U.S. patent application number 11/539949 was filed with the patent office on 2007-08-02 for apparatus and method for monitoring tissue vitality parameters.
This patent application is currently assigned to CRITISENSE LTD.. Invention is credited to Avraham Mayevsky, Eliahu Pewzner.
Application Number | 20070179366 11/539949 |
Document ID | / |
Family ID | 11074672 |
Filed Date | 2007-08-02 |
United States Patent
Application |
20070179366 |
Kind Code |
A1 |
Pewzner; Eliahu ; et
al. |
August 2, 2007 |
Apparatus and Method for Monitoring Tissue Vitality Parameters
Abstract
Apparatus for determining the oxygenation state of at least one
tissue element, comprising: a) illumination means for illuminating
said tissue element with an illuminating radiation at a
predetermined wavelength via at least one illumination location
with respect to said tissue element, said illuminating radiation
being at a wavelength within the NADH excitation spectrum or the Fp
excitation spectrum; b) radiation receiving means and detection
means for measuring the intensity of the corresponding NADH
fluorescence or Fp fluorescence emitted by the tissue element at at
least two predetermined wavelengths within the range of wavelengths
comprised within the corresponding fluorescence emission
spectrum.
Inventors: |
Pewzner; Eliahu; (Modiin
Ilit, IL) ; Mayevsky; Avraham; (Ramat-Gan,
IL) |
Correspondence
Address: |
WOLF, BLOCK, SCHORR & SOLIS-COHEN LLP
250 PARK AVENUE
NEW YORK
NY
10177
US
|
Assignee: |
CRITISENSE LTD.
11 Ben-Gurion Street
Givat-Shmuel
IL
54101
|
Family ID: |
11074672 |
Appl. No.: |
11/539949 |
Filed: |
October 10, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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10381383 |
Mar 25, 2003 |
7130672 |
|
|
PCT/IL01/00906 |
Sep 25, 2001 |
|
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11539949 |
Oct 10, 2006 |
|
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Current U.S.
Class: |
600/310 |
Current CPC
Class: |
A61B 5/14546 20130101;
A61B 5/14553 20130101; A61B 2562/0242 20130101; A61B 5/0261
20130101 |
Class at
Publication: |
600/310 |
International
Class: |
A61B 5/00 20060101
A61B005/00 |
Foreign Application Data
Date |
Code |
Application Number |
Sep 25, 2000 |
IL |
138683 |
Claims
1. Apparatus for determining the oxygenation state of at least one
tissue element, comprising: a) illumination means for illuminating
said tissue element with an illuminating radiation at a
predetermined wavelength via at least one illumination location
with respect to said tissue element, said illuminating radiation
being at a wavelength within the NADH excitation spectrum or the Fp
excitation spectrum; b) radiation receiving means and detection
means for measuring the intensity of the corresponding NADH
fluorescence or Fp fluorescence emitted by the tissue element at at
least two predetermined wavelengths within the range of wavelengths
comprised within the corresponding fluorescence emission
spectrum.
2. Apparatus as claimed in claim 1, wherein the detection means
comprises a spectrometer.
3. Apparatus as claimed in claim 1, further comprising suitable
means adapted for comparing the intensities measured in (b) to
provide an estimate of the relative levels of oxygenated blood to
deoxygenated blood in said tissue.
4. Apparatus as claimed in claim 1, wherein said receiving means
and said detection means are adapted for measuring said intensity
at said at least two fluorescent wavelengths such that one of said
fluorescent wavelengths is at an oxy-deoxy fluorescence emission
isosbestic point in the fluorescent emission spectrum.
5. Apparatus as claimed in claim 1, wherein said receiving means
and said detection means are adapted for measuring said intensity
at said at least two fluorescent wavelengths, such that one of said
fluorescent wavelengths is higher than, and another one of said at
least two fluorescent wavelengths is smaller than, a wavelength
corresponding to an oxy-deoxy fluorescence emission isosbestic
point in the fluorescent emission spectrum.
6. Apparatus as claimed in claim 1, wherein said predetermined
wavelength of said illuminating radiation corresponds to an
isosbestic point of the blood oxy de-oxy absorption spectrum and
being within the fluorescence excitation spectrum of NADH or
Fp.
7. Apparatus as claimed in claim 1, wherein said illumination
location is provided by at least one excitation optical fiber
capable of being brought into registry with said tissue
element.
8. Apparatus as claimed in claim 7, wherein said radiation
receiving means comprises at least one suitable receiving optical
fiber capable of being brought into registry with said tissue
element.
9. Apparatus as claimed in claim 8, comprising at least one probe
which houses said at least one excitation optical fiber, and said
at least one receiving optical fiber.
10. Apparatus as claimed in claim 9, wherein the at least one
probes comprise a plurality of probes, each probe housing at least
one of the at least one excitation fibers and at least one of the
at least one receiving fibers for determining the oxygenation state
of one of the at least one tissue elements.
11. Apparatus as claimed in claim 10, wherein the illumination
means comprises a UV monochromatic light source, adapted to provide
said illuminating radiation to the plurality of probes, in pulses
of predetermined duration and intensity.
12. A system as claimed in claim 11, further comprising suitable
control means for controlling the frequency of pulsing of said
pulses.
13. A system as claimed in claim 12, wherein said control means is
adapted for selectively directing discrete said pulses to any one
of said probes.
14. Apparatus as claimed in claim 12, further comprising coupling
means for optically connecting said at least two probes to said
detection means, wherein said control means is operatively
connected to said detection means.
15. A system as claimed in claim 14, wherein said control means is
selectively responsive to previously detected signals corresponding
to the detection of said fluorescence detected by said detection
means of a prior monitoring cycle.
16. Method for determining the oxygenation state of a tissue
element, comprising: (a) providing an illuminating radiation to the
tissue element, said illuminating radiation being at a wavelength
within one of the NADH fluorescence excitation spectrum or the Fp
fluorescence excitation spectrum; (b) measuring the intensity of
the corresponding NADH fluorescence or Fp fluorescence,
respectively, emitted by the tissue element at least at two
predetermined wavelengths within the range of wavelengths comprised
in the corresponding fluorescence emission; (c) comparing the
intensities measured in (b) to provide an estimate of the relative
levels of oxygenated blood to deoxygenated blood in said tissue
element.
17. Method as claimed in claim 16, wherein one of said wavelengths
in step (b) is chosen to lie at a suitable NADH oxy-deoxy
fluoresence emission isosbestic point or Fp oxy-deoxy fluoresence
emission isosbestic point, respectively.
18. Method as claimed in claim 16, wherein in step (c) the
fluorescence intensity at each wavelength in (b) is normalised by a
normalising fluorescence intensity measured at a suitable
corresponding NADH or Fp oxy-deoxy fluorescence emission isosbestic
wavelength within the fluorescence emission spectrum.
19. Method as claimed in claim 16, wherein said illuminating
radiation is provided at a suitable NADH oxy-deoxy fluoresence
excitation isosbestic wavelength or a suitable Fp oxy-deoxy
fluoresence excitation isosbestic wavelength, respectively within
the corresponding NADH fluorescence excitation spectrum or Fp
fluorescence excitation spectrum, respectively.
20. Method as claimed in claim 18, wherein said illuminating
radiation in step (a) is within the range of wavelengths comprised
in the NADH excitation spectrum, and said normalising fluorescence
intensity is measured at an NADH oxy-deoxy fluorescence emission
isosbestic wavelength within the fluorescence emission spectrum of
about 455 nm.+-.5 nm.
21. Method as claimed in claim 18, wherein said illuminating
radiation in step (a) is within the range of wavelengths comprised
in the Fp excitation spectrum, and said normalising fluorescence
intensity is measured at an Fp oxy-deoxy fluorescence emission
isosbestic wavelength within the fluorescence emission spectrum of
about 530 nm.+-.5 nm.
Description
RELATED APPLICATIONS
[0001] The present application is a divisional application of U.S.
application Ser. No. 10/381,383, filed on Mar. 25, 2003, which is a
U.S. national application of PCT Application No. PCT/IL01/00906,
filed on Sep. 25, 2001.
FIELD OF THE INVENTION
[0002] The present invention relates to apparatuses and methods for
enabling simultaneous, pseudo-simultaneous, or individual
monitoring of a plurality of tissue vitality parameters,
particularly in-vivo, with respect to an identical tissue layer
element or volume. In particular, such parameters include blood
flow rate, NADH concentration, blood volume, blood oxygenation
state and flavoprotein concentration.
BACKGROUND
[0003] Mammalian tissues are dependent upon the continuous supply
of oxygen and glucose needed for the energy production. This energy
is used for various types of work, including the maintaining of
ionic balance and biosynthesis of various cellular components. The
ratio or balance, between oxygen supply and demand reflects the
cells' functional capacity to perform their work. In this way, the
energy balance reflects the metabolic state of the tissue. In order
to assess the tissue energy balance, it is necessary to monitor the
events continuously using a multiparametric system in
real-time.
[0004] The integrated system of energy supply and demand can be
understood by considering the various components thereof. [0005]
O.sub.2 supply: The blood carries the oxygen and other essential
substances to the cells. Therefore, monitoring of blood flow rate,
blood volume and blood oxygenation will reflect the supply of
O.sub.2 to the tissue for the purpose of energy formation therein.
[0006] Energy production and demand: In an inner compartment of the
cells, called the mitochondria, the glucose and O.sub.2 are
transformed into ATP, a form of energy which can be used by the
cells for various types of activities. The ATP production rate is,
in normal states, regulated by rate of consumption of ATP, and is
increased when cellular activity rises. In most pathological
states, the limiting factor for this process is O.sub.2
availability.
[0007] The process of energy (ATP) production and consumption can
be determined through monitoring of Nicotineamide adenine
dinucleotide (NADH) redox state. The NADH and NAD molecules can be
correlated with the process of ATP production. The concentration of
the reduced form of the molecule (NADH) rises when the rate of ATP
production is low, and is unable to meet the demand in the tissue
or cells.
[0008] A complementary indicator of energy production, other than
NADH, is the concentration of flavoproteins (Fp). Flavoprotein
molecules are also linked to the production of ATP in the
mitochondria. Fp concentration drops when the rate of ATP
production is reduced, and is unable to meet the demand in the
tissue or cells.
[0009] There is direct correlation between energy metabolism of the
cellular compartment and the blood flow in the microcirculation of
the same tissue. In a normal tissue, any change in the O.sub.2
demand will be compensated by a corresponding change in the blood
flow to the tissue. By this mechanism, the O.sub.2 supply remains
constant if there is no change in the O.sub.2 consumption. Any
change in the abundance of O.sub.2 in the tissue, in other words a
change in energy state, will be reflected by the NADH and Fp
level.
[0010] It is important to monitor both supply and demand in order
to be able to detect pathological situations in which the balance
is disrupted, and one component of the system reacts abnormally
with respect to the other.
[0011] The parameters used in the art for the assessment of tissue
vitality include: A--Blood Flow Rate; B--Mitochondrial Redox State
via the NADH level; C--Blood Volume; D--Blood Oxygenation State;
E--Mitrochondrial Redox State via flavoprotein level.
A--Blood Flow Rate
[0012] The blood flow rate relates to the mean volume flow rate of
the blood and is essentially equivalent to the mean velocity
multiplied by the number of moving red blood cells in the tissue.
This parameter may be monitored by a technique known as Laser
Doppler Flowmetry, which is based on the fact that light reflected
off moving red blood cells (RBC) undergoes a small shift in
wavelength (Doppler shift) in proportion to the cell's velocity.
Light reflected off of stationary RBC or bulk stationary tissue, on
the other hand, does not undergo a Doppler shift.
[0013] By illuminating with coherent light, such as a laser, and
converting the intensities of incident and reflected light to
electrical signals, it is possible to estimate the blood flow from
the magnitude and frequency distribution of those signals (U.S.
Pat. No. 4,109,647; Stem, M. D. Nature 254, 56-58, 1975).
B--Mitochondrial Redox State or the NADH Level
[0014] The level of Nicotineamide adenine dinucleotide (NADH), the
reduced form of NAD, is dependent both on the availability of
oxygen and on the extent of tissue activity. Referring to FIG. 1,
whilst NADH absorbs UV light at wavelengths of about 300 nm to
about 400 nm and fluoresces at wavelengths of about 400 nm to about
550 nm, the NAD does not fluoresce. The NADH Level can thus be
measured using Mitochondrial NADH Fluorometry. The conceptual
foundations for Mitochondrial NADH Fluorometry were established in
the early 50's and were published by Chance and Williams (Chance
B., & Williams G. R., Journal of Biological Chemistry, 217,
383-392, 1955). They defined various metabolic states of activity
and rest for in-vitro mitochondria.
[0015] An increase in the level of NADH with respect to NAD and the
resulting increase in fluorescence intensity indicate that
insufficient Oxygen is being supplied to the tissue. Similarly, a
decrease in the level of NADH with respect to NAD and the resulting
decrease in fluorescence intensity indicate an increase in tissue
activity.
C--Blood Volume
[0016] The blood volume parameter refers to the concentration of
the blood in the tissue. When tissue is irradiated, the intensity R
of reflection of the excitation wavelength light from the tissue is
informative of the blood volume. The intensity R of the reflected
signal, also referred to as the total backscatter, increases
dramatically as blood is eliminated from the tissue as a result of
the decrease in haemoglobin concentration. Similarly, if the tissue
becomes more perfused with blood, R decreases due to the increase
in the haemoglobin concentration.
D--Blood Oxygenation State
[0017] The blood oxygenation state parameter refers to the relative
concentration of oxyhaemoglobin to deoxy-haemoglobin in the tissue.
It may be assessed by the performance of photometry measurements.
The absorption spectrum of oxyhaemoglobin HbO.sub.2 is considerably
different from the absorption spectrum of deoxy-haemoglobin Hb
(Kramer R. S. and Pearlstein R. D., Science, 205, 693-696, 1979).
The measurement of the absorption at one or more wavelengths can
thus be used to assess this important parameter. Blood oximeters
are based on measurement of the haemoglobin absorption changes as
blood deoxygenates (Pologe J. A., Int. Anesthesiol. Clin., 25(3),
137-53, 1987). Such oximeters generally use at least two light
wavelengths to probe the absorption. One known method uses one
wavelength at an isosbestic point and another wavelength at a point
that exhibits absorption changes due to variation in oxygenation
level. Another technique uses wavelengths at both sides of an
isosbestic point in order to increase measurement sensitivity. The
wavelengths used in commercial pulse oximeters are typically around
660 nm in the red region of the spectrum, and between 800 to 1000
nm in near-infrared region (Pologe, 1987).
[0018] Isosbestic point as referred to herein is a wavelength at
which the intensity of absorption of oxyhaemoglobin HbO.sub.2 is
the same as the intensity of absorption of deoxy-haemoglobin Hb; to
such isosbestic points are indicated as IP.sub.A and IP.sub.B in
FIG. 10. Similarly, there is an isosbestic range marked IR in FIG.
10 where these two functions are substantially coincident. FIG. 10
is based on Anderson, R. R., Parrish, J. A. (1981) Microvasculature
can be selectively damaged using dye lasers: a basic theory and
experimental evidence in human skin. Lasers Surg. Med. 1,
263-276.
[0019] For monitoring the oxygenation levels of internal organs,
fiber-optic blood oximeters have been developed. These fiber-optic
devices irradiate the tissue with two wavelengths, and collect the
reflected light. By analysis of the reflection intensities at
several wavelengths the blood oxygenation is deduced. The
wavelengths used in one such system were 585 nm (isosbestic point)
and 577 nm (Rampil I. J., Litt L., & Mayevsky A., Journal of
Clinical Monitoring, 8, 216-225, 1992). Another blood oximeter
measures and analyzes the whole spectrum band 500-620 nm (Kessler
M. & Frank K., Quantitative spectroscopy in tissue pp. 61-74.
Verlagsgruppe GmbH, Frankfurt au Main, 1992). These devices are
relatively complicated and susceptible to interference from ambient
light, as well as various electronic and optic drifts. Two light
sources are required, and the light sources and the detection
system also incorporate optical filters that are interchangeable by
mechanical means.
E--Flavoprotein Concentration
[0020] In order to determine the metabolic state of various tissues
in-vivo it is also possible to monitor the fluorescence of another
cellular fluorochrome, namely Flavoproteins (Fp). Referring to FIG.
12, Fp absorbs light at wavelengths of about 400 nm to about 470 nm
and fluoresces at wavelengths of about 490 nm to about 580 nm. The
Fp level can thus be measured using Fp Fluorometry. The conceptual
foundations for Fp Fluorometry were established in the late 1960's
and were published in several papers as will be referenced
hereinafter. Simultaneous monitoring of NADH and Fp from the same
layer or volume of tissue provides better interpretation of the
changes in energy production and demand.
[0021] Chance et al. (B. Chance, N. Graham, and D. Mayer. A time
sharing fluorometer for the readout of intracellular
oxidation-reduction states of NADH and Flavoprotein. The Review of
Scientific Instruments 42 (7):951-957, 1971) used a time-sharing
fluorometer to record intracellular redox state of NADH and Fp.
They showed a very clear correlation between the two chromophores
to changes in O.sub.2 supply to the perfused liver. Using a time
sharing fluorometer reflectometer simultaneous monitoring of NADH
and Fp was performed from the surface of the rat's brain (A.
Mayevsky. Brain energy metabolism of the conscious rat exposed to
various physiological and pathological situations. Brain Res.
