|Publication number||US3890959 A|
|Publication date||Jun 24, 1975|
|Filing date||Apr 26, 1973|
|Priority date||Mar 22, 1971|
|Publication number||US 3890959 A, US 3890959A, US-A-3890959, US3890959 A, US3890959A|
|Inventors||June N Barker, Theobald Reich, Myron Youdin|
|Original Assignee||Univ New York|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (7), Referenced by (16), Classifications (15)|
|External Links: USPTO, USPTO Assignment, Espacenet|
United States Patent Youdin et al.
METHOD FOR DETERMINING DISCRETE LOCALIZED BLOOD FLOW IN A LARGE VOLUME OF TISSUE  Inventors: Myron Youdin, Flushing; June N.
Barker; Theobald Reich, both of New York, all of N.(.
 Assignee: New York University, New York,
 Filed: Apr. 26, I973  Appl. No.: 354,489
Related US. Application Data  Continuation-impart of Ser. No. 126,770, March 22,
1971, Pat. No. 3,769,966.
 US. Cl. l28/2.05 F; [28/2 A; 250/303  Int. Cl A6lb 5/02; A61b 6/00  Field of Search 128/205 F, 2.05 R, 2.05 A, 128/205 V, 2 A, 2 G, 2 R; 250/303, 363, 364
 References Cited UNITED STATES PATENTS 3,268,728 3/1966 Stoddart et al. 128/2 A X 3,405,233 10/1968 Anger 250/715 S 3,418,471 12/1968 Gydesen 128/2 A X 3,432,660 3/1969 Anger 250/715 S 3,591,806 7/1971 Brill et al.. 250/715 S 3,658,054 4/1972 lberall 128/205 R 3,769,967 11/1973 Jones et a1 128/2 A X OTHER PUBLICATIONS Reid, W. B. et al., Intern. .lourn. of Applied Radiation & Isotopes, 1958, Vol. 3, pp. 1-7. Miraldi, F. et al., Radiology, Vol. 94, Mar. 1970, pp. 513-520. Hindel, R. et al., Nucleonics, 1967, Mar., Vol. 25, No. 3, pp. 52-57.
1 51 June 24, 1975 Bender, M. A. et al., Intern. Journ. of Applied Radiation & Isotopes, 1959, Vol. 4, pp. 154-161.
Primary E.raminer-Kyle L. Howell Attorney, Agent, or Firm-Hubbell, Cohen and Stiefel  ABSTRACT A method for determining discrete localized blood flow in a large volume of tissue, such as local cerebral emission flow in a discrete radiation emission detection zone of a patient's brain. The brain is subdivided into a plurality of detection zones, such as 144, by means of a headgear apparatus which defines these detection zones by means of a multiplanar array of an equal number of detectors, each detector having an associated collimator to limit the field of view of the detector. A radioactive gas, such as one having two gamma energy levels, such as Xe or a gaseous mixture of two different gamma energy level isotopes, is inhaled by the patient, such as in a closed inhalationexhalation breathing system, over a predetermined data collection period during which the concentration of the gas in the blood is periodically increased. Determinations of arterial blood radiation concentration and tissue radiation concentration are made during this period, and the local cerebral blood flow is subsequently determined from these determinations. The arterial blood radiation concentration is preferably determined by determining the radiation concentration of the gas in the last part of an exhalation, while the tissue radiation concentration is preferably determined from the radiation emissions quantified by the detector array such as by detecting both gamma energy levels, the field of view of the detector depending on the gamma energy level being detected.
59 Claims, 59 Drawing Figures RADIATION DETECTION SIGNAL PROCESSING CIRCUIT RADIATION DETECTION SIGNAL PROCESSING CIRCUIT TO EXHAUST HOOD J Z HEADGEAR 7 APPARATUS (.0 GAS L 0 GAS ANALYZER ANALYZER J iro 80 EXHAUST r END- TIDAL HOOD as SAMOLER 62 r I J 60 IMINICOMPL/TER! -r BUFFER f [8 ROOM 3: 3': TO EXHAUST HOOD I i 1 1 AIR 55 5 J 1 20 16 [a 22 42 GENERAL INHALATION 1 ms PURlexnAcAriona xENoN' oAs PURPOSE CHAMBER '7 FtCATION CHAMBER nwzcnom COMPUTER UNIT 4 SYSTEM 28 it 35 32 34 30 g-so l LOCAL ceases/u.
BLOOD FLOW DATA OUTPUT SHEET PATENTEDJUN 2 4 I975 RADIATION DETECTION -"67 SIGNAL PROCESSING CIRCUIT RADIATION DETECTION 4 SIGNAL PROCESSING CIRCUIT FIG. I.
AUSTHOOD W 0 GAS LYZER MINICOMPUTER BUFFER TO EXH HEADGEAR APPARATUS II ANALYZER -ANA 8 EXHAUST END- TIDAL I. L n R m w E WP T. E w MM mm E D D- 'M-w F O I N R M DA EUO MOT GPc oA &0 a P I. 6 )S m N 3M BUE m 5 NJ w EM X 2 8 4k 44 N @R ma LM m w XC E 4 +2 PN R0 U H PAW SC A 2 G ,6 :7 N 3 mm TB A AI I L H C W /8 2 TO RADIATION DETECTION SIGNAL PROCESSING CIRCUIT167) I34 88 I TO RADIATION DETECTION SIGNAL PROCESS- ING CIRCUIT (67) KG W s S Q N 00 Du CD. N UN 6 S CIRCUIT (67) N w T m D A R m GLABELLAR'SELLAR PL AN E PATENTEIIJIJII 24 I975 SHEET TO RADIATION DETECTION SIGNAL F' PROCESSING cIRcuIT (67) CIRCUIT (67) TO MINI COMPUTER BUFFER (60) TO RADIATION DETECTION SIGNAL PROCESSING 951:
SINGLE CHANNEL ANALYZER w Io W U C R C G N S a n W R 5 I m D. a I.- A E N U F. I I II IIIS IIIIIIIIII rm N u m N 8 T wn ZWNMR I 2 I I I IIIIIIIII m I. C RTA E .I O .I Mm a M CEm m l L I I I l I I I I I II MN m IITL MD 2 H ER H% 8 m Aow T I G D l B III A N/SCM A R 5 a g DEE S 0 am TI N T I IIIIIIIIII mwmw Q N 009m ".0 M33; 0 WU M D H QEQU 20 .F -E.2mUZOU UEOPOQ DEM MT M A m a GM PATENTEDJUN 24 ms sum 3 .890.959
PLANE ABELLAR- SELLAR PLANE 8! I a PARIETAL BONE I I I GLABELLA MANDIBLE FIG. 6.
HIGH VOLTAGE ,L POWER SUPPLY I46) THALLIUM- TEN STAGE ACTIVATED PHOTOMULTI' SODIUM PLIER WITH 23m IODIDE INDIVIDUAL CRYSTAL VOLTAGE DETECTOR DIVIDING ASSEMBLY NETWORK :53 18 0 '3 VARIABLE SUMMING PULSE GATE PRE- HEIGHT AND AMPLIFIER DETECTOR PULSE STRETCHER THRESHOLD TO LEVEL MINICOMPUTER CONTROL sumsmsol PATENTEDJUN 24 I975 SHEET 4 CLT COT CRT BLT BOT BRT C PLANE L'B PLANE PLANE a ALT AOT ART A 2-49 2-50 2-5| 2-62 2-63 2-54 PLANE 265256257263259260 zoo -iAxls OF vaEw cLo CO0 O C PLANE E-BI 2-62 2-63 2-64 2-65 2-66 BLO BOO BRO 2-67 2-66 2-69 2-70 2-7: 2-72 ALO A00 ARO APLANE '4 2-25 2-26 2-27 2-23 2-29 2-30 81c 2-3: 2-32 2-33 2-34 2-35 2-36 CLU cou CRU BPLANE L 2-37 2-36 2-39 2-40 2-4l 2-42 BLU sou BRU c PLANE 2-43 2-44 2-45 Z-46 2-47 2-46 ALU AOU ARU B PLANE A PLANE 2-2 2-3 2-4 2-5 2-6 2-7 \2-9 2-l0 2-u 2-l2 A PLANE k 2-13 2-14 2-15 Z-&6 2-17 Z-l8/ 2-19 2-20 2-2: 2-222-23 2-24 DIRECTION OF VIEW METHOD FOR DETERMINING DISCRETE LOCALIZED BLOOD FLOW IN A LARGE VOLUME OF TISSUE CROSS-REFERENCE TO RELATED APPLICATIONS This application is a continuation-in-part of US. Ser. No. l26.770, filed Mar. 22, l97l, now US. Pat. No. 3,769,966, and entitled Apparatus for Determining Local Cerebral Blood Flow".
