Patent Application
20050043606
The present invention relates to apparatuses and methods for enabling simultaneous or individual monitoring of a plurality of tissue vitality parameters, particularly in-vivo, with respect to an identical tissue element, such parameters including blood flow rate, Mitochondrial Redox State via NADH or flavoprotein concentration, blood volume and blood oxygenation state. In particular, the present invention relates to such apparatuses and methods based on a single illuminating laser radiation.
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.
The integrated system of energy supply and demand can be understood by considering the various components thereof.
O2 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 O2 to the tissue for the purpose of energy formation therein.
Energy production and demand: In an inner compartment of the cells, called the mitochondria, the glucose and O2 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 O2 availability.
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.
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.
There is a 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 O2 demand will be compensated by a corresponding change in the blood flow to the tissue. By this mechanism, the O2 supply remains constant if there is no change in the O2 consumption. Any change in the abundance of O2 in the tissue, in other words a change in energy state, will be reflected by the NADH and Fp level.
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.
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
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.
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; Stern, M. D. Nature 254, 56-58, 1975).
B—Mitochondrial Redox State or the NADH Level
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
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
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
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 HbO2 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 nm to 1000 nm in near-infrared region (Pologe, 1987).
Isosbestic point as referred to herein is a wavelength at which the intensity of absorption of oxyhaemoglobin HbO2 is the same as the intensity of absorption of deoxy-haemoglobin Hb; such isosbestic points are indicated as IPA and IPB in
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
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
The five tissue viability parameters described above represent various important biochemical and physiological activities of body tissues. Monitoring these parameters 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, the better and more accurate an understanding of the functional state of the tissue that may be obtained.
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.
The prior art teaches a wide variety of apparatuses/devices which monitor various parameters reflecting the viability of the tissue, for example, in U.S. Pat. No. 4,703,758 and U.S. Pat. No. 4,945,896.
A particular 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 et al, 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. Additionally since the same optical fiber is utilised for transmission of excitation light and for transmission of the collected light the excitation and the collection point is the same one. This imposes relatively low penetration depth as can be learned from the paper of Jakobsson and Nilsson (Jakobsson and Nilsson, 1991). More importantly, though, two different wavelengths are used for illuminating the organ.
None of these prior art documents suggest monitoring Fp level, with or without any of the other parameters, and less so in an integrated apparatus.
In two earlier patents which have a common inventor with the present invention, U.S. Pat. No. 5,916,171 and U.S. Pat. No. 5,685,313, the contents of which are incorporated herein in their entirety, a device is described that is directed to 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.
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:
Israel Patent Application No. 138683 filed by Applicants, the contents of which application are incorporated herein in their entirety, further addresses some of these problems by using two separate illumination radiation sources, one for determination of blood flow rate, and the other for determination of at least one tissue viability parameter such as NADH, blood volume and blood oxygenation state. By separating the light sources, the problem of having a single source capable of satisfactorily enabling the determination of blood flow rate as well as the other three tissue vitality parameters is avoided.
While U.S. Pat. No. 5,916,171 and U.S. Pat. No. 5,685,313 ostensibly teach a single illumination radiation for laser Doppler flowmetry and NADH monitoring in the substantially identical tissue volume, on closer scrutiny it is not at all obvious for a man of the art to do so at illuminating wavelengths within the range of between about 370 nm to about 400 nm. There is also absolutely no suggestion whatsoever that the illuminating wavelength should be within the Fp excitation range, i.e., about 400 nm to about 470 nm, and in fact these references teach away from this, as NADH cannot be monitored at all at the higher wavelengths. These patents exemplify a radiation source generating electromagnetic radiation at a wavelength at 366 nm or 325 nm. The reason is twofold. On the one hand, and as illustrated in
Furthermore, at the time when these US patents were filed, and indeed until very recently, there were no suitable lasers available capable of generating electromagnetic energy in the wavelength range 370 nm to 400 nm, or indeed in the range 400 nm to about 470 nm with sufficiently low Relative Intensity Noise factor (RIN). The two lasers that were then available were a 325 nm Helium Cadmium laser and a 355 nm “3rd Harmonic of Nd-Yag” laser. The He-Cd laser is a large gas laser, having relatively large power consumption, being generally unsuitable for the applications where small size and power consumption are important considerations. Furthermore, this laser generates a great deal of optical noise, having a Relative Intensity Noise factor (RIN) of about 1% to 2%. There is also a small but significant spectral spread at the operating wavelength, typically comprising about eleven discrete wavebands bundled thereabout, further diminishing the efficiency of operation. While this laser enables single illumination radiation for laser Doppler flowmetry and NADH monitoring, the sensitivity is very low, and operation of such a laser raises many safety issues, since operating at a wavelength of 325 nm carries potential risk of DNA damage to the tissue. The Nd-Yag laser provides radiation at a higher wavelength of 355 nm. However, it generates a great deal of optical noise when operating in continuous wave (CW), resulting in poor quality measurements. While this laser generates less noise in pulse mode, no useful measurements may be made for Doppler Flowmetry using pulsed lasers, since it is very difficult to ensure uniformity between the pulses generated.
Furthermore, there would be little motivation for a man of the art to use a laser at the illuminating wavelength range of 370 nm-400 nm, or indeed in the range of about 400 nm to about 470 nm for laser Doppler Flowmetry, even if one existed, for a number of reasons. Laser Doppler Flowmetry as applied to a tissue volume substantially comprising capillarial blood flow is substantially different from the laminar flow Laser Doppler flowmetry methods used with large tissues having veins or arteries.
In the laminar flow laser Doppler flowmetry the laser ray is split and then converged again in order to produce interference fringes in the path of fluid flow. As fluid particles pass through these fringes they produce an alternating light signal, which can be analysed to provide a measure of particle velocity and fluid flow rate. In such a method, severe constraints are imposed on the laser spectral bandwidth that is acceptable for the task. Broad laser bandwidth causes blurring of the interference fringes, thereby decreasing the quality of the measurements.