113:327-338, 1976). The kinetics of the responses to anoxia or
decapitation were identical for the NADH and Fp indicating that the
NADH signal comes from the same cellular compartment as the Fp--the
mitochondrion.
[0022] The five tissue viability parameters described above
represent various important biochemical and physiological
activities of body tissues. Monitoring them can provide much
information regarding the tissues' vitality. For the monitoring of
different parameters to have maximum utility however, the
information regarding all parameters is required to originate from
substantially the same layer of tissue, and preferably the same
volume of tissue, otherwise misleading results can be obtained. In
general, the more parameters that are monitored from the same
tissue volume or layer, the better and more accurate an
understanding of the functional state of the tissue that may be
obtained.
[0023] There are several techniques that relate to the simultaneous
in-vivo measuring of multiple parameters in certain tissues, which
can be used for the various pathological situations arising in
modem medicine.
[0024] The prior art teaches a wide variety of apparatuses/devices
which monitor various parameters reflecting the viability of the
tissue. For example, U.S. Pat. No. 4,703,758 teaches the use of an
apparatus for monitoring blood flow by using a light source to emit
a beam of light, and a light detector that measures the light
received. This provides the value of the intensity of the
transmitted light, which inter alia depends upon the blood flow in
the path of the light.
[0025] U.S. Pat. No. 4,945,896 teaches the use of a multiprobe
sensor, using independent microelectrodes implanted inside the
brain tissue, for measuring various parameters indicative of the
function of the brain. This device includes a laser Doppler flow
probe for measuring cerebral blood flow, and a probe for monitoring
redox state (NADH). These probes can be mounted sequentially, i.e.,
one after another in the same housing, or they can be placed side
by side. These devices suffer from a major drawback however. Tissue
viability is not merely a reflection of various values of
parameters measured at different times in one place, or different
places at one time. The complex biochemical mechanisms that
determine tissue viability are such that short time deviations
between measurement at short distances between points of
measurement can provide inaccurate or even misleading information.
Thus, while the values of blood flow and redox state (NADH) must be
monitored simultaneously at the same location, with the monitoring
being for the same layer of tissue, this is not performed in the
reference.
[0026] Another drawback encountered in NADH measurements is the
Haemodynamic Artifact. This refers to an artifact in which NADH
fluorescence measurements in-vivo are underestimated or
overestimated due to the haemoglobin present in blood circulation,
which absorbs radiation at the same wavelengths as NADH, and
therefore interferes with the ability of the light to reach the
NADH molecules. The haemoglobin also partially absorbs the NADH
fluorescence. In particular, a reduction of haemoglobin in blood
circulation causes an increase in fluorescence, generating a false
indication of the true oxidation reduction state of the organ. U.S.
Pat. No. 4,449,535 teaches, as means to compensate for this
artifact, the monitoring of the concentration of red blood cells,
by illuminating at a red wavelength (805 nm) simultaneously and in
the same spot as the UV radiation required for NADH excitation and
measuring the variation in intensity of the reflected red
radiation, as well as the fluorescence at 440-480 nm, the former
being representative of the intra-tissue concentration of red blood
cells. Similarly Kobayashi et al (Kobayashi S., Nishiki K., Kaede
K., Ogata E. J. Appl. Physiol. 31, 93-96, 1971) used ultraviolet
(UV) illumination at 366 nm for NADH excitation, and red light at
720 nm for reflection measurements. However, U.S. Pat. No.
4,449,535 has at least two major drawbacks; firstly, and as
acknowledged therein, using a single optical fiber to illuminate
the organ, as well as to receive emissions therefrom causes
interference between the outgoing and incoming signals, and certain
solutions with different degrees of effectiveness are proposed.
More importantly, though, two different wavelengths are used for
illuminating the organ. FIG. 2 (based on Eggert & Blazek, 1987,
.COPYRGT. the Congress of Neurological Surgeons, Lippincott
Williams & Wilkins) illustrates the penetration depth profile
for various tissues of the human brain as a function of
illuminating radiation wavelength, showing a plateau of relative
insensitivity of penetration depth (PD) with wavelength, for a
wavelength range between about 360 nm and about 440 nm. For
illuminating wavelengths greater than 440 nm, the penetration depth
increases sharply with wavelength. Similar characteristics are
found with other organs of the body. Thus, as may be seen from FIG.
2, the use of light radiation at the red end of the spectrum in
accordance with U.S. Pat. No. 4,449,535 or as proposed by
Kobayashi, to correct for blood haemodynamic artifacts in the NADH
signal introduces inaccuracies into the measurements due to
differences in penetration depths and therefore in the actual
sampling volumes. Even though both radiation wavelengths are
incident on the same spot, since detection is also at the same
point, effectively two different elements of tissue volume are
being probed since the different radiation wavelengths penetrate
the tissue to different depths. This results in measurements that
are incompatible one with the other, the blood volume measurement
relating to a greater depth of tissue than the NADH measurement.
Therefore, the device disclosed by this reference does not enable
adequate compensation of NADH to be effected using the
simultaneous, though inappropriate, blood volume measurement. There
is in fact no recognition of this problem, much less so any
disclosure or suggestion of how to solve it. Further, there is no
indication of how to measure other parameters such as blood flow
rate or blood oxygenation level using the claimed apparatus.
[0027] In earlier patents; U.S. Pat. Nos. 5,916,171 and 5,685,313
(which have a common inventor with the present invention), a device
is described that enables the monitoring of microcirculatory blood
flow (MBF), the mitochondrial redox state (NADH fluorescence) and
the microcirculatory blood volume (MBV), using a single source
multi-detector electro-optical, fiber-optic probe device for
monitoring various tissue characteristics to assess tissue
vitality. During monitoring, the device is attached to the
fore-mentioned tissue. The probe/tissue configuration enables
front-face fluorometry/photometry. The two most important
parameters involved in that fiber arrangement are the Optical
Penetration Depth (PD) and the Averaged Sample Depth (SD), the PD
parameter being dependent on both the tissue-type and on the
irradiation wavelength; the SD parameter being dependent on the PD
parameter and the distance between the ends of the excitation and
collection fiber in contact with tissue.
[0028] Although U.S. Pat. No. 5,916,171 and U.S. Pat. No. 5,685,313
represent an improvement over the prior art, they nevertheless have
some drawbacks: [0029] (i)The oxidation level of the blood will
introduce artifacts, affecting both the Mitochrondrial Redox State
measurement (NADH fluorescence) and the microcirculatory blood
volume (MBV) since these patents do not specify how to compensate
for the oxygenation state of the blood in the tissue, i.e., the
relative quantities of oxygenated blood to deoxygenated blood in
the tissue. This problem is substantially overcome in the present
invention by performing the NADH and blood volume measurements at
an isosbestic point of the oxyhaemoglobin--deoxyhaemoglobin
absorption spectrum. [0030] (ii)There is no facility included for
measurement of the oxyhaemoglobin--deoxyhaemoglobin level, i.e. the
Blood Oxygenation State, which is also an important tissue
viability parameter, worthy of monitoring. [0031] (iii) In these
two US patents, the same tissue volume needs to be monitored for
all parameters, and the same light source and wavelength is used
for the illumination needed for monitoring all three parameters. To
measure both the NADH level and the blood flow rate, a relatively
powerful UV laser is used. Using a relatively high intensity UV
laser illumination source as proposed raises safety issues,
especially for long-term monitoring. [0032] (iv) The blood flow
measurements impose several requirements on the UV laser source. In
particular, the UV laser should have a high coherence length and
very low intensity optical noise. Such lasers at these wavelengths
are also not standard components and are indeed quite difficult to
come by, which might lead to supply problems. [0033] (v) There is
no suggestion of monitoring Fp level, with or without any of the
other parameters.
[0034] It is an aim of the present invention to overcome the above
deficiencies in the prior art.
[0035] Particularly, it is an aim of the present invention to
provide a method and apparatus enabling the simultaneous in-vivo
monitoring of blood flow rate (i.e. intravascular mean velocity
times the number of moving red blood cells) and at least one, and
preferably all, of the following: NADH concentration by
fluorescence, total blood volume (i.e. concentration of red blood
corpuscles) by reflectometry, blood haemoglobin oxygenation (i.e.
the oxy/deoxy haemoglobin ratio) by fluorescence, flavoprotein
concentration by fluorescence; for the same body tissue, in
substantially the same layer within the same region. These
parameters, which represent different biochemical and physiological
activities of the tissue, are used to assess the tissue vitality in
said layer and tissue region.
[0036] It is another aim of the invention to provide flexibility in
design of apparatus for simultaneous measurement of four or five
different parameters with reference to the same tissue layer.
[0037] It is another objective of the present invention to provide
a method and apparatus for enabling the blood oxygenation of a
tissue to be measured, which overcomes the deficiencies of the
prior art.
[0038] It is another aim of the present invention to provide a
method and apparatus for enabling the blood oxygenation of a tissue
to be measured where prior art absorption methods cannot be
used.
[0039] It is a further aim of this invention to enable the
concurrent monitoring of blood parameters in different regions of
the same organ.
[0040] It is a further aim of this invention to enable the
concurrent monitoring of blood parameters in the same or different
region of a number of different organs of the same type, for
example the kidney of a number of patients.
[0041] It is a further aim of this invention to enable the
concurrent monitoring of blood parameters in different organs
belonging to the same or different patients.
[0042] Other objects and advantages of the invention will become
apparent as the description proceeds.
[0043] These and other objectives are realised by the present
invention by a revolutionary approach to tissue viability
measurement, directed at a common tissue layer concept rather than
based on necessarily using the same excitation wavelength for all
parameters. The same tissue layer measurements can be achieved, as
explained further on, by utilizing several wavelengths that are all
confined within a well defined wave-band, or alternatively by using
even very different wavelengths and making adequate compensation
for variable penetration depths, rather than being restricted to
using a single radiation illumination as taught by U.S. Pat. No.
5,916,171 and U.S. Pat. No. 5,685,313.
[0044] Thus, the NADH fluorescence, blood volume and blood
haemoglobin oxygenation state are measured using the same
monochromatic illumination wavelength and the same detection
fibers, ensuring that the same tissue volume is monitored for these
three blood parameters. The illumination point is not coincident
with the detection point, and the spacing between these points may
be chosen according to the average sample depth that is desired.
The wavelength of the monochromatic light is chosen to lie at one
of the isosbestic points of the extinction coefficient vs.
wavelength curves for oxyhaemoglobin and deoxyhaemoglobin; wherein
the NADH or blood volume measurements will be substantially
unaffected by the oxygenation state of the blood.
[0045] The blood flow rate may be measured by Laser Doppler
Flowmetry (LDF), typically using coherent light (a laser
radiation), the illumination being applied at the same point on the
tissue as for the above three parameters. Furthermore, this laser
radiation can also be used for excitation of Fp fluorescence which
enables the monitoring of flavoprotein concentration, which is an
important physiological parameter, as discussed above.
[0046] However, the location of the detection fibers with respect
to the illumination fibers, specifically the distance between their
ends, is set to a different value to compensate for the different
penetration depth of the two illuminating wavelengths, and thus to
ensure that the same layer of tissue is monitored for the blood
flow parameter and flavoprotein fluorescence, as is monitored for
the other blood and tissue parameters. This distance will vary;
both as a function of the tissue type being monitored and as a
function of the selected wavelengths of the two illuminations.
While it is generally preferable to monitor NADH and LDF over
exactly the same tissue volume, not just layer, to achieve this
aim, the excitation wavelengths used for the parameters being
monitored should be confined within predefined wave-band for which
penetration depth is substantially insensitive to wavelength.
[0047] Nonetheless, for any given type of tissue, there exists in
general, a range of wavelengths with substantially the same
penetration depth for each tissue type. For example, as illustrated
in FIG. 2 for brain tissues, this plateau in penetration depth as a
function of wavelength, extends from about 360 nm to about 440 nm,
with some indication from other sources, that the plateau extends
to even lower wavelengths. Similarly some other tissues feature
similar plateaus at these or other wavelengths. If the monitored
tissue is radiated using different wavelengths over the appropriate
range, the penetration depth will be similar, and substantially the
same volume of tissue may be monitored. In such cases, the same
detection fibers may be used for both illuminating wavelengths.
[0048] Thus, the present invention also provides a method and
apparatus for the measurement of blood oxygenation level based on
fluorescence measurements, rather than reflection measurements.
Essentially, a single radiation at a particular wavelength
illuminates a tissue such as to stimulate the emission of
fluorescence by the tissue. The intensity of the fluorescent
radiation at two or more wavelengths (within the fluorescent
radiation band) is measured, and the level of oxygenation is
derived from these measurements. Such a method and apparatus is
advantageously incorporated within the apparatus of the invention
in which a number of tissue viability parameters are determined.
Alternatively, a stand-alone device and corresponding method may be
provided for the measurement of blood oxygenation level.
[0049] As discussed above, U.S. Pat. No. 5,916,171 and U.S. Pat.
No. 5,685,313 are directed at the use of a single radiation at a
single wavelength for monitoring a number of tissue viability
parameters, including blood flow rate and NADH level. On the other
hand, EP 442011 describes a sensor for non-invasive measurement of
a single parameter, oxygen saturation in a tissue. In one
embodiment, shown in FIG. 2 thereof, a carrier means has mounted
therein a single light transmitter emitting electromagnetic waves
of different wavelengths, and two receivers at different distances
from the transmitter, each receiver being sensitive to a different
one of these wavelengths reflected from the tissue.
[0050] Returning to the references U.S. Pat. No. 5,916,171 and U.S.
Pat. No. 5,685,313, a single optical fiber guide carries the single
illuminating radiation to the tissue and receives light from the
tissue via another fiber, and the received light is then directed
into two separate channels. A single illuminating radiation is used
to ensure that the same tissue volume is being considered for all
the tissue vitality parameter measurements. Thus, not only would
there be no motivation for a man of the art to consider these
documents when desiring to provide an apparatus with two
illuminating radiations at different wavelengths, these references
actually teach against using more than one illuminating radiation
source, and more so at different illuminating wavelengths. On the
other hand, EP 422011 is directed exclusively at the measurement of
a single parameter, oxygen saturation in a tissue, and does not
consider in any shape or form the measurement of multiple tissue
viability parameters such as blood flow rate and NADH--in fact it
is not concerned with the measurement of two or more parameters,
but rather uses both receivers to measure a single parameter. Thus,
there would be no motivation for a man of the art to combine EP
442011 with U.S. Pat. No. 5,916,171 or U.S. Pat. No. 5,685,313 when
seeking to provide a device according to the present invention.
Moreover, even if the sensor of EP 442011 were to be combined with
the apparatus of U.S. Pat. No. 5,916,171 or U.S. Pat. No.
5,685,313, the combination would not yield the present invention.
For example, the apparatus of U.S. Pat. No. 5,916,171 and U.S. Pat.
No. 5,685,313 does not provide radiation at a range of wavelengths,
and the radiation is provided directly from a remote radiation
source via optical fiber. This enables a relatively small tissue
area to be monitored as the cross-section of the probe can
therefore be quite small. On the other hand the sensor of EP 442011
has the transmitter itself (in the form of an LED) mounted onto the
carrier, which therefore needs to be large enough to accommodate
the same, and which has power leads, rather than optical fibers
connecting the carrier to an external power source. Thus, these two
devices--the apparatus and the sensor--are not compatible with each
other, and very significant modifications to the two would be
required to enable the sensor of EP 422011 to be incorporated into
the apparatus of U.S. Pat. No. 5,916,171 and U.S. Pat. No.
5,685,313. This still leaves the question of how to configure the
combination so that each of the receivers of EP 422011 is coupled
to a different measuring channel of the apparatus of U.S. Pat. No.
5,916,171 and U.S. Pat. No. 5,685,313. More importantly, though,
the sensor of EP 422011 is characterised in that the two receivers
are mounted on the carrier in distances selected such that the
lengths of the light paths through the tissue at the two different
wavelengths are substantially equal. In such a case, by definition,
the two different wavelengths must be directed to two different
tissue layers, not to mention entirely different tissue volumes.
Thus, not only would the fact that different tissue volumes are
targeted by EP 422011 teach away from considering this reference in
combination with U.S. Pat. No. 5,916,171 and U.S. Pat. No.
5,685,313 in the first place, such a combination still does not
provide the apparatus of the present invention in which a single
tissue layer is targeted by both radiations. In the present
invention, the relative location of the detection fibers with
respect to the illumination fibers, specifically the distance
between their ends, is set to such as to compensate for the
different penetration depths of the two illuminating wavelengths,
and thus to ensure that the same layer of tissue is monitored for
the blood flow parameter, as is monitored for the other blood and
tissue parameters. Clearly, far from providing this arrangement, EP
422011 teaches away therefrom.
[0051] Regarding the determination of oxygenation level of a tissue
according to the present invention, such a method and corresponding
device are not disclosed or suggested in the prior art.
[0052] For example, WO 99/02956 uses a laser induced fluorescence
method for assessing the levels of ischemia and hypoxia in a
tissue, rather than blood oxygenation level. The method comprises
the steps of (a) measuring the fluorescence spectra at two
different points on the tissue; (b) calculating the tissue
absorption spectrum from these measurements; and (c) calculating
the intrinsic fluorescence spectrum. Thus, in order to perform the
calculations for determining the absorption spectra, two different
points on the tissue need to be considered, and thus the device
requires two different detectors coupled to two corresponding
measurement channels, in contrast to the present invention in which
a single point on the tissue suffices for obtaining fluorescence
measurements therefrom, which are in the form of intensity
measurements. Further, there is no disclosure or suggestion of
using the intensity of the fluorescent radiating at two or more
wavelengths for determining ischemia or hypoxia, and less so for
determining blood oxygenation level.