BACKGROUND OF INVENTION 1. Field of the Invention The present invention relates to a method for deter mining localized blood flow in large volumes of tissue such as local cerebral blood flow.
2. Description of the Prior Art The ability to measure localized blood flow in various organs of the body such as cerebral blood flow within the human brain is an important medical diagnostic tool as variations in a normal flow, such as in the brain, can indicate actual, probable, or impending cerebral medical deficiencies, such as possible stroke or brain trauma. Since the human brain is the vital command center for the body, it is important that any possible malfunctions which affect the brain be discovered as rapidly as possible. Prior art diagnostic techniques for determining total or regional cerebral blood flow have involved the use of a radioactive diagnostic agent dissolved in the blood. However, such prior art techniques have only been able to determine a total or regional blood flow, which is a mean rate of blood flow through a region, such region normally being defined by a large cone" oftissue. One such prior art device which might provide a means of estimating a qualitative indication of regional blood flow utilizes a uniplanar detector array. However, such a device has a poor energy and spatial resolution and cannot detect small blockages of blood flow within the brain; even large reductions in local flow are not easily discernible.
Other prior art techniques of this type utilize a multiplicity of probes in a random array to view" a large randomly located cone shaped region in the brain for determining a mean rate of cerebral blood flow in this single random location. However, no provision is made for relating the amount of radioisotope contained within the intercepted tissue with the observed count rate in a quantitative manner. In addition, if there is a variation in cerebral blood flow from that ofa previous measurement. there is no way of pinpointing the exact location of the flow reduction other than that there is one in the overall region.
Some of these prior art techniques have utilized the inhalation or intra arterial injection of a radioisotopic gas as the diagnostic agent in conjunction with a re gional blood flow measurement. However, the prior art inhalation technique requires the patient to breathe the radioactive gas until equilibration and subsequent elimination occurs. The equilibration period is normally seven to ten minutes using a closed breathing circuit of a device such as a basal metabolism or anesthesia machine followed by a washout period of greater than minutes. The washout of the radioisotopic gas is then measured with an external radiation detector directed toward the region. However, once again, only a mean rate of blood flow, or regional blood flow,
through the region, or cone of tissue if that is the defined region, is provided.
The technique wherein the radioactive diagnostic agent is injected into the internal carotid artery is considerably more dangerous to the patient and cannot be used except under serious medical indications for the test. While this latter technique obviates the contribution of extracerebral radioisotope, the time limitations for measurement are equally long and the method still only provides a regional blood flow measurement.
One serious limitation of these prior art regional blood flow determination methods, each of which utilize radioisotopic indicators, is that the extra cranial radiation detectors measure radiation activity from an imprecisely defined cone of tissue without being able to distinguish the depth from which the radiation activity emanates. Therefore, quantization of blood flow in brain substructures within the cone is not possible other than in the cerebral cortex which can be monitored with tangentially placed detectors. However, since nonuniformity in the distribution of cerebral blood flow during health as well as during disease is normally present, there being as much as ten to twentyfold difference between blood flow in the cortical white matter and in the inferior colliculus, or, during periods of hyperperfusion or hypoperfusion, the various parts of the brain manifesting strikingly different changes in the associated blood flow, some changes being large while others being quite small, such prior art regional techniques have not been satisfactory as a diagnostic tool, such as in attempting localization of local cerebral ischemia, the essential lesion of stroke. Furthermore, these prior art techniques have not been satisfactory where quantitative assessment of blood flow and ultimately of metabolic rate in many specific loci of the brain simultaneously is desired.
Since these prior art techniques are slow, of low resolution, and provide only generalized regional mean rate blood flow data as opposed to localized, or direct discrete measurements of localized blood flow in a large volume of tissue, such as cerebral blood flow in a defined discrete localized portion of the brain, these techniques have not gained wide acceptance for other than cases in which such generalized data is acceptable.
These disadvantages of the prior art are overcome by the present invention.
SUMMARY OF THE INVENTION A method is provided for determining discrete localized blood flow in a large volume of tissue, such as local cerebral blood flow in a discrete defined radiation emission detection zone of a patients brain. The brain is subdivided into a plurality of detection zones, which are preferably identical, by means of a headgear apparatus which defines these detection zones. The headgear apparatus is preferably aligned on the patients head with respect to three mutually perpendicular reference planes; namely, the glabellansellar plane as a reproducible basal plane, the mid-saggital plane as a zero reference plane, and a reference plane tangent at the forehead. This headgear apparatus preferably includes an array of I44 or more radiation emission detectors arranged in a multiplanar array about the brain so as to provide an equal number of defined viewing or detection zones in the brain. The detectors, such as semiconductors, thalliumactivated sodium iodide crystals, ultrapurified germanium detectors or proportional counter tubes each have associated therewith a collimator which has a lumen of predetermined configuration. preferably square. in order to present a predeter mined portion of the plurality of detection zones to the associated detector.
A radioisotopic gas. such as one having two gamma energy levels. such as Xe or Kr. or a gaseous mi tture of two different gamma energy ievel isotopes, is inhaled by the patient in a closed inhalation-exhalation breathing system over a predetermined data collection period. such as 1 minute. During this period. the concentration of this gas in the blood is increased due to periodic injection of a specified additional amount of the gas into the breathing system. Measurements are made of the arterial blood radioisotope concentration. such by a means separate from the 144 detectors. which determines the radioisotope concentration of the gas in the last part of an exhalation ofthe patient during the data collection period. A determination is also preferably made of the quantity of radioisotope concentration associated with each discrete detection zone based upon measurements of resultant radiation emission by the detector array, such as by detecting both gamma energy levels, the field of view of the detector depending on the gamma energy Ievel being detected, and a predetermined relationship between the plurality of detection zones which pro ides a plurality of solvable simultaneous equations relating the contributions of various detection zones in the plurality to each quantified resultant radiation emission. These equations are solved to provide the tissue radioisotope concentration for each discrete detection zone. Of course. if desired. only one gamma energy level can be detected and the radioisotope concentration can still be determined.
however. the detection of two gamma energy levels having different depths of penetration facilitates the obtaining of the desired results and is preferred over the use of single energy level detection. at least with respect to the detection of cerebral blood flow. The local cerebral blood flow for a discrete detection zone is pro vided from the tissue and arterial radioisotope concentrations for the zone. which concentrations are preferably repeatedly determined at It) second intervals during the l minute data collection period. in accordance with another predetermined relationship between these factors and cerebral blood flow, such as by applying the well known Ficit-l\'ety expression.
BRIEF DESCRIPTION OF DRAWINGS FIG. I is a diagrammatic illustration ofthe preferred system for use in the method of the present invention;
FIGS. 2A and 2B are diagrammatic illustrations of the headgear apparatus portion of the system for use in the method of the present invention illustrating an alternative arrangement of detectors for detecting single gamma energy radiation emission in accordance with the method of the present invention;
FIG. 3 is a block diagram of the radiation detection signal processing circuitry portion for one detection channel of the headgear apparatus;
FIG. 4 is a graphical illustration of isotopic concentration utilized in describing the method of the present invention:
FIG. 5 is a diagrammatic illustration of the human skull from the side utilized in explaining the preferred orientation of the headgear apparatus for use in the method of the present invention;
FIG. 6 is a block diagram partially in schematic, similar to FIG. 3, of an alternative embodiment ofthe signal processing circuitry for one detection channel;
FIG. 7A is an exploded perspective view of a mathe- 5 matical model utilized in describing an alternative method of the present invention wherein a single gamma energy level is detected;
FIG. 7B is an exploded perspective view of a mathematical model. similar to FIG. 7A, utilized in describing the alternative method of the present invention;
FIG. 8 is a diagrammatic illustration utilized in explaining the field of view of a single detector in accordance with the alternative method of the present invention;
FIGS. 9A-9E is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of one gamma energy level in a two-by-two portion ofa typicai quadrant of the detection zone array;
FIGS. l0A-10E is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 9A9E in a different two-by-two portion of the same typical quadrant as in FIGS. 9A-9E,
FIGS. lIA-IIE is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy as in FIGS. 9A9E in still a different tw0-by-tw0 portion of the same typical quadrant as in FIGS. 9A9E;
FIGS. 12A12D is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 9A--9E in another two-bytw0 portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. 13A-13E is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 9A-9E in still another two-by-two portion of the same typical quadrant as in FIGS. 9A-9Ii;
FIGS. I4AI4D is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 9A9E in a different two-by-two portion of the same typical quadrant as in FIGS. 9A9E;
FIGS. 15A-I5C is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of a different gamma energy level than in FIGS. 9A-9E in a three-zone-deep portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. 16A and 16B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 1SA15C in a different threezone-deep portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. I7A-I7C is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. I5A-1SC in still a different threezone-deep portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. 18A and 18B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. ISA-C in a different threeby-three portion of the same typical quadrant as in FIGS. 9A-9E'.
