SYSTEMS AND METHODS FOR CARDIOVASCULAR IMAGING, CARDIOVASCULAR FUNCTIONAL ANALYSIS, AND DETECTING CORONARY ARTERY DISEASE

Abstract: One aspect of the present invention relates to screening patients for an early stage of coronary artery disease. According to this method, a patient is screened based on the time-activity curve for a radioactive tracer passing through a left ventricle region of the patient's body. According to another aspect of the invention, an array of gamma particle detectors is employed to obtain data for a region of interest taht is larger than and encompasses a left ventricle region of the patient's body. An analysis of the data identifies the subset of the region of interest that corresponds to the left ventricle region. According to a further aspect of the present invention, a second technique is employed to locate the left ventricle region. A still further aspect of the present invention relates to high temporal resolution gamma cameras.

Description: TITLE Systems and Methods for Cardiovascular Imaging, Cardiovascular Functional Analysis, and Detecting Coronary Artery Disease. TECHNICAL FIELD The present invention relates generally to medical diagnostic and screening systems and methods. In particular, the present invention relates to non-invasive cardiac imaging and functional analysis systems and methods. BACKGROUND ART While cardiac imaging and functional analysis is the largest single nuclear medical imaging application, there remains a tremendous unmet need for improved cardiac imaging and functional analysis systems and methods. This need is exemplified by the fact that historically, for 30%-50% of those stricken with coronary artery (occlusive) disease (CAD), the first symptom of the disease is death. This has motivated considerable effort to develop diagnostic methods and apparatus to detect CAD prior to the onset of fatal symptoms and assist in the development and implementation of preventive measures. Two strategies are presently used to reduce morbidity and mortality from CAD. The first involves screening for modifiable cardiac risk factors, such as hypertension, elevated serum cholesterol, cigarette smoking, physical inactivity, and diet. The second involves early detection of CAD. The principal tests for detecting CAD include resting and exercise ECGs, which can reveal the presence of myocardial infarctions and inducible myocardial ischemia. Tc-99m myocardial perfusion and computed tomography (CT) classification scoring can provide visual evidence of plaques in the coronary arteries. Thallium-201 scintigraphy, exercise echocardiography, and ambulatory ECG (Holter monitoring) are less commonly used for screening purposes. None of these strategies has produced a solution to the high incidence of death due to undetected CAD. Accordingly, there remains an unsatisfied need for improved cardiac imaging and functional analysis systems and methods for reducing morbidity and mortality from CAD. First-Pass Radionucleide Angiography (RNA) is used to provide information about cardiac function such as regional ventricular wall motion just at the peak of maximum exercise stress, left and right ventricular ejection fraction, regional contractivity, and left to right shunt quantitation. Presently, the nuclear medicine health care sector can perform a first-pass RNA diagnostic protocol using the commercially available dedicated cardiac imager multicrystal scanner model SIM400 from Picker International of Cleveland, Ohio, or a single crystal Anger gamma camera. The single crystal Anger camera has numerous disadvantages when used for first-pass RNA related to its low count rate capability including pulse-pileup and low count density. A low-count density image limits the ability to define changes in wall motion. The SIM400 scanner uses a 2.54 cm thick Nal(TI) crystal divided into 400 detector elements in a 20 x 20 matrix. Surrounding each detector element is reflective material. Every second row is partially cut through the crystal to create a scintillation bridge between two adjacent elements. The scintillation light is detected by 115 3/4" diameter bialkali tubes and each photomultiplier tube (PMT) detects scintillation from two optically decoupled detector elements. Variations in pulse output from adjacent PMTs viewing the doublet detector element provides the photon spatial information. In practice, clinical interpretations of the SIM400 scanner treadmill stress acquisitions are difficult to analyze. Patient motion during peak stress acquisition can cause image artifacts that can trigger an incorrect diagnosis. To counteract this problem Picker International uses an Am-241 source placed on the patient's chest. Dual energy windows are utilized during treadmill stress acquisitions and the Am-241 energy window is used to correct for patient motion. Unfortunately, this method only corrects for up/down/side planar motion (i.e., motion in a plane, and motion in any direction outside the defined plane is uncorrected). If the patient rotates along the z-axis, the planar patient correction can magnify image artifacts. Gamma cameras with compact photomultiplier tubes (PMTs) are described in literature. However, the prior art instruments were slow, designed primarily for relatively low rate (up to 5 kHz) breast imagers. The array concept was conceived and first tested in a small laboratory prototype by Dr. Roberto Pani and his group in University "La Sapienza" in Rome, Italy. At the time of this application, there are three companies in the United States that are developing products based on this mature and reliable concept: Gamma Medica Instruments ( " www.gammamedica.com " ) of Northridge, California, Dilon Technologies (www.dilon.com) of Newport News, Virginia, and PEM Technologies, of Bethesda, MD. Initially, imaging of breast cancers was emphasized, taking advantage of the small size of the cameras that allows flexibility in positioning the detector for better localization and visualization of breast lesions. One company, Digirad (www.digirad.com) of Phoenix, Arizona, produces heart imagers based on not yet technically mature and expensive solid state technology (CsI(Tl) scintillator and PIN diode arrays). The only high count-rate dedicated heart imager (developed by Proportional Technologies, Inc., of Houston, TX) is based on a high pressure wire chamber concept and therefore has its energy range practically limited to less than 100 keV. Also, its rate capability is in fact limited by the physical nature of radiation interaction with a gas detector medium in a two-step process. Related teachings and prior art include the following: [1] Dr. Roberto Pani's papers on multi-PSPMT imagers: a) Single Photon Emission Imaging by Position Sensitive PMT. Pani R., Pellegrini R., Soluri A., De Vincentis G., Scafe R., Pergola A, Nucl. Instr. & Meth. 1998. A 409: 524-528. b) Multi PSPMT Scintillating camera. Pani R., Soluri A., Scafe R.,Pergola A., Pellegrini R., De Vincentis G.,G. Trotta, Scopinaro F.. 1997 IEEE Nuclear Science Simposium, Conference Record vol.2 pg. 1068-1072, November 9-15,1997 Albuquerque, New Mexico USA. c) Multi PSPMT Scintillating camera. Pani R., Soluri A., Pergola A., Pellegrini R., Scafe R., De Vincentis G., Scopinaro F. IEEE Trans, on Nuclear Science 46 N.3 ; 702-708, 1999. [2] New Developments in Portable Gamma Cameras. Joyce Ward, ADVANCE for Radiologic Science Professionals, page 16-17, August 3, 