[0053] In U.S. Pat. No. 5,318,022 and in WO 98/44839, oximetry
techniques are described, wherein in each case an excitation source
of several wavelengths is used, and the intensity of the reflected
radiation for each wavelength is measured, wherein the appropriate
ratios of oxygen saturation are determined. In contrast, the
present invention uses only a single wavelength, and the intensity
of the fluorescent radiation emitted as a result thereof is
measured at two or more fluorescent wavelengths, from which blood
oxygenation level is determined.
SUMMARY OF THE INVENTION
[0054] The present invention relates to an apparatus for
selectively monitoring a blood flow rate tissue viability parameter
and at least one second tissue viability parameter corresponding to
at least a substantially identical layer of tissue element, the
apparatus comprising:-- [0055] illumination means for illuminating
at least said layer of tissue element with a first illuminating
radiation at a first wavelength and with a second illuminating
radiation at a second wavelength via at least one common
illumination location with respect to said tissue element; [0056]
first radiation receiving means for receiving a first radiation
from said layer of tissue element as a result of an interaction
between said first illuminating radiation and said layer of tissue
element, said first radiation being correlated to said blood flow
rate tissue viability parameter, said first radiation receiving
means being displaced from said illumination location by a first
displacement; [0057] second radiation receiving means for receiving
at least a second radiation from said layer of tissue element as a
result of an interaction between said second illuminating radiation
and said layer of tissue element, said second radiation being
correlated to said at least one second tissue viability parameter,
said second radiation receiving means being displaced from said
illumination location by a second displacement.
[0058] Typically, the blood flow rate tissue viability parameter is
provided by the Doppler shift of said first radiation received by
said first radiation receiving means with respect to the said first
illuminating radiation. First detection means are provided for
detecting said first radiation received by said first radiation
receiving means.
[0059] The second tissue viability parameter may be NADH
concentration, and said corresponding second radiation received by
said second radiation receiving means is an NADH fluorescence
emitted by the tissue in response to illumination thereof by said
second illuminating radiation, said at least one second tissue
viability parameter being provided by the intensity of said NADH
fluorescence. Second detection means are provided for detecting
said second radiation received by said second radiation receiving
means.
[0060] Additionally or alternatively, the second tissue viability
parameter is blood volume within said tissue element, and said
corresponding second radiation received by said second radiation
receiving means is a reflection from the tissue element in response
to illumination thereof by said second illuminating radiation; the
said at least one second tissue viability parameter being provided
by the intensity of said reflection. Third detection means for
detecting said second radiation received by said second radiation
receiving means.
[0061] Additionally or alternatively, the second tissue viability
parameter is blood oxygenation ratio within said tissue element,
and said corresponding second radiation received by said second
radiation receiving means is a fluorescence emitted by the tissue
in response to illumination thereof by said second illuminating
radiation, said at least one second tissue viability parameter
being provided by the intensity of said fluorescence at least at
two fluorescent wavelengths. Optionally, one of said at least two
fluorescent wavelengths is chosen to lie at an oxy-deoxy
fluorescence emission isosbestic point. Alternatively, one of said
at least two fluorescent wavelengths is higher and another one of
said at least two fluorescent wavelengths is smaller than a
wavelength corresponding to an oxy-deoxy fluorescence emission
isosbestic point. Fourth detection means for detecting said second
radiation received by said second radiation receiving means.
[0062] The common illumination location is typically provided by at
least one excitation optical fiber capable of being brought into
registry with said tissue element. The ratio of said first
displacement to said second displacement may be substantially
correlated to a ratio of said second wavelength to said first
wavelength. In one embodiment, the first radiation receiving means
comprises at least one suitable first receiving optical fiber
capable of being brought into registry with said tissue element,
and the second radiation receiving means comprises at least one
suitable second receiving optical fiber capable of being brought
into registry with said tissue element. Preferably, the at least
one excitation optical fiber, said at least one first receiving
optical fiber and said at least one second receiving optical fiber
are housed in a suitable probe head.
[0063] In another embodiment, the ratio of said first displacement
to said second displacement is substantially unity. Optionally, but
not necessarily, the first wavelength is substantially the same as
said second wavelength, and the first wavelength and said second
wavelength have substantially similar penetration depths with
respect to said tissue element. In this embodiment, the first
radiation receiving means and said second radiation receiving means
are comprised of at least one common third optical fiber capable of
being brought into registry with said tissue element. Preferably,
the at least one excitation optical fiber and said at least one
common third optical fiber are housed in a suitable probe head.
[0064] For all embodiments, the first illuminating radiation of
said first wavelength is typically provided by a first coherent
light source, which is preferably a laser light source. Optionally,
the laser light source is adapted to provide said first
illuminating radiation of said first wavelength in pulses of
predetermined duration and intensity. Optionally, the apparatus may
further comprise suitable control means for controlling the
frequency of pulsing of said pulses. The control means may be
further adapted to provide said pulses in packages of pulses, each
package comprising at least one pulse and separated from a
preceding or following package by a predetermined time period. This
predetermined time period is greater than the time interval between
consecutive pulses within a package. Preferably, this time period
is controllably variable, and the number of pulses within each
package is also controllably variable. Advantageously, the control
means is operatively connected to at least one of said first
detection means, second detection means, said third detection means
and fourth detection means. The control means is preferably
selectively responsive to previously detected signals corresponding
to the detection of said second radiation detected by means of any
one of said second detection means, said third detection means or
said fourth detection means, of a prior monitoring cycle.
[0065] For all embodiments, the said second wavelength is chosen to
lie at a suitable isosbestic point of the NADH. This isosbestic
point is within a substantially isosbestic range of the
oxyhaemaglobin--deoxyhaemaglobin absorption vs. wavelength curves.
Typically, the said second wavelength is within the range of
wavelengths corresponding to the NADH excitation spectrum, and is
typically between about 300 nm and about 395 nm, and preferably
between about 300 nm and about 340 nm. More specifically, the said
second wavelength is within about .+-.5 nm of any of the following
wavelengths; 325 nm, 337 nm, 349 nm, 355 nm, 366 nm, 370 nm, 385 nm
or 390 nm. The radiation of said second wavelength may be provided
by any suitable UV light source.
[0066] In some embodiments such as in the first embodiment, the
first wavelength may be substantially different from said second
wavelength. In other embodiments, such as in the second embodiment,
the first wavelength may be substantially similar to or even the
same as the second wavelength, wherein the penetration depths
associated with the first and second wavelengths are substantially
the same.
[0067] In the first embodiment, the first wavelength is typically
about 440 nm or greater than this. In the second embodiment, the
first wavelength is typically between about 300 nm and about 440
nm.
[0068] Typically, then, the first wavelength may be either in the
range 410.+-.30 nm or within about .+-.5 nm of any one of the
following wavelengths; 355 nm, 430 nm, 440 nm, 455 nm, 460 nm, 490
nm, 532 nm or 805 nm.
[0069] In the third and fourth embodiments of the invention, the
apparatus is adapted for further monitoring a flavoprotein
concentration tissue vitality parameter, comprising fifth detection
means for detecting a portion of said first radiation received by
said first radiation receiving means, said portion of said received
first radiation being a flavoprotein fluorescence emitted by the
said tissue element in response to illumination thereof by said
first illuminating radiation, said flavoprotein tissue viability
parameter being provided by the intensity of said flavoprotein
fluorescence. In these embodiments, the said first wavelength is
within the range of wavelengths corresponding to the flavoprotein
excitation spectrum. Typically, the first wavelength is between
about 400 nm and about 470 nm, and preferably within about 440 nm
and 460 nm.
[0070] The present invention also relates to a system for
selectively monitoring at least two tissue viability parameters at
a plurality of tissue elements; said system comprising a plurality
of monitoring probes, each said probe comprising an apparatus as
described hereinbefore. At least two said probes may be adapted for
monitoring said tissue viability parameters of tissue elements
within the same organ. Additionally or alternatively, at least two
said probes are adapted for monitoring said tissue viability
parameters of tissue elements within different organs. The
different organs may be different organs within the same organism,
and/or different organs within different organisms, and/or include
donor organs.
[0071] The second illuminating radiation of said second wavelength
for each said probe may be provided by a common suitable light
source, which is typically a UV monochromatic light source or the
like, which in turn may be adapted to provide said second
illuminating radiation of said second wavelength in pulses of
predetermined duration and intensity. The system optionally further
comprises suitable control means for controlling the frequency of
pulsing of said pulses.
[0072] Similarly, the first illuminating radiation of said first
wavelength for each said probe may be provided by a common suitable
light source, which is typically a laser light source, which in
turn may be adapted to provide said first illuminating radiation of
said first wavelength in pulses of predetermined duration and
intensity. The system optionally further comprises suitable control
means for controlling the frequency of pulsing of said pulses. The
control means may be further adapted to provide said pulses in
packages of pulses, each package comprising at least one pulse and
separated from a preceding or following package by a predetermined
time period. The predetermined time period may be set to be greater
than the time interval between consecutive pulses within a package.
Preferably, this time period is controllably variable. The number
of pulses within each package may also be controllably variable.
The control means may be adapted for selectively directing discrete
said pulses to any one of said probes.
[0073] The system optionally further comprises any one and
preferably all of the following:-- [0074] (a) suitable first common
detection means for detecting said first radiation received by said
first radiation receiving means of at least two said probes, and
coupling means for optically connecting the at least two said
probes to said first common detection means; [0075] (b) suitable
second common detection means for detecting said second radiation
received by said second radiation receiving means of at least two
said probes, and coupling means for optically connecting the at
least two said probes to said second common detection means,
wherein said second tissue viability parameter corresponding to
said second radiation is NADH concentration; [0076] (c) suitable
third common detection means for detecting said second radiation
received by said second radiation receiving means of at least two
said probes, and coupling means for optically connecting the at
least two said probes to said third common detection means, wherein
said second tissue viability parameter corresponding to said second
radiation is blood volume; [0077] (d) suitable fourth common
detection means for detecting said second radiation received by
said second radiation receiving means of at least two said probes,
and coupling means for optically connecting the at least two said
probes to said fourth common detection means, wherein said second
tissue viability parameter corresponding to said second radiation
is blood oxygenation ratio; [0078] (e) suitable fifth common
detection means for detecting a portion of said first radiation
received by said first radiation receiving means of at least two
said probes, said portion of said received first radiation being a
flavoprotein fluorescence emitted by the said tissue element in
response to illumination thereof by said first illuminating
radiation, said flavoprotein tissue viability parameter being
provided by the intensity of said flavoprotein fluorescence, said
system further comprising coupling means for optically connecting
the at least two said probes to said fifth common detection
means.
[0079] The control means may be operatively connected to the first
common detection means. Optionally, the control means may also be
operatively connected to any one of, and preferably all of, the
first common detection means, the second common detection means,
the third common detection means, the fourth common detection means
and the fifth common detection means.
[0080] The control means may also be selectively responsive to
previously detected signals corresponding to the detection of said
second radiation detected by means said second, third, fourth or
fifth common detection means of a prior monitoring cycle,
respectively.
[0081] In a another aspect of the invention, the present invention
is directed to a method for determining the oxygenation state of a
tissue element, comprising: [0082] providing an illuminating
radiation to the tissue element, said illuminating radiation being
at a wavelength within the NADH excitation spectrum or within the
flavoprotein (Fp) excitation spectrum; [0083] measuring the
intensity of the corresponding NADH or Fp fluorescence,
respectively, emitted by the tissue element at least at two
predetermined wavelengths within the range of wavelengths comprised
in the corresponding fluorescence emission; [0084] comparing the
intensities measured in (b) to provide an estimate of the relative
levels of oxygenated blood to deoxygenated blood in said tissue
element.
[0085] One of said wavelengths in the second step may be chosen to
lie at a suitable second isosbestic point, which is a suitable NADH
oxy-deoxy fluorescence emission isosbestic point, or an Fp
oxy-deoxy fluorescence emission isosbestic point, respectively.
[0086] The intensity of fluorescence emitted by the tissue element
at two target wavelengths within the range of wavelengths comprised
in the fluorescence may be measured, wherein a first target
wavelength thereof is chosen at below the wavelength corresponding
to said second isosbestic point, and wherein a second said target
wavelength thereof is chosen at above the wavelength corresponding
to said second isosbestic point. Preferably, the first target
wavelength and the second target wavelength are chosen such as to
correspond to the maximal change in fluorescence intensity
occurring at wavelengths below and above, respectively, the second
isosbestic point.
[0087] The intensity of fluorescence emitted by the tissue element
may be measured over a range of wavelengths within a specified
window of the range of wavelengths comprised in the
fluorescence.
[0088] In the third step of the method, the fluorescence intensity
at each wavelength in the second step is normalised by a
"normalising" fluorescence intensity measured at a suitable
corresponding NADH OR Fp oxy-deoxy fluorescence emission isosbestic
wavelength within the corresponding fluorescence emission
spectrum.
[0089] The method may be adapted for NADH, wherein the illuminating
radiation in step (a) is within the range of wavelengths comprised
in the NADH excitation spectrum, and in which case the
"normalising" fluorescence intensity is measured at an isosbestic
wavelength of about 455 nm.+-.5 nm.
[0090] Alternatively, the method may be adapted for Fp, wherein the
illuminating radiation in step (a) is within the range of
wavelengths comprised in the Fp excitation spectrum, and in which
case the "normalising" fluorescence intensity is measured at an
isosbestic wavelength of about 530 nm.+-.5 nm.
[0091] The illuminating radiation is typically, but not necessarily
provided at a suitable NADH oxy-deoxy fluorescence excitation
isosbestic wavelength or a Fp oxy-deoxy fluorescence excitation
isosbestic wavelength, respectively, within the corresponding NADH
or Fp fluorescence excitation spectra, respectively.
[0092] According to this aspect of the invention, an apparatus is
provided for determining the oxygenation state of a tissue element,
comprising: [0093] illumination means for illuminating said tissue
element with an illuminating radiation at a predetermined
wavelength via at least one illumination location with respect to
said tissue element; [0094] radiation receiving means for receiving
a fluorescence emitted from said tissue element as a result of an
interaction between said illuminating radiation and said tissue
element, [0095] suitable detection means for detecting said
fluorescence received by said radiation receiving means.
[0096] The detection means is adapted for detecting the intensity
of the fluorescence received by said radiation receiving means at
least at two fluorescent wavelengths. One of said at least two
fluorescent wavelengths may be chosen to lie at an isosbestic point
in the fluorescent emission spectrum. Alternatively, one of said at
least two fluorescent wavelengths is higher and another one of said
at least two fluorescent wavelengths is smaller than a wavelength
corresponding to an isosbestic point in the fluorescent emission
spectrum. Preferably, the said predetermined wavelength of said
illuminating radiation corresponds to an isosbestic point of the
excitation spectrum. The illumination location is typically
provided by at least one excitation optical fiber capable of being
brought into registry with said tissue element. The radiation
receiving means typically comprises at least one suitable first
receiving optical fiber capable of being brought into registry with
said tissue element. The at least one excitation optical fiber and
the at least one receiving optical fiber are preferably housed in a
suitable probe head. Typically, the predetermined wavelength of
said illuminating radiation is within the range of wavelengths
corresponding to the NADH excitation spectrum, and the fluorescence
emitted from said tissue as a result of an interaction between said
illuminating radiation and said tissue element is an NADH
fluorescence emission. Alternatively, the predetermined wavelength
of said illuminating radiation is within the range of wavelengths
corresponding to the Fp excitation spectrum, and the fluorescence
emitted from said tissue as a result of an interaction between said
illuminating radiation and said tissue element is an Fp
fluorescence emission.
[0097] Thus the present invention relates to an apparatus or device
for in-vivo monitoring of NADH level, Fp level, microcirculatory
blood volume, microcirculatory blood flow and blood haemoglobin
oxygenation, in the same layer of tissue in substantially the same
location over the same time period, and corresponding methods for
performing said monitoring.
[0098] More particularly, the invention relates to a single probe
device that measures these parameters, indicative of the function
of the tissue in the identical layer and substantially the same
location of the tissue; determining the ratios of those parameters,
and storing and retrieving said information to enable long-term
monitoring.
[0099] The NADH level, microcirculatory blood volume and blood
haemoglobin oxygenation being determined by monitoring the
fluorescence and reflectance resulting from UV monochromatic light
irradiation at an isosbestic wavelength of the oxy-deoxy
haemoglobin absorption spectrum. The microcirculatory blood flow
and Fp level being measured by Laser Doppler Flowmetry (LDF), with
irradiation by a suitable laser light source.
[0100] The present invention enables the monitoring of the
metabolic state of a tissue by NADH fluorometry. Using the device
of the invention, it is possible to monitor NADH levels in a
certain tissue volume and to correct efficiently for the
Haemodynamic Artifact. Reflection measurements are used to monitor
the blood volume that present in the same volume element, and to
resolve the interference of haemoglobin with the NADH measurements.
The reflection measurements are taken at the same wavelength, and
from the same tissue sample as the NADH measurements. This enables
monitoring of corrected NADH fluorescence in same volume element.
