.S. Pat. No. 6,445,491 to Sucha et al. and U.S. Pat. No.
5,936,739 to Cameron et al. teach the use of optical parametric
processes to amplify signals in imaging systems. Neither of these
patents teaches the use of upconversion to produce a signal which
is necessarily free from background contamination from for example
fluorescence processes or Raman scattering. Neither of the patents
teaches the use of the very fast non-linearities found in fiber
Optical Parametric Amplifiers to yield time-gated information in a
straightforward manner.
[0091] U.S. Pat. No. 5,451,785 to Faris teaches the use of
upconversion processes in a transillumination imaging system.
[0092] U.S. Pat. No. 6,665,557 to Alfano et al. teaches
spectroscopic and time-resolved optical methods for imaging tumors
in turbid media where time gating of the ballistic and
near-ballistic photons is used to improve the reconstruction of the
image. The more diffusely scattered photons are rejected in this
technique and no attempt is made to localize the interaction using
ultrasound.
[0093] U.S. Patent Appl. No. 2004/0122300 A1 Boas et al., U.S.
Patent Appl. No. 2004/0116789 to Boas et al., U.S. Pat. No.
6,332,093 to Painchaud et al., U.S. Pat. No. 5,630,423 to Wang et
al., U.S. Pat. No. 5,424,843 to Tromberg et al. and U.S. Pat. No.
5,293,873 to Fang teach variations on the theme of Photon Migration
Spectroscopy, Photon Migration Imaging (PMI), Diffuse Optical
Tomography (DOT), or Diffuse Imaging, where photons from a source
diffuse through the target and are detected using detectors placed
at various distances from the source launch point. The
characteristics of the diffusing photons are interpreted to yield
functional and structural information about the medium they have
diffused through. No attempt is made to "tag" these photons to
localize the region of interaction. No attempt is made to time-gate
the detected signal. Embodiments of the present invention are
superior to Photon Migration Imaging (PMI, DOT etc) in that they
allow accurate depth and location localization of the target.
[0094] Embodiments of the present invention are also superior to
speckle based imaging techniques because they are insensitive to
the speckle decorrelation time of the tissue being probed. This
speckle decorrelation is very fast in larger vascular structures
with flowing blood inside, preventing use of speckle-based
techniques in the types of vessels the current invention aims to
address.
[0095] Embodiments of the present invention can also be designed in
such a way as to be insensitive to the presence of epidermal
melanin (unlike many of the wavelengths used in PMI/DOT and
ultrasound tagged spectroscopy and imaging). Embodiments of the
present invention can also be designed in a manner that will not
suffer from significant solar or environmental background light
contamination.
[0096] Embodiments of the present invention do not require the
development of sophisticated single frequency lasers and
interferometric detection techniques. As a result embodiments of
the present invention will be simpler to implement and more
technologically robust in a clinical setting. Apparatus according
to embodiments of the present invention can use proven
telecommunication-based fiber-based technology to yield a robust,
small, and efficient product.
[0097] Embodiments of the present invention do not require 2-D
imaging arrays or cameras (for example CCD cameras), and in
particular do not require infra-red detector arrays such as InGaAs
CCDs. These devices need to be cooled to achieve low noise
conditions, further complicating the experimental/clinical
implementation. Apparatus according to embodiments of the present
invention can use proven single element silicon detectors which do
not need to be cooled and which do not need extensive computational
support.
[0098] FIG. 2A is a schematic block diagram of a diagnostic
apparatus 200 according to an embodiment of the present invention.
The apparatus 200 generally includes an optical source 202, launch
optics 204, an ultrasound transducer 206, collection optics 208, an
optical detector 210, associated electronics such as a filter 212
and an optional display 214. The optical source 202 provides pulsed
electromagnetic radiation. The launch optics 204 may include one or
more optical fibers 205 that couple the electromagnetic radiation
from the optical source 202 to a body 201. Similarly the collecting
optics 208 collect optical signals reflected from within the body
201. The collecting optics 208 may also include one or more optical
fibers 209 that couple signals scattered electromagnetic radiation
to the optical detector 210. The optical source 202 may supply a
timing signal (which may be either optical or electronic) to
trigger a detector source 211 that provides an optical signal used
in detection of the scattered radiation.
[0099] In some embodiments the launch optics 205, ultrasound
transducer 206, and collecting optics may be mounted together in a
handpiece to form a combined ultrasound optical sensor 203. In
other embodiments, the detector 210 may be part of the sensor 203
without the need for collecting optics. In some embodiments, the
optical source 202, optical detector 210, detector source 211,
filter 212, display 214 and an ultrasound generator 207 may be part
of a remote unite 213 coupled to the sensor 203 by fiberoptics 205,
209 and electrical cables. The remote unit 213 may include a system
controller 215. The system controller 215 may include a central
processor unit (CPU) and a memory (e.g., RAM, DRAM, ROM, and the
like). The controller 215 may also include well-known support
circuits, such as input/output (I/O) circuits, power supplies
(P/S), a clock (CLK), Field Programmable Gate Arrays (FPGAs) and
cache. The controller 215 may optionally include a mass storage
device such as a disk drive, CD-ROM drive, tape drive, or the like
to store programs and/or data. The controller may also optionally
include a user interface unit to facilitate interaction between the
controller 215 and a user. The user interface may include a
keyboard, mouse, joystick, light pen or other device. The preceding
components may exchange signals with each other via a controller
bus. In addition, the optical source 210, detector source 211,
filter 212, display 214 and an ultrasound generator 207 may
exchange signals with the controller 215 via the system bus
216.
[0100] The controller 215 typically operates the optical source,
202, ultrasound generator 207, optical detector 210, detector
source 211, filter 212 and display 214 through the I/O circuits in
response to data and program code instructions stored and retrieved
by the memory and executed by the processor. The program code
instructions may implement embodiments of the diagnostic technique
described herein. The code may conform to any one of a number of
different programming languages such as Assembly, C++, JAVA,
Embedded Linux, or a number of other languages. The CPU forms a
general-purpose computer that becomes a specific purpose computer
when executing program code. Although the program code is described
herein as being implemented in software and executed upon a general
purpose computer, those skilled in the art will realize that the
method of pulsed pumping could alternatively be implemented using
hardware such as an application specific integrated circuit (ASIC)
or FPGA or other hardware circuitry. As such, it should be
understood that embodiments of the invention can be implemented, in
whole or in part, in software, hardware or some combination of
both.
[0101] Operation of the apparatus 200 may be understood with
respect to the close-up schematic diagram depicted in FIG. 2B. An
embedded target structure within the body 201 such as an artery AR
or vein VE can be identified by ultrasound imaging.
[0102] The ultrasound generator 207 and transducer 206 can be used
to do both the ultrasound imaging and the target modulation. Once a
target has been located, the apparatus 200 switches between a
regular ultrasound mode (imaging) and a radiation pressure
modulation mode, firing tone bursts to modulate the target. The
basic approach is first to image to choose a location to deliver
radiation pressure and then to apply the appropriate phase to the
array elements of the transducer 206 to have a focus at the
location of interest. The radiation pressure is supplied by
applying a tone burst (many cycles of electrical signal at the
frequency of operation of the array) from the ultrasound generator
207 to the elements of the array in the transducer 206. The
repetition rate at which the tone burst is applied is the frequency
at which the radiation pressure is applied. This repetition rate is
constrained at the upper end by the fundamental frequency of the
ultrasound transducer 206, i.e. the tone burst cannot have a higher
repetition rate than the fundamental frequency of the transducer
itself. By way of example, the ultrasound transducer 206 can
operate at fundamental frequencies in the range 2-50 MHz, and
preferably from 2-15 MHz. The tone bursts may produce radiation
pressure modulation occurring at frequencies between 50 Hz and 750
kHz.
[0103] The sensor 203 is then placed proximate to a tissue boundary
TB of the boundary 201. The target structure is then vibrated using
radiation pressure from the transducer 206 and illuminated with a
diffuse photon field with a characteristic frequency
.omega..sub.INJ delivered from the optical source 202 via the
launch optics 204. The radiation-pressure modulation of the target
is detected by its effect on the emerging photon field at the
detector (e.g., via the collecting optics 208). In the example
depicted in FIGS. 1 and 2A, it may be possible to measure both
venous and arterial oxygenation separately by illuminating and
modulating the vein and then separately illuminating and modulating
the artery. In the case where the target is the internal jugular
vein, the corresponding arterial structure is the carotid artery.
This method, when it can be used, will implicitly provide a
calibration signal. Cardiac output can then be calculated from the
Fick Principle, as described above.
[0104] To make the measurement a biological structure within the
body 201, such as the pulmonary artery, descending branch of the
aorta, internal jugular, or external jugular, is located in a
standard manner with medical imaging. Once found the combined
ultrasound/optical sensor 203 can be positioned proximate to the
targeted structure. This can either be external dermal placement,
e.g., on the neck in the case of the internal jugular vein, or an
inserted catheter, either endotracheally for direct access to the
left pulmonary artery and thoracic aorta, or trans-esophageally for
access to the right pulmonary artery. The sensor 203 is preferably
positioned such that the distance between the emitting tip of the
launch optics 204 and the lumen of the targeted vessel is
approximately minimized.
[0105] The ultrasound transducer 206 is used to physically modulate
(vibrate) the selected target using ultrasound radiation pressure.
The ultrasound transducer 206 is designed to focus its acoustic
output into the target at various modulation frequencies. Examples
of ultrasound transceivers that can provide such focused output
include phased array ultrasound transceivers and single element
ultrasound transducers with imaging designs. Phased array
transducers typically have an array of ultrasound transducer
elements that are narrow and have a wide acceptance angle so that
energy from various angles is collected, and so that several
elements (if not all) in the array contribute to the focusing at a
certain location. To generate a beam, the various transducer
elements are pulsed at slightly different times. By precisely
controlling the delays between the transducer elements, beams of
various angles, focal distance, and focal spot size can be
produced. Furthermore, for a given point within the targeted tissue
a unique set of delays will maximize the constructive interference
of acoustic signals from each of the transducer elements. Such
transducers can therefore selectively modulate particular
structures within the target without modulating surrounding
structures. Beam forming in ultrasound refers to the signal
processing scheme used to focus the signals from various
transducers. The energy is preferentially deposited using focusing
to allow the application of radiation pressure at the location of
interest with a relatively low level of input signal.
[0106] Examples of suitable ultrasound transducers include, for
example, the GE Logiq 7 BTO3 made by General Electric of Fairfield,
Conn., or the Aspen.RTM. Echocardiography System made by Siemens
(Acuson) of Mountain View, Calif. Other suitable array transducers
are made by Philips (The Netherlands), or Hitachi (Japan). It is
best to choose an instrument that is used commonly in hospitals say
to image the heart.