1998. [3] US Patent 6,091,070: Semiconductor Gamma-Ray Camera and Medical Imaging System, Lingren C. L. et al., July 18, 2000. [4] US Patent 4,458,688: Method and Apparatus for Cardiac Nuclear Imaging, Von Behren P. L., July 10, 1984. [5] US Patent 4,999,501 : High speed multi-wire photon camera, Lacy J. L., March 12, 1991. [6] US Patent 5,753,917: Dual crystal scintillation camera, Engdahl J. C. May 19, 1998. [7] US Patent 5,431,161 : Method and apparatus for information acquisition, processing, and display within a medical camera system, Ryals Carl J. et al., July 11, 1995. [8] US Patent 5,377,681 : Method of diagnosing impaired blood flow, Drane Walter E., January 3, 1995. [9] US Patent 5,249,124: Multi-isotope imaging using energy-weighted acquisition for, e.g., myocardial perfusion studies, DeVito; Raymond P., September 28, 1993. [10] US Patent 5, 199,438 : Measurement of cardiac performance, Pearlman Andrew L. April 6, 1993. [11] Scintillation products from Saint Gobain Crystals and Detectors (formerly: Bicron Corporation), Newbury, Ohio. [12] Compact and Flat Panel photo-multipliers from Hamamatsu Corporation, Bridgewater, New Jersey. [13] Model 85001 000 photo-multiplier from Burle Industries, Inc, Lancaster, Pennsylvania. There remains an unsatisfied need for high-performance heart imaging technology. In particular, there is an unsatisfied need for dedicated nuclear cardiac imaging equipment that is simpler in design, easier to manufacture, and easier to use than the prior art. A need also exists for such simple, low cost equipment that can perform first pass imaging, positron imaging, quantitative myocardial perfusion measurements, multi-gated equilibrium pooled blood (MUGA) imaging and coronary transit-time screening. Further, there is a need for portability of such a device. DISCLOSURE OF INVENTION The following presents a simplified summary of the invention in order to provide a basic understanding of some of its aspects. This summary is not an extensive overview of the invention and is intended neither to identify key or critical elements of the invention nor to delineate its scope. The primary purpose of this summary is to present some concepts of the invention in a simplified form as a prelude to the more detailed description that is presented later. One aspect of the present invention relates to a method of screening patients for ari early stage of CAD. According to this method, a patient is screened based on the time-activity curve for a radioactive tracer passing through a left ventricle region of the patient's body. The time-activity curve can be analyzed by comparing it to a time-activity curve for a region of the patient's body downstream of the left ventricle, such as the patient's ascending aorta or the patient's brain. A quick transit through the patient's left ventricle region indicates an early stage of CAD wherein the free volume of the patient's coronary artery system is reduced, but the patient does not necessarily yet suffer from inducible myocardial ischemia. Another aspect of the present invention relates to a method for obtaining a time-activity curve for a radioactive tracer passing through a left ventricle region of the patient's body. According to this method, an array of gamma particle detectors is employed to obtain data for a region of mterest which is larger than and encompasses a left ventricle region of the patient's body. An analysis of data from the detector array identifies the subset of detection data that corresponds to the left ventricle region. This method avoids difficulties associated with accurately locating the left ventricle region prior to a testing or screening procedure. A further aspect of the present invention relates to another method for obtaining a time-activity curve for a radioactive tracer passing through a left ventricle region of the patient's body. According to this method, a second technique, such as ultrasound, MRI, x-ray, computed tomography (CT), planar nuclear medicine, positron emission spectroscopy (PET), single photon emission computed tomography (SPECT), or a second radioactive tracer, is employed to locate the left ventricle region. A time activity curve is then obtained using a gamma particle detector or detector array positioned according to the determined location of the left ventricle region. Potential errors resulting from failure to accurately locate the left ventricle region are thereby avoided. A still further aspect of the present invention relates to a method of obtaining images of a patient's heart. According to this method, the images are obtained with a high temporal resolution gamma camera. A high temporal resolution gamma camera provides diagnostically meaningful images of the patient's heart even when the high temporal resolution comes at the expense of spatial resolution. Imaging is further improved with the use of a second gamma camera at an angle to the first gamma camera. Two gamma cameras can be used to provide three-dimensional structural information, measure heart motion in three dimensions, and be used to correct for bulk movements of the patient motion in three dimensions. A still further aspect of the present invention relates to a dedicated cardiovascular imaging and functional analysis systems. The systems provide dedicated fast, sensitive, compact and economical gamma cameras that are especially suited for heart imaging. Features include pixellated scintillation detectors, fast compact photo-multipliers, a matching high efficiency light guide system, and very fast readout and data processing systems. Cardiovascular imaging systems of the present invention can be used as dedicated nuclear cardiology small field of view imaging cameras. Embodiments of the present invention for dedicated nuclear cardiographic imaging can compete with magnetic resonance imaging (MRI), computed tomography (CT), and echocardiography; all of which are slowly adopting nuclear cardiology practice. Systems of the invention can image physiology, yet also have the potential to be inexpensive and portable, unlike current nuclear medical, MRI, CT, and echocardiography systems. Cardiovascular imaging system of the present invention employ a basic modular design that can be used for cardiac imaging with one or more radionucleide tracers. The detector can be positioned in close proximity to the chest and heart from several different projections, making it possible to accumulate data rapidly for first-pass analysis and ejection fraction of both right and left ventricles. The portability of the instrument would also make it very desirable for imaging in the cramped quarters of intensive care units in the cases when it would be not practical to transport the patients to the nuclear cardiology department. Furthermore, preferred embodiments of the present mvention include at least a second camera head, routinely placed at approximately a right angle to the first camera head to provide additional out-of-plane views. The additional camera angle permits the system to account for patient motion in all planes, and eliminating artifacts caused by such patient motion. This is especially useful in traditional stress testing. Other advantages and novel features of the invention will become apparent from the following detailed description of the invention and the accompanying drawings. The detailed description and drawings provide certain illustrative examples of the invention. These