In contrast, the laser Doppler blood flowmetry method used in the present invention relates to the measurement of blood perfusion through tiny capillaries. The flow is not laminar but random in velocity and direction. The illumination of the tissue is done through a single optical fiber and the collection performed by at least one collection fiber. The most intuitive way to understand the blood laser Doppler measurements is by specklemetery. The laser light is shone onto a tissue containing the random network of tiny capillaries through which the Red Blood Cells (RBC) are flowing. Both the direction and velocity of the RBC flow are random. As the laser radiation penetrates the tissue bulk, some small part of the excitation light is reflected after numerous scatterings inside the tissue. This reflected light produces a random interference pattern referred to as a speckle pattern on the tissue surface. The RBC movements inside the capillaries cause random changes in this speckle pattern. The collection optical fiber, which is placed near to the excitation fiber, delivers to an appropriate detector the changes in speckle intensity due to the blood flow, and an electrical signal corresponding to the changing intensity is generated. The Doppler signal is thus represented by small fluctuations of the total light intensity, i.e. the AC ripple in relation to the DC total intensity, which represents the reflection signal that is correlated to blood volume. From analysis of the detector power spectrum the blood flow parameter can be deduced according to the algorithm described by Bonner and Nossal [Bonner and Nossal, 1990]. Thus, the flowmetry method is based on measuring the perturbations in the intensity of the electromagnetic radiation received from the tissue in relation to the mean intensity of the radiation, in other words, the ratio between the AC signal to the DC signal of the radiation received by the monitoring probe. With lasers operating at large wavelengths, for example red lasers, the penetration of the laser into the tissue is relatively large, and therefore the AC signal is proportionally larger, since more capillaries interact with the illuminating laser radiation. In other words, as the laser wavelength increases, or as the corresponding laser bandwidth narrows, the speckle pattern becomes more defined and the intensity fluctuations became higher. It follows that with lasers operating at lower wavelengths the opposite is true, and at lower illuminating radiation wavelengths, including the range 370 nm to 470, and in particular in the range 370 nm to 400 nm, the penetration of the laser radiation into the tissue would be less, providing a lower AC signal relative to the mean DC signal, which drastically lowers sensitivity.
However, there are further problems associated with using an illuminating radiation wavelength in the range 370 nm to 470 nm that teach away from using such a laser wavelength for Doppler flowmetry:—
All the above problems individually, and more so in combination, teach away from considering the use of a laser in the 370 nm to 470 nm range for measuring blood flow rate together with blood NADH or with Fp, since the combined inefficiencies reduce the possibility of providing meaningful flowmetry results. However, even if the problem of decreased sensitivity is resolved, there are yet another two problems that dissuade the use of such lasers in the present context.
Firstly, it is by far most convenient to provide an illuminating radiation at an isobestic wavelength to mimimise the effect of the Haemodynamic Artifact on measurements. For NADH fluorescence, an isosbestic wavelength in the range 370 nm to 400 nm exists at a wavelength of about 390 nm. Similarly, an isosbestic point exists in the Fp excitation spectrum between about 400 nm and 470 nm wavelengths, at about 455 nm. Until recently, no laser diodes capable of operating at these isosbestic wavelengths were available. Recently, though, a range of laser diodes by Nichia Chemical Industries Ltd., Anan, Japan capable of operating in continuous mode within the range 385 nm to 440 nm has become available, including the violet laser diode such as the NLHV500. However, even these lasers are still subject to the above problems.
Secondly, even the existence of such a laser in and of itself does not render its use obvious in the context of laser Doppler flowmetry and NADH monitoring. For example, if such a laser were to be used in combination with the device disclosed in U.S. Pat. No. 5,916,171 and U.S. Pat. No. 5,685,313, the device would still be incapable of providing meaningful Doppler flowmetry measurements. The reason for this is as follows. While lasers generate electromagnetic radiation nominally at a single wavelength, in practice, this is not achieved, and two or more discrete narrow wavebands are generated. This plurality of wavebands are herein referred to as longitudinal multi-mode radiation, and is different conceptually and practically to the “transverse multi-mode” radiation commonly encountered with many lasers. The phenomenon of longitudinal multi-mode radiation generation occurs as a result of more than one stable wavelength being generated by the laser in general and also the laser diode itself due to the physical constraints imposed by the laser cavity. While very close in wavelength, these wavebands are nonetheless discrete. For example, in the NHLV500 diode which operates nominally at 405 nm wavelength has several longitudinal modes separated by about 0.05 nm or 95 GHz, as illustrated in
A further problem associated with longitudinal multi-mode radiation is the phenomenon of mode competition, in which the actual wavelength of the illuminating radiation randomly switches from one of the discrete modes or wavebands to another, which dramatically increases the level of RIN. Thus, for good flowmetry measurements a very low noise laser with preferably a single longitudinal mode and as narrow as possible bandwidth associated therewith is needed, and thus there is no motivation for a man of the art to combine off-the-shelf lasers of wavelength between about 370 nm to about 400 nm with the device of U.S. Pat. No. 5,916,171 or U.S. Pat. No. 5,685,313. As mentioned above, there is even less motivation for a man of the art to combine lasers of wavelength 400 nm to 470 nm with the device of these patents. Finally, even if such a laser diode were to be configured to generate radiation in nearly longitudinal single mode by critical choice of current and temperature, which is in itself far from self-evident, factors such as temperature and current drifts may cause regression to multi-mode operation. Furthermore such single mode operation still has intrinsically a very broad bandwidth in order of 400 Mhz, which of itself is still problematic for laser Doppler flowmetry. Thus, all the above factors would tend to teach away from employing such a laser configuration for multiparamater monitoring In fact, even a grating-stabilized laser diode that nominally operates in a single longitudinal mode exhibits intensity instability. Generally, this happens when, during operation, the operating parameters of the laser system change due to aging thermal drift etc. Due to such changes the laser system gradually drifts to highly unstable multimode operation accompanied with very high optical noise caused by mode competition.