FIGS. 19A and 19B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. 9A-9E in a different two-zonedeep portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. A-20C is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of both the gamma energy levels of FIGS. 9A-9E and FIGS. ISA-15C in another three-by-three portion of the same typical quadrant as in FIGS. 9A-9E;
FIGS. 2IA and 21B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. ISA-15C in two different three-deep-zone portions of the same typical quadrant as in FIGS. 9A-9E;
FIGSv 22A and 22B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGSv I5AlSC in a different substantially three-zone-deep portion of the same typical quadrant as in FIGS. 9A-9E', and
FIGS. 23A and 23B is a mathematical model utilized in describing the preferred method of the present invention illustrating the detection of the same gamma energy level as in FIGS. ISA-15C in still a different substantially threezonedeep portion of the same typical quadrant as in FIGS. 9A-9E.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS Ge neral Description Referring now to the drawings in detail and especially to FIG. 1 thereof, the system of the present invention, generally referred to by reference numeral 10, for practicing the method of the present invention is shown. The preferred method of the present invention shall be described, by way of example, for determining local cerebral blood flow, which is the blood flow within a discrete preselected portion of a patients brain although the preferred method of the present invention may be utilized, if desired, to determine discrete localized blood flow in any large volumn of tissue such as in the myocardium, the lung, the kidney, the liver, the spleen or any other body organ, or to resolve the metabolic rates of component regions of the brain and other organs.
It should be understood that FIG. 1 is merely a diagrammatic illustration of the system 10 for purposes of explanation, and is not meant to indicate the critical dimensions of the system, such as preferred requirement of minimum dead space in the closed breathing system or the critical arrangement and orientation of radiation detectors in the headgear apparatus. The system 10 of the present invention preferably includes a breathing mask 12 which is preferably removably mounted on the face of the patient during the performance of the local cerebral blood flow determination so as to isolate the breathing passages of the patient from the surrounding environment and provide a closed inhalationexhalation breathing system, which will be described in greater detail hereinafter. The mask 12, which is preferably a conventional breathing mask, is preferably formed of a material which is pliable enough to provide a comfortable but close fit on the face of the patient in the mounted position so as to reduce apprehension during a data collection period which is associated with the local cerebral blood flow determination.
The mask 12 includes a conventional type of mouthpiece breathing port 14 which is provided with an inlet or inhalation conduit 16 and an outlet or exhalation conduit 18. Conduits l6 and 18, respectively, supply the substance or gas to be inhaled to the mask 12 and remove the substance exhaled from the mask 12. Preferably, the inhalation conduit 16 and the exhalation conduit 18 each include a conventional demand valve 20 and 22, respectively, disposed within the conduit flow path so as to close off the respective associated conduits I6 or 18 unless the patient is either demanding the next breath (inhalation) or exhaling his previous breath (exhalation).
Inhalation demand valve 20 is preferably interposed in inhalation conduit 16 between the mouthpiece breathing port I4 and a small gas reservoir radiationshielded spirometer or inhalation chamber 28 from which the substance or gas to be breathed by the patient is provided. Exhalation demand valve 22 is preferably interposed in exhalation conduit 18 between the mouthpiece breathing port 14 and an exhalation chamber 30 which is preferably equal in volume to the inhalation chamber 28. The inhalation chamber 28 and the exhalation chamber 30 are preferably connected together in the closed inhalation-exhalation breathing system through a conventional gas purification unit 32 having an inlet conduit 34 connected to the exhalation chamber 30 outlet, and an outlet conduit 36 connected to the inhalation chamber 28 inlet. The gas purification unit 32 operates in a conventional manner to remove excess carbon dioxide from the breathing system, add oxygen to the breathing system, and filter the air passing therethrough. The gas purification unit 32 is preferably operated responsive to the beginning of a breath by the patient in a manner to be described in greater detail hereinafter.
In practicing the preferred method of the present invention, as will be explained in greater detail hereinafter, in order to introduce a diagnostic agent into the cerebral blood, the patient breathes a radioactive gas which most preferably has two distinct gamma energy levels such as the radioisotopic gas Xenon which is the preferred diagnostic agent, although a gaseous mixture of two different radioisotopic gases, each having different gamma energy levels could be utilized as the diagnostic agent. In addition. if desired, alternatively, a radioisotopic gas having only a single gamma energy level or one having more than one level with only one level being detected could be utilized if the improved resolution of the above is not desired. As will also be explained in greater detail hereinafter, a high concen tration of this gas is desirable during the performance of the local cerebral blood flow determination and, most preferably, this concentration increases during the quantification of the data necessary for the local cerebral blood flow determination. For purposes of explanation, the preferred method of the present invention, as well as an alternative method thereto shall be de scribed with the use of radioactive Xenon diagnostic agent.
In order to accomplish the previously described desired condition, a radioactive gas injection system 40 is connected through the valve 42 via a conduit 44 to the inlet of the exhalation chamber 30 for injecting a predetermined quantity of the radioactive Xenon gas into the exhalation chamber 30 each time valve 42 is opened. in addition, gas injector 40 also preferably includes a conduit 48 which is connected to the inlet of the inhalation chamber 28 via another valve 50 which is preferably only opened at the end of the patients next breath after the beginning of the data collection period so as to also inject this predetermined quantity of radioactive gas into inhalation chamber 28 as well as exhalation chamber 30 at this time. Preferably, valve 50 remains closed throughout the remainder of the data collection period and gas injector 40 only injects the radioactive gas into the exhalation chamber 30 where it is mixed, such as by an impeller (not shown), with the gas contained therein and passed through the gas purification unit 32 to the inlet of the inhalation chamber 28. If desired, other variations of this system may be utilized, such as by locating the purifier 32 between valve 22 and chamber 30, so that fresh radioisotope need not pass through purifier 32.
Preferably, for purposes of gas evacuation at the completion of the data collection period, another valve 52 is interposed in inhalation conduit 16 between demand valve and the mouthpiece breathing port 14, and another valve 54 is interposed in conduit 18 between the exhalation demand valve 22 and the mouthpiece breathing port 14. Valves 52 and 54 are each, respectively, connected to conduits S6 and 58. Conduit 56 is preferably open to the surrounding environment to provide room air therethrough, and conduit 58 preferably opens to an exhaust hood (not shown) for evac' uating the gas from the system 10. Valve 52 preferably blocks conduit 56 during the data collection period while leaving conduit 16 open to demand valve 20 and, similarly, valve 54 blocks conduit 58 during this period while leaving conduit 18 open to demand valve 22. As will be explained in greater detail hereinafter, at the conclusion of this period, valves 52 and 54 are operated, such as electrically via a signal from a conventional programmable minicomputer 60, such as a Honeywell 316, which may be utilized to coordinate the data collection and the resultant cerebral blood flow determination. The actuation of valves 52 and 54 changes them from their normal mode to the gas evacuation mode wherein conduits l6 and 18 are closed off to demand valves 20 and 22 and inhalation chamber 28 and exhalation chamber 30, respectively, and open to conduits 56 and 58, respectively.
In this manner, a closed inhalation-exhalation breathing system is provided during the data collection period for the local cerebral blood flow determination from inhalation chamber 28, via inhalation conduit 16 to mouthpiece breathing port 14 of the mask 12, via mouthpiece breathing port 14 to exhalation conduit 18, via exhalation conduit 18 to the exhalation chamber 30, via exhalation chamber 30 to the gas purification unit 32, and via gas purification unit 32 back to the inhalation chamber 28; with the radioactive gas being injected periodically during this period into the closed inhalation-exhalation breathing system from gas injector 40.
gas as the The portion of the system utilized in making a quantitative local cerebral blood flow determination will now be described. As will be explained in greater detail hereinafter, a conventional general purpose computer 61, such as an lBM 360, may preferably be utilized in conjunction with, or in place of, the minicomputer in order to process the data collected during the data collection period, if desired, in order to enhance the speed in which results are obtained. However, if desired, any other conventional method of processing such data may be utilized. For purposes of illustration, it shall be assumed that the general purpose computer 61 is utilized in conjunction with minicomputer 60 to process this data.
A conventional end-tidal sampler 62, such as a Falk- Kupferman end-tidal sampler, preferably has an inlet which is connected to exhalation conduit 18 for sampling the alveolar or endtidal gas concentrations at the end of each breath, and an outlet which is connected to a radioactive gas analyzer 63, to be described in greater detail hereinafter, which preferably includes a coiled passageway 64 of known volume and a radiation emission detector assembly 65. As will be described in greater detail hereinafter, the end-tidal sampler 62 contains a relay mechanism (not shown) which is normally actuated at the beginning and end of each breath in a conventional manner. This relay is preferably utilized to activate the gas injector 40 at the end of each breath and the gas purification unit 32 at the beginning of each breath in a manner to be described in greater detail hereinafter.