examples are indicative of but a few of the various ways in which the principles of the invention can be employed. BRIEF DESCRIPTION OF DRAWINGS Fig. 1 is a flow chart illustrating a method of screening patients for an early stage of CAD in accordance with one aspect of the present invention. Fig. 2 a high level schematic of a gamma particle detector. Fig. 3 illustrates, with a view from the feet, a gamma particle detector measuring the time variation in radioactivity from a left ventricle region of a patient. Fig. 4 is a bottom view of a patient illustrating a preferred position for a gamma particle detector in a method of screening patients for an early stage of CAD. Fig. 5 is a side view of a patient illustrating a preferred position for a gamma particle detector in a method of screening patients for an early stage of CAD. Fig. 6 is a graph show time-activity curves for the heart and brain regions of a cardiac patient. Fig. 7 is an illustration of a gamma particle detector array. Fig. 8 is an illustration of a gamma particle detector array with a flat position sensitive photomultiplier tube. Fig. 9 illustrates the use of two gamma particle detector arrays to detect and locate positron emitters. Fig. 10 is a schematic illustration of electronics for parallel processing of photo-multiplier data according to one aspect of the present invention. Fig. 11 is a schematic illustration of electronics for parallel processing of photo-multiplier data according to another aspect of the present invention. Fig. 12 is an illustration of a system that employs two gamma particle detectors. Fig. 13 illustrates a gamma particle detector array according to another aspect of the present invention. Fig. 14 provides another view of the gamma particle detector array of Fig. 13. Fig. 15 illustrates a three detector system according to a still further aspect of the present invention. Fig. 16 provides a front view of a patient illustrating the positioning of the three detectors in the system of Fig. 15. Fig. 17 provides a side view of a patient illustrating the positioning of the three detectors in the system of Fig. 15. Fig. 18 provides a bottom view of a patient illustrating the positioning of the three detectors in the system of Fig. 15. Fig. 19 illustrates a dual energy gamma probe according to yet another aspect of the present invention. MODE(S) FOR CARRYING OUT THE INVENTION Figure 1 provides a flow chart illustrating a method 100 of screening patients for an early stage of CAD in accordance with one aspect of the present invention. The method involves action 101, injecting a patient with a radioactive bolus, action 103, measuring the time variation in radioactivity from the bolus in a left ventricle region of the patient's body, and action 105, analyzing the time variation to screen the patient for an early stage of CAD. The early stage of CAD detected by method 100 is a significant occlusion of the coronary artery system that occurs before the onset of inducible myocardial ischemia. Inducible myocardial ischemia occurs when one or more major coronary arteries is restricted to a point where coronary circulation is substantially reduced. Blood flows through the coronary circulation system due to a pressure differential created by the heart. The flow passes first through the major coronary arteries, which have the largest diameters, then through minor coronary arteries, which have smaller diameters, and then through capillaries, which have the smallest diameters. In a healthy patient, the capillaries are the bottleneck of the system and primarily limit the flow rate. For inducible myocardial ischemia to occur, one or more of the largest diameter vessels, the major coronary arteries, must become almost completely occluded. The early stage of CAD detected by the present invention involves a substantial reduction in the free volume of the major coronary arteries, which may be considerably less than the degree of occlusion required to cause inducible myocardial ischemia. Of course, method 100 will also detect more advanced stages of CAD and the invention can be used to screen for more advanced stages of CAD. A significant occlusion of the coronary artery system would be one that a reasonable medical practitioner would consider as requiring close monitoring and/or treatment. Action 101 is injecting a patient with a bolus containing a radioactive in a suitable carrier medium. The volume of the bolus is, for example, from about 1 to about 10 cc. The bolus is injected at a suitable point, for example, the antecubital vein of the right or left arm. A suitable radioactive tracer decays to produce gamma particles or alternatively positrons, which rapidly combine with electrons to produce gamma particles. Examples of suitable radioactive tracers include Tc-99m DTP A, 1-131 labeled hippuran, 1-123, In- 111, In- 113m, Tl-201, F-18 labeled fluorodeoxyglucose, Cu-62-PTSM, and 0-15 labeled water. The half-lives of these tracers vary from a few minutes to a few hours. Dosages are selected with due consideration for patient health and government regulations. For example about 1 to about 20 milliCurie of Tc99m (140 keV) DTPA can be used or about 50 to about 300 microCurie of hippuran labeled with 1-131 (360 keV). Higher energy tracers are preferred in that absorption correction is much less important and smaller doses can be used. The bolus is preferably injected in such a way as to provide a single slug of tracer that circulates through the bloodstream. Although the tracer gradually disperses as it flows through the bloodstream, it is desirable to maintain as tight a dispersion as possible at the time of injection. A fractionated bolus, that is one that is separated into two or more slugs, is preferably avoided. According to another aspect of the invention, a gamma particle detector is used to observe the shape of the injected bolus upstream of a region of interest. For example, where the region of interest is a left ventricle region, the bolus can be observed as it passes through the thorax. If the bolus shape is fractionated or otherwise unsuitable, the procedure can be aborted and/or later repeated. Action 103 is measuring the time variation in radioactivity from the bolus in a left ventricle region of the patient. Measuring the time variation in radioactivity involves the use of one or more gamma particle detectors. Figure 2 provides a high level schematic of an exemplary gamma particle detector 200. Gamma particle detector 200 includes scintillation crystal 201, photo-multiplier 203, and signal processing electronics 205. Gamma particles interact with scintillation crystal 201 to produce flashes of light that are detected by photo-multiplier 203. When photo-multiplier 203 detects a flash of light, it provides a signal to electronics 205. Electronics 205 can include any combination of suitable components, such as sealers, amplifiers, and logic circuits. Sealers are typically employed to provide an output that represents number of detection events per time interval. Figure 3 illustrate gamma particle detector 200 in use to measure the time variation in radioactivity from a left ventricle region of a heart 303 of patient 305. The left ventricle region substantially includes left ventricle 307 of the heart 303 while substantially excluding the right ventricle 309. Radiation 301, which is a portion of the radiation produced by the radioactive tracer, leaves the heart 