In any case, 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”.
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.
Both the above standards permit a maximum of 1 mW/cm2 irradiance for 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. These values correspond to laser wavelengths in the range of between about 315 nm to about 400 nm. At higher wavelengths than about 400 nm, the (MPE) values are higher.
It is an aim of the present invention to overcome the above deficiencies in the prior art.
Particularly, it is an aim of the present invention to provide an 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 set: Mitrochondrial Redox State via 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 NADH fluorescence; or alternatively blood flow rate and at least one of and preferably all of Mitrochondrial Redox State via flavoprotein concentration by fluorescence, optionally including total blood volume by reflectometry and/or blood haemoglobin oxygenation by fluorescence, based on flavoprotein fluorescence; for the same body tissue, in substantially the same tissue element. These parameters, which represent different biochemical and physiological activities of the tissue, are used to assess the tissue vitality in said tissue element.
It is another aim of the present invention to provide a device or apparatus capable of monitoring blood flow rate and at least one other tissue vitality parameter including NADH level, in a substantially identical tissue volume, using a single illuminating radiation having a wavelength in the range from about 370 nm to about 400 nm, and more particularly at about 390 nm.
It is another aim of the present invention to provide a device or apparatus capable of monitoring blood flow rate, NADH level, blood volume and blood oxygenation state in a substantially identical tissue volume, using a single illuminating radiation having a wavelength in the range from about 370 nm to about 400 nm, and more particularly at about 390 nm.
It is another aim of the present invention to provide a device of apparatus capable of monitoring blood flow rate and at least one other tissue vitality parameter including flavoprotein level, in a substantially identical tissue volume, using a single illuminating radiation having a wavelength in the range from about 400 nm to about 470 nm, and more particularly at about 440 nm or 455 nm.
It is another aim of the present invention to provide a device or apparatus capable of monitoring blood flow rate, flavoprotein level, blood volume and blood oxygenation state in a substantially identical tissue volume, using a single illuminating radiation having a wavelength in the range from about 400 nm to about 470 nm, and more particularly at about 440 nm or 455 nm.
It is another aim of the present invention to provide such a device or apparatus that provides a single illuminating radiation at longitudinal single mode having a single bandwidth.
It is another aim of the present invention to provide such a device or apparatus that provides a single illuminating radiation at longitudinal multi-mode comprising three or less bandwidths in non-competing modes.
It is another aim of the present invention to provide such a device or apparatus that provides for the stabilisation of such a longitudinal single mode or such a longitudinal non-competing multi-mode.
It is another aim of the present invention to provide such a device or apparatus that conforms to the relevant laser safety standards.
It is another aim of the present invention to provide such a device or apparatus that is of a convenient size, weight and power consumption such as to enable the same to be portable and/or installable within regular operating theaters.
These and other aims are achieved in the present invention by providing a device or apparatus capable of generating an illuminating laser radiation characterised in one embodiment in having a nominal wavelength in the range of about 370 nm to about 400 nm, and in particular about 390 nm, adapted for monitoring NADH level, blood volume and blood oxygenation state as well as blood flow rate, and in another embodiment in having a nominal wavelength in the range of about 400 nm to about 470 nm, and in particular about 440 nm and more particularly 455 nm, adapted for monitoring flavoprotein level, blood volume and blood oxygenation state as well as blood flow rate. The invention is further characterised in providing means to filter out most of unwanted bandwidths generated naturally by the laser, and thus provide a longitudinal single mode illuminating radiation to the tissue, or a typically two or three non-competing multi-mode illuminating radiation to the tissue, such as to enable blood flow rate and at least one of, and preferably all of the set comprising NADH level, blood volume and blood oxygenation state or at least one of and preferably all of the set comprising flavoprotein level, blood volume and blood oxygenation state to be determined for the substantially identical tissue volume. The invention further provides for the stabilisation of the illuminating radiation wavelength, such as to prevent regression to a competing multi-mode situation. Typically one or a bundle of optical fibers are provided for illuminating the tissue at nominally a single location thereon, together with one or bundle of detection fibers. The detection fibers are all substantially equi-distant from the illuminating fibers, thereby ensuring that the substantially identical tissue volume is the subject of all the measurements.
Other purposes and advantages of the invention will appear as the description proceeds.
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 a substantially identical tissue element, the apparatus comprising:—
In preferred embodiments, the laser radiation is generated in stable single longitudinal mode, wherein said nominal wavelength comprises a single waveband element, and the waveband element typically comprises a bandwidth of about 4 MHz.
In other embodiments, the said laser radiation is generated in two or three stable longitudinal non-competing modes, wherein said nominal wavelength comprises two or three, respectively, discrete waveband elements. The waveband elements each typically comprise a bandwidth of about 4 MHz.
The illumination location is provided by at least one excitation optical fiber having a free end capable of being brought into registry with said tissue element. The radiation receiving means comprises at least one suitable receiving optical fiber having a free end capable of being brought into registry with said tissue element. The at least one excitation optical fiber and said at least one receiving optical fiber are preferably housed in a suitable probe head, wherein said free end of said at least one excitation fiber and said free end of said at least one receiving fiber are comprised on a contact face of said probe. Preferably, the at least one excitation fiber comprises a suitable first connector at an end thereof opposed to said free end thereof, said first connector capable of selectively coupling and decoupling said excitation fiber from the rest of the said apparatus. Similarly, the at least one collection fiber also preferably comprises a suitable second connector at an end thereof opposed to said free end thereof, said second connector capable of selectively coupling and decoupling said collection fiber from the rest of the said apparatus. The probe may be disposable and/or sterilisable.