All fluorescence excitation measurements are performed at low
irradiation intensities to avoid photo-bleaching of the measured
NADH chromophores. In addition, NADH excitation wavelengths are
specified to resemble the isosbestic points of the Haemoglobin
oxy-deoxy absorption spectrum. This prevents artifact resulting
from the changes in the oxygenation state.
[0101] The reflection of the light at the excitation wavelength is
used to monitor the Blood Volume in same volume element examined.
The reflection signal is measured in the excitation wavelength,
which is at an isosbestic point of the absorption spectra of
oxyhaemoglobin and deoxyhaemoglobin and at NADH excitation
wavelengths. The irradiation intensity is maintained as low as
possible in order to eliminate possibility of photo-damage to the
irradiated tissue.
[0102] The blood flow rate is monitored using Laser Doppler
Flowmetry. The light intensity is optimized to give a high
signal-to-noise ratio, with the choice of wavelength and intensity
being specified to minimise the possibility of photo-damage to the
irradiated tissue. The laser irradiation signal is modulated by
suitable chopping means to limit the exposure of the tissue to
laser light to within acceptable limits. By utilizing a separate
light source for monitoring the blood flow from that used for
monitoring the other parameters, it is possible to monitor the
other parameters constantly over extended periods of time whilst
only irradiating the tissue with the laser light required for
measuring the blood flow on an intermittent basis. The LDF system
can be activated periodically by a clock triggering mechanism, or
it may also be switched to actively monitor the volume blood flow
during periods of activity, being triggered by a change in one of
the other parameters.
BRIEF DESCRIPTION OF THE DRAWINGS
[0103] The present invention will be more clearly understood from
the detailed description of the preferred embodiments and from the
attached drawings in which:
[0104] FIG. 1 shows the excitation fluorescence spectrum
(F.sub.EXT) and emission fluorescence spectrum (F.sub.EMS) for
NADH, in terms of the corresponding fluoresence intensities (IF) as
a function of wavelength (WL).
[0105] FIG. 2 illustrates typical penetration depth characteristics
for human brain tissues as a function of illumination
wavelength.
[0106] FIG. 3(a) illustrates, in transverse cross-sectional view, a
probe according to a first embodiment of the first aspect of the
present invention.
[0107] FIG. 3(b) shows in greater detail the tissue-abutting end of
the probe of FIG. 3(a).
[0108] FIG. 4(a) illustrates in end view the embodiment of FIGS.
3(a) and 3(b) taken along X-X.
[0109] FIG. 4(b) illustrates in end view a second embodiment of the
probe of the present invention.
[0110] FIG. 5 illustrates schematically the main components of the
first embodiment of the first aspect of the present invention.
[0111] FIG. 6(a) illustrates schematically the normalised
fluorescence intensity (IF) emitted by a tissue as a function of
wavelength (WL) and oxygenation level of the blood contained in the
tissue.
[0112] FIG. 6(b) illustrates schematically the ratio of
fluorescence intensities at two wavelengths with respect to the
fluorescence intensity at an isosbestic point of FIG. 6(a).
[0113] FIG. 7(a) shows the main clock sequence that enables the
transmission of the light from light source (102) by the
acousto-optic modulator (AOM).
[0114] FIG. 7(b) shows the output voltage of the detector in
response to the light modulated by the AOM.
[0115] FIG. 7(c) shows the clock sequence applied to the sample and
hold (S/H) circuitry (440) and (470) as shown in FIGS. 9(a) to
9(c).
[0116] FIG. 7(d) shows the clock sequence that is applied to the
reference sample and hold (S/H) circuitry (450) as shown in FIGS.
9(a) to 9(c).
[0117] FIG. 7(e) shows the light signal as it appears at the output
of (S/H) circuitry (450) and (470) of FIGS. 9(a) to 9(c).
[0118] FIG. 7(f) shows the sequence train of pulses as provided
during state II operation of the device.
[0119] FIG. 7(g) shows the sequence train of pulses as provided
during state III operation of the device.
[0120] FIG. 8 illustrates schematically the main components of the
second embodiment according to the first aspect of the present
invention.
[0121] FIG. 9(a) schematically illustrates a circuit diagrams for a
signal detector optionally used with the embodiments of FIG. 5 and
FIG. 8.
[0122] FIG. 9(b) schematically illustrates a circuit diagrams for
another signal detector optionally used with the embodiments of
FIG. 5 and FIG. 8.
[0123] FIG. 9(c) schematically illustrates a circuit diagrams for
another signal detector optionally used with the embodiments of
FIG. 5 and FIG. 8.
[0124] FIG. 10 illustrates light absorption of blood
oxy-haemoglobin and blood deoxyhaemoglobin in terms of an
Extinction parameter (E) as a function of wavelength (WL).
[0125] FIG. 11 illustrates a fifth embodiment of the present
invention comprising a plurality of probes.
[0126] FIG. 12 shows the excitation fluorescence spectrum
(FP.sub.EXT) and emission fluorescence spectrum (FP.sub.EMS) for
Fp, in terms of the corresponding fluoresence intensities (IF) as a
function of wavelength (WL).
[0127] FIG. 13 illustrates schematically the main components of the
third embodiment of the first aspect of the present invention.
[0128] FIG. 14 illustrates schematically the main components of the
fourth embodiment of the first aspect of the present invention.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
[0129] The present invention is defined by the claims, the contents
of which are to be read as included within the disclosure of the
specification, and will now be described by way of example with
reference to the accompanying figures.
[0130] In the description to follow, the following illustrative
apparatuses and methods are described, it being understood that the
invention is not limited to any particular form thereof, and the
following description being provided only for the purposes of
illustration.
[0131] The present invention is directed to an apparatus for
simultaneously monitoring at least one tissue viability parameter
from each of two sets of tissue viability parameters, from at least
a substantially identical layer of tissue element, and preferably
from the same volume of tissue element. In particular, one of these
parameters of the first set is the blood flow rate corresponding to
the tissue layer, and a dedicated radiation at a particular
wavelength is used for monitoring this parameter in conjunction
with a laser Doppler flowmetry method (LDF). The second set of
tissue viability parameters includes at least one of, and
preferably more than one of, and most preferably all of, at least
NADH concentration, blood volume, blood oxygenation corresponding
to the tissue layer, and this parameter or plurality of parameters
are measured using another dedicated radiation at a desired
wavelength, which may be chosen to be the same as or different from
the wavelength of the radiation used for monitoring the blood flow
rate. The first set of tissue viability parameters also includes a
fifth tissue viability parameter, flavoprotein concentration, can
be monitored using the same illumination source that is used for
the LDF measurements.
[0132] Thus--this being a great advantage of the present
invention--the blood flow rate measurement is conducted totally
independently from the monitoring of the second set of tissue
viability parameters, providing a great deal of flexibility in
terms of configuration and design of the monitoring apparatus, as
well as in the method of use, as will be evident from the following
description.
[0133] The wavelength of the illumination radiation that is
required for any of the second set of blood viability
parameters--NADH, blood volume and oxy-deoxy state--is determined
by the absorption and fluorescence characteristics of the NADH in
the tissue, as exemplified in FIG. 1 for human brain tissue (the
absorption and fluorescence characteristics of NADH are similar for
other tissues). The fluorescence wavelength band is typically
between about 400 nm and about 550 nm, and an excitation wavelength
between about 300 nm and about 400 nm needs to be provided.
[0134] Preferably, an excitation wavelength is chosen such as to
simplify correction for the haemodynamic artifact. The haemodynamic
artifact arises from the absorption of the NADH fluorescence
emission and excitation light by the blood haemoglobin. A change in
blood volume cause misleading changes in apparent NADH
fluorescence. Since blood haemoglobin has two oxygenation states
namely oxy-haemoglobin and deoxy-haemoglobin each one with its
distinct absorption spectrum, as shown in FIG. 10, the precise
correction of the haemodynamic effect can become extremely complex.
The problem is considerably simplified when the wavelength chosen
for NADH fluorescence excitation corresponds to one of the
isosbestic points, since at these wavelengths the absorption of
both haemoglobin species is identical. At these isosbestic
excitation points, fluorescence changes are due substantially to
changes in total blood volume, and, of course, to changes in NADH
concentration only. Thus, by suitably choosing the excitation
wavelength, correction for the haemodynamic artifact is
significantly simplified. Even at wavelengths within the isosbestic
range of from about 300 nm to about 340 nm, or at the isosbestic
point of 390 nm, the haemodynamic artifact requires correction.
Suitable algorithms for this purpose are described on the prior art
(Koyabashi et al., 1971; Renault G., et al. American Journal of
Physiology, 246, H491-H499, 1984; Mayevsky A. and Chance B., Brain
Res. 65, 529-533, 1974; Harbig et al., J. Appl. Physiol. 41,
480-488, 1976; U.S. Pat. No. 4,449,535). The most widely used
correction algorithm (Jobsis et al. Neurophysiology 3465, 735-749,
1971) utilizes the value of the reflection at the NADH excitation
wavelength as an indicator for blood changes. The corrected NADH
fluorescence values are calculated by subtraction of the reflection
signal from the fluorescence signal.
[0135] At the same time, the wavelength of the illumination
radiation that is required for the monitoring of the first tissue
viability parameter--i.e., the blood flow rate--may be
independently chosen from that required for the other parameters.
For the various soft body tissues, there are ranges of wavelengths
over which the penetration depths of incident light is relatively
independent. This is illustrated by way of example for brain tissue
in FIG. 2. Here the penetration depth into tissue is relatively
independent of the incident radiation wavelength for wavelengths
between 360 nm to about 440 nm, while above this range, the
penetration depth increases dramatically with wavelength.
[0136] The Optical Penetration Depth (PD) may be defined as the
depth where the total optical power of incident light is reduced to
37% of the incident power. The PD parameter is useful for
estimation of light depth penetration into tissues. If two
different wavelengths are used in the same instrument it is
important to compare the penetration depth for both wavelengths in
order to estimate differences in collected light intensities. In
general the PD is proportional to 1/.mu..sup.1/2, where .mu. is the
total absorption coefficient, given by:--
.mu.=.mu..sub.a+.mu..sub.s where
[0137] .mu..sub.a is the apparent absorption coefficient due to
light absorption; and .mu..sub.s is the apparent absorption
coefficient due to light scattering.
[0138] The average sample depth (SD) parameter is dependent on the
tissue, the wavelength used and on the separation distance of the
excitation and collection fibers in the probe. The SD can be
approximated by the equation (Taitelbaum H., OSA Proceeding on
Advances in Optical Imaging and Photon Migration (ed. by Alfano R.
R.) 21, 305-309, 1994): SD.apprxeq.0.4r.sup.1/2PD.sup.1/2 where r
is the distance between the excitation and collection fiber. This
coefficient depends on both the fiber separation and on the PD.
[0139] Apparatuses that incorporate a laser light source are
generally required to comply with relevant laser safety standards.
The two relevant standards which deal with exposure of human tissue
to laser radiation are the ANSI Z136.1-2000 "American National
Standard for Safe Use of Lasers" and the IEC60825-1-1994
International Standard called "Safety of laser products".
[0140] These standards define the Maximum Permissible Exposure
(MPE) values. These standards relate to laser irradiation of
external tissues such as skin and eye and not of the internal
organs, in contrast to typical applications of the present
invention. Still they are the only known, well established
references to safe irradiation values for tissues, and any laser
device that is intended to perform nondestructive measurements
should comply with these in the absence of a more appropriate full
damage test being performed on specific tissue type with specific
light irradiation.
[0141] Both the above standards permit a maximum of 1 mW/cm.sup.2
irradiance for the UVA wavelengths region (about 315 nm to about
400 nm) for an exposure time larger then 1000 sec. This requirement
implies a severe limitation on the light intensity emitted by the
distal tip of the fiber optic probe, particularly when shorter
wavelength, higher intensity radiation is used.
[0142] In the present specification, the magnitudes of wavelengths
specified herein may be varied by about .+-.5 nm, and even up to
about .+-.10 nm without significantly affecting operation of the
apparatus of the invention.
[0143] The first embodiment of the present invention, according to
a first aspect thereof, is directed to an apparatus in which the
wavelength of the illumination radiation that is required for
monitoring the first set of tissue viability parameters is outside
the range of relative independence of penetration depth with
wavelength. By way of example for brain tissue, this illumination
would have a wavelength substantially higher than about 440 nm,
while the second embodiment of the present invention is directed at
an apparatus in which the wavelength of the illumination radiation
that is required for the first set of tissue viability parameters
is within said range of relative independence, being between 300 nm
and about 440 nm if the brain is the tissue under examination. As
will become clearer hereinbelow, the question of whether the first
set of tissue parameters includes the blood flow rate and
flavoprotein fluorescence, or only the blood flow rate, will depend
on whether or not the first illuminating radiation is of a
wavelength within the flavoprotein excitation spectrum,
respectively.
[0144] The second aspect of the present invention is directed
towards a method and corresponding apparatus for the monitoring of
blood oxygenation state using fluorescence measurements. Such
method and apparatus may also be advantageously incorporated in the
apparatus according to the first aspect of the present
invention.
[0145] Thus, in the first aspect of the present invention, and
referring to FIGS. 3(a), 3(b), 4(a) and 5, the apparatus according
to the first embodiment, generally designated by the numeral (100),
is directed to the monitoring of blood flow rate of the first set
of tissue viability parameters, and any combination of the three
tissue viability parameters of the second set of tissue viability
parameters. Thus, the apparatus (100) may be in the form of a probe
(2) having at the distal tip thereof contact face (12) for making
contact with the surface of the tissue (25) being monitored. In its
simplest form, the probe (2) has a single fiber (201) for directing
two radiations of different wavelengths to the same point (15) on
the tissue (25). Alternatively, a bundle of fibers may replace a
single fiber (201). The two radiations may come from first and
second sources, (22) and (24) respectively, and are coupled to the
fiber (201) by any suitable optical coupler. Referring to FIG. 2,
other than at the plateau of wavelengths between 300 nm and about
440 nm; for a given illumination wavelength substantially above
this range, the further the detection means is displaced from the
illumination means, the greater the effective depth into the tissue
that may be monitored by the detection means. Thus, if a detection
means is provided at location A displaced by a distance R1 from the
illumination fiber (201), the radiation from first source (22)
having wavelength (WL1) will penetrate to a depth (D1) as
illustrated in FIG. 3(b). If the detection means is at position
(B), displaced from the illuminating fiber (201) at a greater
distance (R2), then the penetration of the radiation corresponding
to this position is much greater (D2). When second source (24)
provides radiation of wavelength (WL2), substantially shorter than
(WL1), if detection is effected from position (B) as well, the data
obtained relating to the second source (24) will be for a much
shallower depth than for the first source (22). Conversely, if the
detection with respect to the higher wavelength (WL1) is performed
at a distance, say (R1), closer to the illuminating fiber (201)
than the detection distance (R2) with respect to the lower
wavelength (WL2), then it is possible to choose (R1) and (R2) such
that the detection corresponding to both wavelengths corresponds to
the same sampling depth (D1). While the actual sampling volumes for
each of the wavelengths in such a case will be different, the
quality of results obtained is substantially insensitive to such
differences since it confined to the same tissue layer, in contrast
to providing results at substantially different depths.
[0146] For most body tissues, the anatomic structure of the blood
supply is generally similar in that there is a definite gradient in
blood oxygenation level with penetration into the tissue. The
oxygenated blood supply to the tissue can be in the general
direction from the monitored surface into the tissue, as in the
brain, or out of the tissue towards the monitored surface as in the
skin or the kidneys. In common to both flow directions however, the
blood carrying capillaries running parallel to the tissue surface
in the superficial layers substantially run randomly throughout the
two dimensional plane. Since the overall direction of blood supply
is thus generally perpendicular to the tissue surface, the changes
in blood flow, blood oxygenation and partial oxygen pressure will
occur along the normal to the tissue surface rather than in the
plane of the tissue surface. On the other hand, the distribution of
cells and mitochondria is very homogeneous along the vertical as
well as the horizontal axis in all tissues. Therefore, monitoring
of blood flow, oxy-deoxy haemoglobin, NADH fluorescence, as well as
flavoprotein fluorescence, will be homogeneous in the plane
parallel to the tissue axis, whereas monitoring along an axis
perpendicular to the tissue surface will provide heterogeneous
results.
[0147] Preferably, the probe (2) comprises plastic flexible housing
in the form of tube (208) to protect the optical fibers, which are
advantageously encapsulated within a stainless steel tube (209) at
its distal tip.
[0148] As described above, the configuration of the excitation
fiber (201) and the collection fibers (202), (203), in particular
their relative positions within the probe (2), is important in
ensuring that at least the same tissue layer is being considered
for all the parameters being monitored. When more than one
detection fiber (202) and more than one detection fiber (203) are
used, this combined plurality of fibers may be arranged in two sets
on concentric circles arranged coaxially with the excitation fiber
(201), as illustrated in FIG. 4(a). The gaps or distances (R2),
(R1), respectively between the excitation fiber (201) and each set
of collection fibers (202), (203) influence both the average sample
depth (SD) and the collected signal intensity. The number of
collection fibers (202), (203), as well as the core diameters of
each type of fiber, also influence both these factors as well as
the signal to noise ratio (S/N) of the laser Doppler measurements.
For example, good results for monitoring the blood flow in brain
gray matter tissues were achieved using a 532 nm laser illumination
source, a 200 micron core excitation fiber (201), used together
with four collection fibers (203), each of 100 micron core
diameter; and a separation gap (R1) of 0.25 mm.