[0107] An ultrasound imaging system can also be used in association
with the ultrasound generator 207 and transducer 206 to locate the
blood vessels in order to orient the delivery of the pulsed
radiation from the optical source 202. The imaging system can be
incorporated into the system controller 215. The transducer 206 can
be a piezo type transducer as used in the above-described
commercially-available ultrasound machines or a cMUT (capacitative
Micromachined Ultrasonic Transducer), see X. Jin, I. Ladabaum, B.
T. Khuri-Yakub. "The Microfabriction of Capacitive Ultrasonic
Transducers", J. Microelectromechanical Systems vol. 7, pp 295-302,
September 1998. and U.S. Pat. No. 6,262,946 to Khuri-Yakub et al,
both of which are incorporated herein by reference. Using the cMUT
will allow a very compact 2-D array to be made. Such compact arrays
are very important for ring-shaped transducers such as that shown
in FIGS. 12C-12D for the trans-tracheal/trans-esophageal
applications.
[0108] Using an array or other beam-forming transducer one can
steer the ultrasound from artery to vein using phase, and
alternately modulate each one, allowing a direct calibration of the
optical signal. For example one can steer the beam from internal
jugular to carotid artery, alternatively sampling 100% oxygen
saturated blood and the venous blood with reduced saturation. The
ultrasound imaging system can also be used to derive the width of
the arteries and veins, and the blood flow velocity using Doppler
shift of the scattered ultrasound. Such a measurement can provide
an estimate of the cardiac output that can be compared to cardiac
output as derived from the use of the apparatus 200. This adjunct
measurement will have additional diagnostic value as discussed
above for the diagnosis of shunts, septic and cardiogenic shock
etc.
[0109] Once the ultrasound transducer 206 and launch optics 204 are
aligned with respect to the targeted vessels, the array of
transducers in the ultrasound imaging system will all be fired,
with appropriate phase delays, with a burst of energy to deliver
radiation pressure at the focus as determined by the phase delays.
The focus of the acoustic signal can be chosen to be inside the
vessel acting on the blood cells, or on the side walls of arteries.
The radiation pressure associated with the acoustic pulse which is
equal to the acoustic intensity divided by the speed of sound in
the medium, will act to impart a movement on the cells or arterial
walls on which it acts. The use of radiation pressure
(alternatively "radiation force") to induce motion in a target
which is then detected by conventional ultrasound techniques has
been described by Nightingale et al "Acoustic Radiation Force
Impulse Imaging: In Vivo Demonstration of Clinical Feasibility",
Ultrasound in Medicine and Biology, 28(2): 227-235, (2002) and in
U.S. Pat. No. 6,371,912 to Nightingale et al, both of which are
incorporated herein by reference. Embodiments of the present
invention are superior to this technique in that they will permit
functional (oxygenation) information to be derived from the target,
whereas in the aforementioned prior art only mechanical information
(stiffness, elasticity etc) is derived.
[0110] In this fashion, the optical signal, which relates to the
oxygen content in the blood cells in the target volume, will be
modulated at the frequency at which the radiation pressure pulse is
applied, a .omega..sub.RPM. For instance, using a 7.5 MHz imaging
system, one can use a burst of say 10 cycles at any repetition rate
up to around 750 kHz as determined by the physical and mechanical
properties of the target and the experimental implementation. It
may be possible to tune the interpulse spacing (the repetition
rate) in the tone burst to resonantly modulate the target depending
on its elastic properties. It may also be possible to tune the
ultrasound fundamental frequency to optimize its interaction with
the desired target (blood cells, vessel walls etc). In this manner
the detector 210 may detect only those photons which have
interacted with the desired target 201.
[0111] The optical source 202 may be configured to deliver the
temporally correlated groups of photons at a repetition rate of
between about 100 kHz and about 500 MHz, preferably between about 1
MHz and about 250 MHZ, more preferably between about 10 MHz and
about 200 MHz. The groups of photons may be in the form of pulses
having pulse widths in the range of about 1 picosecond to about 1
nanosecond, preferably, about 1 to 100 picoseconds, more preferably
about 5 to 50 picoseconds. The photons may be characterized by
wavelengths between about 650 nm and about 1175 nm, preferably
between about 650 nm and about 930 nm or between about 1020 nm and
about 1150 nm.
[0112] The optical source 202 provides temporally correlated
photons at two or more different wavelengths. For example radiation
from a pulsed laser may be incident on a device that converts
radiation at the fundamental frequency of the laser into a pair of
photons at two different predetermined frequencies. Such a device
could be a nonlinear crystal causing Spontaneous Parametric Down
Conversion (SPDC) as for example described by Shi and Tomita,
"Highly efficient generation of pulsed photon pairs with bulk
periodically poled potassium titanyl phosphate", J. Opt. Soc. Am.
B. 21(12) 2081-2084 (2004), or a highly non-linear fiber source as
described by Rarity et al., "Photonic crystal fiber source of
correlated photon pairs", Opt. Exp. 13(2), 534-544 (2005).
[0113] Alternatively the optical source 202 may include a
non-linear crystal phased matched to act as an optical parametric
oscillator (OPO) to provide a temporally correlated photon pair. An
OPO takes a fundamental electromagnetic wave at frequency
.omega..sub.P1 and converts it to two new frequencies called the
signal and idler, .omega..sub.SIG and .omega..sub.IDL related by
the equation
.omega..sub.P1=.omega..sub.sig+.omega..sub.idl
where the signal and idler waves are emitted in temporal
coincidence.
[0114] The OPO may be driven by the second harmonic of a pulsed
laser operating at a fundamental frequency .omega..sub.P1 to create
two new frequencies called the signal and idler, .omega..sub.SIG
and .omega..sub.IDL related by the equation
2.omega..sub.P1=.omega..sub.sig+.omega..sub.idl
where 2.omega..sub.P1 is the second harmonic of the fundamental
frequency. For example the drive laser may be a mode-locked or
Q-switched Nd:YAG laser operating at 1064 nm, giving a second
harmonic wave at 532 nm. This wave in turn is used to drive the
OPO. In this manner three clinically useful, temporally coincident
photon waves at 1064 (.omega..sub.P1), 1030 (.omega..sub.sig) and
1100 (.omega..sub.idl) may be generated. The nonlinear crystal may
be selected from a variety of substances, for example BBO, LBO,
KTP, KTA, RTP, or periodically poled materials such as periodically
poled lithium Niobate (PPLN), periodically poled stoichiometric
lithium tantalate (PP-SLT) and the like. Such materials are
described, e.g., in the freeware program SNLO distributed by Sandia
National Laboratories, Albuquerque, N. Mex.
[0115] By way of example, the optical source 202 may include a
pulsed solid state laser, for example a picosecond mode-locked
laser such as the picoTRAIN.TM. series compact, all-diode-pumped,
solid state picosecond oscillator manufactured by High-Q Lasers of
Kaiser-Franz-Josef-Str. 61 A-6845 Hohenems Austria. The source may
also be a mode-locked fiber laser, such as the picosecond version
of the Femtolite.TM. D-200 from IMRA America Inc., Ann Arbor Mich.
48105. Alternatively, a picosecond pulsed diode such as the PicoTA
amplified picosecond pulsed laser diode heads manufactured by
Picoquant, of Berlin, Germany, may be used as the optical source
202. The optical fibers 205 coupling the optical source 202 to the
launch optics 204 may be, e.g., single mode fiber optic, such as
the P1-980A-FC-2-Single Mode Fiber Patch Cable, 2m, FC/PC
manufactured by Thorlabs, Inc. of Newton, N.J. Radiation coupled
from the optical source 202 to the target 201 via the launch optics
204 is used, e.g., to illuminate the lumen of a selected blood
vessel with pulses of radiation at two or more different
wavelengths carefully chosen to have deep penetration into tissue,
to have differing affinities for oxy-and deoxy-hemoglobin, or for
oxy-hemoglobin and met-hemoglobin, but to have substantially
similar scattering cross-sections and anisotropy parameters.
[0116] Some of the radiation scatters from the target 201 and is
collected by the collecting optics 208 and/or detector 210. By
detecting pairs or multiplets of photons at different wavelengths
returning from the target tissue in substantial temporal
coincidence, it can be inferred that the coincident photons have
traveled approximately the same path length in the tissue. This is
the main difference between making measurements in clear
transparent media where the Beer-Lambert law may be presumed to
apply, and making measurements in turbid media where elastic
scattering causes a substantial and generally indeterminate
pathlength increase, as discussed by Okui and Okada, "Wavelength
dependence of cross-talk in dual-wavelength measurement of oxy- and
deoxyhemoglobin", J. Biomed. Opt. 19(1), 011015 (2005).
[0117] The detector is coupled to a filter 212 that selects
coincident photon signals having modulation at the radiation
pressure modulation frequency or a harmonic thereof. The filter 212
may be coupled to the display 214, e.g. a CRT screen, flat panel
screen, computer monitor, or the like, that displays the results of
the aforementioned process in a manner readily interpretable, e.g.,
in the form of text, numerals, graphical symbols or images.
[0118] By detecting arrival rates of pairs or multiplets of photons
at the frequency of the radiation pressure modulation or a harmonic
of the radiation pressure modulation frequency, one can infer that
these photons interacted with the radiation-pressure-modulated
target 201. If the target 201 contains the oxygenated or
deoxygenated forms of hemoglobin (Hb), the detected pair or
multiplet coincidence rate will be altered depending on how the
wavelengths were selected. The extent to which the detection rate
is altered can be correlated to the oxygenation level of the target
or to the pH in the target. The met-hemoglobin absorption spectrum
is dependent on pH as shown in Zijlstra et al., "Visible and Near
Infrared Absorption Spectra of Human and Animal Haemoglobin,
1.sup.st ed. Utrecht: VSP Publishing; 2000, page 62. Thus a
non-invasive probe of met-Hb absorption may be used to probe the pH
of the structure being targeted.
[0119] There are many possible configurations for the optical
source 202 of FIG. 2A. For example, FIG. 3 is a schematic diagram
of a three-wavelength pulsed optical source 300 that emits three
laser pulses at the three wavelengths with temporal coincidence.
This could be the OPO source described above. Alternatively the
source 300 generally includes a pulsed laser 302, a seed source
304, and a highly non-linear fiber (HNLF) 306. Optics, 308 such as
one or more lenses couple pump radiation at a drive frequency
.omega..sub.p to the HNLF 306. A 2.times.2 coupler 310 couples seed
radiation at a frequency .omega..sub.s from the seed source 304
into the HNLF 306. When .omega..sub.p and .omega..sub.s are
properly chosen, the HNLF 306 acts as an optical parametric
amplifier (OPA) that produces three temporally correlated
electromagnetic waves at three frequencies: pump radiation at
.omega..sub.p, amplified seed radiation at .omega..sub.s and idler
radiation at an idler frequency .omega..sub.idl given by:
.omega..sub.idl=2.omega..sub.P-.omega..sub.s.