In preferred embodiments, the illumination means comprises a suitable external cavity laser diode system, typically based on a suitable violet laser diode having an operating wavelength in the range of between about 370 nm and about 470 nm. The external cavity laser diode system may be configured according to the Littrow design or according to the Metcalf-Littman design. Preferably, the external cavity diode laser system comprises a laser stabilisation control system for ensuring stable single mode operation of the said external cavity laser diode system. Typically, the laser stabilisation control system is adapted for monitoring the laser intensity of the said external cavity laser diode system at a predetermined input current to said external cavity laser diode system and providing an electrical signal representative of said intensity, for varying the said input current within a predetermined range to provide corresponding electrical signals correlated to the resulting laser intensities generated, for identifying the corresponding electrical signal providing minimum RIN noise levels, and for adjusting the said input current such as to provide and maintain said electrical signal providing minimum RIN noise levels.
In preferred embodiments, the blood flow rate tissue viability parameter is provided by applying a laser Doppler flowmetry technique to said radiation received by said radiation receiving means, and the apparatus further comprises first detection means for detecting said received radiation received by said radiation receiving means.
Preferably, the illumination means is adapted to provide said illuminating radiation in pulses of predetermined duration and intensity by correspondingly chopping the illuminating radiation generated by said illuminating means. The apparatus 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 greater than the time interval between consecutive pulses within a package, and the time period and/or the number of pulses within each package may be controllably variable.
Preferably, the control means is operatively connected to said first detection means.
In the first preferred embodiment, the nominal wavelength is at wavelength within the NADH excitation spectrum, preferably at a suitable isobestic wavelength within the NADH excitation spectrum, and more preferably is about 390 nm±5 nm.
In this embodiment, one of the second tissue viability parameters is NADH concentration, wherein said radiation received by said radiation receiving means comprises an NADH fluorescence emitted by the tissue in response to illumination thereof by said illuminating radiation, said at least one second tissue viability parameter being provided by the intensity of said NADH fluorescence. The apparatus thus may comprise suitable second detection means for detecting said received radiation received by said radiation receiving means. Further, the control means may be operatively connected to said second detection means, and the control means may be selectively responsive to previously detected signals corresponding to the detection of said received radiation detected by means of said second detection means of a prior monitoring cycle.
In the second preferred embodiment, the nominal wavelength is at wavelength within the Fp excitation spectrum, preferably at a suitable isobestic wavelength within the Fp excitation spectrum, and more preferably is about 440 nm±5 nm.
In this embodiment, one of the second tissue viability parameters is Fp concentration, wherein said radiation received by said radiation receiving means comprises an Fp fluorescence emitted by the tissue in response to illumination thereof by said illuminating radiation, said at least one second tissue viability parameter being provided by the intensity of said Fp fluorescence. The apparatus thus may comprise suitable second detection means for detecting said received radiation received by said radiation receiving means. Further, the control means may be operatively connected to said second detection means, and the control means may be selectively responsive to previously detected signals corresponding to the detection of said received radiation detected by means of said second detection means of a prior monitoring cycle.
The first and second preferred embodiments may, in addition or in lieu of NADH or Fp concentration, respectively, include monitoring of another second tissue viability parameter, namely blood volume within said tissue element, and in this case said corresponding radiation received by said radiation receiving means comprises a reflection from the tissue element in response to illumination thereof by said illuminating radiation, the said at least one second tissue viability parameter being provided by the intensity of said reflection. Thus, the apparatus further comprises third detection means for detecting said received radiation received by said radiation receiving means. The control means may be operatively connected to said third detection means, and the control means may be selectively responsive to previously detected signals corresponding to the detection of said received radiation detected by means of said third detection means of a prior monitoring cycle.
Similarly, another second tissue viability parameter may be monitored, in addition to or in lieu of either to the other parameters, i.e., blood oxygenation ratio within said tissue element, and said corresponding radiation received by said radiation receiving means is a fluorescence emitted by the tissue in response to illumination thereof by said illuminating radiation, said at least one second tissue viability parameter being provided by the intensity of said fluorescence at least at two fluorescent emission wavelengths.
For the first preferred embodiment, one of said at least two fluorescent wavelengths is chosen to lie at an isosbestic point of the NADH fluorescence emission spectrum. Alternatively, and preferably, 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 of the NADH fluorescence emission spectrum. Preferably, blood oxygenation ratio parameter is provided by normalising said fluorescent intensities at said two wavelengths with respect to the fluorescent emission intensity at said isosbestic point of said NADH fluorescence emission spectrum. The wavelength corresponding to one such isosbestic point is about 455 mm±5 nm. Preferably, the apparatus further comprises fourth detection means for detecting said received radiation received by said radiation receiving means. The control means is operatively connected to said fourth detection means, and the control means is selectively responsive to previously detected signals corresponding to the detection of said received radiation detected by means of said fourth detection means of a prior monitoring cycle. For the second preferred embodiment, one of said at least two fluorescent wavelengths is chosen to lie at an isosbestic point of the Fp fluorescence emission spectrum. Alternatively, and preferably, 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 of the Fp fluorescence emission spectrum. Preferably, blood oxygenation ratio parameter is provided by normalising said fluorescent intensities at said two wavelengths with respect to the fluorescent emission intensity at said isosbestic point of said Fp fluorescence emission spectrum. The wavelength corresponding to one such isosbestic point is about 530 nm±5 nm. Preferably, the apparatus further comprises fourth detection means for detecting said received radiation received by said radiation receiving means. The control means is operatively connected to said fourth detection means, and the control means is selectively responsive to previously detected signals corresponding to the detection of said received radiation detected by means of said fourth detection means of a prior monitoring cycle. The present invention is also directed to a system for selectively monitoring at least two tissue viability parameter at a plurality of tissue elements. The system comprises a plurality of monitoring probes, each said probe being substantially similar to that comprised in the apparatus according to the first or second embodiments.