The end-tidal samples passed through end-tidal sampler 62 are preferably drawn through the coiled passageway 64, or another type of sampling chamber if desired; of analyzer 63, which is preferably shielded against external radiation. For efficiency and convenience, the coiled passageway 64 and detector assembly 65 are preferably located near the patients head adjacent a headgear apparatus 66 which is mounted in a predetermined position on the patients head preferably by means of a universal mounting means 68. A more detailed description of a typical headgear apparatus 66 will be presented hereinafter in connection with FIGS. 2A and 28, although the detectors illustrated therein are illustratively arranged and oriented with respect to their associated fields of vision in accordance with the alternative method of the present invention, the arrangement, orientation and field of view of these detectors differing in accordance with the preferred method of the present invention, as will be described in greater detail with reference to FIGS. 9-23.
Suffice it to say at this point that the headgear apparatus 66 preferably provides radiation emission detection signals which are each fed through an associated radiation detection signal processing circuit 67, one circuit being provided for each detector although only one such circuit and signal path are shown for purposes of clarity. Circuit 67 will be described in greater detail hereinafter. The output of circuit 67 is in turn preferably fed to the minicomputer 60, which, preferably acts as a buffer in order to more efficiently handle the data collected during the data collection period as considerable expense is involved in utilizing a large scale general purpose computer 60 such as the IBM 360 on line" throughout the data collection period. However, if this is not a consideration, the minicomputer buffer 60 may be omitted and the data fed directly to the general purpose computer 61.
The radiation emission detector assembly 65, which is to be described in greater detail with reference to the headgear apparatus 66 which preferably contains similar assemblies, is mounted in analyzer 63 in relation to coiled passageway 64 so as to quantify the radiation emission associated with the sampled gas contained in the coiled passageway 64. Detector assembly 65, which preferably provides a signal output in response to a gamma radiation input, is connected to a radiation detection signal processing circuit 74, which is preferably similar to detection circuit 67, whose output is in turn also preferably connected to the minicomputer buffer 60.
Preferably, the gas analyzer coiled passageway 64 is also operatively connected, in a manner to be described in greater detail hereinafter, to an end-tidal sample analyzer system 75, to be described in greater detail hereinafter, in order to determine the concentrations of carbon dioxide and oxygen in the patients breath. Both of these substances, and especially carbon dioxide, are influential upon the rate at which the blood flows through the human brain and should, therefore, preferably be taken into consideration. Normally, the patients respiration is adjusted to maintain these gases at fairly constant levels. However, some patients may have additional diseases which may influence the level of these gases; while other patients may ovenbreathe during the data collection period because of tension or apprehension induced by the data collection proce dure. If the patient overbreathes during this period and thus reduces his carbon dioxide pressure, his cerebral blood flow will be slowed and of an altered pattern even though his cerebrovascular system is normal. However, the relationship between the reduction in this blood flow for a given reduction in carbon dioxide pressure is known and can be easily compensated for in the final determination of cerebral blood flow if this reduc tion is known. Similarly, a patient having impaired ventilation may have a higher rate of cerebral blood flow which can also be predicted from the elevated carbon dioxide pressure if it is known. in addition, low arterial oxygen tension during impaired ventilation/profusion can also cause a faster rate of cerebral blood flow. It is therefore both preferable and desirable to know the carbon dioxide and oxygen pressures in order to compensate for their effects on the cerebral blood flow determination or errors may be introduced in the final determination.
The end-tidal sample analyzer system 75, preferably includes a conventional through-flow carbon dioxide gas analyzer 76, such as a Capnograph which is an infrared carbon dioxide gas analyzer, and a conventional through-flow oxygen analyzer 77, such as a Rappox manufactured by Instrumentation Associates, which analyzers 76 and 77 are preferably connected together in a serial through-flow path, although they may be connected in parallel through-flow paths if desired. The inlet to the carbon dioxide gas analyzer 76 is preferably connected to the coiled passageway 64 outlet via a conduit 78 having a valve 79 disposed therein for preferably directing the end-tidal sampled gas leaving the coiled passageway 64 either through conduit 78, from which it is drawn through the carbon dioxide analyzer 76 and therefrom through the oxygen analyzer 77 from where it is exhausted to the closed circuit breathing system or through another conduit 80, through which it is exhausted directly to the exhaust hood. Preferably, the valve 79 only passes the end-tidal sampled gas through the analyzer system prior to the initiation of the actual data collection period associated with the local cerebral blood flow determination, the valve 79 directing the gas through conduit during this period, in order to quantify or measure the oxygen and carbon dioxide gas pressures present in the end-tidal sample as well as the period oftime in which the carbon dioxide gas present in the end-tidal sample is on the plateau portion of its exhalation curve, which curve is similar to that for Xenon gas. The end-tidal sampler 62 may then preferably be normalized or calibrated to operate on this portion of the exhalation curve during the data collection period.
Furthermore, the oxygen and carbon dioxide pressure information may be utilized to determine the time at which the actual data collection period is to commence, which is preferably when these pressures are stabilized. At that time, the data collection may proceed without introduction of any significant errors in the cerebral blood flow determination due to carbon dioxide or oxygen gas pressure variations. If desired, these measurements may be displayed visually and the data collection begun manually by an operator by actuation of a start switch (not shown) when these pressures are stabilized, or in the alternative, these values may generate corresponding electrical signals which when fed to a comparator (not shown) may provide a begin data collection" signal to the minicomputer 60 at stabilization if the minicomputer 60 is utilized as an overall control mechanism for the data collection procedure.
As will be explained in greater detail hereinafter, the normalized end-tidal sampler 62 is preferably initially activated from minicomputer 60 at the time when the actual data collection period is to commence so as to synchronize the selection and radiation emission detection times of the end-tidal gas samples with the radiation emission detection times of the headgear apparatus 66, the end-tidal sampler 62 being activated in a conventional manner thereafter. if the start signal is manually initiated, then this signal merely puts the minicomputer 60 on-line so that it may provide the end-tidal sampler 62 synchronization signal; whereas, if the minicomputer 60 is utilized as an overall control, then it can be programmed to only provide this synchronization signal after it has received the start signal from the comparator.
If desired, the carbon dioxide and oxygen gas analyzers 76 and 77, respectively, may be electrically connected to the minicomputer 60 in order to provide dynamic pressure information during the data collection period instead of, for example, merely providing a start signal and, in this instance, the computer 61 may then be preferably programmed to compensate for blood gas tension variations in the resultant local cerebral blood flow determination. In such instance, the valve 79 would be open so as to direct the end-tidal gas sample to the calibration systen 75 throughout the actual data collection period, as opposed to only prior thereto.
HEADGEAR APPARATUS Now referring to FIGS. 2A and 2B and describing the headgear apparatus 66 in greater detail, the system 10 of the present invention, as will be explained in greater detail hereinafter, preferably subdivides" the patients brain into a predetermined plurality of discrete detec tion zones for purposes of determining the rate of cerebral blood flow in each of these zones, although, if desired, the final determination may be limited to any number of zones in the plurality, such as one. These detection zones. which are shown by way of example by dotted lines in FIGS. 2A and 2B are preferably represented by three-dimensional cubes 81-81 which are each of equal volume and dimensions. It should be noted that these detection zones 81 are merely mathe matical models of sub-volumes defined within the over all brain volume by means of the headgear apparatus 66 and are preferably correlated to a brain map so as to pinpoint their exact location within the brain. Most preferably. 144 such detection zones are utilized to define the overall brain volume, although other numbers might be employed. As will be explained in greater detail hereinafter, during the data collection period. each of these zones 8] will be provided with an associ ated radioisotope concentration and corresponding radiation emission due to the presence of the radioactive gas in the cerebral blood. Hence, these zones 81 are termed radiation emission detection zones.