303 and reaches gamma particle detector 200. Gamma particle detector 200 provides the number of gamma particles detected per unit time, which is a measure of the amount of radioactive tracer present in the left ventricle region of the heart 303. The count rate of gamma particles in excess of background levels provides the measure of radioactivity from the bolus. This measure varies over time as the bolus passes through the region of interest. Measurements are made over a period that is generally about 30 seconds or less. The left ventricle region of the patient generally encompasses the left ventricle, although it may include only a portion of the left ventricle. As a practical matter, the left ventricle region also includes portions of the patient outside the left ventricle. The left ventricle region is selected such that the radiation detected is primarily due to radioactive tracer passing through the left ventricle and the coronary circulation system. Radioactive tracer passes through the right heart prior to passing through the left heart. Detection of radioactive tracer passing through the right heart is preferably minimized in action 103. Scintillation crystal 201 is made with an appropriate scintillation material, for example, Nal(Tl), LSO, GSO, CsI(Tl), YSO, or CsI(Na). The scintillation material is is selected to permit detection of gamma photons with energies of at least about 60 keV. Preferably, the thickness permits detection of gamma photons with energies of about 511 keV, whereby gamma photons produced by positron-electron interactions can be detected. For example, crystals of about 2.5 cm thick Nal(Tl), about 2-3cm thick Gadolinium Oxyorthosilicate (GSO), or about 2 cm thick Luthetium Oxyorthosilicate (LSO) can be used. GSO and LSO scintillators are faster and have higher stopping power for 511 keV photons than Nal(Tl) but are also more expensive and have lower energy resolution for lower energy photons, such as 140 keV from Tc-99m. In some cases, a gamma particle detector can be directed toward the patient's left ventricle using for orientation and guidance external anatomical features, such as the locations of ribs. A preferred position of the detector 200 is approximately 30 degree right anterior oblique and in close proximity to the patient, as illustrated in Figures 4 and 5. This generally gives a view peφendicular to the long axis of the cardiac chamber. However, due to anatomical variations that occur among individuals, this is not always the case. Observations of external anatomical features alone do not permit consistent accurate location of the left ventricle. According to one aspect of the invention, an array of detectors is employed, whereby data can be obtained for an array of locations that is greater than the left ventricle region or other region of interest. The array of locations can be chosen so that there is a high degree of certainty that the left ventricle region is within the array. In this context, an array of detectors includes a plurality of detectors physically arranged to collect data over a contiguous area greater than that observed by any one detector in the array. The data from the detector array can be analyzed to determine a subset of detectors in the array that substantially correspond to the left ventricle region. Detector data corresponding to the left ventricle can be identified from characteristics of the time-activity curve. Figure 6 illustrates the shape of a time-activity curve 601, which corresponds to a portion of the heart, and time-activity curve 603, which corresponds to the brain. The heart curve 601 has a characteristic double peak structure. The left peak results from flow through the right heart, while the right peak results from flow through the left ventricle. There are a variety of methods by which time-activity curves corresponding to the left ventricle can be identified. According to one method, the left ventricle time-activity curves are those having a double peak structure and in which the height of the left peak is less than or equal to a certain fraction, about 80% for example, of the height of the right peak. More generally, left ventricle time activity curves can be identified by their functional form. Another method screens the curves according to a learning probabilistic model, such a neural net. Such a model is trained by having an expert manually review a plurality of time activity curves and identifying which do and which do not correspond to the left ventricle. While the left ventricle region can be taken as that which corresponds to the detector that provides the data most definitively produced from the left ventricle, it is preferred that the left ventricle regions include all the detectors corresponding to the left ventricle. Using all the data increases the signal to noise ratio. Some clustering or grouping of the detector data often takes place prior to screening for the data corresponding to the left ventricle. For example, if the detector array provides data in a 16x16 square array with 256 elements, the data may be grouped to form a 4x4 array with only 16 elements. The grouped elements have greater signal to noise ratio but lower spatial resolution. The optimal grouping can depend on the particular data set, including the location of the left ventricle within the array. Thus it can be desirable to use an iterative algorithm wherein, the grouping of the array elements is redefined after a preliminary determination of the left ventricle region. Figure 7 is a schematic illustration of an exemplary detector array 700. Detector array 700 is also a gamma camera. Detector array 700 includes collimator 710, scintillator array 720, light guides 730, photo-multipliers 740, electronic circuitry 750, and tungsten shell 760. Gamma particles passing through collimator 710 strike scintillator array 720. Scintillator array 720 responds to the gamma particles by producing scintillation light. Light guides 730 guide the scintillation light to photo- multipliers 740. Photo-multipliers 740 produce signals in response to the light that are processed by electronic circuitry 750. Detector array 700 provides count rates for a plurality of spatially segregated positions, the spatial resolution generally being limited by collimator 710. Scintillator array 720 generally comprises a pixellated scintillation crystal. Alternatively, a scintillator array can be constructed using an array of individual scintillation crystals. The number of elements in the array is selected according to the desired spatial resolution of the detector array 700. Light guides 730 are employed to guide light from scintillator array 720 to photo-multipliers 740. Light guides 730 can serve to capture light that would otherwise strike between elements of a photo-multiplier array. Generally, light guides 730 are optional. Photo-multipliers 740 produce signals corresponding to light from the scintillator array 720. The point of origin for the light within scintillator array 720 can be determined using a center of gravity calculation applied to individual signals from photo-multipliers 740. Photo-multipliers can be provided in one-to-one correspondence with scintillation array elements, in which case light guides 730 can be used to guide light from scintillator array elements to corresponding photo-multipliers. While a detector can employ a plurality of conventional photo-multipliers, the detector can be