In the system, at least two said probes are adapted for monitoring said tissue viability parameters of tissue elements within the same organ. Alternatively, at least two said probes are adapted for monitoring said tissue viability parameters of tissue elements within different organs, wherein different organs are different organs within the same organism, or wherein different organs are different organs within different organisms. The different organs may comprise include donor organs.
The illuminating radiation for each said probe may be provided by a common suitable light source, which is a laser light source as in the first and second embodiments.
The laser light source may be adapted to provide said first illuminating radiation of said first wavelength in second pulses of predetermined duration and intensity. The system may further comprise suitable control means for controlling the frequency of pulsing of said second pulses. The control means may be further adapted to provide said second pulses in packages of pulses, each package comprising at least one second pulse and separated from a preceding or following package by a predetermined time period the predetermined time period may be greater than the time interval between consecutive pulses within a package. Preferably, the time period may be controllably variable, and the number of second pulses within each package may be controllably variable. The control means may be adapted for selectively directing discrete said second pulses to any one of said probes.
The present invention will be more clearly understood from the detailed description of the preferred embodiments and from the attached drawings in which:
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.
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.
The present invention is directed to an apparatus for simultaneously monitoring at least two tissue viability parameters from a substantially identical volume of tissue element. In particular, one of these parameters is the blood flow rate corresponding to the tissue volume, and the other tissue viability parameter includes at least one of, and preferably more than one of, and most preferably all of, the set of parameters comprising at least NADH concentration or flavoprotein concentration, blood volume and blood oxygenation state corresponding to the tissue volume. A laser radiation source provides a single illumination radiation at a particular excitation wavelength that is used for monitoring these parameters, as will be described in detail hereinbelow. In particular, means are provided to ensure that the excitation radiation is at a single mode bandwidth, or at least at three or less stable bandwidths, i.e., that are not in competition with each other. Thus the blood flow rate measurement is conducted concurrently with the monitoring of the other tissue viability parameters, providing simplicity 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.
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.
According to embodiments of the present invention in which the tissue vitality parameters being monitored (other than blood flow rate) are based on the NADH parameter, the excitation wavelength is in the range between about 370 nm and about 400 nm. According to other embodiments of the present invention in which the tissue vitality parameters being monitored (other than blood flow rate) are based on the flavoprotein parameter, the excitation wavelength is in the range between about 400 nm and about 470 nm.
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
Similar considerations apply to the haemodynamic artifact arising from Fp fluorescence emission and excitation light by the blood haemoglobin, mutatis mutandis.
As is explained in greater detail herein, in the present invention the problems normally associated with using an illuminating radiation wavelength in the range 370 nm to about 470 nm, are addressed as follows.
(a) Low sensitivity in the measurement of the AC/DC ratio and (b) safety issues. In the present invention, a relatively high intensity laser is used together with a chopping technique. This enables high intensity irradiation while mean irradiation intensity is hold below safety limit for the tissue. Additionally, high gain low noise detectors are also employed. (c) Optical noise generated by the laser. In the present invention, a specially stabilized low noise laser is used. (d) Sensitivity to detecting the AC component in the wavelength range 370-470 nm. In the present invention, special UV—Blue enhanced photodiodes with high sensitivity at UV or blue region are used. (e) Difficulty in detecting speckle pattern. In the present invention, thin collection fibers are used in order to enhance sensitivity to intensity fluctuations caused by speckle movement into and out the collection fiber area. (f) Poor optical efficiency in the 370 nm to 470 nm range In the present invention this is overcome by using, for example, silica high OH optical fibers, which have better transmission at UV.
Referring in particular to
Normally, laser diodes are sold as OEM modules such as PPM(400-5) from Power Technology, Little Rock, Ark., USA. These modules provide laser diode temperature stabilization, laser diode current driver and a collimating lens. Although the experience with laser Doppler flowmeters based on red laser diodes shows that the regular current driver and temperature stabilization provided by such module can be sufficient for Doppler measurements, this is not the situation with the violet laser diodes. In particular, the NLHV500 laser diode with regular current driver and temperature stabilization was found by the present inventors to be problematic for Doppler flowmetry, and therefore in its current form unsuitable for the present invention.
There are two other aspects of the violet laser diode radiation that where also found by the present inventors to be problematic in the context of the present invention, and may be present in other laser diodes of the same operating frequency. The first aspect is that of laser light amplitude noise. The NLHV500 laser diode emits several longitudinal modes as can be seen on
In the present invention, the broad mode bandwidth and global bandwidth obtained with such diodes is effectively and dramatically reduced by utilizing an external cavity system (Nakamura S. and Kaenders W. Market-ready blue diodes excite spectroscopists. Laser Focus World, April 1999). As illustrated in
An external cavity laser diode system as the DL 100 laser, particularly based on the NLHV500 diode, can operate at various radiation modes according to temperature and current conditions. In general four discrete operation states can be defined: (A) stable single longitudinal mode operation; (B) two stable longitudinal modes; (C) several longitudinal modes competing with each other; and (D) broad band operation. When considering the laser noise and the speckle visibility, which are the most important parameters for LDF measurements, the most appropriate operational state is state (A), with state (B) being less preferable though nevertheless possible. An intermediate state similar to state (B) can also exist, wherein 3 stable non-competing longitudinal modes are generated is less desirable. At operational states (C) and (D) the laser RIN noise is high, and therefore these operational states must be avoided during the laser Doppler flowmetry measurements.
In order to ensure that an external cavity laser diode system such as DL100 only operates in operation state (A) a special stabilization system is needed. This stabilization system implemented in the laser stabilization control system (LSC) (7), which is described in detail hereinbelow. This LSC system is an essential feature for the smooth and long-term continuous operation of the present invention, since the ECDL system (100) without the stabilization provided by the LSC system (7) will gradually drift between various states (A) through (D), as described above, and the laser Doppler flowmetry measurements obtained therefrom will thus be correspondingly unstable and unreliable.