As was previously mentioned, the arrangement and orientation of the detectors, which will be discussed in greater detail hereinafter, illustrated in FIGS. 2A and 2B, and their associated illustrated fields of view, are in accordance with an alternative method of the present invention, to be described in greater detail hereinafter, wherein only a single gamma energy level of the radioisotopic gas is detected. Before describing this alterna tive method and arrangement which differs from the preferred arrangement utilized in accordance with the preferred method of the present invention, the preferred arrangement shall be described with reference to FIGS. 9-23 wherein two different gamma energy levels of the radioisotopic gas are detected, these levels, by way of example, being 31 Kev and 81 Kev for Xenon As will be explained in greater detail hereinafter, the headgear apparatus 66 should preferably bear a specific and reproducible relationship to the patients brain. Therefore, in aligning the headgear apparatus 66, a reproducible basal plane, which defines the angle of viewing to which all other planes will be parallel, perpendicular, or angulated at 45 degrees, must be chosen. In the preferred system 10 of the present invention the glabellar-sellar plane (see FIG. 5), which is the plane which delimits the lower surface of the human forebrain, is preferably chosen as a reference plane. The mid-saggital plane (see FIG. 5), which is the plane perpendicular to the basal plane and midway between its lateral limits, is preferably chosen as the second plane oforientation or reference in the preferred alignment position of the headgear apparatus 66. The third plane of orientation or reference for the headgear apparatus 66, which thereby defines the exact alignment or orientation of the headgear apparatus 66 with respect to the brain of the patient, is the plane perpendicular to the basal plane and the mid-saggital plane and tangent to the forehead (see FIG. 5). When the headgear apparatus 66 is aligned in these three planes, with the midsaggital plane preferably being the zero reference plane, the patients brain volume is subdivided into the plurality of detection zones 81, which are pref erably cubic or rectangular parallelepiped blocks in volume, arranged in a rectangular six-by-six array in 12 the glabellarsellar plane and four blocks high in the midsaggital plane, as shown in FIG. 5 by the dotted lines.
Referring now to FIGS. 923, the array of detectors is preferably arranged to view a block-shaped volume (FIG. 5) of sufficient size to include a typical adult human head although, if desired, the size could be varied to include any size head such a childs. The headgear apparatus 66 preferably includes a substantially rectangular hollow helmet housing 82 having a front portion 84, a rear portion 86, a top portion 88, side portions and 92, and a bottom portion 94 through which the head is inserted, the housing 82 forming a rectangular parallelepiped. By way of example, and not limitation, typical internal dimensions for a typical housing 82 which should accommodate approximately I00 percent of the adult female heads and approximately 99 percent of adult male heads are approximately 8% inches from anterior to posterior, 6% inches from side to side, and 5 2l/32 inches from superior to inferior, the total volume being subdivided into I44 smaller rectangular blocks or detection zones 81 of equal volume, each being, in the example given above, preferably 36 28.6 36mm. The large rectangular block encompassing the brain volume is preferably subdivided into six divisions from side to side, six from front to back and four from top to bottom, resulting in the 144 detection zones 81 preferably requiring, as will be described in greater detail hereinafter, the solution of I44 unknowns representing the quantity of radioi sotope in each of the I44 zones 81, an equal number of detectors (I44) preferably being provided spaced about the brain volume within the helmet housing 82.
Preferably, the radioisotopic gas, such as Xenon 133,
- has a limited penetration for the gamma photons emitted thereby which enables the I44 block or zone 81 volume to be preferably divided into four equal quad rants of 36 blocks each, each quadrant being 3 3X4 blocks, for purposes of determining the local cerebral blood flow in each detection zone 81 in accordance with the preferred method of the present invention, 36 detectors being associated with each quadrant. Of course, if desired, the entire I44 zone block may be treated as a whole rather than as four equal quadrants, although when the quadrants can be treated separately (depending on gamma energy) as preferred errors in measurement which may inadvertently occur in one quadrant cannot affect the results in the other quadrant. Preferably, the block or zone sizes are chosed so that natural attenuation will limit the source of the lower gamma energy (31 Kev for Xenon I33) photons almost entirely to the two superficial zones in depth of the array and the higher gamma energy (8i Kev for Xenon 133) photons to three zones in depth (skull and scalp attentuation included). The presently preferred arrangement ofdetectors on a typical 36 zone quadrant (zones Z-I through 21-36) is illustrated in FIGS. 9-23 wherein arrows labeled with Q and P subscripts, respectively, represent the position of the low energy and high energy detectors, respectively, in the detector array, the illustrated quadrant representing the frontright quadrant.
As will be described in greater detail, each detector is preferably fabricated from p type high purity germanium and has an associated collimator to define the field of view illustrated in FIGS. 923 by the dotted lines for the Q or low energy level detectors and the P or higher energy level detectors. Preferably each of the detectors has an associated single channel signal analyzer (see FIG. 3), each detector channel preferably having the capability of quantitating and discriminating more than one photon or gamma energy level within its field of view, the energy of the photon determining its tissue penetration. For purposes of explation, separate detectors are illustrated (O and P) for low energy and high energy photons, the associated field of view being selected in accordance with the setting of the single channel analyzer window; however, if desired. detectors may be utilized which quantitate both photon energies simultaneously. Preferably, a plurality ofindividual detector assemblies (similar in structure to the alternative detector illustrated in FIG. 8 but differing in fietd of view) equal in number to the plurality of radiation emission detection zones 81, which in the preferred method of the present invention is 36 for each quadrant, are arrayed about each quadrant of the housing in various orientations as illustrated by the P and Q arrows in FIGS. 9-23. Each of the detector assemblies is preferably substantially identical in structure only differing in the field of view provided by the associated collimator (approximately two zones 81 deep for the Q detectors and approximately three zones 81 deep for the P detectors as illustrated in FIG. 21B by way of example) and the setting of the single channel analyzer window to detect low energy photons (e.g. 31 Kev) for the group of detectors and high energy photons (eg 81 Kev) for the P group of detectors, the field of view for all P group detectors preferably being substantially equal in volume with the variations therefrom illustrated in FIGS. 18B, 22B and 23B, and the field of view for all Q group detectors preferably being substantially equal in volume with the variations therefrom illustrated in FIGS.9E,10E,1]E, 12D, 13D. 14C and C, each of these 36 individual detector assemblies preferably yielding an independent linear simultaneous equation.
As shown and preferred in FIGS. 9-23, each of the detector assemblies represented by the P and Q arrows is preferably located at the center of the point of entry into the adjacent detection zone 81 so as to be sub stantially symmetrical about the longitudinal viewing axis of the detector field of view. Thus, as shown and preferred in FIGSv 9-23, the P and Q groups of detector are arranged at 26 positions at the center of the associated adjacent viewing surface defined by the corresponding detection zone 81 in the quadrant and at 10 positions substantially along a 45 degree line at the center of the corner ofthe associated adjacent viewing surface defined by the corresponding deteetion zone 8] in the quadrant for a total of 36 different positions inclusive of all 36 zones 81 in the quadrant yielding the preferred 36 independent simultaneous linear equations containing all 36 unknowns representative of the quantity of radioisotope in each of the 36 zones 81 in the quadrant, this procedure being repeated for each of the other quadrants to yield the preferred 144 independent simultaneous linear equations containing all l44 unknowns representative of the quantity of radioisotope in each of the I44 zones 81 in the entire brain volume.
Preferably, as was previously mentioned, the headgear apparatus 66 is designed for the largest typical adult human head likely to be encountered and, therefore, variations in the air space between the detector assemblies 95 and the scalp of the patient may occur.
This variable air space is due to variable cranial curvatures associated with different patients as well as variations in the size of the head ln order to compensate for these variations in air space, if desired. which variations affect the quantity of the radiation emissions detected by the detector assemblies, a gauge-type means may preferably be included in the helmet housing 82 substantially at the position of each of the detector assemblies which measures the size of the air space between the associated detector assembly and the scalp of the patient. These gauges (illustrated by way of example in the alternative arrangement of FIGS. 2A and 28 by ref erence numerals 130, 132, I34, 136 and 138) may be calibrated rods or, if desired, may be conventional position transcucers so as to provide a contemporaneous measurement of the air space between the associated detector assembly and the scalp of the patient during the data collection period. If desired, a separate headgauge, comprising solely position transducers or cali brated rods located at each of the detector assembly positions in an identical array with that of the helmet housing 82, may be utilized prior to the use of the head gear apparatus 66 so as to provide these air space measurements prior to the actual data collection period. ln this instance, the headgear apparatus 66 is preferably not provided with a gauge at each detector assembly position, but rather is preferably merely provided with only one calibrated rod or position transducer gauge in an identical position with that of such gauge in the head-gauge in each of the portions or planes ofthe helmet housing 84 through 92, inclusive. illustratively shown by gauges through 138, inclusive, so as to insure that the helmet housing 82 is aligned in the identical position in which the head-guage was aligned when the air space measurements were made. The universal mounting means 68, as was previously mentioned. is preferably capable of aligning the headgear apparatus 66 in this desired alignment or orientation position and preferably includes locking means (not shown) for retaining it in this position during the data collection period. If a separate head guage device is utilized to measure the variable air space, as was previously mentioned. then this device should preferably be similarly oriented in the manner previously described with reference to the orientation of the headgear apparatus 66 prior to the taking of these measurements.