[0120] For example, if .omega..sub.p corresponds to a vacuum
wavelength of 1064 nm and .omega..sub.s corresponds to a vacuum
wavelength of 1100 nm, .omega..sub.idl corresponds to a vacuum
wavelength of about 1030 nm.
[0121] The fiber 306 preferably has a non-linearity that is high
enough to allow non-linear optical effects to occur efficiently in
a reasonable length of fiber, and where the non-linearity is
sufficiently fast to create the required temporal synchronization
between the pump, seed and idler waves. Such fibers may be obtained
from Crystal Fibre of Birkenrod, Denmark, for example the
NL-5.0-1065 type. The non-linear optics underlying the conversion
have been described by for example, Ho et al., "Narrow-linewidth
idler generation in fiber four-wave mixing and parametric
amplification by dithering two pumps in opposition of phase", J.
Lightwave. Tech. 20(3),469-476 (2002), which is incorporated herein
by reference. The drive frequency .omega..sub.p may be provided by
a high repetition rate mode-locked picosecond laser, such as the
picoTRAIN.TM. series compact, all-diode-pumped, solid state
picosecond oscillator manufactured by High-Q lasers of
Kaiser-Franz-Josef-Str. 61 A-6845 Hohenems Austria or a mode-locked
fiber laser, such as the picosecond version of the Femtolite.TM.
D-200 from IMRA America Inc., Ann Arbor Mich. 48105.
[0122] In the source 300 the seed source 304 may be a distributed
feedback (DFB) or DBR (Distributed Bragg Reflector) laser, for
example the EYP-DBR-1063-00100-2000-SOT02-0000 diode laser
manufactured by Eagleyard Photonics, Berlin Germany. There are a
number of different possible configurations for the pulsed laser
302. Generally, the pulsed laser 302 should be capable of providing
picosecond pulses of pump radiation to the HNLF 306. FIG. 4 is a
schematic diagram of an all-electronic optical source 400 of
picosecond pulses which could be used as the pulsed laser 302 of
FIG. 3. The source 400 generally includes a diode laser 402 an
electro-optic modulator (EO) 404 a Faraday isolator 406 and a doped
fiber amplifier 408. The diode laser 402 provides radiation at
.omega..sub.p which is modulated by the EO modulator 404 to create
weak picosecond radiation pulses 401, which are coupled to the
fiber amplifier 408. The Faraday isolator 406 transmits pulses to
the fiber amplifier 408 and blocks radiation from being reflected
back towards the EO modulator. A fiber pump source 410 provides
fiber pump radiation (e.g., at a vacuum wavelength of 980 nm) to
the cladding or core of the fiber amplifier 408. The fiber
amplifier may include a dump for the pump laser so that fiber pump
radiation does not oscillate through fiber amplifier 408. The
amplifier 408 amplifies the weak pulses 401 to create amplified
pulses 403 that can be fed to the HNLF 306.
[0123] By way of example, the diode laser 402 is a continuous wave
(CW) tunable DFB or DBR diode laser, such as the
EYP-DBR-1063-00100-2000-SOT02-0000 diode laser manufactured by
Eagleyard Photonics, Berlin Germany. The EO modulator 404 may be a
Model 4853 6.8/9.2-GHz Modulator from New Focus (Bookham) San Jose,
Calif. The Faraday isolator 406 may be a model 411055 from
Electro-Optic technology, of Traverse City, Mich. The fiber
amplifier 408 may be doped with Ytterbium or Neodymium, such as the
DC-225-22-Yb made by Crystal Fibre (Birkerod, Denmark). The fiber
pump may for example be a model 4800, 4 W, Uncooled, Multi-Mode
pump module from JDS Uniphase, of San Jose, Calif.
[0124] As discussed above, the optical source 202 may include
produce the correlated photons by optical parametric oscillation.
FIG. 5 is an example of such an optical source 500. The source 500
generally includes a pulsed laser 502, a second harmonic generator
(SHG) 504, a dichroic mirror 506 and an optical parametric
oscillator (OPO) 508. The pulsed laser produces pump radiation at a
frequency .omega..sub.p. The second harmonic generator interacts
with the pump radiation to produce second harmonic radiation at
double the frequency of the pump radiation, i.e., at
2.omega..sub.p. The SHG 504 may be less than 100% efficient at
doubling the frequency of the pump radiation. The dichroic mirror
506 deflects pump radiation that makes it through the SHG 504. In
the OPO 508, some of the second harmonic radiation is converted to
signal and idler radiation, respectively at frequencies
.omega..sub.sig and .omega..sub.idl that are related by:
2.omega..sub.P=.omega..sub.sig+.omega..sub.idl
[0125] The pulsed laser 502 may be of any of the types described
above. The second harmonic generator may be a non-linear crystal of
any of the types described above phased matched for second harmonic
generation. The OPO 508 may be a non-linear crystal of any of the
types described above phased matched for optical parametric
oscillation. The source 500 has the advantage of being tunable by
virtue of the OPO phase matching. The phase matching is
typically-tuned by adjusting e.g., the angle of the non-linear
crystal used in the OPO, or by changing its temperature.
Alternatively the poling period may be adjusted in periodically
poled materials to phase match at different wavelengths.
[0126] Radiation pulses from the source 200 may be affected by
traveling through tissue. For example, FIG. 6 is a schematic
diagram of the signal expected at the tissue boundary TB shown in
FIG. 2B. Injected pulses of radiation at frequency .omega..sub.INJ
with a short pulse widths (e.g., about 1 to 50 picoseconds) are
delivered into the body 201 at the tissue boundary TB. An injected
pulse interacts with tissues in the body and emerges as a signal
pulse at an optical frequency .omega..sub.SIG, which may be
slightly different from .omega..sub.INJ as a result of interaction
with the ultrasound pulse. However any frequency shift occurring as
a result of interaction between the optical pulses and the
ultrasound pulses will be insignificant compared to the natural
linewidth of the picosecond laser pulse as a result of the
time-bandwidth constraint which derives directly from the
Heisenberg Uncertainty Principle. Furthermore detection of this
ultrasound-induced frequency shift is not required in the proposed
embodiments of the invention, distinguishing this technique from
those in the prior art. The signal pulse is typically broadened
(e.g., to a pulse width of several hundred picoseconds to several
nanoseconds) compared to the injected pulse due to the random-walk
nature of photon propagation in turbid media, as shown by Turner et
al., "Complete-angle projection diffuse optical tomography by use
of early photons", Opt. Lett. 30(4), 409-411 (2005). This random
walk increases the effective pathlength considerably. The time at
which the photon arrives at the tissue boundary may be related to
its approximate pathlength through mathematical relationships, for
example the diffusion approximation or the Transport Equation.
[0127] The pulse spreading described above must be taken into
account in time-gated detection of the signal pulse. One possible
approach to taking such pulse spreading into account utilizes a
technique referred to herein as time gated upconversion. FIG. 7 is
a schematic diagram illustrating the principle of time gated
upconversion. The broadened signal pulse at .omega..sub.SIG
emerging from the tissue boundary TB with a pulse width .DELTA.T
of, e.g., a few nanoseconds, is mixed with a short mixing pulse
(e.g., pulse width .delta.t of order several picoseconds) of
radiation at an optical frequency .omega..sub.P2. A master
oscillator or a secondary slave oscillator may provide the short
mixing pulse at .omega..sub.P2. The mixing takes place in an
upconverter such as a fiber OPA or a mixing crystal. Mixing can
only occur when the two pulses are temporally and physically
overlapped, so by strobing the mixing pulse through the emerging
signal pulse it is possible to time gate the signal that is to be
detected. This upconversion process may be accomplished in a manner
that is highly efficient as described by Langrock et al.,
"Sum-frequency generation in a PPLN waveguide for efficient
single-photon detection at communication wavelengths", Stanford
Photonics Research Center Annual Report (2003) D-19-D-21, which is
incorporated herein by reference.
[0128] FIG. 8 is a schematic diagram illustrating an alternative
optical signal generation and detection apparatus 800 for use with
embodiments of the present invention. The apparatus 800 includes
first and second pulsed optical sources 801, 802 that respectively
produce pulsed optical signals at optical frequencies
.omega..sub.P1 and .omega..sub.P2. The first source 801 serves as a
master oscillator for timing purposes and its output is used in one
of the aforementioned processes to create two or more pulses of
light at two or more wavelengths selected per the criteria
described above. A timing signal .PHI. is derived from the first
source 801 and used to trigger the second source 802, which
operates at substantially the same pulse repetition rate as the
first source 801, but with an adjustable delay (phase angle)
between the two pulse trains. The pulse train from the second
source 802 is mixed in an upconversion apparatus 804 with the
emerging signal at optical frequency .omega..sub.SIG from a tissue
boundary 807 and the time delay between the two sources is adjusted
to temporally gate the resulting signal, which is detected at a
detector 806. This permits background-free, time-gated analysis of
the emerging signal. The resulting upconverted signal may have an
optical frequency .omega..sub.UC of .omega..sub.P2+.omega..sub.SIG
or 2.omega..sub.P2-.omega..sub.SIG depending on the nature of the
upconversion apparatus 804. The two signals may be mixed, e.g.,
using a relay fiber 808 coupled to collection optics 810 and a
2.times.2 coupler 812 coupled to the relay optics and the second
source 802.
[0129] In some embodiments, the upconversion apparatus 804 may
include a local oscillator, e.g., a laser for time-gated
upconversion. For example, as depicted in FIG. 9A, the signal pulse
at .omega..sub.SIG and mixing pulse at .omega..sub.P2 are combined,
e.g., using a 2.times.2 coupler 902. Upconversion as described
above may then be used to create a new signal photon wave at either
.omega..sub.P2+.omega..sub.SIG or 2.omega..sub.P2-.omega..sub.SIG.
A local oscillator laser 904 produces a pulsed wave at an optical
frequency row and a repetition rate correlated to the ultrasound
tone burst that is mixed in a mixing stage 906 with the new signal
pulses before detection, generating a composite wave at optical
frequency .omega..sub.UC given by either
(.omega..sub.P2+.omega..sub.SIG+.omega..sub.LO) or
(2.omega..sub.P2.omega..sub.SIG+.omega..sub.LO) that is coupled to
the detector 210. The mixing stage 906 may be a waveguide of for
example a PPLN or PP-SLT, or a crystal of KTP or other material
with high optical non-linearity. In this manner a signal may be
generated that is temporally selected for an effective pathlength
in the tissue. Upconverting the signal from the near-IR (around 1
micron) to the visible (400-700 nm) in this manner allows the use
of silicon-based detector technology that has several advantages
over InGaAs technology as discussed by Langrock et al. For example
benefits include greater receiver sensitivity and lower dark counts
from the detector.