Thus, in preferred embodiments of the present invention, and referring to
Referring to FIGS. 4, 8(a) and 8(b), the probe (2) comprises at the distal tip thereof a 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 a laser radiation at a target excitation wavelength to nominally a point (15) on the tissue (25). Alternatively, a bundle of fibers may replace the single fiber (201). The laser radiation comes from a suitable source, such as violet laser diode (101) coupled to the fiber (201) by any suitable optical coupler that preferably enables selective coupling and decoupling of the probe (2) with respect to the rest of the apparatus (99), such as SMA connector (205). Preferably, the probe (2) comprises a 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.
The probe further comprises one, and preferably a plurality of, collection fibers (202) for collecting light from the tissue. When more than one detection fiber (202) is used, this plurality of fibers may be arranged on a circle, and in each case the fibers (202) are distanced (R1) from the excitation fiber (201) (or coaxially with the geometric center of the corresponding bundle of excitation fibers, where appropriate), as illustrated in
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 (99) 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 (99) itself, or of the medical costs associate with a procedure using said apparatus (99), 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 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.
Radiation at nominally one wavelength, as will be further described hereinbelow, is delivered from the LSU (1) 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) 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.
In one preferred embodiment of the present invention, the apparatus (99) is directed to the monitoring of blood flow rate, and at least one of and preferably all of NADH concentration, blood oxygenation state and blood volume pertaining to an identical tissue volume, and is described in detail hereinbelow. In another preferred embodiment of the present invention, the apparatus (99) is directed to the monitoring of blood flow rate, and at least one of and preferably all of Fp concentration, blood oxygenation state and blood volume pertaining to an identical tissue volume.
Thus, in the first preferred embodiment of the present invention, 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 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 give a measure of the blood flow rate, and such changes are detected via one or more said collection fibers (202).
Referring to
The laser radiation from the laser head (100) is reflected by a suitable mirror (111) towards an acousto-optic (AO) modulator (103). In the preferred embodiment, the AO modulator (103) enables fast chopping/modulation of the laser light, as will be described in more detail hereinbelow. The CW laser radiation is chopped by the AO modulator at a repetition rate of 4 KHz to provide a stream of substantially identical pulses of laser light. In the preferred embodiment, each cycle duration is 250 microsec with a duty cycle of 1:10, i.e., with the ON period duration being 25 microsec, and the OFF period being 225 microsec. The chopping sequence is generated by the clock (403), which is comprised in the EU (4). The chopped radiation appears at the 1st order of the (AO) modulator (103). This order is spatially filtered by a circular diaphragm (not shown). The modulated radiation from the AO modulator (103) is reflected by dielectric mirror (106) and coupled to excitation fiber (201) of probe (2) by the lens (104) mounted on a suitable coupler adapter (105). A small portion (typically about 1%) of the excitation radiation passes through the dialectric mirror (106) towards the photodiode (107). This photodiode (107) enables the excitation intensity that reaches the probe (2) to be monitored. For enhanced safety a mechanical safety shutter (113) may be provided for preventing the laser radiation from reaching the probe (2) when the apparatus (99) is not in measurement mode or when the probe connector (205) is disconnected from the coupler (105). Preferably, operation of the safety shutter may be controlled by means of computer (5) via A/D (401).
The laser stabilization controller (LSC) (7) controls the temperature and current of the laser diode (101) in order to ensure that whole ECDL system (100) will operate in stable single longitudinal mode (state (A)), though if necessary or desired, the LSC (7) may be configured to enable operation in state (B). Thus, the LSC (7) unit comprises all sub-components needed for operation of the laser diode (101), and especially the laser head (100), which typically comprises a ECDL system, as described herein. The LSC (7) comprises a temperature controller (701) that includes the feedback of temperature data so as to enable the target to be heated or cooled, as appropriate to maintain a nominal temperature, and a highly stable current controller (702) for the diode laser (101). Additionally the LSC (7) further comprises a micro-controller, typically a microprocessor (704), with Analog to Digital (A/D) (703) and Digital to Analog (D/A) (705) converters. The A/D converter (703) receives the laser intensity proportional voltage from the photodiode (102). This analog voltage is converted to digital form and is fed into the microprocessor (704) as digital data. After processing the laser intensity information according to a predefined algorithm, the microprocessor (704) determines the precise value for the laser current and temperature that is required, and provides corresponding analogue signals, through the D/A converter (705), to the current controller (702), which in turn powers the diode (101). Various algorithms may be devised for finding and maintaining the operational state (A) for the ECDL system laser head (100). Perhaps the simplest such algorithm may be based on the measurement of the RIN parameter of the ECLD system laser head (100). As discussed above, the RIN is in essence the ratio between the (AC) fluctuations of the laser intensity to the total laser intensity (DC). Therefore, the RIN parameter is easily available from the measurement of the AC and DC components of the output voltage of the photo-detector (102). In order to maintain stable operational state (A), the microprocessor (704) controller initialises a short sweep of the laser current dI around the predefined nominal current I0. For example the NLHV500 operates at a nominal current I0=40 mA, and the current sweep dl may be about ±1 mA, for example. During this current sweep the ECDL system of the laser head (100) passes through all Operating States (A) through (D). Since the desired state (A) has the smallest RIN the current regimes associated with each state may be easily determined. After determination of the current regime associated with state (A), the center of the regime can be calculated and the laser current to laser diode (101) can be changed accordingly.
The ECLD system stabilization procedure can be initialized immediately at system power ON and with every calibration procedure of measured parameters as will be described hereinafter. In general after ECDL system is forced to operate at a stabilized state, it will in general remain there for several tens of minutes. During this period the system should preferably be re-stabilized periodically in order to force it to remain in such stabilized state in the long term. This re-stabilization can be performed during the OFF periods of the laser pulses in state II or III as will be described later. The duration of the ECLD system stabilization procedure is less then 0.4 sec therefore it can be easily performed in between the trains of pulses of state II or III.