[0149] If the second group or set of parameters, (i.e. NADH
fluorescence, blood volume and blood oxygenation state) is measured
by 390 nm excitation light, for example, the collecting fibers
(202) should be placed at a separation gap (R2) three times the
separation gap (R1) in order to ensure measurement from the same
tissue layer. This requirement originates in the fact that the PD
of the 532 nm light is about 0.6 mm for brain gray matter (see FIG.
2) while the PD for the 390 nm light is about 0.2 mm. In other
words, at an excitation wavelength of about 390 nm the PD is three
times lower than for excitation wavelengths of 532 nm. Since the
sampling depth determined by the relationship SD.apprxeq.0.4
(R*PD).sup.1/2 it is clear that in order to maintain the same SD
for both wavelengths the separation gap (R2) should be three times
larger for NADH fluorescence collecting fiber (202). Therefore the
fiber (202) should be placed at about 0.75 mm from the excitation
fiber (201). The diameter of this collecting fiber (202) may be
larger than the diameter of collection fibers (203), since the
collected intensity will be lower at such large distance.
[0150] For monitoring other body tissues, the gaps (R1), (R2)
between the excitation fiber (201) and the collection fibers (203),
(202), the number of each of the collection fibers (202), (203)
provided in the probe (2), and the core diameters of the excitation
fiber (201) and of the collection fibers (202), (203) may be
individually optimized for each tissue type and for each type of
excitation wavelengths.
[0151] Referring in particular to FIG. 5, in the preferred
embodiment of the present invention the first radiation source (22)
is a laser light source (102) for laser Doppler Flowmetry (LDF),
and the second radiation source (24) is a monochromatic light
source (101) of specified wavelength for monitoring at least one of
NADH, blood volume and oxy-deoxy haemoglobin levels. Both light
sources (101), (102) are comprised in a light source unit (LSU),
shown at (1).
[0152] The probe (2) is preferably disposable, but may be
semi-disposable or non-disposable. The term "disposable" in the
present application means that the probes are designed (in
corresponding embodiments) to be disconnected from the rest of the
apparatus (100) and thrown away or otherwise disposed off after one
use with only negligible economic loss. Negligible economic loss
herein means an economic loss per probe which is substantially less
than that of the apparatus (100) itself, or of the medical costs
associate with a procedure using said apparatus (100), or indeed of
the costs associated with sterilising and reconditioning the probe
for a single subsequent use. The term "semi-disposable" herein
means that while the probe is disposable, it may nevertheless be
used a limited number of times, with appropriate sterilising and
reconditioning thereof between uses. The term "non-disposable"
herein means that the probe (2) is designed for multiple use, and
is only disposed of when sterilisation and reconditioning thereof
is no longer possible or economic. Thus, the probe (2) is typically
designed for once-only use for minimising risk of cross-infection,
for example. Optionally, though, the probe (2) may be adapted for
sterilisation using an ETO or any other suitable sterilization
technique, enabling the probe to be semi-disposable or
non-disposable. In any case, the probe (2) is also typically made
from biocompatible materials.
[0153] The probe (2) is operatively connected, in addition to the
LSU (1), to the following:--a detection unit--(DTU) (3), a signal
processing and conditioning electronics unit--(EU) (4), a suitable
computer (PC) with dedicated software (5), and a suitable power
supply--(PS) (6), which are also comprised in apparatus (100).
[0154] Radiation of two wavelengths (WL1), (WL2) from the LSU (1)
is delivered to the tissue (25) to be monitored via a single
optical or excitation fiber (201) (or bundle thereof). The
excitation fiber (201) and the collecting fibers (202) and (203)
are placed in direct contact with tissue (25) in order to maximize
the portion of light signal that penetrates the tissue and is
subsequently collected from the tissue.
[0155] The photons of the penetrating light undergo scattering and
absorption as they interact with the body tissue matter. The
scattering of excitation light is mainly due to interaction with
stationary tissue and with the red blood cells. The absorption of
the excitation light of the second wavelength (WL2) is mainly due
to tissue and blood haemoglobin, and to a lesser extent is due to
NADH molecules. Some of the energy that is absorbed by NADH is
re-emitted by NADH molecules as fluorescence photons, a small
portion of whom eventually reaches the tissue surface, and are
collected by one or more collection fibers (202) and transmitted to
the DTU (3). Doppler shift changes in the radiation of first
wavelength (WL1) give a measure of the blood flow rate, and such
changes are detected via one or more collection fibers (203).
[0156] The DTU (3) comprises appropriate optical filters and
detectors for converting the collected light intensities to
electronic signals from which the four tissue vitality parameters
may be monitored. The converted signals from the DTU are fed into
the EU (4) for processing.
[0157] The EU (4) serves as a conditioning and signal processing
system. It also converts the analogue signals to digital data that
feeds into computer (5). The acquired data is processed by suitable
software and may displayed by any suitable means and form, such as
for example on the computer screen as charts and in digital form.
The PS unit (6) provides each of the components of the apparatus
(100) with the required electrical power.
[0158] Excitation fiber (201), and collecting fibers (202) and
(203), are provided with optical connectors (205), (206) and (207),
respectively, to enable convenient coupling of the fibers to the
LSU (1) and to the DTU (3) respectively.
[0159] The light from collecting fiber (203) is collimated by a
suitable lens (313) within the DTU (3), and the collimated light
passes through a long-wavelength pass filter (not shown). The said
filter blocks out any reflections at the excitation wavelengths
(WL2) and enables the passage of the longer wavelength reflection
at excitation wavelength (WL1). The light that passes through this
filter is then channeled towards a low-noise, fast photodiode
detector (315). Preferably, a suitable condensing lens (314) is
used in order to fill the photo-detector active area. The signal
thus obtained from the photodiode detector (315) is used to perform
Laser Doppler Flowmetry measurements in the usual manner, to
determine the blood flow rate tissue viability parameter.
[0160] The light from collection fiber (202) is also collimated by
a suitable lens (306), passing first through a cut-off filter (not
shown) which enables the shorter wavelengths, including reflection
at (WL2) and the NADH fluorescence to pass through, while blocking
the higher wavelengths such as reflection at wavelength (WL1). The
resulting collimated and filtered light beam is then split by a
series of dichroic mirrors or beam splitters (302), (303),
(304).
[0161] The light collected by the collection fibers (202) consists
mainly of reflected light at the excitation wavelength, but it also
comprises much lower intensity NADH fluorescence light at higher
wavelengths. The portion of the collimated beam comprising the
reflected light will thus have the lowest wavelength, corresponding
to the excitation wavelength, while at the same time having the
highest intensity of the radiation collected by the collecting
fiber (202). Thus, the first dichroic mirror (302) splits off light
at the excitation light wavelength from the collimated beam,
channeling this portion of the beam towards a low-noise, fast
photodiode detector (301). Preferably, a condensing lens (305) is
used in order to fill the photo-detector active area. The dichroic
beam splitter (302), therefore reflects most of the light at
excitation wavelength and while permitting transmission
therethrough for most of the higher wavelengths in the collimated
beam, and thus provides enough filtration for the photodiode
detector (301), with no additional filtration being generally
needed. The signal from the photodiode detector (301) is used to
perform reflection measurements to determine the blood volume
tissue viability parameter. The remainder of the collimated light
beam continues towards the second dichroic mirror (303).
[0162] Light of wavelengths higher than the excitation wavelength
passes through the dichroic mirror or beam splitter (302) and is
incident on a second dichroic beam-splitter (303), which is
selected to reflect wavelengths lower then about 440 nm and to
transmit all higher wavelengths. The reflected light beam is passed
through a suitable filter (307), preferably a 435 nm (10DF) filter,
and is then fed into a first photo-multiplying tube (PMT) (308).
The light transmitted through the second dichroic beam-splitter
(303) is subjected to additional splitting by a third dichroic
beam-splitter (304) that reflects wavelengths lower then 460 nm,
but is transparent to higher wavelengths. The reflected light from
the third dichroic beam splitter (304) is filtered by a suitable
filter (309), preferably a 455 nm (10DF) filter, and is then
incident on a second photo-multiplying tube (PMT) (310). This
wavelength is close to an oxy-deoxy isosbestic point, so the
fluorescence intensity as measured by this PMT (310) correlates
directly with the NADH fluorescence. The light that passes through
the third dichroic beam-splitter (304) is subsequently filtered by
a suitable filter (311), preferably a 475 nm interference filter
(DF10), and the filtered light is incident on a third
photo-multiplying tube (PMT) (312). The precision of all
above-mentioned filters are .+-.5 nm.
[0163] The fluorescence intensity measurements provided by the
first, second and third PMTs (308), (310) and (312) respectively,
are used to determine the blood oxygenation state, i.e., the ratio
of oxygenated blood to deoxygenated blood, within the tissue
element, according to the second aspect of the present
invention.
[0164] According to a second aspect of the present invention, a
fluorescence-based method is used for determining the blood
oxygenation state of a tissue element. The method may be
incorporated in a dedicated probe-like device, for example, or with
the other monitoring apparatuses known in the art, but particularly
with the apparatus (100) according to the first aspect of the
present invention. According to the second aspect of the invention,
the tissue element is illuminated by an illumination radiation, and
the wavelength of said illumination radiation being is preferably,
but not necessarily, chosen to correspond to a suitable isosbestic
point, and such that fluorescence is emitted from the tissue
element. The intensity of the fluorescence emitted, as a function
of wavelength, will vary according to the blood oxygenation state
of the tissue element.
[0165] According to the second aspect of the present invention, the
NADH fluorescence intensities at two points on either side of the
fluorescence isosbestic wavelength of about 455 nm are each
normalised with respect to the fluorescent intensity at this
isosbestic point. Alternatively, the Fp fluorescence intensities at
two points on either side of the fluorescence isosbestic wavelength
of about 530 nm are each normalised with respect to the fluorescent
intensity at this isosbestic point. In either the NADH or Fp case,
at the corresponding said isosbestic point, the absorption by the
blood does not change as the blood changes its oxygenation state.
Similarly, the whole fluorescence spectra may be normalised with
respect to the intensity at this isosbestic point, and such
normalisation of the fluorescent intensities provides a better
representation of the fluorescence spectra, as it is independent of
the actual fluorescent intensities and actually compensates for any
changes in corresponding NADH or Fp concentration or
instrumentation factors such as fiber to tissue coupling, for
example. Since the normalisation renders the absolute fluorescent
intensity unimportant in itself in the oxygenation state
determination, it is also possible to provide the illumination
radiation at a wavelength which is not at an isosbestic point of
the absorption spectrum of the oxy-deoxy haemoglobin Indeed if the
excitation wavelength is not the isosbestic one, this will cause
some changes in the fluorescence intensity, since these changes are
only in total intensity but not in the functional form of the
fluorescence spectrum as function of wavelength. Such changes
disappear in the corresponding normalized spectrum. For example,
excitation at non-isosbestic wavelength at about 355 nm by means of
a 3.sup.rd harmonic Nd-Yag laser, will generally result in changes
in the fluorescence intensity due to oxygenation changes at
excitation wavelength, additionally to the changes of the form of
the emission curve. Referring to the NADH case by way of example,
by normalizing the fluorescence emission spectrum by the
fluorescence intensity at the isosbestic wavelength of about 450
nm, changes due to oxygenation changes at the excitation are
effectively cancelled out.
[0166] Thus, referring to FIG. 6(a), curve (P) represents the
intensity of the NADH fluorescence (IF), emitted for the full range
of wavelengths (WL) of the emission when the blood in the tissue
element is fully oxygenated, while curve (R) shows the
corresponding (IF)-(WL) relationship for the fully deoxygenated
condition. Curve (Q) represents an intermediate condition in which
the blood is partially oxygenated and partially deoxygenated. All
the curves pass through point (T), which is herein referred to as
an "oxy-deoxy" fluorescence emission isosbestic point,
corresponding to a wavelength of (WLIP). This oxy-deoxy emission
isosbestic point will be at the same wavelength as the isosbestic
point of oxy-deoxy haemoglobin absorption spectrum namely 455 nm as
shown in FIG. 10. The actual absolute values of the fluorescence
intensities will depend on the illuminating radiation wavelength,
but this dependence disappears by normalising the emission spectrum
by the value of fluorescence intensity at the oxy-deoxy emission
isosbestic wavelength (WLIP) of about 455 nm as shown in FIG. 6(a).
Thus, normalised fluorescence emission spectra obtained at a
plurality of illuminating radiation wavelengths will go through a
common normalised point (T), regardless of the illuminating
radiation wavelength.
[0167] By measuring the ratio of the intensity of the fluorescence
at a wavelength below (WLIP), say at (WLL) to the fluorescence
intensity at (WLIP) and also at a higher wavelength, the ratio of
the intensity at say (WLH) to the intensity at (WLIP), the actual
blood oxygenation state can be determined.
[0168] Thus, and referring to FIG. 6(b), if the normalised
fluorescence intensity, i.e., ratio IF(WLL)/IF(WLIP) is increased,
then the blood is substantially mostly oxygenated, while the
converse is true if the intensity ratio decreases. An increase in
the normalised fluorescence intensity, ratio IF(WLH)/IF(WLIP),
indicates that the blood has become more deoxidized. By suitable
calibration, particularly of the maximum and minimum normalised
intensities (IF) at these wavelengths, the actual relative
percentages of oxygenated to deoxygenated blood may be determined
given the fluorescence intensities measured at these points. In
order to maximize the sensitivity and precision of the method the
(WLL) and (WLH) should be chosen in such a way that the change of
the ratio IF(WLL)/IF(WLIP) and IF(WLH)/IF(WLIP) will be maximised
with respect to oxy-deoxy relative concentration variations.
Therefore, the wavelengths where this change is maximal,
(WLL).sub.MAX and (WLH).sub.MAX, should be used. Indeed, greater
sensitivity to even minor changes in blood oxygenation may be
achieved by monitoring the ratio of the aforementioned ratios
IF(WLL)/IF(WLIP):IF(WLH)/IF(WLIP) which is, of course, equivalent
to the ratio of IF(WLL)/IF(WLH).
[0169] The electronic and electro-optic components described herein
are given by way of example. There are many alternative methods of
realizing the current invention. For example, although the
monitoring of the three parameters is accomplished with PMT
detectors, optical filters and dichroic splitters in the embodiment
described herein, it is possible to replace all these components by
using a grating spectrometer and appropriate detector such as a CCD
or by using a multianode PMT with a Multi-band interference filter
such as Hamamatsu R5900F-L16. These solutions could potentially
monitor intensity ratios with even higher precision, but at current
prices, are not economical options.
[0170] Thus, the present invention is also directed, in the second
aspect thereof, to a suitable apparatus for determining the blood
oxygenation level of a tissue element, the apparatus comprising an
illumination means capable of illuminating the tissue element with
radiation at a particular wavelength, preferably but not
necessarily corresponding to a suitable isosbestic point of the
absorption spectra for oxyhaemoglobin and deoxyhaemoglobin, and
detection means for detecting the intensity of the fluorescence
emitted by the tissue, at least at two wavelengths. These two
wavelengths are any two different wavelengths chosen from the
oxy-deoxy isosbestic wavelength (WLIP), a wavelength (WLH) higher
than (WLIP), and a wavelength (WLL) lower than (WLIP). Preferably,
the detection means detect the intensity of the fluorescence
emitted by the tissue, at all three wavelengths. Preferably, the
wavelengths (WLL) and (WLH) must be chosen to be at the point where
maximal change in fluorescence intensity occur and at two different
sides of the isosbestic point (WLIP). Typically, the fluorescence
emission spectrum is normalised by the fluorescence intensity at
the oxy-deoxy fluorescence isosbestic wavelength (WLIP), which for
NADH is about 455 nm.+-.5 nm.
[0171] Similarly, the method and apparatus according to the second
aspect of the present invention adapted with respect to the Fp
fluorescence is similar to that described herein with respect to
that adapted with respect to NADH fluorescence, mutatis mutandis,
with the main difference that normilisation of the fluorescence
intensities is with respect to the fluorescence intensity at a
corresponding Fp oxy-deoxy fluorescence emission isosbestic
wavelength, typically about 530 nm.+-.5 nm.
[0172] Returning now to the first embodiment according to the first
aspect of the present invention, blood oxygenation level is
provided by the first, second and third PMTs, (308), (310) and
(312) respectively, in accordance with the second aspect of the
present invention, and wherein the second PMT (310), in which
fluorescence intensity is measured at an isosbestic point, also
provides the NADH parameter. Thus, the ratio of the fluorescence
intensity measured by the first PMT (308) to the intensity measured
by second PMT (310), generally increases as the blood becomes more
oxygenated, while the ratio of the fluorescence intensity as
measured by the third PMT (312) to the intensity measured by the
second PMT (310) under the same conditions will decrease.
Conversely, as the blood becomes more de-oxygenated, the
fluorescence intensity ratios measured by the first PMT (308) and
by the third PMT (312) relatively to the intensity measured by
second PMT (310) generally will decrease and increase,
respectively. The measured fluorescence ratios can be calibrated to
actual levels of oxy-deoxy haemoglobin using measurements by other
known methods, such as pals-oximetery. Thus relative levels of
oxygenated blood to deoxygenated blood within the tissue element
may be determined.