[0130] The signal may be further selected for a temporal
relationship to the modulating ultrasound tone burst from the
transducer 206 by triggering the local oscillator 904 with an
appropriate reference signal from the ultrasound source 207. For
example by triggering the local oscillator 904 at twice the
repetition rate of the tone burst, one can make a direct on/off
comparison between the signal coming back from the tissue in the
presence of, and absent the effect of the mechanical
modulation.
[0131] Alternatively, the upconversion apparatus 804 may provide
background free time gated amplification of the signal pulse. This
may alternatively be accomplished using fiber Optical Parametric
Amplification, e.g., as depicted in FIG. 9B. In an OPA-based
background-free time-gated upconversion detector 910, optical
signals at optical frequency .omega..sub.SIG emerging at a tissue
boundary 907 are coupled into a relay fiber 912 by collection
optics 914. The emerging optical signals at .omega..sub.SIG are
then mixed (e.g., using a 2.times.2 coupler 916) into a Highly
Non-Linear Fiber (HNLF) 918 with a drive pulse at optical frequency
.omega..sub.P2 from a pump source 920. The drive frequency p may be
provided by a high repetition rate mode-locked picosecond laser,
such as the picoTRAIN.TM. series compact. all-diode-pumped, solid
state picosecond oscillator manufactured by High-Q lasers of
Kaiser-Franz-Josef-Str. 61 A-6845 Hohenems Austria or a mode-locked
fiber laser, such as the picosecond version of the Femtolite.TM.
D-200 from 1 MRA America Inc., Ann Arbor, Mich. 48105. The signal
at .omega..sub.SIG is converted to a detected signal at
.omega..sub.DET by an Optical Parametric Amplification (OPA)
process in the fiber 918. The OPA process creates the detectable
signal .omega..sub.DET, e.g., through a four-wave mixing process
described by:
.omega..sub.DET=2.omega..sub.P2-.omega..sub.SIG
[0132] Since the upconversion process only happens when the drive
pulse at .omega..sub.p2 is present the upconversion can be time
gated. It should be noted that the frequency .omega..sub.DET of the
detected signal is higher than either the signal or drive
frequencies respectively. This means that the signal detected at
frequency .omega..sub.DET will be substantially free of
contaminating signals, e.g., from tissue autofluorescence (which
always occurs to longer wavelength than the excitation wavelength),
inelastic scattering internal to the fiber (Raman scattering for
example) which is also always to longer wavelength than the
fundamental, and other non-linear inelastic processes. By delaying
the onset of the mixing or upconversion pulse used in the detection
stage (802, 920), and then lengthening it in time using for example
the all-electronic source shown in FIG. 4, we may adjust the
detector to:
[0133] a) eliminate signal from photons which could not have
interacted with the target, and
[0134] b) include all possible contributions from photons which
could have interacted with the target. This is equivalent to
applying a Heaviside (step) function to the detected signal.
[0135] The aforementioned detection method may be more efficient
compared to slowly moving a short upconversion 1 mixing pulse
through the temporally broadened signal (FIGS. 6 and 7) by varying
the delay as this latter technique implicitly selects a small
subset of the photon trajectories, ignoring other possible
contributions.
[0136] The detected signal may be amplified in a time gated manner
by selecting a delay between the signal at .omega..sub.P1 from the
tissue boundary to be amplified and the drive pulse at
.omega..sub.P2. The drive pulse may be part of the beam from a
master laser or may preferably be produced by a second pulsed laser
operating at similar repetition rate and pulsewidth to the master
oscillator. The amplification of a particular segment of the
returning signal may also be selected by overlapping the two
signals in time using a variable delay line. Using this technique,
the signal at .omega..sub.P1 will also be amplified by gains of for
example 10-60 dB, as described by Ho et al. and references therein,
allowing very weak signals to be detected.
[0137] Other background-free time-gated upconversion detection
schemes can be implemented. For example FIG. 10 depicts an
alternative background-free time-gated upconversion detector 1000.
In the detector 1000 a master oscillator 1002 produces a first
master oscillator pulse at an optical frequency .omega..sub.P1. The
first master oscillator pulse is used to generate temporally
correlated photons (e.g., as described above) that are scattered
from a target tissue 1003 within a body 1001 to provide a signal.
Signal photons at an optical frequency .omega..sub.SIG emerging
from a tissue boundary 1007 are coupled into a fiber 1004, e.g.,
via relay optics 1006. After amplification in a doped section of
the fiber 1004, the signal photons are mixed (e.g., using a
2.times.2 coupler 1005) in a non-linear crystal 1008 with a second
time-delayed master oscillator pulse having an optical frequency
.omega..sub.P2. The non-linear crystal 1008 is phase matched for
frequency mixing of the signal photons and the second oscillator
pulse. The resulting upconverted signal is characterized by an
optical frequency .omega..sub.UC given by:
.omega..sub.UC=.omega..sub.P2+.omega..sub.SIG.
[0138] A temporal delay between the first and second oscillator
pulses is adjusted such that the time evolution of the signal
emerging from the tissue boundary can be probed. This allows early
arrival photons, which could not have interacted with the target by
virtue of their arrival time, to be gated out.
[0139] It should be understood that the signals referred to above
generally include two or more signal photons of different
wavelengths that are detected in coincidence. Coincidence detection
of the two signal photons can be accomplished by balanced
photoreceivers, for example New Focus (Bookham) model 1807 and
1817, San Jose, Calif. The wavelengths of interest can be isolated
by interference filters such as the RazorEdge.TM. and MaxLine.TM.
Laser and Raman filters from Semrock, Inc. of Rochester, N.Y.
Alternatively coincident photon pairs or multiplets can be detected
using high speed analog and digital electronics, for example time
correlated single photon counting equipment such as the SPC-134
from Becker and Hickl GmbH, Berlin, Germany, or boxcar integrators
such as the Model SR200 Boxcar from Stanford Research Systems,
Sunnyvale, Calif.
[0140] The time-gated amplified signal is analyzed to reveal the
component being modulated at the radiation-pressure modulation
frequency .omega..sub.RPM. This can be accomplished using lock-in
detection using for example a lock-in amplifier (e.g., a Stanford
Research Systems SRS Model 844) as the filter 212 in FIG. 2A.
[0141] The remaining signal by virtue of the above generation and
detection techniques must have:
[0142] a) Interacted with the target structure being modulated by
the radiation pressure field,
[0143] b) Been generated by photons at each of the two or more
selected wavelengths which traveled approximately the same path
length from the launch site, through the target being modulated,
and back to the detector.
[0144] The two or more wavelengths of the correlated photons
provided by the optical source 202 may be selected to have
different affinities for the various states of hemoglobin (oxy-Hb,
met-Hb, deoxy-Hb). The arrival of correlated photons at the
different wavelengths therefore can be interpreted to indicate for
example the oxygenation level or pH of the blood in the modulated
target structure. For example, if one radiation-pressure modulates
a blood vessel and its contents, and illuminates the area with two
wavelengths of light, one selectively absorbed by oxy-hemoglobin
and one substantially less selectively absorbed, the arrival rate
of correlated photon pairs will be higher if they traverse a
radiation-pressure-modulated vascularized area containing high
levels of deoxy-Hb (because one of the pair will be selectively
more absorbed in areas of higher oxygen saturation). By way of
example 1030-nm radiation is absorbed more strongly by
oxy-hemoglobin than 1064-nm radiation. Similarly, 11OO-nm is more
strongly absorbed by oxy-hemoglobin than 1064-nm radiation. These
three wavelengths may be conveniently generated as shown above.
They also have the added attraction of having substantially similar
elastic scattering coefficients, which will lead to a
simplification in calculation of the effective pathlength each
traverses. They also have substantially similar absorption in
water, leading to a simplification in assessing the potential
contribution for error in the measurement caused by non-hemoglobin
related absorption of the probe wavelengths.
[0145] FIG. 11 is a graph showing the absorption of oxyhemoglobin
(diamonds) and water (solid curve) in the range 700-1200 nm, the
nominal variation of the scattering coefficient as a function of
wavelength (squares), and the expected difference in absorption
between an artery with fully oxygen-saturated blood (SaO.sub.2=100)
and a representative vein where the oxygen saturation is 55%
(asterisks--Delta AV 55). The points at which the difference curve
crosses Y=0 are known as isosbestic points. There are two
isosbestic points in the absorption spectra of oxy-hemoglobin and
deoxyhemoglobin, one around 810 nm and one around 1135 nm. At these
wavelengths the absorption of blood in the vessel is independent of
oxygen saturation. These points are known to be useful for internal
reference calibration, for example to exclude the effects of volume
changes in the absorption resulting from pulsatile flow from the
heart.
[0146] The wavelength range 1025-1135 nm is characterized by having
reduced absorption as the venous oxygen saturation decreases. This
means that the signal derived as described in the embodiments of
the present invention will increase with decreasing saturation in
this wavelength range. The gradient of the absorption change with
respect to oxygen saturation at the 1135 nm isosbestic point is
also very steep, much more so than at 810 nm, making it of
significant potential value. Around this wavelength range, we may
make sensitive measurements at two or more wavelengths on each side
of the isosbestic point. The sign of the absorption change will
change from one side of the isosbestic point to the other.
[0147] The scattering function in FIG. 11 varies as the inverse
fourth power of the wavelength. This means that longer wavelengths
(for example from 1025 nm-1150 nm are not as affected by scattering
as shorter wavelengths from for example 700-930 nm). This
translates to a smaller increase in the effective pathlength
resulting from elastic scattering. The scattering function in the
1025 nm-1150 nm also does not vary significantly, indicating that
if we probe the target using wavelengths in this range we may
regard scattering as a secondary effect and model it as a
perturbation. This is not true in the 700-930 nm range, where the
scattering function varies by more than a factor of three.
[0148] The wavelength range 1025-1150 nm has rich structure in the
difference spectrum, has much lower scattering than the visible and
near-IR wavelength ranges, and has relatively modest water
absorption. This region offers several convenient and readily
available laser sources (Nd:YAG, Yb:Fiber lasers) which are known
from dermatology to have excellent penetration properties into
tissue.
[0149] It is possible to bias the selection of wavelengths to
enhance the diagnostic value of the measurement. For example, fetal
oxygenation levels are known to be substantially lower than the
conjugate maternal levels. Thus, the selection of wavelengths can
be biased to probe the fetus preferentially over the mother.