The light from collecting fibers (202) is coupled to the DTU (3) via optical connector (267). The light from the connector (267) is collimated by lens (306) within the DTU (3).
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), such as a Hamamatsu S5973-02 detector. 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, and to perform Doppler flowmetry measurements to determine the corresponding blood flow rate. The remainder of the collimated light beam continues towards the second dichroic mirror (303).
Thus, radiation 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 than 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.
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 method described in co-pending Israel Patent Application No. 138683 filed by Applicants, the contents of which application are incorporated herein in their entirety, mutatis mutandis.
Thus, as far as blood oxygenation measurements are concerned, any suitable illumination wavelength may be used, and fluorescence that is emitted from the tissue element is then monitored. When combining the blood oxygenation measurements with NADH measurements, and particularly with blood volume measurements, the wavelength of the illumination radiation is advantageously chosen to correspond to a suitable isosbestic point. The intensity of the fluorescence emitted, as a function of wavelength, will vary according to the blood oxygenation state of the tissue element. Thus, referring to
Thus, blood oxygenation level is provided by the first, second and third PMTs, (308), (310) and (312) respectively, 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.
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 NADH, blood volume and blood oxygenation state, 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.
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) comprises 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 circuits for all 3 PMT detectors are identical.
The signal output of the PMT detector (312) is fed to the signal 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.
The monitoring of all parameters, LDF, NADH fluorescence, blood volume and blood oxygenation level involve excitation light at about 390 nm, which is in the UVA spectral region. Exposure to UVA radiation should be minimized as it is considered to be potentially dangerous even at low irradiation levels. In order to reduce the radiation safety problem, the option is provided in the present invention to chop the excitation light with a duty cycle of {fraction (1/10)} (ON Time/Total Cycle). Additionally use of chopped light enables the employment of synchronous detection techniques that enable better signal detection and recovery from noise.
The bandwidth of the laser Doppler flowmetry signals is from several tenths Hz to several KHz. This bandwidth imposes the lowest suitable chopping frequency according to the Nyquist principle. In practice the present inventors found that a 4 KHz-chopping rate is sufficient for typical laser Doppler flowmetry measurements.
Still after applying the chopping technique in order to reduce the tissue irradiation during long procedures or during the use of the apparatus in Intensive Care Unit (ICU), the total amount of irradiation applied to the tissue can be even more reduced by using the adaptive chopping technique.
There are many clinical conditions such as at long procedures or during ICU hospitalization where true real-time on-going measurements are unnecessary. At these stable state periods the measurements can be performed for only second set of parameters namely NADH fluorescence, blood volume and blood oxygenation level, since for measurement of these parameters only few sampling events are needed. The measurements of LDF must be performed by many sampling events at high repetition rate, since it requires a measurement of signals of a bandwidth of at least several KHz. This relatively demanding measurement of LDF can be omitted during the steady state period.
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.
Thus, in order to reduce the tissue irradiation the apparatus (99) 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:
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.
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.
Thus, according to the present invention the apparatus (99) may be used according to an adaptive chopping procedure. In such an adaptive chopping procedure, the radiation provided source (101) 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. It is important that this form of “pulsing” is different from the “pulsed” laser mode commonly encountered. The “pulsing according to the present invention provides a great degree of uniformity between the pulses generated as a result of the chopping procedure. Regular pulsed lasers cannot generally provide such a high level of uniformity between pulses generated thereby. Furthermore, the apparatus (99) 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 source (101) may be of a higher permitted intensity than would normally be allowable, albeit for shorter duration. 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.
As described in greater detail hereinbelow, 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, once a second, which while sufficient for monitoring NADH concentration, blood oxygenation state and blood volume, are too low for blood flow measurements, and thus minimise exposure to the laser radiation. Indeed the second set parameters—NADH concentration, blood oxygenation state and blood volume—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 (99) 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.
The outputs of the photodiode detector (301) 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.
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.
The clock (403) sub-unit provides the appropriate timing for the AOM (103) and the signal conditioner (402).
Referring now to FIGS. 10(a), 10(b) and 10(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.
The first type of detector as shown in
The second type of detector, illustrated schematically in
The third type of detector, illustrated in
Referring to FIGS. 11(a) to 11 (d), the trigger signal timing in
The gain of the detectors is defined automatically by the accompanying software in computer PC (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. 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 PMT detector. 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.
The gain of the second detector type is set by changing the gain of the second stage of the pre-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).
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.
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) and the detector (301) in
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
State I: The clock signal consists of a train of pulses in
State II: State I is additionally chopped by an ON/OFF adaptive duty cycle which enables and disables the light pulses train of
State III: The clock generates a sequence of five cycles of the state I like the pulses shown in
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 observed the clock is switched to state II, having a short OFF period t′off. After the changes became more moderate, the OFF period t′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.
Of course, the apparatus (99) can be configured to operate only in State I, either permanently, or whenever desired, without resolving the tissue irradiation safety issues in particular regarding internal tissue sensitivity to the UVA radiation. The current laser safety standards define only standards for the skin and eyes, but information is still lacking regarding the limiting values for the irradiation of internal tissues.
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.
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.
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 three parameters are measured i.e. NADH fluorescence, Oxygenation and 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.
The system or apparatus (99) 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.
At the end of the calibration, the computer switches the clock to state II.
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′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 25-microsecond 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 2% change is 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.
Advantageously, an ECLD system stabilisation procedure can be initialized immediately at system power ON and with every calibration procedure of the measured parameters. In general after the ECDL system is forced to operate at a stabilized state, by means of the LSC (7), it will in general remain at this stabilised state for several tens of minutes. During this period the system should preferably be re-stabilised periodically in order to force it to remain in such stabilized state in the long term. This re-stabilization can be performed during the OFF periods of the laser pulses in state II or III, described above. The duration of the ECLD system stabilisation procedure is typically in the order of less then 0.4 sec, and therefore such stabilisation can be easily performed in between the trains or packages of pulses of state II or state III.