[0173] By way of example, a suitable component for the PMT detector
modules (308), (310) and (312) is the Hamamatsu 6780 PMT. Each of
the PMT detector modules (308), (310), and (312) comprise a PMT
tube and all electronics necessary for the PMT gain control. These
modules are supplied with the operation voltage and each module has
gain control input and signal output connections. The electronics
circuit for all 3 PMT detectors is identical so only PMT detector
(312) will be described.
[0174] The signal output of the PMT detector (312) is fed to the
conditioner (402) input. There are several ways of accomplishing
the signals processing which are well known in the art. All
detectors in the proposed system are synchronous detectors. The
appropriate electronic circuit is described below.
[0175] Thus, in the first embodiment, the same fiber (201) is used
for illumination by both the first wavelength (WL1), typically a
laser incident light wavelength, and the second wavelength (WL2),
typically a UV monochromatic wavelength. Two sets of detection
fibers (202) and (203) are situated at specific distances, R2 and
R1, from the illumination fiber (201) to ensure that the tissue
monitored by both the UV monochromatic and the laser incident light
wavelengths is from the same layer (D1) of tissue (25).
[0176] In the first embodiment, particularly when used for
monitoring brain tissue, the second light source (101) provides
monochromatic UV light with wavelength of about 390.+-.5 nm for
monitoring the NADH, blood volume and oxy-deoxy haemoglobin level
(blood oxygenation state). This light source (101) may be a
filtered spectral lamp such as mercury or xenon lamp, a light
emitting diode--LED, or a suitable laser such as laser diode. The
specified wavelength for (WL2) thus complies with the two important
properties hereinbefore discussed:--it falls in the absorption
spectrum of the NADH molecule, and it is at an isosbestic point of
the haemoglobin oxy-deoxy absorption spectrum. It is thus a
preferred wavelength for both NADH fluorescence excitation and for
blood volume measurements by reflection. The first light source
(102) may be provided by any suitable laser of suitable intensity,
coherence length and optical noise, having a wavelength greater
than about 440 nm.
[0177] As hereinbefore described, the first embodiment of the
present invention employs two separate radiations at different
wavelengths for illuminating the tissue element, and thus the blood
flow rate measurement may be conducted totally independently from
the monitoring of the other tissue viability parameters, albeit
within the same tissue layer (and preferably the same tissue
volume) providing a great deal of flexibility in terms of
configuration of the monitoring apparatus, as well as in the method
of use.
[0178] In general, it is preferable that all four parameters,
namely the first set for measuring blood flow, and the second set,
for measuring NADH, blood volume and blood oxygenation level, are
monitored with irradiation and detection occurring with high
sampling frequency, however this may give rise to safety issues
regarding the excitation radiation used for the measurements. In
particular, the monitoring of the second set of tissue viability
parameters--NADH fluorescence, blood volume and blood oxygenation
level, requires use of excitation wavelengths between 300 nm and
400 nm, which lie in the UVA spectral region; the exposure to which
should be minimised as it is considered to be potentially dangerous
even at low irradiation levels. The monitoring of the first
parameter namely the blood flow by laser Doppler technique also
raises safety issues, especially where the laser Doppler utilizes
irradiation inside UVA region, since a higher irradiation intensity
is required for Doppler measurements than that needed for the
measurements of the second set of parameters, and even when the
laser Doppler wavelength is in the visible spectral region, there
are still safety concerns due to the relatively high irradiation
intensity required. In order to minimise the problem, the option is
provided in the present invention to chop the excitation light with
a duty cycle of 1/20 (ON/OFF). Additionally there are many clinical
conditions where continuous monitoring with frequent updating is
unnecessary, and this constraint may therefore be relaxed. Thus,
whereas during critical parts of a surgical operation procedure the
output data should be renewed at least at the rate of two data
points per second, there are however, many cases where the
patient's condition is stable, so that a data sampling rate of
only, say, once every two seconds is required, for example.
[0179] According to the present invention the apparatus (100) may
be used according to an adaptive chopping procedure. In such an
adaptive chopping procedure, the radiation provided by each one of
sources (101) and/or (102) may be chopped to provide corresponding
pulses of radiation at the appropriate wavelengths, the pulses
being provided at a preferably variable frequency of pulsation,
i.e., chopping frequency. Furthermore, the apparatus (100) may be
further adapted such that packages of pulses may be provided as and
when required or desired. Such packages may each comprise a
variable number of pulses, and the time interval between packages
of pulses may also be independently controlled. Thus, at periods
where relatively little monitoring is required, few packages
containing a few pulses each may be transmitted with large "OFF"
intervals in-between packages (i.e., where no radiation is
provided), while at other, more intense periods, the packages of
pulses may be sent with little or no intervals between successive
packages. By pulsing, and by also packaging the pulses as
described, the radiations provided by sources (101) and also by
(102) may be of a correspondingly higher permitted intensity than
would normally be allowable, albeit for shorter durations. This
results in better signal-to-noise ratios of the signal, as well as
to safer radiation levels for both the patient and the operators of
the apparatus and equipment.
[0180] Using the concept of adaptive chopping, it is possible to
entirely stop the laser Doppler measurements after this parameter
has reached a steady state. The remaining three parameters, the
second group, may be measured by providing short packages of pulses
at a frequency of, say, twice a second. Indeed the second set
parameters will also be in steady state until some change occurs.
If the change originates in the blood flow rate, it will
immediately induce a change in the other, actively monitored
parameters, such as the blood volume. The apparatus (100) may then
be configured such that when such a change is detected, the Laser
Doppler measurements automatically restart and continue until at
least the next steady-state condition is reached.
[0181] Thus, referring to FIG. 5 and also FIG. 7(a), the laser
radiation or light from source (102), typically a stabilized laser
diode, is chopped by an Acousto-Optic Modulator (AOM) (103). A
clock (403) that is part of the EU (4) generates the chopping
sequence. The chopped light appears at the 1.sup.st order of the
modulator. This order is spatially filtered by a circular diaphragm
(not shown) and coupled to an excitation fiber by the lens (104)
mounted on a suitable adapter (105). The excitation light is split
by a beam splitter (106) and a small portion of it is directed
towards a photodiode (107). This photodiode provides tracking of
the excitation intensity of light source (102).
[0182] A similar beam-splitter (108) may also be utilized with the
other light source (101), typically monochromatic 390.+-.5 nm light
source, and a small portion of signal is directed towards
photodiode (109). This photodiode provides tracking of the
excitation intensity of light source (101).
[0183] The radiations originating from the two light sources (102)
and (101) are combined to be colinear by the cube beam combiner
(111). Preferably, the polarisations of the two radiations are
mutually perpendicular, providing advantages in their transmission
efficiencies.
[0184] The outputs of the photodiode detectors (301), (315), (107)
and (109) and the outputs of the three PMT detectors (308), (310),
(312) are connected to the signal conditioner (402). The signal
conditioner (402) receives synchronization signals that correspond
to the chopping sequence from the clock (403). The signal
conditioner features three groups of `channels` or synchronous
detector circuits, which will be described below.
[0185] The signal conditioner (402) of the EU (4) converts the
chopped signals into continuous wave (CW) signals. These are
converted by the A/D unit (401) into digital data, which is then
fed into the computer (5) through the analog input output (AIO)
ports. The A/D sub-unit (401), besides digitizing the analog
measured signals, also enables the receiving of digital commands
from the computer (5) via the digital input output (DIO) ports.
[0186] The clock (403) sub-unit provides the appropriate timing for
the AOM (103) and the signal conditioner (402).
[0187] In the first embodiment, the source (102) may consist of a
single mode laser having a wavelength of 532 nm for laser Doppler
Flowmetry measurements, while the source (101) may provide, for
example, a 325 nm, 337 nm or 390 nm monochromatic excitation
wavelength used for NADH fluorescence, the resulting 415-470 nm
emission wavelengths being used for NADH fluorescence and blood
oxygenation measurements. Such an arrangement is suitable for brain
gray matter, wherein the penetration depth (PD) for a 532 nm
wavelength is 0.6 mm while for 325 nm or 337 nm or 390 nm
wavelengths, the PD is only 0.2 mm, as illustrated in FIG. 2. Since
the ratio of the PD of these two wavelength groups is 3, in order
to maintain the same average sample depth (SD) for both groups of
tissue viability parameters, there should be a 1/3 ratio for the
distance between the excitation to the collection fibers of the two
groups, i.e. for R2/R1.
[0188] While the first embodiment is used preferably with the
source (102) providing a radiation of wavelength substantially
higher than 440 nm, it may also be used with a source (102)
providing radiation of wavelength between 300 nm and 440 nm. Since
the wavelength of the other source (101) is also within this band,
the ratio R1:R2 is close to unity, since the penetration depths for
such a range of wavelengths is about constant (see FIG. 2). Thus,
rather than having two separate collection fibers (202) and (203),
the functions of these fibers may be accomplished by single fiber,
or indeed a plurality of such single fibers, each of which may have
the combined functions of fiber (202) and (203). (Of course,
separate fibers (202), (203), or pluralities thereof, may also be
used.)
[0189] Thus, a second embodiment of the present invention,
illustrated in FIG. 8 and FIG. 4(b), comprises the same structural
elements as the first embodiment, with the exception of the said
collection fibers (202), (203) (including said optical connectors
(206), (207), and lens (313), condensing lens (314) and fast
photodiode detector (315)), as hereinbefore described, mutatis
mutandis.
[0190] In the second embodiment of the invention, the choice of
laser light for source (102) as used for LDF monitoring is limited
to the 300-440 nm range. Across this wavelength band, the
penetration depth is almost constant as taught by Eggert (Eggert,
H. R. & Blazek, V., Neurosurgery, 21, 459-464, 1987), and as
shown in FIG. 2. In this case, all measurements are made
simultaneously with the same excitation and the same collection
fibers, and virtually the same volume element of tissue is
monitored for the 4 parameters--blood flow rate, and the three
parameters of the second group, i.e., NADH, blood volume and blood
oxygenation state. Many other tissues also display similar
penetration depth plateaus at various wavelength ranges, and one
should appreciate that suitable embodiments could be devised for
applying this inventive concept for monitoring blood parameters for
such tissues, mutatis mutandis. In this second embodiment, a light
source (101) such as, for example, filtered spectral lamp, light
emitting diode LED, laser diode or pulsed laser, is used for
excitation of NADH fluorescence. The light source wavelength can be
any wavelength that is inside the absorption spectrum of the NADH
molecule, that is having a wavelength of from about 315 to about
395 nm. To avoid the Haemodynamic artifact, and to enable
measurement of oxy-deoxy haemoglobin, this wavelength is preferred
to be at one of the haemoglobin oxy-deoxy isosbestic points in this
bandwidth. There is no requirement for laser Doppler measurements
to be performed at one of these isosbestic wavelengths.
[0191] Referring to FIGS. 8 and 4(b), only one (or a plurality of)
collection fiber (223) and connector (267) are required instead of
the individual collection fibers (202), (203) (or corresponding
pluralities thereof), and the connectors (206), (207) that are
required for the first embodiment. The signal from photodiode
detector (301) is then used for LDF blood flow rate monitoring as
well as the reflection measurements for blood volume parameter.
[0192] As in the first embodiment, the light from the collecting
fiber (223) enters the DTU (3) via optical connector (267). The
collimating lens (306) collimates the light towards the first
dichroic mirror (302), which splits off the excitation light
wavelength, which is then channeled towards a low-noise, fast
photodiode detector (301), a condensing lens (305) being typically
used in order to fill the photo-detector active area. The signal
from the photodiode detector is used to perform both Doppler and
Reflection measurements.
[0193] Thus the reflection at both detection wavelengths (WL1) and
(WL2) will, in this embodiment, pass through the same collecting
fiber (223). The separation of these different signals is achieved
by time-sharing. The light sources (102) and (101) are working in
chopping mode therefore each source is ON only for a short time
period as is described hereinafter. Each light source has a low
duty cycle, therefore there is plenty of time to turn ON one light
source while the other one in still in the OFF period. Each one of
the corresponding detectors (107) and (109) is synchronously
sampled at the correct timing in order to get intensity information
regarding the corresponding excitation source. The detector (301)
is sampled twice, once for the measurement of reflection at WL1 and
corresponding laser Doppler measurements and a second time, for
reflection measurement at WL2. Both reflections at WL1 and WL2 have
substantially lower wavelengths than the NADH fluorescence and
therefore the same dichroic beam splitter (302) can be used for
separating these two relatively strong reflections from the weaker
fluorescence signal.
[0194] In embodiments where the Doppler excitation wavelength WL1
lies in the range 420 nm to 440 nm, in either the first or second
embodiment, the resulting NADH fluorescence will not pass through
the dichroic beam splitter (302) designed to split these
wavelengths towards detector (301). To overcome this problem at
these wavelengths, a simple 1:20 beam-splitter should be
substituted for the fore-mentioned dichroic beam splitter. It
should be noted, that the strong WL1 reflection will not interfere
with the NADH fluorescence, despite the two signals having similar
wavelengths, because the signal are separated in time.
[0195] While the appropriate illumination wavelengths for sources
(101), (102) referred to hereinbefore are particularly suited for
the monitoring of brain tissue in-vivo using the first and second
embodiments, corresponding illumination wavelengths may be
determined for any other tissue enabling the apparatus (100) to be
used with such tissues, mutatis mutandis.
[0196] In order to reduce the tissue irradiation the apparatus
(100) according to the present invention may be operated in any one
of several irradiation modes, and corresponding to these modes are
several data acquisition modes. There are two basic concepts behind
these operation modes:
[0197] The first concept relates to monitoring that is perceived to
be continuous by the clinical personnel. In general, all vitality
signals data should be presented to the medical personnel in
real-time. That is, the device display should be updated at the
rate that reflects the real physiology events as they evolve in the
patient. This means that if for example the patient is in a
critical stage of the surgery and there are a lot of fast changes
in the physiological conditions, the screen update rate should be
fast i.e. about two data points per second. However, where the
patient is in a more stable condition such as at the beginning of
the surgery, at its final stage or in the intensive care unit
(ICU), the vital parameters will generally tend not to change very
fast, and therefore a much slower screen update rate can be
utilized. In such cases the update rate can be for example one data
point every 2 seconds.
[0198] The second concept is that actually all vital parameters are
mutually connected and inter-related. Therefore a change in one
parameter should immediately trigger a change in at least one other
parameter. Especially any change in the blood flow will be
accompanied by a change in at least one of the other parameters:
blood volume, blood oxygenation or the NADH fluorescence. This
means that if the patient's state is steady, such as in ICU, the
monitoring of the blood flow can be stopped for long periods whilst
all the other tissue vitality parameters are monitored. Where any
significant change in the value of any one of these parameters is
detected the system will automatically start monitoring of all
parameters including the blood flow, until a steady state is again
reached.
[0199] Referring now to FIGS. 9(a), 9(b) and 9(c), In a specific,
non-limiting example of the preferred embodiments brought for
illustrative purposes, three types of optical detectors with
corresponding electronics circuits are used.
[0200] The first type of detector as shown in FIG. 9(a), is a
photon multiplier tube (PMT) detector. This type is suitable for
use as components (308), (310), (312) shown in FIGS. 5 and 8. These
detectors are used for NADH fluorescence measurements. The detector
is build around a PMT module from Hamamatsu H6780. This integrated
module consists of PMT tube, a high voltage power supply and all
necessary control electronics. One need only to supply the
operating voltage and the control voltage for the gain control, and
the module itself changes the high voltage of the PMT accordingly.
The gain of such detector may be controlled by the PC (5) through
the A/D (401) unit. The output of this PMT module is fed to the
inverter (410), since the module produces negative output relative
to ground. The output of the inverter is feed to dual sample and
hold (S/H) circuits built around S/H such as Analog Devices AD781
(440) and (450). The dual S/H circuit enables subtraction of the
dark current and background light that might interfere with the
desired measurements.
[0201] The second type of detector, illustrated schematically in
FIG. 9(b), is a fast photodiode detector such as (301) and (315)
(see FIG. 5) and (301) (see FIG. 8). These kinds of detectors are
used for reflection and Doppler measurements. This type of detector
is build around Hamamatsu S5973 photodiode (417) connected to
trans-impedance amplifier such as Analog Device AD713 (415). The
output of the trans-impedance amplifier (415) is fed to dual sample
and hold (S/H) circuits build around S/H such as Analog Devices
AD781 (440) and (450). The dual S/H circuit enables subtraction of
the ambient light that might interfere with the desired
measurements. This circuit is identical to the circuit used for PMT
based detector and will be described later.
[0202] The third type of detector, illustrated in FIG. 9(c), is a
fast photodiode detector. This is suitable for use for components
such as (107) and (109), which are shown in FIG. 5 and FIG. 8.
These types of detectors may be used for light source intensity
measurements. The light source intensity information is used at the
final stage of data processing to normalize the reflection and
fluorescence intensities according to the changes in the light
source intensities. This type of detector is built around a
Hamamatsu S5973 photodiode (422) connected to transimpedance
amplifier such as Analog Devices AD713 (420). The output of the
transimpedance amplifier (420) is fed to a single sample and hold
(S/H) circuit build around S/H such as Analog Devices AD781 (470).
Since the light to the detector is collected from an internal
source there is no need for subtraction of the background light as
in previous cases, therefore a single S/H circuit can be used.
[0203] The S/H circuits for the first two types of detectors as
illustrated in FIGS. 9(a) and 9(b) are substantially identical. The
output of the trans-impedance amplifier (415) or the inverter (410)
is connected to the first sample and hold S/H circuit (440). The
S/H circuit is triggered by the clock (403). Referring to FIGS.