Furthermore, if it is desired to detect the pH of the blood in the
ultrasound-modulated target, one can inject probe photons at a
frequency known to be selective for met-hemoglobin absorption. For
example in the wavelength range from 800-1350 nm met-hemoglobin has
much stronger absorption than either oxy-hemoglobin or
deoxy-hemoglobin as shown in Kuenster J. T and Norris K. H.
"Spectrophotometry of human hemoglobin in the near infrared region
from 1000 to 2500 nm", J. Near Infrared Spectrosc. 259-65 (1994).
The wavelength range 1000-1300 nm and especially from 1100-1250 nm
is particularly sensitive to met-hemoglobin absorption. The
absorption spectrum of met-hemoglobin is known to be sensitive to
pH, as shown for example in Zijlstra et al., "Visible and Near
Infrared Absorption Spectra of Human and Animal Haemoglobin,
1.sup.st ed. Utrecht: VSP Publishing; 2000, page 62, and one may
thus infer the pH of the target from the coincidence arrival rate
of appropriately chosen photon pairs or triplets or higher
multiplets.
[0150] Embodiments of the present invention are distinguishable
from Diffuse. Optical Tomography, where the signal detected has
subsumed within it all possible absorbers in the path of the field
and no attempt is made to localize the absorber location. The
present technique is further distinguished from the various
practices of ultrasound-tagged optical spectroscopy because it does
not detect small frequency shifts or speckles on the emerging
photons. Instead, the present technique detects the modulation
imparted by physical motion of the target, which in turn affects
the optical absorption cross-section. The present invention is
insensitive to the very short speckle decorrelation time caused by
blood flow in the vessel, which would otherwise severely complicate
the detection of modulated photons in interferometric or
frequency-domain techniques. The present modulation technique
occurs at much higher frequency than other motion artifacts, for
example pulsatile flow from the heart beat, allowing it to be
decoupled in the signal analysis. This is important when, for
example, the technique is used to perform trans-abdominal fetal
oxygenation measurements where it is desirable to distinguish the
maternal and fetal oxygenation systems.
[0151] There are many possible designs for sensors that may be used
in embodiments of the present invention. For example, FIG. 12A
depicts an example of a sensor 1200 for transdermal measurements.
The sensor 1200 generally includes a substrate 1202, which may be
of a flexible plastic or similar material. A thin ultrasound
transducer 1204 is mounted on or embedded within the substrate. The
transducer 1204 receives power from an ultrasound transmitter and
sends return signals through a cable 1205. Optical signals are
transmitted and received through an optical fiber bundle 1206
containing launch and receive fibers terminated with coupling
optics 1208. The launch/receive fibers and coupling optics 1208 may
be mounted to or embedded with a substrate 1202, proximate the
transducer 1204. The launch/receive fibers may be used to both
transmit and receive optical signals. The fibers and coupling
optics 1208 are distributed in a more or less planar fashion. This
type of sensor may be used for transdermal measurements.
[0152] FIG. 12B depicts an alternative sensor 1210 that is a
variation on the sensor shown in FIG. 12A. A transducer 1214,
launch fibers and optics 1218, collection fibers and optics 1219
are mounted to or embedded within a substrate 1212 in a more or
less planar fashion. In this example, the transducer 1214 is
disposed between the launch fibers and the collection fibers. The
transducer 1214 receives power from an ultrasound transmitter and
sends return signals through a cable 1215. The launch fibers and
optics 1218 receive optical radiation from a source via a fiber
bundle 1216. The collection fibers and optics 1219 transmit signals
to a detector via another fiber bundle 1217.
[0153] Other sensor configurations may be useful for
trans-esophageal or trans-tracheal measurements. For example, FIGS.
12C-12D depict a sensor 1220 that may be inserted into the
esophagus or the trachea. The sensor 1220 includes a ring-shaped
substrate 1222 made of a bio-compatible material. Two or more
ultrasound transducers (or transducer arrays) 1224 are mounted to
the substrate 1222. The transducers are arranged to emit ultrasound
in an outward fashion as indicated by the arrows depicted if FIG.
12D. The transducers receive and transmit signals through a cable
1225. Arrays of launch/receive fibers 1228 are disposed on or
embedded within the substrate 1222 proximate the transducers 1224.
The launch/receive fibers 1228 receive or transmit optical signals
via a fiber bundle 1226. The ring-shaped sensor 1220 may be placed
in the esophagus. Alternatively, the sensor 1220 may be placed in
the left or right bronchus, through the trachea, e.g., at the end
of a tube that provides oxygen to the patient. Alternatively, the
sensor 1220 may be implanted into the patient's trachea and
providing a read out to small portable monitoring unit for
continuous ambulatory monitoring.
[0154] Use of the sensors and apparatus described above for the
monitoring of blood oxygenation can be accomplished in a variety of
different ways.
[0155] For example, FIG. 13 illustrates a simple case of
transdermal measurements of oxygenation in the interior or exterior
jugular vein of a patient. A sensor 1300, e.g., of the type
depicted in FIG. 12A or FIG. 12B may be placed against the
patient's neck in the vicinity of the spot marked with an X. The
sensor 1300 may be coupled to a remote unit of the type described
above with respect to FIG. 2A. Venous oxygen saturation in the
jugular vein can be measured using the ultrasound/optical technique
described above while arterial oxygenation can be measured using
standard pulse oximetry. Cardiac output can then be calculated from
the Fick principle as described above. Alternatively arterial
saturation may be measured by radiation-pressure modulating the
carotid artery instead of the internal jugular vein. Although a
single sensor 1300 is depicted on one side of the neck, two or more
such sensors (or one large sensor) may be placed on the dermis
simultaneously on the left and right side of the neck over both
internal jugular veins.
[0156] There are a number of different targets within the body that
are suitable for blood oxygen monitoring using embodiments of the
present invention. These can be understood with reference to the
anatomical diagrams of FIG. 14 and FIG. 15. For example, both right
and left internal jugular veins are potential targets as described
above. Measuring both simultaneously would probably be a superior
method. FIG. 16 illustrates three other possible sensor placements
may be used in conjunction with embodiments of the present
invention. First, a sensor A may be inserted using a bronchoscope
between two ribs (an intercostal space) next to the sternum. In
this case the sensor could be positioned right up against the
pulmonary artery (probably away from the aorta). This is the
optimum place to make the measurement of venous oxygen saturation
assuming that there are no defects in the heart. For example, if
there is an acquired ventricular septal defect, in which blood is
short-circuited from left ventricle to right ventricle, the oxygen
saturation of the pulmonary artery is abnormally high (e.g., about
80, whereas the incoming blood from the jugular vein may be around
50). Such a condition would result in a false reading for the
cardiac output measured using the Fick principle. However an
alternative prove site on the internal jugular vein gives an
adjunct measure of the cardiac output independent of heart defects.
So the two measurements would be complimentary.
[0157] Alternatively, as shown in FIG. 16, a sensor B may be placed
in the esophagus. The sensor B may be of the planar type depicted
in FIG. 12A or FIG. 12B or the ring type depicted in FIGS. 12C-12D.
A sensor C may also be placed in the left bronchus via the trachea.
These two probes will also sample the pulmonary arteries. The
trans-esophageal probe will sample the right pulmonary artery. The
trans-tracheal (bronchial) sensor C will potentially be able to
simultaneously probe the oxygen saturation in both the left
pulmonary artery (the venous saturation) and the descending
thoracic aorta (arterial saturation). This would eliminate the need
for external pulse oximetry to measure the arterial oxygen
saturation. Positioning of a sensor D within the left bronchus or a
sensor E within the right bronchus is illustrated in the dorsal
pull-away view of FIG. 17. Such trans-tracheal sensors may be the
ring-shaped sensor of the type depicted in FIGS. 12C-12D. The
sensors, A, B, C, D, E may be coupled to a remote unit of the type
described above with respect to FIG. 2A. Optical and ultrasound
signals can probe the chemistry of the cardiovascular system in the
manner described above.
[0158] Embodiments of the present invention also have application
to monitoring of neonatal blood oxygenation. Monitoring of neonatal
blood oxygenation is particularly useful in the cases of neonatal
heart defects as illustrated in FIGS. 18A-18C. FIG. 18A depicts an
example of a normal heart. Certain patients exhibit a heart defect
known as Patent Ductus Arteriosus (PDA). As illustrated in FIG.
18B, PDA is the persistence of a normal fetal structure (indicated
by the arrow) between the left pulmonary artery and the descending
aorta. Persistence of this fetal structure beyond 10 days of life
is considered abnormal. Other patients exhibit a defect known as
Patent Foramen Ovale (PFO). As shown in FIG. 18C, PFO is a
persistent opening in the wall of the heart (indicated by the
arrow) which did not close completely after birth. The opening is
required before birth for transfer of oxygenated blood via the
umbilical cord. This opening can cause a shunt of blood from right
to left, but more often there is a movement of blood from the left
side of the heart (high pressure) to the right side of the heart
(low pressure). Normally this opening closes in the first year of
life; however in about 30% of adults a small patent foramen ovale
is still present. Diagnosis of both PDA and PFO may be helped by
measurement of venous oxygen saturation.
[0159] In newborn infants (neonates) the distance across the thorax
may be small enough that in addition to trans-esophageal and
trans-tracheal, and trans-dermal for the internal jugular, it may
be possible to operate the diagnostic apparatus transdermally with
a sensor placed directly on a neonate's chest surface. The sensor,
e.g., of the planar type depicted in FIGS. 12A-12B, is placed
proximate the heart or a blood vessel of interest. The target area
is a neonatal cardiovascular system. As illustrated in the
cross-sectional diagram of FIG. 19 the measurement may be made in
either a reflection mode or trans-illumination mode (in one
side--out the other). In the reflection mode, optical signals are
transmitted and received via a common sensor 1902. In the
trans-illumination mode a transmitter unit 1904 sends optical
signals through an infant's thorax. Scattered photons of radiation
from these signals are collected by one or more sensors 1906, 1908
that are positioned to probe radiation scattered from particular
structures within the neonatal anatomy such as the pulmonary
artery. The sensors 1906, 1908 may be coupled to a remote unit of
the type described above with respect to FIG. 2A. Optical and
ultrasound signals can probe the chemistry of the neonatal
cardiovascular system in the manner described above.