made substantially more compact by employing position-sensitive photo-multipliers, particularly flat position sensitive photo-multipliers. Figure 8 illustrates an exemplary detector array 800 employing a flat position-sensitive photo-multiplier 840. In addition to flat position-sensitive photo-multiplier 840, detector array 800 includes collimator 810, scintillator array 820, light guide 830, electronic circuitry 850, and tungsten shell 860. A position-sensitive photo-multiplier is a photo-multiplier that provides information relating to the location within the photo-multiplier's field of view where light was detected. For example, Hamamatsu Model Nos. R7600-C8, R7600-C12, R7600-M4, and R8520-C12 are suitable for use in the present invention. An array of position-sensitive photo-multipliers can also be employed. For example, a 2x2, 4x4, 6x6, 8x6, or 8x8 square array. Whenever an array of photo-multipliers is used, it is preferable that the photo-multipliers have a square, hexagonal, or other shape that allows close packing, whereby the array is compact. Detector arrays 700 and 800 do not necessarily have high spatial resolution. Spatial resolution can be sacrificed in favor of high signal to noise ratio and/or high temporal resolution. In one embodiment, a detector array provides a spatial resolution of about 1 cm or larger. Where the radioactive tracer produces positrons, spatial resolution can be achieved without the use of collimators. When positrons combine with electrons, two gamma particles are emitted in opposite directions. By detecting both gamma particles using two gamma particle detector arrays located opposite one another, and by comparing the locations within the arrays where the gamma particles are detected, points of origin for the gamma particles can be determined. Figure 9 illustrates a patient 910, lying on a table 920, with gamma particle detector arrays 930 and 940 facing each other and both oriented towards a left ventricle region of the patient 910, whereby positrons in the left ventricle region can be detected. Alternatively, a single detector array can be used to detect positrons, although a high-energy collimator, such as one made of tungsten, is required in order to obtain spatial resolution. According to another aspect of the invention, a gamma particle detector is positioned for use in method 100 according to the location of a left ventricle region as determined by a suitable detection method. In this aspect of the invention, a single detector is often used instead of a detector array. The location of the left ventricle can be determined by any suitable technique, including for example ultrasound, MRI, x-ray, computed tomography, planar nuclear medicine, positron emission spectroscopy, single photon emission computed tomography, or a second radioactive tracer. For example, a small amount of Tl-201 (on the order of 50 microCurie) can be injected and the left ventricle region located based on a detector position that provides the maximum count rate. The injection used to locate the left ventricle region is generally performed 5-10 minutes before the main bolus injection. Regardless of how the left ventricle region is located, or whether a detector array is used, method 100 is preferably carried out with one or more high temporal resolution gamma particle detectors. In one embodiment of the invention, the gamma particle detector has a temporal resolution of about 0.1 seconds or better. In another embodiment, the gamma particle detector has a temporal resolution of about 0.03 seconds or better. In a further embodiment, the gamma particle detector has a temporal resolution of about 0.01 seconds or better. Preferably, the gamma particle detectors realize a high signal to noise ratio. The magnitude of the signal is limited in part by the amount of radioactive tracer contained in the bolus. In one embodiment, one or more detectors have a signal to noise ratio giving a measurement error of less than about 10%. In another embodiment, the measurement error is less than about 3%. In a further embodiment, the measurement error is less than about 1%. The detector head is preferably encased with 0.63 to 1.25 cm thick tungsten shielding to reduce background noise. Where a detector array is employed, the measurement error can be reduced at the cost of spatial resolution. At lower spatial resolution, each separately observed spatial region within the patient produces a sufficient number of detected radiation events within a measurement interval to produce a measurement with the required statistical accuracy. For example, if the required temporal resolution is 0.01 seconds and 1000 events per interval are required to achieve the desired signal to noise ratio, the detected event rate for a single detector must reach about 100 kHz. The gamma particle detector, or detector array, used to obtain measurements from the region of interest is preferably portable and preferably dedicated to heart diagnosis. A unit dedicated to heart diagnosis generally takes measurements over an area about the size of the heart. In one embodiment, the detector or array takes measurements over an area of about 40 cm 2 or less. In another embodiment, the detector or array takes measurements over an area of about 20 cm 2 or less. In a further embodiment, the detector or array takes measurements over an area of about 10 cm 2 or less. The required data can generally be obtained in less than about 30 seconds, preferably less than about 15 seconds. Where data is sought while the patient is under stress, the short time required significantly reduces the burden on the patient as compared to other methods. Action 105 is analyzing the data to screen the patient for an early stage of CAD. The data provides a time-activity curve for the concentration of tracer in the left ventricle region. This time- activity curve contains two contributions, one from tracer flowing through the left ventricle and another from tracer flowing through the coronary circulation system. The time activity curve can be represented by the equation: A τ = V LV C LV (t) + V c {D)C c (t, D) (i) wherein A τ is the measured activity, V LV is the volume of the left ventricle, C LV is the concentration of tracer in the left ventricle, which depends on time, t, V c is the volume of the coronary circulation system, which depends on the disease state, D, and C c is the concentration of tracer in the coronary circulation system, which depends on both time, t, and the disease state, D. Shortly after the tracer enters the left heart, the first term in equation (1) dominates because V LV is always larger than V c . However, the residence time of the tracer in the left ventricle is much shorter than the residence time of the tracer in the coronary circulation system, typically being six to seven times shorter. Thus at longer times, the contribution of the second term becomes more apparent in the time activity curve. Functional analysis of the time activity curve, particularly at longer times, can be used to analyze the relative contributions of the two terms in Equation (1) and evaluate the disease state of the patient. In a diseased patient, the second term of Equation (1) is diminished relative to the first term. The direct effect of the disease is that the volume, V c , becomes smaller. The concentration C c , is also thought to become smaller, particularly at longer times. The reason is that in the