7(a) to 7(e), the trigger signal timing 7(c) provided by the clock
(403) is correlated with the end of the light ON period in 7(a)
when the output voltage 7(b) of the detector is at a maximum,
enabling the S/H circuit to sample the maximum available signal.
The S/H circuit (440) holds this voltage value 7(e) until a new
trigger signal 7(c) arrives from the clock. The second S/H (450) is
also connected to the same signal input as S/H (440). This S/H
circuit receives a delayed hold signal 7(d), so that the sampling
occurs between the two pulses of 7(a). This delay results from the
delay circuit that is an integrated part of the clock (403). The
sampled intensity (not shown) bears information on the detector
dark current and the ambient light which both interfere with the
measurement. A difference amplifier (460) such as Texas Instruments
TL082 subtracts the output of one S/H from the other, the output of
this differential amplifier is the measured signal.
[0204] The output of the S/H (440) consist of the desired signal
along with noise such as the detector dark current, shot
(electronic) noise and the ambient light. The output of the S/H
(450) contains all these types of noise, but not the desired
signal. The outputs of S/H (440) and SH (450) are fed into a
difference amplifier (460) which subtracts the two signals, the
output being the net light signal. Of course, since the shot noise
is substantially random, it cannot subtracted.
[0205] The third detector circuit as shown in FIG. 9(c), comprises
a single S/H (470). The output of the trans-impedance amplifier
(420) which is identical to the amplifier (415) is connected to a
single sample and hold S/H circuit (470). The S/H circuit is
triggered by the clock (403). The trigger signal timing FIG. 7(c)
is correlated with the end of the light ON period in FIG. 7(a) when
the output voltage of the detector FIG. 7(b) is at a maximum, so
the S/H circuit samples the maximum available signal. The S/H
circuit (470) holds this voltage value until a new trigger signal
as shown in 7(c) arrives from the clock (403). This circuit is thus
very similar to the two previous ones described with respect to
FIGS. 9(a) and 9(b), but here there is no background light
subtraction by a second S/H since the light reaches this detector
from an internal source, and is thus free from external light
interferences.
[0206] The gain of the detectors is defined automatically by the
accompanying software in computer (5), according to the detected
light intensity values. If the detected light signal is too small,
the software provides an appropriate signal to increase the
detector gain as described below. There is a difference in the gain
management of the three types of the detectors as described
above.
[0207] The gain of the first detector type, the PMT, is set by
changing the control voltage (413) of the PMT module (412). This
actually changes the sensitivity of the detector PMT. The gain of
the inverter amplifier (410) is constant. The setting of the
control voltage is performed by the software that runs on the PC
(5) through the analog to digital converter (A/D) module (401) of
the electronics unit (EU) (4). This A/D and D/A module can be any
one of the variety of cards produced by National Instruments and
other manufacturers.
[0208] The gain of the second detector type is set by changing the
transimpedance amplifier (415) gain rather then by changing the
sensitivity of the photodiode detector itself. The setting of the
control voltage is performed by the software that is adapted to run
on the PC (5) through the A/D module (401) of the electronics unit
(EU) (4).
[0209] The gain of the third detector type is constant since this
detector measures light source intensity having a predefined value
that suits the constant dynamic range of the detector.
[0210] The gain setting procedure is initiated by the calibration
command from within the device software. The calibration signal
arrives from the computer (5) via digital to analog converter D/A
(401). At the beginning of the calibration procedure the gain
control voltage of the first and second detector type is reduced to
zero, and then, the gain gradually begins to increase whilst the
intensity of the output signal is monitored. With reference to the
output of the detectors (308), (310), (312) in FIG. 5 and FIG. 8
and the detectors (301) and (315) in FIG. 5 and the detector (301)
in FIG. 8, each detector gain is set separately. When the output
voltage reaches about 2V, the gain is locked to the current value.
This gain value is monitored by the software through the analog to
digital converter A/D (401). From then onwards, any change in
collected light intensity is monitored by the circuit and is
transformed to digital information by (A/D) (401). Since the gain
value, is known, the actual light intensity may be calculated and
displayed on the screen by the software.
[0211] The clock sub-unit (403) typically comprises a programmable
clock. According to computer input via bus (404) the clock output
will be in one of the following states (with particular reference
to FIGS. 7(a)-7(g)):-- [0212] State I: The clock signal consists of
a train of pulses in FIG. 7(a). The ON period t.sub.on of the cycle
is 10 microsec, while the whole cycle t.sub.cycle is 250 microsec
i.e. the repetition rate is 4 KHz. Therefore the duty cycle is
0.04. This sequence shown in 7(a) is used for enabling the light
source (102) by the triggering of the AOM (103), and also is used,
after appropriate delays, to trigger the signal S/H circuit (440)
and reference S/H (450) by sequences 7(c) and 7(d) respectively.
The sequence 7(c) is correlated to the end of each pulse shown in
7(a), while the sequence 7(d) is delayed by t.sub.cycle/4 in order
to enable to pick up the external light interference. The sequence
7(c) is also used to trigger the light source sensing detector
circuit S/H (470). Another second sequence identical to 7(a) with
its two delayed sequences is used for enabling light source (101)
and the appropriate detectors (109), (301), (308), (310) and (312)
(see FIG. 5 and FIG. 8).
[0213] This second sequence is delayed by t.sub.cycle/2 relative to
the first sequence in order to separate in time the measurements of
the two said sets of parameters. [0214] State II: State I is
additionally chopped by an ON/OFF adaptive duty cycle which enables
and disables the light pulses train 7(f). During the ON period
t'.sub.on (0.1 sec) of the adaptive duty cycle, 400 pulses 7(a) of
10 microsec each are generated. The OFF period t'.sub.off of the
adaptive duty cycle is controlled by the computer via bus (404).
The OFF period can be 0.4 sec for relatively fast-changing
conditions and can be prolonged to as much as 5 sec for slow
changing conditions. The t'.sub.off is determined automatically by
the software to minimize the total tissue irradiation. [0215] State
III: The clock generates a sequence of ten cycles of the state I
like the pulses shown in FIG. 7(a). These ten pulses are used for
enabling light source (101) alone and the appropriate detectors
(109), (301), (308), (310) and (312) on FIG. 5 and FIG. 8.
Therefore measurements of only the second set of parameters are
enabled.
[0216] The device software controls the tissue sampling and
irradiation. At measurement initialization the clock is in state I,
enabling the correct setting of the gain for all detectors, and the
normalization of the output signals. After a short time, if fast
changes in any one or more parameters are observable the clock is
switched to state II, having a short OFF period t'.sub.off. As the
changes became more moderate, the OFF period t'.sub.off becomes
longer. After cessation of the changes as steady state is achieved,
the system switches to state III in order to minimize the tissue
irradiation. Detection of changes causes the system to switch back
to state II.
[0217] In state III, only 10 pulses of WL2 are supplied to the
tissue during t''.sub.on see FIG. 7(g). This enables quick
measurement of the second set of parameters using only very limited
irradiation. The t.DELTA..sub.off of the sequence is adjusted by
software according to the condition of the patient, and the total
measurement time needed, so, for example, during an operation, this
OFF period can be a mere 0.5 sec in order to rapidly detect any
changes whereas for a patient in intensive care, this OFF period
may be as much as 5 sec since there are no fast changes and, since
the monitoring could be over several days, the total irradiation
should be strictly controlled.
[0218] The PS (6) typically comprises an on-line medical grade
power supply with an insulating transformer as required by Standard
IEC 601-1 for electrical medical equipment.
[0219] The PC (5) typically comprises a Pentium II or higher system
running Windows 95/98/NT or higher. The dedicated Computer and
Power Supply are specified to meet EMC and other requirements for
medical apparatus.
[0220] The probe (2) is typically adapted for sterilisation using
an ETO or any other suitable sterilization technique, and is also
typically made from biocompatible materials. Optionally, the probe
(2) may be designed for once-only use for minimising risk of
cross-infection, for example.
[0221] The dedicated software for the PC (5) is preferably based on
the National Instruments LabView platform. The Doppler module
calculates the blood flow according to well-established algorithms.
The Exposure Tracking module calculates the total and the mean
exposure. It also decides in which of the three possible clock
modes the system will operate. When stable signals are detected for
all measured parameters, the system will switch to State III. In
that mode the tissue receives extremely low exposure. Only the
three parameters of the second set are monitored i.e. NADH
fluorescence, blood oxygenation state, and blood volume via
reflection. The blood flow rate is not actively monitored. If a
change is detected in the value for any one of the measured
parameters, this module switches the system to State II where all
four parameters are actively measured. When calibration is
initiated the system is switched to State I where all four
parameters are measured at high sampling rate.
[0222] The system or apparatus (100) may be operated as follows: At
the beginning of the measurements the user places the probe (2) on
the tissue (25) and activates the system via a terminal of the
computer. This automatically initiates a calibration sequence that
lasts about 1 sec. During the calibration sequence the gain of the
detectors are established and fixed. During calibration sequence,
the clock generates pulses according to state I.
[0223] At the end of the calibration, the computer switches the
clock to state II.
[0224] When switched to state II the OFF period is set to 0.4 sec
so that the system measures all parameters at the rate of 2 data
points per second. If after 10 readings, (i.e. 5 sec) there is no
substantial change in any of the parameters, the OFF period
t'.sub.off is gradually increased to a maximum of 5 sec. If a
steady state is attained, the clock is switched to state III. In
state III ten 10 usec pulses are generated according to state I.
Although this low number of pulses is insufficient for
laser-Doppler measurement, it is sufficient for reflection,
fluorescence and oxy-deoxy measurements. The pulse packets of state
III are initiated every 0.5 sec to 6 sec depending on the
monitoring mode, or until a physiologically significant change,
such as, say, a 2% change in the value of any of the three
parameters monitored. This change being measured relative to the
value of the parameters as measured in the last state II event.
After leaving state III, the system switches to state II with an
OFF period of 0.4 sec.
[0225] In routine clinical use the system is preferably used in
states II and III, with the mean irradiation being typically less
than 0.5 mW/cm.sup.2.
[0226] The Adaptive Duty Cycle enables the reduction of the tissue
irradiation to a value significantly below the maximum limit
imposed by the various standards, whilst still producing a high
signal to noise ratio.
[0227] This is illustrated by the following calculation brought by
way of example:
[0228] The light intensity emitted from the distal end of probe (2)
may be in the order of about 1 mW. The limiting aperture of the
probe may be typically about 0.1 cm therefore the area is 0.00785
cm.sup.2, so the irradiation during the ON part of the pulse is 127
mW/cm.sup.2. With the first duty cycle set at 0.04, and the second
duty cycle at 0.1/5=0.02, the mean irradiation will be reduced by
factor 0.04*0.02=0.0008, or in other words 1250 times less than the
peak value. The mean irradiation will be 127/1250=0.1
mW/cm.sup.2.
[0229] If fast changes are detected, the mean irradiation might be
as much as 1 mW/cm.sup.2. Although this is higher, it is still well
within allowed limits. Thus, with irradiation during the ON part of
the pulse being 127 mW/cm.sup.2, the first duty cycle being 0.04,
and the second duty cycle now being 0.1 sec/0.5 sec=0.2, the total
irradiation is then reduced by a factor of 0.04*0.2=0.008. In other
words, the mean irradiation will be 127*0.008=1 mW/cm.sup.2, and
this value of irradiation can be delivered for 30,000 sec, or
nearly nine hours, whilst remaining within the radiation limits of
the relevant standard.
[0230] Thus, tissue may be irradiated with chopped light to provide
important advantages, such as improving the accuracy in the
measurements for all four parameters. Chopping enables the peak
illumination intensity to be increased while holding constant the
average intensity of the excitation. It allows the average
excitation intensity to be reduced to within safe limits with
respect to photo-damage. This can be achieved without significant
loss of reasonable signal to noise levels.
[0231] "Chopped light" may be produced by chopping the excitation
light illumination, and this may be achieved, for example, by an
Acoustic Optic Modulator (AOM), though a fast rotating chopper
wheel or any other chopping device may also serve this purpose.
Similarly, direct modulation of the light source current could be
used to generate the chopping effect.
[0232] In the context of this specification the duty ratio (DR) of
the pulsed excitation is defined as the ratio of the duration of
each pulse to the total cycle time. When the duty ratio is
decreased, the signal to noise ratio is increased by factor
(DR).sup.-1 for a parameter whose measurement is limited by
background noise and by factor of (DR).sup.-1/2 for a parameter
which signal quality is limited by white noise generated in
detection apparatus (Hodby J., J. Physics E: Scientific
Instruments, 3, 229-233, 1970).
[0233] The ambient light interference and the dark current noise
can be compensated for. This may be done by simply measuring the
detector output during the OFF period and subtracting its value
from the value during the ON period as described above. Since the
chopping produces Amplitude Modulation (AM) of the measuring
signals, all the drift fluctuations and 1/f noise are canceled.
[0234] The third and fourth embodiments of the present invention
are directed to further measuring the flavoprotein parameter, in
addition to the other four parameters that are measured in the
first and second embodiments, respectively.
[0235] As illustrated in FIG. 12, in order to elicit a flavoprotein
fluorescent spectrum from a tissue element, the illuminating
radiation must be within the flavoprotein excitation spectrum,
typically at a wavelength of between about 400 nm and 470 nm.
[0236] The third embodiment of the present invention comprises
similar components as previous embodiments, viz LSU (1) probes (2),
DTU (3), EU (4) PC (5), PS (6) as described with respect to the
first and second embodiments, in particular the first embodiment,
mutatis mutandis, with the following exceptions. The DTU (3) of the
fourth embodiment, as shown in FIG. 13, though substantially
similar to the DTU of the first embodiment (FIG. 5), further
comprises the additional feature that the fluorescent radiation
from the tissue is passed through beam splitter (335) filter (331)
and fourth PMT detector (330). Appropriate modification to the
conditioning electronics of the EU (4) and the software running on
the PC (5) as described for the first embodiment is required for
the third embodiment.
[0237] Thus, the third embodiment enables measurement of the
flavoprotein concentration, i.e., Fp tissue viability parameter, in
additional to the four above-mentioned parameters of the first
embodiment. While in the first embodiment of the invention, the
choice of laser light for source (102) as used for LDF monitoring
is typically of a wavelength above 440 nm as described above, in
order to measure the Fp fluorescence by the same light source (102)
its wavelength should be limited, rather, to be within the
excitation spectrum of the Fp and not beyond the same. In the third
embodiment the excitation wavelength of light source (102), which
used for Doppler measurements, is chosen to lie within the
excitation spectrum of Fp (i.e., in the range of about 440 nm to
about 470 nm) and preferably around the Fp absorption peak, which
is in the range of about 440 nm to about 455 nm. As with the first
embodiment, the two sets of tissue viability measurements are made
using the same excitation and two corresponding groups of
collection fibers, and the same layer element of tissue is
monitored for both sets of parameters--the first set including
blood flow and Fp, and the second set including NADH, blood volume
and blood oxygenation state.
[0238] In the third embodiment, a similar light sources (101) and
(102) as in the first embodiment may be used. As in the first
embodiment, the wavelength for light source (101) can be any
wavelength that is inside the absorption spectrum of the NADH
molecule, i.e. from about 315 nm to about 395 nm. To avoid the
Haemodynamic artifact, and to enable measurement of oxy-deoxy
haemoglobin, this wavelength is preferred to be at one of the
haemoglobin oxy-deoxy isosbestic points in this bandwidth, as with
the second embodiment. There is no requirement for laser Doppler
measurements or for the Fp fluorescence measurements to be
performed at one of these isosbestic wavelengths, since both of
these measurements are normalized to the total reflection. The
Doppler signal is normalized according to the Doppler algorithm,
the Fp signal is normalized or corrected by the reflection
measurement similarly to the NADH correction as described above.
Therefore the excitation light source for (102) may be any suitable
laser diode that operates within the aforesaid 400 nm to 470 nm
range.
[0239] Essentially, the Fp measurement is very similar to that of
NADH. The Fp excitation is by monochromatic light at a wavelength
within the Fp absorption spectrum. In the present invention, this
monochromatic light is provided by, and at the wavelength of, the
laser light source (102). The Fp fluorescence is measured by
measurement of fluorescence intensity of the fluorescence emission
at single wavelength, which is within the emission fluorescence
spectrum.
[0240] As with the NADH fluorescence parameter, problem of
haemodynamic artifact is also relevant to Fp measurements, and
compensation for this artifact is similar to that for the NADH
measurements. For the Fp parameter, reflection is measured at the
wavelength of the excitation of the Fp fluorescence. This
wavelength, in the present invention, is also the wavelength of the
Doppler LDF measurement. In the embodiments described herein, the
same detector that measures Doppler LDF also measures the
reflection at the same wavelength since it is the intrinsic Doppler
measurement that consist of measurement of AC signal that is
superimposed on the DC reflection signal. This reflection value is
subtracted from the Fp fluorescence value (in the same manner as in
NADH measurements) in order to get corrected Fp fluorescence
values. This typifies the compensation procedure.
[0241] As with the NADH parameter, it is preferable to measure the
Fp emission (fluorescence) at oxy-deoxy isosbestic points such as
530 nm or 546 nm or 570 nm. Otherwise the fluorescence value will
be influenced by the blood oxygenation.