[0160] Further embodiments of the invention include using
diagnostic apparatus of the type described herein for fetal
monitoring. For example, as depicted in FIG. 20, one or more
sensors 2002A, 2002B, 2002C, e.g., planar sensors of the type
depicted in FIGS. 12A-12B, may be placed on a pregnant woman's
abdomen to probe the fetal cardiovascular system. The sensors
2002A, 2002B, 2002C may be coupled to a remote unit of the type
described above with respect to FIG. 2A. Optical and ultrasound
signals can probe the chemistry of the fetal cardiovascular system
in the manner described above. In this case, the target area is the
fetal oxygen exchange system, including the placenta, placental
vasculature, fetal heart and major fetal blood vessels. Such
trans-abdominal fetal monitoring can provide information about
fetal blood oxygenation levels in a minimally invasive or
non-invasive manner. Fetal oxygenation levels are known to be
substantially lower than the conjugate maternal levels. The
selection of wavelengths used can be biased to probe the fetus
preferentially over the mother.
[0161] A further embodiment of the invention will now be described
relating generally to non-invasive optical systems and methods for
determining the trends of the ratio of oxygenated and deoxygenated
hemoglobin and other parameters, such as pH, in blood vessels such
as the internal jugular vein (IJ), and/or the carotid artery.
[0162] Near-infrared spectroscopy has been used for non-invasive
measurement of various physiological properties in animal and human
subjects. The basic principle underlying the near-infrared
spectroscopy is that physiological tissue is a relatively highly
scattering medium and is a relatively low absorbing medium to the
near-infrared waves. Many substances in a medium may interact or
interfere with the near-infrared light waves propagating
therethrough. Human tissues, for example, include numerous
chromophores such as oxygenated hemoglobin, deoxygenated
hermoglobin, wate, lipid, etc, where the hemoglobin is the dominant
chromophores in the spectrum range of 700 nm to 900 nm.
Accordingly, the near-infrared spectroscope has been conventionally
applied to measure the trend in oxygen levels of blood in the
physiological medium.
[0163] Various techniques have been developed for the near-infrared
spectroscopy. Examples include: time-resolved spectroscopy (TRS),
phase modulation spectroscopy (PMS), and continuous wave
spectroscopy (CWS). The conventional NIRS spectroscopy techniques
have been developed for measuring average blood oxygen saturation
within a tissue volume which is the mixture of arterial, venous,
and capillary blood. For the applications of monitoring cardiac
output, blood saturation can be measured within the internal
jugular vein. This is advantageous in many circumstances since the
internal jugular vein is a major blood vessel surrounded by highly
scattering tissue medium in the neck.
[0164] Advantageously, the present embodiment can separate or
distinguish blood flow in the internal jugular vein from
surrounding tissue medium, and can therefore more accurately
measure the trends in internal jugular vein blood oxygen
saturation. Additionally, the present invention can provide such
measurements non-invasively using optical systems and determine the
trends of the ratio of oxygenated and deoxygenated hemoglobin in
internal jugular vein blood vessel.
[0165] FIG. 21 shows a system for making relative measurements
relating to blood oxygenation according to an embodiment of the
invention. As shown in FIG. 21, the system includes a patch sensor
2120 which can be reusable or disposable. Patch 2120 is made from a
bio-compatible material, such as a suitable bio-compatible rubber,
and is preferably attached to the skin of the patient using a
bio-compatible adhesive. In some applications, the adhesive should
be less permanent, so the patch can be repositioned. In other
applications, the adhesive should be more permanent, to make
measurements over longer periods of time. As described in further
detail below, patch 2120 includes one or more electromagnetic
radiation transmitters and one or more electromagnetic radiation
detectors. The radiation from patch 2120 transmits into the body
2100, such as the tissues of the neck. The body 2100 includes a
target blood vessel 2102 in which measurements related to blood
oxygen saturation are taken. The target blood vessel 2102 can be
more than 1 cm below the surface of the skin (or other tissue
boundary) at the location where patch 2120 is engaged, and in many
cases, such as where the target blood vessel is the internal
jugular vein in an adult patient, vessel 2102 is typically about 2
cm below the skin. The crescent-shaped pathways 2122 and 2124 of
the radiation transmitted by patch 2120 scattered through tissues
of body 2100 and collected by the detectors has it shown. The
transmitters and detectors are configured, arranged and/or
positioned preferably such that the pathways include at least two
different penetration depths. As shown in FIG. 21, pathway 2122 is
shallower and does not include blood vessel 2102, whereas pathway
2124 is deeper and includes blood vessel 2102. Note that the two
different depths are achieved in the system of FIG. 21 preferably
by two different transmitter/detector pair separation
distances.
[0166] Optical fiber cables or electronic wire cables 2130 and 2136
connect the patch 2120 to either a main station box 2112, or to a
portable unit 2132 which sends out data through wireless
communication to station box 2112 as illustrated by arrow 2134. In
communication with station box 2112, display 2110 shows both time
course trend and digits of oxygen saturation in the blood vessel(s)
of interest, e.g. the internal jugular vein and/or carotid artery,
as wells as oxygen consumption rate. Portable unit 2132 is
preferably dimensioned and sized such the patient can carry the
portable box for extended periods.
[0167] FIG. 22 is a schematic block diagram of a monitoring system
according to an embodiment of the invention. The apparatus 2200 is
similar in many respects to apparatus 200 as shown in FIG. 2A and
FIG. 2B and described above. Apparatus 2200 generally includes an
optical source 2202, launch optics 2204, collection optics 2208 and
2232, optical detectors 2210 and 2211, associated electronics such
as a filter 2212, an optical display 2214, optional wireless
transmitter 2250 and optional battery 2252. The optical source 2202
provides pulsed or continuous electromagnetic radiation. The launch
optics 2204 may include one or more optical fibers 2205 that couple
the electromagnetic radiation from the optical source 2202 to
tissues within tissue boundary TB. Similarly, the collecting optics
2208 and 2232 collect optical signals reflected from within the
tissues. Shallow pathway 2242 shows the region of radiation travel
from source 2204 to collecting optics 2232. Deep pathway 2240 shows
the region of radiation travel from source 2204 to collecting
optics 2208. Note that shallow pathway 2242 substantially excludes
blood vessel 2201, and preferably does not include blood vessel
2201, whereas deep pathway 2240 includes blood vessel 2201. The
collecting optics 2208 and 2232 may include one or more optical
fibers 2209 and 2230 respectively, that couple signals scattered
electromagnetic radiation to the optical detectors 2210 and 2211
respectively.
[0168] Launch optics 2204 and collecting optics 2008 and 2232 are
preferably mounted together on a patch 2203. Note that patch 2203
is shown having the arrangement of the launch optics and optical
detectors such that pathways 2240 and 2242 are perpendicular to
blood vessel 2201. Alternatively, patch 2203 can be engaged such
that the radiation pathways are parallel to the blood vessel, as
shown for example in FIGS. 21 and 23, and in some cases such
parallel engagement has been found to be preferable. Referring
again to FIG. 22, in some embodiments, the optical source 2202,
optical detectors 2110 and 2232, filter 2212 and display 2214 may
be part of a station box 2213 coupled to the patch 2203 by
fiberoptics 2205, 2209, 2230 and electrical cables. The station box
2213 may include a system controller 2215. The system controller
2215 may include a central processor unit (CPU) and a memory (e.g.,
RAM, DRAM, ROM, and the like). The controller 2215 may also include
a well-known support circuits, such as input/output (I/O) circuits,
power supplies (P/S), a clock (CLK), Field Programmable Gate Arrays
(FPGAs) and cache. The controller 2215 may optionally include a
mass storage device such as a disk drive, CD-ROM drive, tape drive,
or the like to store programs and/or data. The controller may also
optionally include a user interface unit to facilitate interaction
between the controller 2215 and a user. The user interface may
include a keyboard, mouse, joystick, light pen or other device. The
preceding components may exchange signals with each other via a
controller bus. In addition, the optical source 2202, detectors
2210 and 2211, filter 212, and display 214 may exchange signals
with the controller 2215 via the system bus 2216. Alternatively,
2213 could be a portable unit that includes battery 2252 and
optional wireless transmitter 2250 for communicating with a
separate station box. By using a portable unit arrangement, the
patient is freer to move about, and the monitoring can take place
over much longer periods of time than would otherwise be
practical.
[0169] The controller 2215 typically operates the optical source,
2202, optical detectors 2210 and 2211, filter 2212 and display 2214
through the I/O circuits in response to data and program code
instructions stored and retrieved by the memory and executed by the
processor. The program code instructions may implement embodiments
of the diagnostic technique described herein. The code may conform
to any one of a number of different programming languages such as
Assembly, C++, JAVA, Embedded Linux, or a number of other
languages. The CPU forms a general-purpose computer that becomes a
specific purpose computer when executing program code. Although the
program code is described herein as being implemented in software
and executed upon a general purpose computer, those skilled in the
art will realize that the method of pulsed pumping could
alternatively be implemented using hardware such as an application
specific integrated circuit (ASIC) or FPGA or other hardware
circuitry. As such, it should be understood that embodiments of the
invention can be implemented, in whose or in part, in software,
hardware or some combination of both.
[0170] FIG. 23 shows a side view of a monitoring patch placed in
close proximity to tissues containing two blood vessels of
interest, according to an embodiment of the invention. Patch 2300
is shown in side view attached, preferably adhered to tissue
boundary 2304. Patch 2300 includes three electromagnetic radiation
transmitters 2312, 2314 and 2316, as well as one electromagnetic
radiation receiver 2310. By positioning the patch 2300 in the
appropriate location, and due to the spacings between each of the
transmitters 2312, 2314, and 2316 on the one hand and receiver 2310
on the other hand, three different pathways of electromagnetic
radiation scattering 2320, 2322 and 2324 are produced within
tissues 2302. Tissues 2302 include two blood vessels of interest
2306 and 2308. Through position of patch 2300 and
transmitter/receiver spacing as described above, pathway 2320
includes tissues but substantially excludes both blood vessels 2306
and 2308. Likewise pathway 2322 includes blood vessel 2306 while
substantially excluding blood vessel 2308, and pathway 2324
includes both blood vessels 2306 and 2308. Through processing
described in more detail below and preferably performed in system
controller 2215 in FIG. 22, relative measurements related to blood
oxygenation in the two blood vessels can be made. Referring again
to FIG. 23, according to a preferred embodiment, tissue 2302 is
neck tissue and tissue boundary 2304 is the skin of the patient's
neck. According to this embodiment, blood vessel 2306 is the
internal jugular vein, and blood vessel 2308 is the carotid artery.
Using processing as described above and more fully below,
monitoring relating to blood oxygenation in the internal jugular
vein and carotid artery can be obtained and cardiac output can be
accurately calculated in a non-invasive manner. Additionally, if
patch 2300 is made of a comfortable compliant material such as
bio-compatible rubber and is adhered to the neck skin, the
oxygenation measurements and cardiac output calculations can take
place over a relatively long period of time. In cases where a
portable box is used such as shown in FIG. 22 and described above,
the patient need not remain at the clinic or hospital location. In
this way, the measurements and calculations can be performed
continuously over a long periods, for example for several days,
while the patient is performing a wide range of normal activities
throughout each day.