early stages of disease, the volumetric flow rate of blood through the coronary circulation system is not substantially affected; the volumetric flow rate is primarily controlled by the bottleneck presented by the capillaries. But if the volumetric flow rate remains constant while the volume of the system is reduced, the system flushes more quickly - the residence time is reduced. Thus the concentration in the coronary circulation system falls off more quickly in a diseased patient than in a healthy patient. While it is possible to carry out Action 105, analyzing the data to screen the patient for an early stage of CAD, using only the data for the left ventricle region collected in Action 103, in a preferred embodiment of the present invention the data for the left ventricle region is analyzed by comparing it to the data for a control region. A control region is a portion of the patient downstream of the left ventricle, such as the ascending aorta or the brain, for example. Data for the control region is generally obtained using an additional gamma particle detector. The control region is selected to provide a time activity curve that is similar to that of the left ventricle apart from the coronary circulation system. As illustrated in Figure 6, the down slope of the time-activity curve 601 for the heart region, when plotted on logarithmic scale, is greater than the down slope of the time-activity curve 603, corresponding to the brain or a similar control region. The ratio of the logarithmic scale down slope (mono-exponential half- life) for the heart region to the down slope for the control region is referred to as the coronary transit index (CTI). A low CTI, particularly one that approaches unity, indicates the presence of coronary artery disease. In one embodiment, only patients with a CTI of about 1.5 or greater pass the screen for early stage of CAD. In another embodiment, only patients with a CTI of about 1.4 or greater pass the screen. In a further embodiment, only patients with a CTI of about 1.3 or greater pass the screen. Another aspect of the invention relates to a high temporal resolution gamma particle detector array, which is a gamma camera. The camera can be used for first pass imaging, positron imaging, quantitative myocardial perfusion measurements, planar hot spot imaging, and planar, gated, non-gated, and multi-gated pooled blood imaging. These images can be used for such puφoses as measuring regional ventricular wall motion, measuring left and right ventricular ejection fractions, measuring blood flow in the myocardium, measuring regional contractivity, measuring synchrony, detecting infarct sites and abnormalities of conduction, and left to right shunt quantification. In a preferred embodiment, the gamma camera is able to operate in the gamma energy range of 60-600 keV with rate capability of at least about 1 MHz. High temporal resolution can be achieved, in part, by configuring the electronic circuitry for parallel processing of data from the photo-multiplier or photo-multiplier array. Figure 10 illustrates one electronic circuitry configuration for achieving such parallel processing. In Figure 10, each of a plurality of photo-multiplier tubes 1010 is coupled to a corresponding amplifier 1020. The amplified signals are provided to corresponding two level discriminators 1030, which distinguish those signals that have sufficient energy to correspond to the radioactive tracer in use. The output of the two level discriminators 1030 is fed to sealers 1040, which count the number of events and periodically supply output to buffer/multiplexer 1050. Buffer/multiplexer 1050 combines the count rates provided by the various sealers into a single output stream that can be provided to a computer. Some additional options for configuring the electronic circuitry for parallel processing are illustrated in Figure 11. Figure 11 shows the signals coming from various anodes 1110 of one or more photo-multipliers. There does not need to be a one-to-one correspondence between parallel channels 1120 and either anodes or photo-multipliers. Rather, the signal processing can be divided in any suitable fashion. The anode signals are supplied through channels 1120 to VME analog-to-digital converter modules 1130. The digital data is accumulated in VME memories 1140 and processed by CPUs 1150. The data can then be shared across a network, such as ethernet 1160. A gamma camera of the present invention can be made compact and portable. Such a camera can be positioned in close proximity to the heart from several different directions, which facilitates first- pass analysis and measuring ejection fractions from both left and right ventricles. A compact camera is particularly useful in intensive care units, where space is at a premium. The camera can be provided with a support gantry/harness to provide co-registration with the patient's body. In imaging the heart, it is often desirable to compensate for patient motion. This is particularly true when heart function is being analyzed while the patient is active. One approach to compensating for patient motion is to place a radiation source on the patient's body as a point of reference. For example, an Am-241 source can be fixed to the patient's chest. The location of the source can be used to correct for patient motion in a plane peφendicular to the gamma camera. The reference source can be distinguished from the tracer, if desired, using dual energy windows. According to a further aspect of the invention, a second camera is placed at an approximately right angle to the first camera head, whereby the system can compensate for patient motion in any direction. Figure 12 illustrates two gamma cameras, first gamma camera 1220 and second gamma camera 1210, arranged in such a manner and oriented toward the left ventricle region of a patient's body 1230. In addition to facilitating compensation for patient motion, a second camera also allows all edges of the ventricular blood pool to be viewed simultaneously. These include the antero-lateral, apical, posterior, septal, and posterior-lateral walls. The second view also increases the accuracy of the diastolic chamber volume calculation and provides information that can be used in correcting for heart movement in first pass test procedures. Wall motion information to correct for heart movement in the first pass test procedure can be provided by placing second gamma camera 1210 approximately 90 degrees from first camera 1220, and therefore imaging at a right angle to the first camera 1220, which is looking directly into the heart of patient 1230, as depicted in Figure 12. In the first-pass test a preferred positioning of the first camera 1220 is approximately 30 degree right anterior oblique. This allows a view peφendicular to the long axis of the cardiac chamber such that valve location can be detected and provides the additional utility of permitting the cardiovascular imaging system to calculate stroke volume. This information becomes available in about 15 seconds. In the blood pool (MUGA) studies, the preferred position of the head is 30 degree left anterior oblique. An approximately -45 degree tilt angle option can be employed to better align the imaged field of view with the heart chambers. Embodiments of the present mvention comprising a two head positron imager can operate in a simple planar tomographic reconstruction mode (such as laminography) to produce