[0242] Regarding fluorescence excitation for Fp, if only Doppler or
only Fp is measured, and there is no need for the reflection
measurements for evaluation of blood volume, then any excitation
wavelength can be used, and does not need to be restricted to an
isosbestic wavelength. Indeed as far as the Fp measurements are
concerned, the reflection is measured, and used for correcting for
the haemodynamic artifact, but the reflection measurements will not
correctly represent blood volume changes since they will be
influenced by blood oxygenation. However, it is important to
provide a reflection that represents the blood volume, and for this
reflectance must be measured when excitation is at an isosbestic
point. Thus, either the NADH excitation is chosen to be at a
corresponding isosbestic point, or the Fp excitation is chosen to
be at an isosbestic point. At least one of these conditions is thus
required, wherein the second parameter may be monitored by using an
excitation wavelength that is not at an isosbestic wavelength.
While there is generally no intrinsic advantage in either one, the
availability of suitable light sources at the desired illuminating
wavelengths generally decides the issue.
[0243] While it is advantageous to measure both NADH and Fp,
providing either only one or only the other for a tissue element is
also valuable.
[0244] Determination of the blood flow rate and of the second set
of tissue viability parameters--NADH, blood volume and blood
oxygenation state--in the third embodiment is as described for the
first embodiment, mutatis mutandis. Further, and referring to FIG.
13, the signal from photodiode detector (301) is used for the
reflection measurements for blood volume parameter. Thus, the light
from the collecting fiber (202) enters the DTU (3) via optical
connector (206). The collimating lens (306) collimates the light
towards the beam splitter or dichroic mirror (302), which splits
off the excitation light wavelength, which is then channeled
towards a low-noise, fast photodiode detector (301), a condensing
lens (305) being typically used in order to fill the photo-detector
active area. The signal from the photodiode detector is used to
perform the reflection measurements for the blood volume
parameter.
[0245] As with the first embodiment, detectors (308), (310) and
(312) enable monitoring of the NADH, blood oxygenation state,
mutatis mutandis. In the third embodiment, a dichroic beam splitter
(335) is provided to split the light provided by collection fiber
(203). Part of the split light is used for the blood flow rate
measurements, conducted via detector (315), as in the first
embodiment, after passing through filter (313) and a condensing
lens (314). The second part of the split light continues is
subjected to additional filtering by a suitable filter (331),
preferably a 530 nm interference filter (such as a DF20, for
example), and the filtered light is incident on a fourth
photo-multiplying tube (PMT) (330). The precision of all above
mentioned filters are about .+-.5 nm.
[0246] The timing of sample and hold synchronous detection of the
detectors involved in DTU (3) may be configured as follows.
[0247] When the light source (102) of WL1 for Doppler and Fp
excitation is ON the detector (315) and (330) are gated. This
enables measurement of Doppler and Fp fluorescence.
[0248] When the light source (101) of WL2 for NADH excitation is ON
the detectors (301), (308), (310) and (312) are gated. This enables
measurement of NADH fluorescence, reflection at NADH excitation
wavelength and blood oxygenation.
[0249] Thus, the third embodiment may be used in a similar way to
that described for the first embodiment, mutatis mutandis.
[0250] The fourth embodiment of the present invention comprises
similar components as previous embodiments, viz LSU (1) probes (2),
DTU (3), EU (4) PC (5), PS (6) as described with respect to the
first and second embodiments, in particular the second embodiment,
mutatis mutandis, with the following exceptions. The DTU (3) of the
fourth embodiment, as shown in FIG. 14, though substantially
similar to the DTU of the second embodiment (FIG. 8), further
comprises the additional feature that the fluorescent radiation
from the tissue is passed through beam splitter (333) filter (331)
and fourth PMT detector (330). Appropriate modification to the
conditioning electronics of the EU (4) and the software running on
the PC (5) as described for the second embodiment is required for
the fourth embodiment.
[0251] Thus, the fourth embodiment enables monitoring of the
flavoprotein concentration, i.e., Fp tissue viability parameter, in
additional to the four above-mentioned parameters of the second
embodiment. While in the second embodiment of the invention, the
choice of laser light for source (102) as used for LDF monitoring
is limited to the 315-440 nm range as described above, in order to
measure the Fp fluorescence by the same light source (102) its
wavelength should be limited, rather, to be within the excitation
spectrum of the Fp and not within the excitation spectrum of NADH.
In the fourth embodiment the excitation wavelength of light source
(102), which is used for Doppler measurements, is chosen to lie
within the excitation spectrum of Fp (i.e., in the range of about
400 nm to about 470 nm therefore the excitation wavelengths for
light source (102) are limited to the range of about 400 nm to 440
nm. As with the second embodiment, all tissue viability
measurements are made using the same excitation and using the
corresponding collection fibers, and virtually the same volume
element of tissue is monitored for both sets of parameters--the
first set including blood flow and Fp, and the second set including
NADH, blood volume and blood oxygenation state.
[0252] In the fourth embodiment, as in the second embodiment, a
light source (101) such as, for example, filtered spectral lamp,
light emitting diode LED, laser diode or pulsed laser, may be used
for excitation of NADH fluorescence. The light source wavelength
can be any wavelength that is inside the absorption spectrum of the
NADH molecule, in the range of from about 315 nm to about 395 nm.
To avoid the Haemodynamic artifact, and to enable measurement of
oxy-deoxy haemoglobin, this wavelength is preferred to be at one of
the haemoglobin oxy-deoxy isosbestic points in this bandwidth, as
with the second embodiment. There is no requirement for laser
Doppler measurements or for the Fp fluorescence measurements to be
performed at one of these isosbestic wavelengths, since both of
these measurements are normalized to the total reflection. The
Doppler signal is normalized according to the Doppler algorithm,
the Fp signal is normalized or corrected by the reflection
measurement similarly to the NADH correction as described above.
Therefore the excitation light source for (102) may be any low
noise CW laser such as Nichia blue laser diode, for example.
[0253] Determination of the blood flow rate and of the second set
of tissue viability parameters--NADH, blood volume and blood
oxygenation state--in the fourth embodiment is as described for the
second embodiment, mutatis mutandis. Further, and referring to FIG.
14, the signal from photodiode detector (301) is used for LDF blood
flow rate monitoring as well as the reflection measurements for
blood volume parameter for Fp and NADH correction. Thus, the light
from the collecting fiber (223) enters the DTU (3) via optical
connector (267). The collimating lens (306) collimates the light
towards the beam splitter or dichroic mirror (302), which splits
off the excitation light wavelength, which is then channeled
towards a low-noise, fast photodiode detector (301), a condensing
lens (305) being typically used in order to fill the photo-detector
active area. The signal from the photodiode detector is used to
perform both Doppler and Reflection measurements.
[0254] Thus the reflection at both detection wavelengths (WL1) and
(WL2) will, in this embodiment, pass through the same collecting
fiber (223). The separation of these different signals is typically
achieved by time-sharing. The light sources (102) and (101) are
typically working in "chopping mode", and therefore each source is
ON only for a short time period as is described hereinafter. Each
light source has a low duty cycle, therefore there is plenty of
time to turn ON one light source while the other one in still in
the OFF period. Each one of the corresponding detectors (107) and
(109) is synchronously sampled at the correct timing in order to
get intensity information regarding the corresponding excitation
source. The detector (301) is sampled twice, once for the
measurement of reflection at WL1 and corresponding laser Doppler
measurements and a second time, for reflection measurement at
WL2.
[0255] In case that the Doppler excitation wavelength WL1 will be
lower then 420 nm, both reflections at WL1 and WL2 have
substantially lower wavelengths than the NADH fluorescence and
therefore the same dichroic beam splitter (302) can be used for
separating these two relatively strong reflections from the weaker
fluorescence signal. On the other hand, if the Doppler excitation
wavelength WL1 will have a wavelength that higher then about 420
nm, the dichroic beam splitter (302) cannot be used since the NADH
fluorescence light will not pass through it. Therefore in such a
case the dichroic beam splitter (302) should be replaced by simple
beam splitter of for example 1:10 ratio in order to split to the
detector (301) a small part of the desired reflection intensity at
the excitation wavelengths. This also enables transmission of the
similar wavelengths of the NADH and Fp fluorescence to all another
detectors namely (308), (310), (312) and (330). It should be
emphasized that the strong intensity of the reflection at WL1 will
not interfere the NADH fluorescence measurements since these
signals are uncorrelated in the time domain. Similarly, there is no
interference with the Fp fluorescence measurements since an
interference filter (314) is used, as described below.
[0256] Light of wavelengths higher than the excitation wavelength
WL1 or WL2 passes through the dichroic mirror or beam splitter
(302) and is incident on a second dichroic beam-splitter (303),
which is selected to reflect wavelengths lower then about 440 nm
and to transmit all higher wavelengths. The reflected light beam is
passed through a suitable filter (307), preferably a 435 nm (10DF)
filter, and is then fed into a first photo-multiplying tube (PMT)
(308). The light transmitted through the second dichroic
beam-splitter (303) is subjected to additional splitting by a third
dichroic beam-splitter (304) that reflects wavelengths lower then
460 nm, but is transparent to higher wavelengths. The reflected
light from the third dichroic beam splitter (304) is filtered by a
suitable filter (309), preferably a 455 nm (10DF) filter, and is
then incident on a second photo-multiplying tube (PMT) (310). This
wavelength is close to an oxy-deoxy isosbestic point, so the
fluorescence intensity as measured by this PMT (310) correlates
directly with the NADH fluorescence. The light that passes through
the third dichroic beam-splitter (304) is subjected to additional
splitting by a fourth dichroic beam-splitter (333) that reflects
wavelengths lower then 485 nm, but is transparent to higher
wavelengths. The reflected light pass additional filtering by a
suitable filter (311), preferably a 475 nm interference filter
(DF10), and the filtered light is incident on a third
photo-multiplying tube (PMT) (312). The light that pass the
dichroic beam-splitter (333) comprises wavelengths higher then 485
nm and comprises the Fp fluorescence emission. This light is
subjected to additional filtering by a suitable filter (331),
preferably a 530 nm interference filter (such as a DF20, for
example), and the filtered light is incident on a fourth
photo-multiplying tube (PMT) (330). The precision of all above
mentioned filters are about .+-.5 nm.
[0257] The timing of sample and hold synchronous detection of the
detectors involved in DTU (3) may be configured as follows.
[0258] When the light source (102) of WL1 for Doppler and Fp
excitation is ON the detector (301) and (330) are gated. This
enables measurement of Doppler, Reflection at Fp excitation
wavelength and Fp fluorescence.
[0259] When the light source (101) of WL2 for NADH excitation is ON
the detectors (301), (308), (310) and (312) are gated. This enables
measurement of NADH fluorescence, Reflection at NADH excitation
wavelength and blood oxygenation.
[0260] Thus, the fourth embodiment may be used in a similar way to
that described for the second embodiment, mutatis mutandis.
[0261] In some clinical procedures it is desirable to monitor the
blood parameters for the assessment of organ tissue vitality in
different regions of the body. In these situations, a multiple
probe system is desirable. By way of example, a fifth embodiment of
the present invention, consisting of a multi-probe system is shown
in FIG. 11. This embodiment as illustrated, uses a plurality of
probes, each probe being substantially the same as those described
in the second embodiment. It will be appreciated however, that a
plurality of the probe according to the first embodiment could be
used, or, alternatively one or more probes according to each one of
the first and second embodiments, mutatis mutandis. Clearly, any
given probe comprised in the system may be adapted to monitor the
same or different parameter to any one of the other probes
therein.
[0262] The chopping feature, which provide advantages in minimising
exposure of the probed tissue to dangerous illumination levels,
also facilitates a diversion of the irradiation light to any one of
a plurality of probes, and subsequent detection of the return
signals therefrom, by effectively time-sharing the detection unit
(DTU) between the probes. In other words, the multiprobe detection
system essentially multiplexes the signals obtained from each of
the plurality of probes, situated in different parts of the tissue
or organs.
[0263] The fifth embodiment of the present invention comprises
similar components as previous embodiments, viz LSU (1) probes (2),
DTU (3), EU (4) PC (5), PS (6) as described with respect to the
first and second embodiments, mutatis mutandis, with the following
exceptions. The LSU (1) of the fifth embodiment, as shown in FIG.
11, though substantially similar to the LSU of the second
embodiment (FIG. 8), further comprises the additional feature that
the excitation light is passed through an acousto-optic deflector
(AOD) (140) before being coupled and deflected to a plurality of
excitation fibers by a corresponding plurality of lenses (104),
each one being mounted on one of a plurality of adapters (105). The
LSU of the fifth embodiment is thus connected by said plurality of
excitation fibers to a corresponding plurality of fiber optic
probes, each probe being coupled via an optical connector (205). In
the third embodiment, the collecting fiber (223) from each probe is
thus connected to the DTU (3) by an optical connector (267) that is
essentially similar to that used in the second embodiment. The
radiation received via the optical connectors (267) is coupled to a
common optical coupler (340) via corresponding optical fibers
(341). The optical coupler (340) and plurality of optical
connectors (267) of this fifth embodiment replaces the single
optical connector (267) described in the second embodiment. From
the optical coupler (340) the light passes through a collimating
lens (306) and on to the detection equipment of the DTU (3) as
described for the second embodiment mutatis mutandis. Appropriate
modification to the conditioning electronics of the EU (4) and the
software running on the PC (5) as described for the second
embodiment is required for the fifth embodiment.
[0264] For embodiments of the multi-probe system comprising a
plurality of probes according to the first embodiment, a
corresponding plurality of connectors (206) and of connectors (207)
(which are in turn appropriately connected to corresponding ones of
a plurality of fibers (202) and (203)) are connected to one of two
optical couplers. The optical coupler that couples the (206)
connectors may then be connected by the existing arrangement of
filters and beamsplitters to the PMTs (301), (308), (310) and (312)
as shown in FIG. 5, and the optical coupler that couples the (207)
connectors may then be connected by the existing arrangement of
optical components to the PMT (315).
[0265] It should also be noted that multi-tissue element monitoring
could also be accomplished by a plurality of probes, each one
having a dedicated light source and associated optical components
(LSU) and detection system (DTU), with each probe unit being
controlled by the same PC and EU units, and being powered using the
same PS.
[0266] The fifth embodiment may be operated in a variety of modes
as required by the clinical situation and diagnostic needs to which
it is applied. Two particular modes of monitoring for which such
multiple probe systems can be usefully applied, are described:
[0267] In the first mode, the mean signal intensities from the
multiplicity of probes is calculated and displayed. This results in
the parameters detected representing an average response of the
multiplicity of tissue volumes probed, and will generally, better
reflect the state of the organ layer as a whole. This mode of
monitoring could be useful in transplantation surgery when better
monitoring of the viability of donated organs are needed.
[0268] In the second mode, by applying one or several of the
plurality of probes to each of several locations on the same organ
or several different locations of different organs, the
quasi-continuous monitoring of these organs over the same time
period can be achieved by multiplexing the signals from the
individual probes, with the parametric response of each organ being
separately monitored and displayed.
[0269] The electronics and the software for the first mode will be
substantially similar to that described with respect to the second
embodiment. The main difference being that the chopping sequences
used, and the sampling rate per probe, are engineered and optimized
depending on number of probes, patient condition, and tissue type
under observation.
[0270] The t.sub.on period per probe remains the same as for the
single probe embodiment, but in the OFF time for aforementioned
probe, additional probes are excited and measured. Accordingly, the
timing of the AOD (140) is correlated with the sequence shown in
FIG. 7(a) so that each subsequent pulse is delivered to the
subsequent probe. After appropriate smoothing, the output signals
from each detector are used to generate a value for the desired
blood viability parameter, corresponding of the mean value for the
plurality of monitored tissue volumes, and thus more representative
of the viability of the organ as a whole.
[0271] The second measurement mode of the fifth embodiment requires
the same chopping sequence as that required by the first mode. The
(S/H) circuits and the accompanying data acquisition system in the
EU (4) are somewhat different however. For tracking the various
monitored parameters for each probe separately, using the same
system, the (S/H) circuits shown in FIGS. 9(a) and 9(b) may be
advantageously performed by fast multi input analog to digital
(A/D) converters. This requires that the fast A/D system (410)
receives the detector signals directly without the signal
conditioning electronics (402) that were used in the second
embodiment. The fast A/D converter (401) will digitize the detector
signal for at each time period and the sampled values for each
probe and each measured parameter will be stored in the temporary
memory of the PC (5). Since the whole information, that is all
signals for each probe is available in the computer, the signal
from each probes can be processed separately, allowing the vitality
parameters of each monitored tissue volume, corresponding to
different organs to be monitored and displayed on the screen.
[0272] A sixth embodiment of the present invention (not
illustrated) comprises all the elements of the fifth embodiment as
described herein, mutatis mutandis, wherein the DTU is modified in
a similar manner to that described for the third embodiment,
mutatis mutandis, enabling the multi-probe system to monitor Fp
fluorescence as well as the blood flow rate and the three
parameters of the second set of parameters, at a plurality of
locations.
[0273] In each one of the first through sixth embodiments, the
signals provided by the various detectors are fed to the PC (5) and
converted to suitable corresponding values of blood flow rate, NADH
fluorescence, blood oxygenation state, blood volume and Fp
fluorescence, via suitable algorithms, correlations, tables and so
on, in a manner known in the art.
[0274] While specific embodiments of the invention have been
described for the purpose of illustration, it will be understood
that the invention may be carried out in practice by skilled
persons with many modifications, variations and adaptations,
without departing from its spirit or exceeding the scope of the
claims.
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