[0171] As used herein, the term "pathway" or "pathways" for
electromagnetic radiation or photons refers to the photon pathway
spatial probability distribution, and the pathways depicted in the
figures herein illustrate the approximate locations having a
probability of 95% or greater.
[0172] As used herein with respect to pathways including or
excluding certain structures such as blood vessels, the terms
"substantial" or "substantially" mean that approximately 50% or
more (or less than 50% in the case of excluding a structure) of the
cross section of the structure at a particular longitudinal
position falls within (or falls outside) the pathway. For example,
the phrase "pathway 2242 substantially excludes blood vessel 2201"
as used herein means that pathway 2242 includes less than 50% of
blood vessel at the position in question. The phrase "pathway 2242
. . . preferably does not include blood vessel 2201" as used herein
means that preferably the photon pathway spatial probability
distribution illustrated by pathway 2242, having a probability of
95% or greater does not include blood vessel 2201, or in other
words, any spatial location within the blood vessel have a less
than 5% photon spatial probability.
[0173] To distinguish between oxy-hemoglobin and deoxy-hemoglobin,
as described above and shown in FIG. 11, at least two different
wavelengths of light between 680 nm to 900 nm are transmitted from
each source position to obtain the absorption properties of the
blood vessel of interest at each wavelengths and to calculate blood
oxygen saturation. FIG. 24 illustrates amplitude modulation of two
different wavelengths of electromagnetic radiation, according to an
embodiment of the invention. The amplitude of two different
wavelengths of light are shown being modulated, for example as by
the pulse of an artery in the case of radiation pathway 2324 in
FIG. 23 being modulated by the pulse in carotid artery 2308. Using
the processing methods and apparatus as described above in the
examples relating to ultrasound modulation, more accurate
measurements relating to arterial oxygenation can be taken by
making use of natural modulations due the pulse in arteries.
[0174] FIGS. 25a and 25b show a top view of sensor patches
according to embodiments of the invention. FIG. 25a shows sensor
patch 2500 including transmitter 2502, and two receivers 2504 and
2506. The spacing between the transmitter/receiver pairs should be
chosen such that the depth of the resulting radiation pathway is
appropriate for the particular application. For example, in some
cases where patch 2500 is placed on the skin of the neck, and the
shallower radiation pathway between transmitter 2502 and receiver
2504 is to include the superficial tissues of the neck but not the
internal jugular vein, a spacing of about 2 cm has been found
suitable, and the deeper pathway between transmitter 2502 and
receiver 2506 is to include the superficial tissues as well as the
internal jugular vein, a spacing of about 5 cm has been found
suitable.
[0175] FIG. 25b shows sensor patch 2510 including two transmitters
2512 and 2514, and one receiver 2516. As described above, the
spacing between the transmitter/receiver pairs should be chosen
according to the particular application. In some cases of
non-invasive neck monitoring of oxygenation in the internal jugular
vein, it has been found that the shorter spacing (transmitter 2514
and receiver 2516) should be about 2 cm, and the longer spacing
(transmitter 2512 and receiver 2516) should be about 5 cm. In
cases, such as shown and described above with respect to FIG. 23,
where the carotid artery will also be monitored, an even longer
spacing should be provided and in some cases it has been found that
a spacing of about 7 cm is suitable for making measurements on the
carotid artery. The spacing for a particular case depends not only
on the type and location of the blood vessel being monitored, but
also in general on the individual size and body type of the
patient. For example a child versus a large adult, and also the
thickness and make up of the superficial tissue layers between the
neck skin and the internal jugular vein and carotid artery. It has
been found that in many cases the useful range of spacing should be
between about 0.5 to 3 cm for a pathway that does not include the
internal jugular vein, should be between; 3 to 7 cm for a pathway
that includes the internal jugular vein but not the carotid artery;
and 5 to 10 cm for a pathway that includes the carotid artery.
[0176] FIG. 26 shows a cross section along I-I' of the sensor patch
of FIG. 25a placed on tissues containing a blood vessel of
interest, according to an embodiment of the invention. Patch 2500
contains a single sensor 2502 and two detectors 2504 and 2506 which
generate two radiation pathways 2616 and 2618 within tissues 2610.
Note that the shallower pathway 2616 includes the superficial
tissues within 2610 but not blood vessel 2612, whereas pathway 2618
includes both the superficial tissues and the blood vessel 2612.
Using the processing as described in greater detail below,
parameters such as relative blood oxygenation in the blood vessel
2612, such as the internal jugular vein, can be monitored.
[0177] FIGS. 27a and 27b show plan views of sensor patches
according to further embodiments of the invention. FIG. 27a shows
patch 2700 including two transmitters 2702 and 2704 and two
receivers 2706 and 2708. Using this arrangement, a total of four
different pathways can be established within the underlying
tissues. FIG. 27b shows a sensor patch 2720 having four
transmitters, 2722, 2724, 2730 and 2732 and four receivers 2726,
2728, 2734 and 2736. Using this arrangement, a total of 16
different pathways can be established within the underlying
tissues. Also, in addition to different pathway depths due to the
different transmitter/receiver spacings, the arrangements shown in
FIGS. 27a and 27b can provide different pathway locations without a
need to re-position the patch on the patient's skin or other tissue
boundary.
[0178] FIG. 28 shows a cross section along II-II' of the sensor
patch of FIG. 27a placed on tissues containing a blood vessel of
interest, according to an embodiment of the invention. Two
exemplary pathways 2816 and 2814 are shown within tissues 2800. As
shown, pathway 2816 does not include blood vessel 2802, and pathway
2814 includes blood vessel 2802.
[0179] FIG. 29 shows a plan view of a sensor patch having an array
of transmitters and receivers, according to an embodiment of the
invention. The sensor patch of FIG. 29 includes 21 transmitters,
shown as the circles marked "S", and "21 receivers, shown as the
squares marked "D". An array arrangement as shown in FIG. 29 has a
number of advantages over smaller sensor patches such as shown
FIGS. 25a and 25b, including repositioning the effective pathways
without detaching the patch from the skin or tissue boundary,
mapping or imaging to more precisely locate structures within the
tissues (such as large veins and arteries), providing several
different depths through the great many transmitter-receiver
spacings. For example, if an array-type sensor patch is used over
several days on adhered to a patient's neck, the array will be able
to adjust itself to accommodate different neck positions which may
cause shifting of the position of the internal jugular vein
relative to the sensor patch. The sensor patch of FIG. 29 can be
made of launch optics and collecting optics and optical fibers as
shown and described in FIG. 22. Alternatively, CCD technology, or
custom made silicon photo detector and laser diode arrays could be
used. In general greater numbers of transmitter and receiver sites
will provide better imaging and freedom to choose optimal pairings.
For some applications, it has been found that providing at least 10
transmitters and at least 10 receivers is suitable. It has been
found that a patch 7 cm.times.5 cm, having adjacent
transmitter/receiver spacing of about 1 cm is suitable for many
applications (about 35 transmitters and receivers total). However,
higher resolution imaging can be accomplished with the same size
patch but with 0.5 cm spacing (about 140 transmitters and receivers
total), or even 0.25 cam spacing (about 560 transmitters and
receivers) is suitable. Although it is not necessary for equal
numbers of transmitters and receivers to be present in an
array-type sensor, roughly equal numbers and distributions are
preferred for imaging purposes.
[0180] FIG. 30 is a flowchart illustrating several steps relating
to measuring cardiac output according to embodiments of the
invention. In step 3010, electromagnetic radiation is transmitted
into the patient's tissues at one or more locations and back
scattered light is detected or received at one or more locations.
The radiation is preferably transmitted and received using the
systems and apparatus as shown and described above with respect to
FIGS. 21-29. As described above, there should be at least two
different pathway depths for the radiation that is detected by one
or more receivers. In the embodiments described above, the
different depths of the pathways is accomplished by providing
different transmitter receiver distances.
[0181] In step 3014, the transmitter-receiver, or source-detector
pair creating the shallower depth pathway is used to measure the
average tissue scattering and absorption properties of the
superficial layer above the blood vessel (preferably the internal
jugular vein). In the foregoing embodiments shown and described
with respect to FIGS. 21, 23, 26 and 28, the shallower or
shallowest pathway should substantially exclude the blood vessel of
interest (in many cases, the internal jugular vein). In order to
provide more accurate calculations for W1, as described below,
preferably spatial locations within the blood vessel should have
less than about 20% photon probability. Even more preferably,
spatial locations within the blood vessel should have less than
about 5% probability for photons traveling between the transmitter
and receiver pair for the shallowest pathway. Finally, it has been
found that in order to further increase the practical applicability
and further increase the accuracy for calculations for W1 the
photon probability should preferably be less than about 1%.
[0182] In step 3018, the transmitter-receiver, source-detector pair
creating the deeper depth pathway is used to measure the average
tissue scattering and absorption properties of both the superficial
tissues as well as the blood vessel of interest.
[0183] In step 3022, for the source-detector with large distance,
the probably for light to pass through superficial layer (W1) and
the layer contains the internal jugular vein (W2) are calculated,
preferably through the photon diffusion equation as discussed in
further detail below, using the average tissue scattering and
absorption properties measured in step 3014. The absorption
properties of the layer containing the internal jugular vein can
then be calculated through the following relationship:
.mu. a , IJ = .mu. _ a - W 1 - .mu. a , suoerfucuak W 2
##EQU00002##
[0184] As mentioned, the values for W.sub.1 and W.sub.2 are
preferably calculated from photon diffusion equation:
- D .gradient. 2 .PHI. ( r , t ) + v .mu. a .PHI. ( r , t ) +
.differential. .PHI. ( r , t ) .differential. t = v S ( r , t )
##EQU00003##
where D is the diffusion constant, V is the speed of light, and
.mu..sub.a is the absorption coefficient of medium to the
light.
[0185] The measurements from the transmitter-receiver or
source-detector pair having the shorter separation distance, or
shallower depth, are used to calculate the probability of photon
distribution inside depth from surface to z.sub.1 (z.sub.1 is
typically 2 cm) which is W.sub.1.
W 1 = .intg. 0 z 1 [ .intg. .intg. .infin. 1 4 .pi. D ( r .fwdarw.
- r .fwdarw. s ) k ( r .fwdarw. - r .fwdarw. s ) 1 4 .pi. D ( r
.fwdarw. d - r .fwdarw. ) k ( r .fwdarw. d - r .fwdarw. ) x y ] z
##EQU00004##
[0186] The measurements from the source-detector pair with larger
separation distance, or greater depth, are used to calculate the
photon probability distribution from depth z.sub.1 to z.sub.2,
which is W.sub.2.