DETECTOR ASSEMBLY Now describing the radiation emission detector assembly in greater detail, the detector assembly preferably being functionally identical with detector assembly 65 of the Xenon gas analyzer 63 as was previously mentioned except differing in the associated field of view and single channel analyzer window or settingv A typical preferred detector assembly 95a or 95c is shown in FIGS. 2A and 8, respectively, this detector assembly being functionally identical with the P and Q group of detectors except differing in the illustrated associated field of view, as was previously mentioned, and the as sociated single channel analyzer window or setting. Accordingly, for purposes of explanation, the structural details of the preferred detector assemblies 95, as well as the P and Q groups of detectors, all of which are preferably functionally identical, will be described hereinafter only with reference to detector assembly 95a. The detector assembly 95a preferably includes a collimator portion 140 made of radiation absorbing material cooperatively associated with a radiation emission detector 142. Preferably. the collimator portion 140 of the detector assembly 9511 is a tube of we cific length and cross-sectional lumen area whose function is to limit the direction from which radiation may enter the detector 142. It should be noted that it is the collimator 140 which primarily determines the degree of spatial resolution of the system 10 of the present in vention. In the preferred system 10 of the present invention, the shape of the collimator lumen [44 is preferably square so as to maximize the area ofthe detector 142 utilized for viewing. although other shapes, such as circular, may be utilized therefor if desired. The collimator tubing 140 is formed of a material which has a high radiation absorption characteristic, such as a composite material composed of 90 percent tungsten plus 10 percent nickel and copper (erg. Kullite 112 or Mallory I000), so as to sharply restrict the radiation emission to the lumen-area of the detector 142.
The length of the collimator 140 determines the volume of tissue observed at the detector 142 through the lumen 144 of a given cross-sectional area and, there fore, determines the number of detection zones 81 observed by the detector 142 in its direction of viewv The preferred square collimator lumen 144 provides a square prism field of view for the detector 142, each detector 142 arrayed about the helmet housing 82 pref erably looking into the face of the cubic detection zone 81 associated therewith, as previously described, and seeing its own prism of tissue. The collimator 140 length is preferably chosen so that the prismatic base of the field of view for the P group of detectors is sub stantially only slightly smaller in area than the area of a cube face ofa detection zone 81 at the distal side of the third detection zone 81 in depth in a straight line based from the detector assembly in that direction or field of view and substantially only slightly smaller in area than the area of a cube face of a detection zone 81 at the distal side of the second detection zone 81 in depth in a straight line based from the detector assembly in that direction or field of view. In addition, with the preferred collimator 140 of the present invention, the square prism field of view for the associated detectors 142 located at an angle to the detection zone viewing surface (FIGS. 9E, 10E, 11E, 12D.13D,14C,18B, C and 228) preferably includes portions of the detection zones 81 in a two-by-two, four-zone matrix array for the Q group of detectors and a three-by-three, nine-zone matrix array for the P group of detectors, with the exceptions illustrated in FIGS. 22B and 238, about the axis of view of the associated detector. The preferred collimator 140, therefore, provides a preferred effective field of view substantially three detection zones in depth for the P group of detectors and two detection zones in depth for the Q group of detectors, with a proportionate amount of such zones 81 which are in a three-by-three matrix array or two-by-two matrix array with the zone at the point of "entry" into the volume for the angled P group of detectors and Q group of detectors. respectively. Therefore, as will be explained in greater detail hereinafter, the radiation emission quantified or detected by a given detector as sembly will be a resultant value dependent on the con tributions to this resultant provided by the respective radiation concentrations of the radioactive gas in the portions of the detection zones 81 contained within the preferred effective field of view.
This preferred optimal choice for the collimator 140 for the P and 0 groups of detectors is based upon the useful depth of view; that is, beyond a given depth of tissue, for the gamma energies of the preferred radioisotope Xe the tissue attenuation and dispersion of radiation emission according to the inverse-square law only allows a negligible fraction of radiation emission to actually arrive at the detector 142 from these depths. Therefore. radiation emissions from beyond three detection zones 81 in depth for the P group (81 Kev) and for the 0 group 31 Kev) are outside the effective fields of view as they do not significantly affect the determinations of local cerebral blood flow. If it is desired to take these negligible amounts into consideration, the collimator 140 may be designed so as to include these additional detection zones in its effective field of view.
As was previously mentioned, the detector 142 of the detector assembly a or the P or Q groups is preferably a semiconductor such as ultra high purity germa nium. Such a solid-state detector 142 is particularly applicable to the detection of gamma radiation, such as the gamma radiation associated with the preferred gas Xenon (3l Kev and 81 Kev) with high energy resolution, good stability and efficiency over the entire (31 Kev to SI Kev) energy range. This is of particular con cern since gamma radiation is the preferred radiation emission to be detected, beta radiation not penetrating the skull of the patient. Such a preferred solid-state detector 142 absorbs the gamma radiation and generates secondary electrons in response thereto which signal is processed in the radiation detection signal processing circuit 67 associated with each of the detector assemblies. If desired, a crystal detector, such as a crystal of thallium-activated sodium iodide may be utilized in place of the solid-state detector. However, as shown and preferred in FIG. 6, when such a crystal is utilized, a photomultiplier 146 should preferably be utilized therewith due to the fact that such a crystal emits a very weak light photon when an energy quantum of the gamma radiation is absorbed by the crystal. the photomultiplier 146 converting this light photon into electri cal energy and amplifying it to a transmittable level in the form of electrical pulses. The crystal detector 142, in such an instance, is preferably interposed between the collimator and the photomultiplier tube 146 so as to receive a gamma radiation emission input from the collimator I40 and provide an output to the photo multiplier tube 146 in response thereto.
SOLID STATE DETECTOR RADIATION DETECTION SIGNAL PROCESSING CIRCUIT Now describing in greater detail the typical radiation detection signal processing circuit 67, shown in greater detail in FIG. 3, which is preferably utilized for each of the detection channels in which the preferred solid state detector is utilized as the detector 142. As was previously mentioned, the same processing circuit 67 configuration is utilized for the Xenon gas analyzer radiation detection signal processing circuit 74. Therefore, a detailed description of this circuit is deemed unnecessary.
Referring now to FIG. 3, the output of the solid state detector assembly, 95 for the alternative arrangement of FIGS. 2A and 2E or for the P or Q group of detectors for the preferred arrangement of FIG. 923, for a typical detection channel or processing circuit 67 is preferably operatively connected through a conventional charge sensitive pre-amplifier stage 148 and conven tional linear amplifier stage 150 to a conventional sin gle channel analyzer or pulse height detector stage 152, such as a circuit comprising transistor-transistor logic (TTL), to preferably provide a TTL type output signal to the minicomputer buffer 60 which is indicative of the detected count in that detection channel, which signal is subsequently preferably fed to the general purpose computer 61. Since the three stages 148, I50 and 152 comprising the preferred radiation detection signal processing circuit 67 are preferably conventional, a detailed description is deemed unnecessary. Suffice it to say that, for the P and O group of detectors, dependent on the energy level (31 Kev or 81 Kev) to be detected the single channel analyzer window is set for this level by conventionally gating at the selected energy level and discriminating against all other levels.
CRYSTAL DETECTOR RADIATION DETECTION SIGNAL PROCESSING CIRCUIT If a crystal detector 142 fabricated of thalliumactivated sodium iodide is utilized in place of the preferred solid-state detector then the radiation detection signal processing circuit configuration shown in FIG. 6 is preferably utilized for each of the detection channels in place of the detection signal processing circuit 67 configuration previously described with reference to FIG. 3.
Referring now to FIG. 6, the photomultiplier tube 146 for a typical detection channel in which a thallium activated sodium iodide crystal is utilized as the detector 142 is preferably a conventional lO-stage photomultiplier having an individual voltage dividing net work, such as is provided by an RCA type 7767 multiplier phototube which has ten electrostatically focused dynode stages with in-line arrangement thereof. In the system of the present invention, such a tube 146 is preferably magnetically shielded. A conventional high voltage power supply 170 is connected to the photomultiplier tube 146 via a conventional gain adjustment potentiometer 172 having an adjustable wiper arm 173 so as to supply the required voltage necessary to excite the photo-cathode to its range of sensitivity to light. If desired, a single high voltage power supply l70 may be utilized to supply power to any number of photomultiplier tubes 146 in the detector assembly array. In such instance, a different gain adjustment potentiometer 172 is preferably provided for each tube 146 so as to properly adjust the voltage for each of the individual detectors 142. The output of the photomultiplier tube 146 is connected through a conventional pre-amplifier 174 and amplifier 176 to a conventional discriminator stage or variable pulse height detector stage 178 whose function is to eliminate all pulses or voltages below and above a specified level so as to eliminate low voltage electrical noise and background counts or emissions outside of the energies to be detected. In addition the discriminator stage 178 preferably is provided with conventional means 179, such as a variable potentiometer, for shifting the discriminator to another threshold level which would be appropriate for a different radioisotopic gas as well as to the optimal levels for the preferred radioisotopic gas Xenon. The output of the discriminator stage 178 is preferably fed through a conventional summing gate and pulse stretcher network 180 whose output is preferably fed to the general purpose computer 61 via the minicomputer buffer 60, if desired.