image slices through the heart. A cardiovascular imaging system of the present invention can employ compact position sensitive photomultipliers to provide dedicated fast rate heart gamma imagers. The system provides highly improved spatial resolution and sensitivity to gamma-rays at a lower cost than current clinical nuclear medicine systems. Because of its compact size, these dedicated instruments are also capable of permitting new imaging geometries for improved visualization of important prognostic aspects of heart function. These capabilities can be achieved with an array of compact position sensitive photo-multiplier tubes (PMTs). Thick pixellated Nal(Tl) scintillator arrays with pixel sizes of 2-5mm by 20-25 mm thick can detect gamma rays in the broad range from 60-80 keV (Tl-201) to 511 keV (F-18). The scintillator array may be coupled to the photomultiplier array via an optical window to form very compact detector heads with an active field of view on the order of 20 cm 2 . The detector heads may be attached to a support gantry/harness to provide co-registration with the patient's body. In an imager system with a single detector head at lower energy range, including 140 keV from Tc-99m, a lead collimator may be implemented, while at higher energies, including 511 keV, a tungsten collimator will typically be preferred. In the positron imaging mode two detector heads will be preferably used in coincidence with no collimators attached, as previously described. Enhancing the modular design of the present invention, a set of interchangeable collimators may be employed depending on the required balance between the detector efficiency and its spatial resolution. The detector heads are preferably encased in a 0.63 to 1.25 cm thick tungsten shield to avoid unwanted counts from background radiation. To enhance further the understanding of the cardiovascular imaging and functional analysis system disclosed herein, the following example of a specific embodiment of the present invention is presented. In this exemplary Non-Invasive Screening Probe System, the first probe is made out of a single crystal scintillator. The size of the first probe is about 2.5 to about 5 cm square and thick enough to detect efficiently gamma photons up to 511 keV. In a preferred embodiment of the exemplary Non-Invasive Screening Probe System, a single Nal(T) crystal of about 3 cm by 3 cm by 3 cm in size is used in the first probe. A collimator is located in front of the probe. The first probe has a single large hole (single bore) or is made of a multitude of smaller holes (e.g., 5-mm in diameter) placed in front of the probe, between the chest wall and the gamma detector/probe. The second probe of the exemplary Non-Invasive Screening Probe System has an approximately 5x5-cm active field of view and is built out of a multitude of smaller-size crystals/pixels. Pixel sizes are generally from 3mm to 20mm, but are typically in the 5 tolO-mm range. A collimator aligns the received gamma radiation before directing it to the scintillator array. As in the case of the first probe, in the second gamma imaging probe, the thickness of the material is selected so as to accommodate the energy range from 60-511 keV. In the specific embodiment depicted, the scintillator array uses a thick Nal(Tl) pixellated scintillator. In this embodiment of the exemplary Non-Invasive Screening Probe System, the scintillator crystal array typically consists of ~5x5x25-mm pixels that are optically separated by thin (0.2-0.3 mm) white diffusive layers. Four approximately 2.5 cm-square Hamamatsu R7600-M4, R7600-C8, or R7600-C12 PSPMTs arranged in a compact 2x2 array can be used to detect and define the position of a scintillation flash from the gamma conversion in the scintillator array. Figure 13 depicts a side view cross section of another embodiment 1300 of the second gamma imaging probe of the exemplary Non-Invasive Screening Probe System. As shown in Figure 13, collimator 1310 directs incoming light to an array of GSO scintillators 1320 arranged as a 4x4 matrix. Each individual scintillator 1320 in the array is coupled to an individual photomultiplier tube 1330 in a PMT module. The coupling of scintillators 1320 with individual PMTs 1330 eliminates the need for a light guide, thus simplifying the exemplary Non-Invasive Screening Probe System. Figure 14 depicts an arrangement of the 4x4 matrix of GSO scintillators 1320 in the gamma imaging probe 1300. When performing a screening test with radiopharmaceuticals labeled with Tc-99m (140 keV gamma emission) or other isotopes emitting lower energy gamma radiation, absoφtion and scattering of gamma radiation on its way to the gamma detector-probe decreases sensitivity for gamma emissions on the further side of the heart away from the second probe. In Figure 15, the two cardiovascular regions 1530 and 1540 of heart 1520 each have quite different positions relative to the standard second probe 1510 as used in the standard two-probe arrangement described above. Probe 1500 is used as the first, reference probe positioned above the aorta, as described above. This effect may limit relative sensitivity of the screening test to the cardiovascular problems in the more distant part of the heart. Therefore, to compensate for this near-far asymmetry in the sensitivity of the screening test, a third probe 1550 may be added. This probe can be identical in construction to the second probe 1510 and is positioned in such a way as to view region 1540 of heart 1520 from a much closer distance than the second probe 1510. The relative external position of each of the three probes is shown in Figures 16, 17, and 18. The locations of the three probes are shown in a frontal view in Figures 16, in a side view in Figures 17, and in a bottom view in Figures 18. In this embodiment, the signals from third probe 1550 can be analyzed separately and this additional information can be added to the information recorded with the second probe 1510. While adding to the complexity and cost of the equipment and the procedure, the method is substantially increases the overall sensitivity of the screening test when performed with Tc-99m. The additional probe 1550 may require position adjustment to view the proper part of the heart (e.g., left ventricle). This can be achieved by guiding the probe by ultrasound, X-ray radiation, or using an additional radiation source. For example, the probe positioning technique may involve injecting small amounts of Tl-201 (on the order of 50 microCurie) and then directing the probe at the LV region by a maximum count rate criterion. In this case one pre-injection would be performed 5-10 minutes before the main screening test procedure. In the case of higher energy isotopes such as 1-131 or In-113m, the above addition is not necessary. Radioisotopes used in nuclear medicine are strictly regulated, thereby restricting the practical range of available doses. Furthermore, lower doses are typically preferred over higher doses, if similar clinical results obtain from differing doses. The results described herein can be achieved with doses of various radioisotopes emitting as little as 50 microCurie. The puφose of the multi-probe (alternatively and interchangeably referred to herein as multi-detector, multi-sensor or multi-pixel) approach of the present invention, and particularly of the exemplary embodiment, is to assure that the measurement is performed on