W 2 = .intg. z 1 z 2 [ .intg. .intg. .infin. 1 4 .pi. D ( r
.fwdarw. - r .fwdarw. s ) k ( r .fwdarw. - r .fwdarw. s ) 1 4 .pi.
D ( r .fwdarw. d - r .fwdarw. ) k ( r .fwdarw. d - r .fwdarw. ) x y
] z ##EQU00005##
where r.sub.s is the position of light source, r.sub.d is the
position of detector, and r is the position of a certain position
inside medium. K is the wave vector which can be derived from the
photon diffusion equation, D is the diffusion constant of
medium.
[0187] From W.sub.1 and W.sub.2 and the absorption property of
tissue at depth from z.sub.1 to z.sub.2 can be derived from
equation:
.mu. a , IJ = .mu. _ a , depth 2 - W 1 .mu. a , depth 1 W 2
##EQU00006##
[0188] In step 3026, steps 3010 through 3022 are repeated for a
second wavelength of electromagnetic radiation. Note that the
transmission of the second wavelength can be alternated with that
of the first wavelength, such that the steps shown in FIG. 30 do
not have to be performed in the strict order shown. Preferably, at
least two different wavelengths of light between 680 nm to 900 nm
are transmitted from each source position to obtain the absorption
properties of the internal jugular vein at each wavelength and to
calculate blood oxygen saturation. If the absorption properties of
the blood (in this example, the internal jugular blood) are
obtained at two or more wavelengths then the concentrations of
deoxygenated hemoglobin (C_Hb) and oxygenated hemoglobin (C_HbO)
can be calculated from:
.mu..sub.a,IJ.sup..lamda.C.sub.Hb,IJ.sub..epsilon..sub.Hb.sub..lamda.+C.-
sub.HbO,IJ.sub.HbO.sub..lamda.
where .epsilon. is the pre-determined absorption of deoxygenated
hemoglobin (Hb) and oxygenated hemoglobin (HbO) per unit
concentration (e.g. grams/liter). The internal jugular blood
saturation can then be calculated as:
S IJ O 2 = C HbO , IJ C Hbo , IJ + C Hb , IJ % ##EQU00007##
[0189] Note that while the present and several of the foregoing
embodiments have been described using the example of blood oxygen
saturation and cardiac output, the invention is also applicable to
monitor other parameters relating to the patient's blood. For
example, blood pH can be monitored using met-hemoglobin as a target
chromophore, as is described in further detail above. Another
example is monitoring water and/or lipid in the blood, using
radiation wavelengths where are selected to be suitable for the
particular chromophore application.
[0190] In step 3030, the oxygen saturation in the blood vessel is
then used for calculation of cardiac output, as discussed above
with respect to the Fick principle:
CardiacOutput = OxygenConsumption ( S a O 2 - S IJ O 2 ) A
##EQU00008##
[0191] In cases where the blood vessel being monitored is a vein,
then S.sub.aO.sub.2, the arterial blood oxygen saturation, can be
measured through conventional methods, for example using pulse
oximetry. The value for A can be calculated as described above as
[Hb].times.1.36 where [Hb] is the blood hemoglobin concentration
and 1.36 is a factor subsuming the oxygen carrying capacity of the
hemoglobin. The calculation and processing steps described with
respect to FIG. 30 are preferably performed in a processor such as
in system controller 2215 in FIG. 22 as described above.
[0192] According to an alternative embodiment of the invention, The
S.sub.aO.sub.2 can be measured without using conventional means,
such as a standard pulse oximeter. According to this embodiment,
S.sub.aO.sub.2 is measured through the same patch sensor as shown
in FIG. 23 as described above, the amplitudes of the optical
signals especially the source-detector pair with largest separation
are modulated by the pulsation by the major artery which is
adjacent to the internal jugular vein, i.e. the carotid artery. The
amplitude of such modulated signals at least two different
wavelengths are used to calculate the oxygen saturation of arterial
blood, as described above.
[0193] Although the above description emphasizes measurement of
blood oxygenation for the purpose of determining venous oxygen
saturation, cardiac output and pH, the invention is not limited to
such applications. The technique described herein can be adapted to
selectively probe tissues within the body to measure the level of a
particular target chromophore within those tissues and derive
diagnostic information about the tissue from the measurement. These
measurements can be made in a manner which is accurate,
reproducible, precise, fast, operator independent, easy to use,
continuous, cost effective, and substantially free of increased
mortality and morbidity. Embodiments of the present invention allow
measurements that used to be made in a highly invasive manner to be
made in a non-invasive or minimally invasive manner. Applications
of the technique include measuring the health of a transplanted
organ to check for signs of rejection, measuring the perfusion of a
skin graft in, for example a burn victim, to determine the health
of the graft, potential ambulatory monitoring of high-risk
cardiovascular patients, and ambulatory monitoring of high-risk
pregnancies.
[0194] FIGS. 31a and 31b show a sensor patch according to a further
embodiment of the invention. FIG. 31 a shows sensor patch 3120 that
includes two electromagnetic transmitters 3122 and 3130, four
electromagnetic receivers 3126, 3128, 3134 and 3136. These
structures are identical or similar to those shown and described
above, such as, for example, FIGS. 25a, 25b, 27a and 27b.
Additionally, sensor patch 3120 includes ultrasonic transducer
3140, which can be similar or identical to the ultrasonic
transducer 206 as shown and described, for example, with respect to
FIGS. 2A and 2B. Ultrasonic transducer 3140 is coupled to and
ultrasound source (not shown). FIG. 31b is a cross-section of the
patch 3120 along in FIG. 31a. As shown, the transmitter-receiver
pairs 3122-2126 and 3122-3128 generate electromagnetic radiation
pathways 3116 and 3118 respectively. Ultrasonic transducer 3140 is
positioned as shown to be used to both image the underlying tissues
3110, for example to precisely locate blood vessel 3112, and to
modulate the blood vessel 3112 to provide for more accurate
measurement, as described further below.
[0195] FIG. 32 shows steps involved monitoring oxygenation and
cardiac output according to a further embodiment of the invention.
In step 3210, and ultrasound transducer is placed in contact with
the patent's skin in the vicinity of the blood vessel of interest.
The ultrasound transducer in this embodiment can be of the type
integrated into the sensor patch as shown and described in FIGS.
31a and 31b, however, the transducer may also be of the
conventional ultrasound imaging type transducer. In step 3212 the
ultrasound transducer is used to make images of the issues beneath
the skin to precisely locate the target blood vessel. A suitable
mark is made on the skin. In step 3214, the optical sensor patch,
such as shown in FIGS. 21, 23, and 25-29 is then accurately placed
on the patient's skin using the mark such that the electromagnetic
pathways are well positioned for accurate measurements. In step
3216, the measurements and calculations are made, such as shown and
described with respect to FIG. 30 above.
[0196] FIG. 33 shows steps involved in monitoring oxygenation and
cardiac output according to a further embodiment of the invention.
In step 3310, ultrasound imaging is used to accurately position the
sensor patch on the patient's skin, as described with respect to
FIG. 32. According to this embodiment, a sensor patch having an
integrated ultrasound transducer, such as shown and described with
respect to FIGS. 31a and 31b is used. In step 3312, the integrated
ultrasound transducer is used to generate radiation pressure
modulations in the target blood vessel. In step 3314,
electromagnetic radiation is transmitted and received through at
least two transmitter-receiver pairs to provide at least two
different pathways having different depths of penetration. In step
3316, the average tissue scattering and absorption properties of
the superficial tissues is measured using the measurements from the
shallowest pathway, as also described with respect to FIG. 30
above. In step 3318, the sensor data relating to the deeper pathway
is processed using processing methods and systems described above
using ultrasonic radiation pressure modulation. The calculations
for relative oxygenation and for cardiac output can have increased
accuracy over prior embodiments since the tissue scatter and
absorption properties of the superficial tissues are known, and the
ultrasound radiation pressure modulation provides for greater
signal to noise ratio.
[0197] FIG. 34 shows steps involved in monitoring oxygenation and
cardiac output according to a further embodiment of the invention.
In step 3410, ultrasound imaging is used to accurately position the
sensor patch. In step 3412, the ultrasound transducer is used to
generate pressure modulations in the target blood vessel, as is
described above. In step 3414, light is transmitted and received
through at least two different pathways in the tissues using at
least two different transmitter-receiver pairs, as described above.
In step 3418, oxygenation in the blood vessel of interest is
measure and calculated both with and without the ultrasound
radiation pressure modulations. After removing the ultrasound
transducer in step 3420, the measurements and/or calculations made
without the ultrasound pressure modulations are calibrated or
corrected using the measurements made in step 3418. In this way, a
combination of ultrasound pressure modulation and multi-depth
multi-pathway optical measurements are used to enhance the accuracy
of measurement, and allow for increased mobility and duration of
the continuous measurement while decreasing the complexity of the
system, since the ultrasound system is removed after the initial
calibration steps.
[0198] Although several of the foregoing embodiments have been
described using the internal jugular vein as a target structure for
monitoring, there are a number of other target structures within
the body that are suitable for blood oxygen monitor using
embodiments of the present invention. Several representative
example applications will now be described. The exterior jugular
vein can be monitored transdermally as shown and described above
with respect to FIG. 13. The right subclavian vein, superior vena
cava, pulmonary artery and other major blood vessels may be
monitored as shown and described above with respect to FIG. 14.
Neonatal blood oxygenation can be monitored as shown and described
above with respect to FIGS. 18A-18C, and 19. Fetal monitoring can
be provided as shown and described above with respect to FIG.
20.
[0199] The techniques described are not limited to the hospital or
medical office setting. Embodiments of the invention could be made
portable and simple to use by virtue of its use of rugged telecom
components and low power-consumption devices which could in turn
allow its use in ambulances. As shown for example in FIGS. 21 and
22, embodiments of the invention may be useful for real-time
monitoring of personnel in high risk situations. For example rescue
workers in chemical plants responding to emergencies, or firemen in
burning buildings could be monitored remotely for signs of physical
distress. Military personnel with ambulatory versions of the
sensors could be monitored on the battlefield, and portable
versions of the device could be used for first-responder
battlefield triage.
[0200] While the above is a complete description of the preferred
embodiment of the present invention, it is possible to use various
alternatives, modifications and equivalents. Therefore, the scope
of the present invention should be determined not with reference to
the above description but should, instead, be determined with
reference to the appended claims, along with their full scope of
equivalents. In the claims that follow, the indefinite article "A",
or "An" refers to a quantity of one or more of the item following
the article, except where expressly stated otherwise. The appended
claims are not to be interpreted as including means-plus-function
limitations, unless such a limitation is explicitly recited in a
given claim using the phrase "means for".
* * * * *