LOCAL CEREBRAL BLOOD FLOW DETERMINATION METHOD General Description Now generally describing the preferred method of the present invention in accordance with which the local cerebral blood flow determination system 10 previously described is preferably utilized. In any local ce rebral blood flow determination there are significant physical factors which affect this determination but which may be considered to be constants for a given patient whose local cerebral blood flow is being determined. However, these significant constants must preferably be initially determined for the patient. In utiliz ing the system 10 in accordance with the preferred method of the present invention, the local cerebral blood flow is ultimately determined in accordance with the well known Fick-Kety expression of which one form is where C,- equals the brain tissue radioisotope concentration preferably in microcuries per gram; (C,,C,.) equals the arterialvenous blood radioisotope concentration difference, preferably in microcuries per gram; m, equals the starting time of the local cerebral blood flow determination, preferably in seconds, F equals the local cerebral blood flow, preferably in microliters per gram per minute; W equals the weight of the tissue comprising the detection zone for which the local cerebral blood flow is being determined; and M1 equals the number of seconds after the beginning of the first inhalation by the patient during the data collection period.
As will be explained in greater detail hereinafter, this Fick-Kety expression is preferably utilized for each of the detection zones 81 in which it is desired to determine the local cerebral blood flow therefor. Furthermore, it should be noted that the local cerebral blood flow determined in accordance with the preferred method of the present invention is preferably a quantitative determination of cerebral blood flow for that detection zone 81 for which the Fick-Kety expression is utilized (as opposed to a mean rate of flow for a region) which, as will be explained in greater detail hereinafter, is possible because the various factors in the FickKety expression can be localized for each detection zone 81 in accordance with the preferred method of the present invention. C,- and C,- are related by the expression C,- (venous concentration) C, (brain tissue concentration)+'r (tissue/blood partition coefficient for the gas), and the value of 1' can be determined. Furthermore, since the specific gravity of brain tissue is approximately equal to one, the weight of the tissue W is approximately equal to the volume of the detection zone 81, which is also a known factor as it is predetermined by the volume of the model which defines each of the detection zones 81. Therefore, the only unknowns involved in applying this Fick-Kety expression to the determination of a local cerebral blood flow is the tissue radiation concentration quantity C,- and the arterial biood radiation concentration quantity C It is in de- 19 termining these quantities for a preselected discrete de tection zone 81 that the preferred method of the present invention is most useful.
DETAILED DESCRIPTION Now describing the preferred method of the present invention in greater detail. in utilizing the system of the present invention in accordance with the preferred method of the present invention, the headgear apparatus 66 is preferably aligned on the patients head so as to provide a specific and reproducible relationship to the patients brain. As was previously mentioned, the headgear apparatus 66 is preferably aligned so as to be oriented as shown in FIG. 5 with the glabellarsellar plane being the basal plane for the detector array of the rectangular helmet 82.
When the headgear apparatus 66 is so aligned, the other two preferred reference planes ofthe three reference planes which define the orientation or alignment of the headgear apparatus 66 are, as was previously mentioned, the midsaggital plane and the plane perpendicular to the basal plane and the mid-saggital plane and tangent to the forehead, and the detection zones 81 of the patient's brain are defined as shown in FIG. 5 by the dotted lines,
As was previously mentioned, the length of air space between the associated detector assembly and the portion ofthe patients scalp opposite that detector assembly affects the radiation emission detected thereby. Therefore, it is both preferable and desirable that the length of air space between a particular detector assembiy and the skull of the patient be known for purposes of the local cerebral blood flow determination. Accordingly, if a separate head-guage is utilized to measure the lengths of these air spaces, as was previously mentioned, the gauges 130 through 138, inclu sive. located in the helmet housing 82 must be checked to insure that the patients head retain exactly the same position during the actual data collection period as during the time when the air space measurements were made.
In addition, compensation should also preferably be made for the thickness of the patients skull and the degree of bone mineralization as these factors may introduce errors in the measurements of radiation emission by the P and Q detector assemblies. These factors may be determined by ultrasonic techniques such as are employed for neurologic echo-scanning. However, if the errors introduced by variations in bone mineralization are tolerable. then no additional compensation need be made for this factor If the errors introduced by variations in skull thickness can be tolerated. no additional compensation need be made for this factor either. Cor rections for mean skull thickness and density are then sufficient.
Once the headgear apparatus 66 is aligned in the preferred predetermined orientation or alignment position, the balance of the local cerebral blood flow determination may be performed in accordance with the preferred method of the present invention. The radioactive gas, which preferably has at least two gamma energy levels, such as Xenon preferably, although other gases such as Krypton 85 or gaseous mixtures of two radioactive gases each having a different gamma energy level may be utilized if desired, is then dissolved in the patients blood by means of the closed inhalation-exhalation breathing system associated with the mask 12. Throughout the actual data collection period, the patient receives his air supply through this closed inhalation-exhalation breathing system. When the actual data collection is to commence, that is preferably when the patient's breathing conditions, such as oxygen and carbon dioxide pressure, are stabilized as indicated by the carbon dioxide and oxygen analyzers 76 and 77, the minicomputer buffer preferably sends a signal to the end-tidal sampler 62 which activates the sampler 62 to turn it ON so as to preferably synchronize the initiation of end-tidal sampling and the resultant radiation emission detection or counting times in the Xenon' gas analyzer detector assembly 65 with that of the p and Q detector assemblies associated with the headgear apparatus 66 detector array. After this initial activation, the end-tidal sampler 62 functions in a conventional manner throughout the duration of the data collection period, and its operation will not be described in greater detail hereinafter. Suffice it to said that the end-tidal sampler 62 relay mechanism (not shown) is actuated at the beginning and end of each breath by airway pressure changes within the closed breathing system resulting from these conditions; the end-tidal sampler 62 functioning to extract end-tidal samples when activated by the gas pressure within the mask 12 reaching a given pressure at the end of the breath and continuing to sample until negative pressure arises at the beginning of the next breath.
Subsequent to the initial turning ON of the endtidal sampler 62, signifying the beginning of the actual data collection period, the actuation of the end-tidal sampler 62 relay at the end of the patients next breath actuates the Xenon gas injector 40 solenoids (not shown) associated with valves 42 and 50. Gas injector 40 thereafter injects a predetermined amount of the radioisotopie Xenon gas into both the inhalation and the exhalation chambers 28 and 30, respectively, through the respective intermittently opened valves 42 and 50, which thereafter close, and associated conduits 44 and 48, respectively. The injected radioisotopic gas mixes with the air already contained in the chambers 28 and 30.
As the patient continues to breath normally, the airway pressure in inhalation conduit 16, and the closed breathing system as a whole, changes and demand valve 20 opens at the beginning of the next breath. The end-tidal sampler 62 relay is again actuated in response to this initial pressure change, which preferably sends an initial signal to the minicomputer 60, or if desired, directly to the measuring devices; namely, the radiation detection signal processing circuits 67 and 74, placing them in the data collection mode. This signal is preferably only sent at this initial time, the circuits 67 and 74 remaining in this mode throughout the balance of the data collection period. In addition. a signal is also sent to the gas purification unit 32 which signal activates the unit 32 so as to refill the inhalation chamber 28 from the exhalation chamber 30 via the gas purification unit 32 which removes excess carbon dioxide, adds oxygen to the closed breathing system, and filters the air contained thereon.
Each subsequent breath of the patient during the data collection period causes the actuation of the end tidal sampler 62 relay mechanism at the beginning of the breath; the resultant operation of gas injector 40 to inject a predetermined quantity of radioisotopie Xenon gas only into the exhalation chamber 30,
valve 50 preferably remaining closed throughout the balance of the data collection period; the actuation again of the end-tidal sampler 62 relay mechanism at the end of the breath; and the resultant operation of gas purification unit 32 to return the injected gas well mixed to the inhalation chamber 28 to be breathed by the patient. In this manner. the radioisotopic gas concentration in the blood is caused to rise steeply since each breath contains a still higher concentration of the radioisotopic gas.
Arterial Blood Radiation Concentration Determination Preferably, as was previously mentioned, in order to determine the arterial blood radiation concentration quantity C,, in accordance with the preferred method of the present invention, the alveolar or end-tidal gas concentration at the end of each breath is sampled in the manner previously described above. This end-tidal gas sample is drawn through the coiled passageway 64 which, as was previouisly mentioned, contains a known total volume of the sampled gas. Detector assembly 65 looks" at the coiled passageway 64 and preferably measures the gamma radiation emission emanating therefrom, which is equivalent to the gamma radiation emission in that known sample volume. The gamma radiation emission entering collimator 140 which is incident on detector 142 is absorbed and causes the gener ation of secondary electrons which are amplified in the radiation detection signal processing circuit 74 to pro vide a pulse or signal which is the equivalent of the resultant radiation emission count of detector assembly 65. Since the lung gas in contact with the blood has essentially the same partial pressure as the gas contained in arterial blood, and since it is this equilibrated gas which appears in the last part of each breath, the gas concentration in the last part of each breath is essentially equivalent to the concentration of the radioisotopic gas in arterial blood. This signal, which therefore is representative of the arterial blood radiation concentration, is then preferably fed to the minicomputer buffer 60 for temporary storage before being passed to the general purpose computer 61 for processing along with the other collected data.