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/cm2.
Thus, tissue may be irradiated with chopped light to provide important advantages, such as improving the accuracy in the measurements for all four parameters that are being monitored. 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.
“Chopped” laser illuminating radiation 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.
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)−1 for a parameter whose measurement is limited by background noise and by factor of (DR)−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).
The apparatus according to the present invention is based on a violet laser diode having moderate power consumption, and may be designed to occupy a reasonable volume such as to be easily and conveniently transportable and also storable within a regular operating theater, for example.
In the second preferred embodiment of the present invention, the apparatus (99) is adapted for monitoring blood flow rate and at least one of flavoprotein level, blood volume and blood oxygenation state based on the flavoprotein parameter, and thus comprises all the components and in the arrangement thereof similar to that of the first preferred embodiment as described above, mutatis mutandis, with the following differences.
Whereas in the first embodiment the illuminating wavelength is between about 370 nm and about 400 nm, and preferably 390 nm, so as to lie within the NADH excitation spectrum, in the second embodiment the illuminating wavelength is chosen to be between about 400 nm and about 470 nm, and preferably about 440 nm, so as to lie within the flavoprotein excitation spectrum, and preferably at an Fp isosbestic wavelength within the flavoprotein excitation spectrum at about 455 nm±5 nm.
Essentially, the Fp-based measurements are very similar to those based on the NADH parameter. The Fp fluorescence excitation is by monochromatic light at a wavelength within the Fp excitation spectrum, preferably at an Fp isosbestic wavelength thereof. In the present invention, this monochromatic light is provided by, and at the wavelength of, the laser light source. The Fp fluorescence is measured by measurement of fluorescence intensity of the fluorescence emission at single wavelength, which is within the emission fluorescence spectrum. This Fp fluorescence intensity provides a measure of the Fp concentration in a similar manner to that the derivation of NADH concentration from NADH fluorescence intensity, mutatis mutandis.
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.
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. Otherwise the fluorescence value will be influenced by the blood oxygenation. Similarly, for blood oxygenation measurements, the intensities of Fp fluorescence at two wavelengths are normalised with respect to the Fp fluorescence intensity at an isosbestic point, typically at a wavelength of about 530 nm.
Regarding fluorescence excitation for Fp, if only blood flow rate and/or only Fp concentration are to be monitored, and there is no need for the blood oxygenation parameter and/or for the blood volume parameter to be monitored, then any Fp 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 wavelength. Thus the Fp excitation radiation is preferably chosen to be at an isosbestic wavelength, typically about 455 nm.
Thus, other than the specific choice of laser illumination wavelength, and the choice of wavelengths for determining the Fp level, reflectance and blood oxygenation state, determination of the blood flow rate and of the set of tissue viability parameters—Fp, blood volume and blood oxygenation state—in the second embodiment is as described for the first embodiment, mutatis mutandis, and thus the LSU (1), DTU(3), LSC (7), and to a lesser extent the EU (4), PC (5) and PS (6) need to be correspondingly adapted accordingly to take into account the different illuminating wavelength and the different range of fluorescent emission and reflection wavelengths obtained.
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 third embodiment of the present invention, is directed to a multi-probe system (99′), illustrated in
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 (2), and subsequent detection of the return signals therefrom, by a corresponding plurality of detection units (DTU) (3). Each DTU (3) may be sampled in a predefined sequence that is correlated with the appearance of excitation light at appropriate probe (2). In other words, the multiprobe system (99′) essentially multiplexes the light source towards each one of the plurality of probes (99′), located on different parts of the tissue or organs.
The system (99′) according to the third embodiment of the present invention thus comprises similar components as previous preferred embodiments, viz LSU (1) probes (2), DTU (3), EU (4) PC (5), PS (6), LSC (7) as described with respect to the first and second embodiments, mutatis mutandis, with the following exceptions. The LSU (1) of the third embodiment, as illustrated in
From the optical coupler (267) the light passes to all the necessary components of the DTU (3), in a similar manner to that described for the first or second embodiment mutatis mutandis. Appropriate modification to the conditioning electronics of the EU (4) and the software running on the PC (5) is described below.
It should also be noted that multi-tissue element monitoring could also be accomplished by a plurality of probes (2), each one having a dedicated light source (LSU) with each probe unit being controlled by the same PC and EU units, and being powered using the same PS.
The EU (4) of the third embodiment is typically very similar to that of first or second embodiments. However, it further enables controlling of plurality of DTU (3) while the AOD (140) provides the excitation illumination each time to the appropriate probe (2). The electronics circuitry of the EU (4) is essentially the same as for the first and second embodiments.
The software running on the PC (5) is also typically very similar in concept to that described for first and second embodiments. However, the software also enables the two operation modes of third embodiment, as described hereinafter.
The third 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:
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 (comprising the tissue volumes) as a whole. This mode of monitoring could be useful in transplantation surgery when better monitoring of the viability of donated organs are needed.
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.
The electronics and the software for the first mode will be substantially similar to that described with respect to the first or 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.
The ton period per probe (2) of the system (99′) remains substantially the same as for the single probe of the first and second embodiments. However, during the OFF period for any probe in the system (99′), other probes (2) of the system may be selectively excited and measured. Accordingly, while the timing of the AOD (140) for each of the probes (2) in the system (99′) may be correlated with the sequence shown in
The second measurement mode of the third embodiment requires the same chopping sequence as that required by the first mode. In that second measurements mode the data acquired by each one of the DTUs will be treated and displayed by the PC separately.
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.
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.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/IL01/00900 | 9/25/2001 | WO |