As will be explained in greater detail hereinafter, each time an end-tidal sample is taken and the resultant arterial blood radiation concentration signal produced in accordance therewith, a point on an arterial concentration curve is created. Most preferably, the data collection period duration is one minute and six determinations at l second intervals m and m,, m m;,,m,, and m as shown in FIG. 4, are made so as to construct an arterial concentration curve for purposes of the local cerebral blood flow determination in accordance with the Fick-Kety expression, this arterial blood radiation concentration quantity being represented by the term C,,.
BRAIN TISSUE RADIATION CONCENTRATION DETERMINATION Now referring to the operation of the system in ac cordance with the preferred method of the present invention so as to provide the brain tissue radioisotope concentration quantity C,- for, preferably, each of the detection zones 81 so that this data may be processed in accordance with the previously stated Fick-Kety expression. The radioactive gas, which as was previously mentioned dissolves in the blood passing through the lungs, is delivered to the brain parts or detection zones 81 in proportion to the local rate of blood flow. The resultant gamma radiation emissions detected by the various P and Q detector assemblies included in the headgear apparatus 66 detector array are a function of the amount of radioactive gas deposited. or radioisotope concentration, in the various detection zones 81 and the effective field of view associated with a given detector assembly.
As was previously described, in the preferred method of the present invention, the I44 blocks or zones 81 which describe the brain volume are preferably subdivided into four equal 36 Zone quandrants of 3X3X4 blocks or zones 81 each which are each preferably similarly solved for the valves of the radioisotope concentration in each of the zones 81 by solving four sets of 26 linear simultaneous equations, one set for each quandrant, yielding the radioisotope concentrations for each of the 144 detection zones 81. In the alternative, if desired, the brain volume can be looked at in its entirety as a whole and 144 simultaneous equations for the radioisotope concentrations in the 144 various zones 81 be solved as a group. The P and Q detector assemblies look at each quandrant of the brain from 36 different positions (or locations), for a total of 144 different positions (or locations), and the degree of resolution or localization observed by each P and Q detector assembly, that is, its effective field of view, is determined by the collimator associated with the particular P or Q detector 142. In order to properly determine the rate of local cerebral blood flow, the radioactive gas must be administered to the patient in sufficiently high concentration to provide a statistically acceptable counting rate, or quantity of detectable radiation emission, in the preferred sampling interval which is 10 seconds, but not so high that the patient is endangered by the radiation. As was previously mentioned, the preferred square lumen collimator 140 has a length which provides the preferred substantially three detec tion zone 81 useful depth of view for the preferred ef fective field of view for a P group detector and two de tection zone 81 useful depth of view for a Q group de tector with a proportionate amount of such zones 81 which are in a three-bythree matrix array or two-bytwo matrix array with the zone at the point of entry" into the volume for the angled P and 0 groups of detectors, respectively. This preferred collimator 140, which preferably differs in length for the p and 0 groups of detectors, respectively, represents the optimal compro' mise between an acceptable level of radioactive gas isotope administration and the collimator size suitable for localization. Each preferred P or Q detector 142, therefore, sees its associated preferred effective field of view of either three or two detection zones in depth, respectively, the total or resultant radiation emission observed by each detector 142, as was previously mentioned, being due to the radioisotope concentration present in the portions of each of the detection zones 81 contained within the effective field of view of the particular detector 142.
This relationship between the resultant radiation emission detected or observed by a given detector 142 and the radioisotope concentration present in the portions of each of the detection zones 81 contained within the preferred effective field of view of the detector 142 may be preferably defined by the following expressions in which we have assumed, for purposes of explanation, that all the detection zones 81 are included within the bone surrounding the patients brain, although it would be obvious to one of ordinary skill how to readily modify the following expressions to take into account the surrounding bone if the detection zones were defined so as to include this bone which, in such instance, could be considered as non-radioactive source bearing material. Before defining these expressions, however, a nomenclature for the detection zones 81 including the effective field of view, preferably must first be established.
For pusposes of explanation, the 36 detection zones for the typical quadrant illustrated in FIGS, 923, which is the front-right quadrant, are labeled 2-] through 2-36, inclusive, with the uppermost horizontal plane 3X3 matrix array labeled Z-l through 2-3 in the outermost row, 2-4 through 2-6 in the next adjacent row and 2-7 through 2-9 in the innermost row. Similarly the next adjacent horizontal plane matrix array is labeled, 2-H] through 2-12 in the outermost row, 2-13 through 2-15 and 2-16 through 2-18; the array in the plane below being labeled 2-19 through 2-21, 2-22 through 2-24, and 2-25 through 2-27, respectively; and the array in the lowermost plane being labeled 2-28 through 2-30, 2-31 through 2-33 and 2-34 through 2-36, respectively. The P and O detectors are each labeled with a subscript containing the number of the zone 81 through which the brain is "looked into or entered and a letter indicating the direction of entry", "G" indicating entry from the top of the 36 zone array, C indicating entry from the right side, F indicating entry on a preferably 45 angle from the right side above towards the left side below, A indicating entry from the front, H indicating entry on a preferably 45 angle from the front above towards the rear below and B indicating entry on 21 preferably 45 angle in the horizontal plane from the right towards the left.
For purposes of explanation, based upon the preferred effective fields of view, and the preferred planes of orientation of the headgear apparatus 66, the given P or Q detector assembly, for which the expressions are defined, shall be assumed to be located, in the most general case, adjacent the center of the outermost surface of the zone 81 through which the appropriate P or Q detector assembly looks into the brain volume, those P and O detector assembly which look into the brain volume at an angle generally being assumed to be located adjacent the center of the line of intersection of the planes of the two normal adjacent outermost surfaces of the zone 81 through which the appropriate angled P or Q dectector assembly looks into the brain volume. Furthermore, as was previously mentioned, the
effective field of view ofeach P or Q detector assembly is preferably symmetrical about the axis of view, the field of view including a proportionate amount of each zone which can be seen. For purpose of explanation, the proportionate amounts of the zones 81 included within the effective field of view ofa Q detector assembly are generally designated for the first zone in depth and F' for the second zone in depth for the nonangled detectors, and "13" for the first zone in depth and Zf' and 'F for the portions of the second zones in depth for the angled detectors Similarly, the proportionate amounts of the zones 81 included within the effective field of view of a P detector assembly are generally designated 5" for the first zone in depth, F for the second zone in depth and E for the third zone in depth for the non-angled d etectors, and Z for the first zone in depth, 2* and f for portions of the second zones in depth and F and F for portions ofthe third zones in depth for the angled detectors. Furthermore, for certain of both the P and Q detector assemblies, a proportionate amount of an additional portion of another zone within the zone adjacent the first zone in depth within the effective field of view is generally designated If. The ggieral letters 5", 5", F, J, and are factors for each appropriate zone 81 related to the efficiency e of the detection system; the fraction 5" of radioisotope remaining since its original assay at the time of the data collection for the local cerebral blood flow determination (loss due to radioactive decay); the factor s expressing the attenuation of radiation by the scalp and skull; the area A,, of the detector; the dimensional constant K", the total attenuation constant u for gamma radiation of a specific energy (in square centimeters/gram); the total thickness r of brain tissue through which the radiation is attenuated (in grams/square centimeter); the distance I from the center of the particular zone to the detector along the central axis of view; the volume of the zone Vf; the collimator factor ,7 for a given zone, which is the projected area of the detector divided by the total area of the detector as seen through the collimator; VP zone or VQ zone represents the integration limit of the total volume of the particular P or Q zone; and l zone", representing the radiation concentration (in microcuries) of radioisotope contained within a given zone; P subscript" and Q subscript representing the detected resultant radiation emission or counting rate resulting from the contributions of all the detection zones 81 within the effective field of view of the particular P or Q detector assembly, where the following general expressions define these general letters as follows with the term C being a constant equivalent to the expression .tPZmE ztzove
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|U.S. Classification||600/431, 250/303|
|International Classification||A61B5/08, G01T1/29, A61B5/0275|
|Cooperative Classification||G01T1/2985, A61B5/6803, A61B5/0813, A61B5/02755, A61B5/416|
|European Classification||A61B5/68B1B, A61B5/41J4, G01T1/29D4, A61B5/0275B, A61B5/08N|