the left ventricular bolus transit accompanying that of the aorta. While positioning of the first (aortic) probe is relatively simple to perform for each patient, and no multiple sensors are necessary, technologists operating the system may have a problem in positioning the second ventricular probe above the heart's left ventricle. To remedy the potential mispositioning, which would thus collect activity from superimposed left and right heart regions, the multi-probe system collects data from a larger region than necessarily encompasses the left ventricle. In some cases it is advantageous to perform a screening test using several radiopharmaceuticals simultaneously in the same patient. For example, the first radiopharmaceutical can be used to "lighten" the cardiac region to enable facilitate positioning of a probe for the dynamic part of the screening procedure. In this case the second radiopharmaceutical compound may be injected while the first compound is still circulating in the blood stream and/or is present in the myocardium or in other parts of the patient's heart. In another possible situation both compounds are injected at the same time because they will have a different role to play in the screening test and their temporal distribution in the blood stream and heart tissue will be different and complementary. There might be many other situations when by injection or by other route of entering, several radiopharmaceuticals adjusted to the same patient will be necessary or beneficial to a screening test and/or to other diagnostic procedure. In some cases the additional radiopharmaceuticals can come from other than cardiac-related procedures. In all these cases it would be beneficial or even necessary to operate the screening probe system in the presence of different radiations emanating from the patient. The Tc-99m radioisotope is used in most nuclear medicine cardiac imaging studies. 1-131 radioactive label is preferred in the screening probe test. 1-131 is easy to use and emits gamma radiation of higher than Tc-99m energy (360 keV vs. 140 keV) that penetrates the patient's body with less scatter and absoφtion on its way to the detector. A clearer picture of flow distribution is therefore obtained. Additionally, a much lower injection dose of about 200 microCurie (vs. many milliCurie) will suffice to perform the screening test. However, 1-131 emits highly ionizing beta radiation and the regulations intended to minimize patient exposure to radiation limit the availability of this radiolabel. Used in the past in diagnostic imaging of kidneys, 1-131 is presently not readily available. It is however possible that the advantages of the screening test will outbalance the harmful effects of its radiation and it will become available again. A short lived (2 minute half-life ) 0-15 positron emitter (emitting 511 keV annihilation gamma rays) used in water is a very good performing and well accepted tracer for blood flow studies. Therefore, the screening probe system of the present invention is designed to accommodate the higher energy gamma emitters in addition to the lower energy of Tl-201 (60-80 keV) and Tc-99m (140 keV). With an efficient and high count rate system, a high temporal resolution is possible to improve precision of the screening test. To separate signal contributions from these compounds in the screening probe system, the mixture of applied radiopharmaceuticals is typically composed of compounds labeled with different radionucleides. The differing radiation emissions from these radionucleides can be distinguished in a probe system of the present invention by the different characteristics of the signals accompanying detection of gamma or beta rays from these radionucleides. For example, the first compound can be labeled with Tc-99m emitting 140 keV gamma rays, while the second compound can include positron emitter such as 0-15 or F-18, which in turn will produce 511 keV annihilation gamma rays detected by the probes. A screening system of the present invention can be equipped to differentiate between the differing gamma radiations emitted by the different components in the radiopharmaceutical mixture. The screening system can also possess sufficient detection efficiency for the potentially broad spectrum of gamma ray energies emitted from the different radionucleides. In a preferred embodiment, the screening system is built with a single material, possessing both high stopping power and high energy resolution. However, in practical situations, even when using the best available high energy resolution solutions, such as solid state Cadmium Zinc Tellurium (CdZnTe), there will always be signal and information overlap from the two or more radiation components from the radiopharmaceutical mixture. In another preferred embodiment, two differing scintillation materials are optically coupled to the same photodetector. An example of such a dual energy probe 1900 is shown in Figure 19. Dual energy probe 1900 includes collimator 1970 positioned at an open end of housing 1960, first scintillation layer 1910, second scintillation layer 1920, optional optical coupling 1930, glass window 1950, position- sensitive photo-multiplier 1940, and electronic circuitry 1980. A preferably first scintillation layer 1910 is a Nal(Tl), CsI(Tl), or CsI(Na) array of approximately 3x3x6mm pixels with an approximately 50mm square active coverage. First scintillation layer 1910 is optically coupled to the second scintillation layer 1920, which includes a heavy scintillator, for example a square slab 1 cm thick and 6 cm on a side of Gadolinium Oxyorthosilicate (GSO). Other examples of the second scintillator material are: Lutetium Oxyorthosilicate (LSO) or a Lutetium-Gadolinium Oxyorthosilicate (LGSO), which is a crystal scintillator produced as a mixture of GSO and LSO. Scintillation light from the first scintillator 1910 produced on detection of gamma rays in its material is transported via the optically transparent medium of the second scintillator 1920 to the following optional light guide (not shown) and then to the photocathode of the photo-multiplier 1940 via the glass window 1950 of the photo-mulitplier. An optical coupling compound and an additional optical window 1930 can be used between the first and second scintillation layers 1910 and 1920. For example, in the Nal(Tl) scintillator case, an enclosure with a transparent glass window is used to prevent corrosion of the scintillation material on contact with air and humidity. A major fraction of scintillation light produced from gamma rays interacting with the second scintillation layer 1920 enters the photo-multiplier 1940 via the optional optical window (not shown) or light guide and the glass photo-multiplier window 1950, finally being stopped and converted to an electronic signal in the photocathode of the photo-multiplier 1940. The thickness of the first scintillation layer 1910 is selected so as to detect efficiently gamma radiation up to about 150 keV, while the second scintillation layer 1920 is intended to be optimized for higher energy gamma rays such as 360 keV from 1-131, 390 keV from In-113m or 511 keV annihilation photons from positron emitters such as 0-15 or F-18. While most lower energy gamma rays will be stopped in the first scintillation layer 1910, a major fraction of higher energy gamma rays will pass undisturbed through the first scintillation layer 1910 and will interact and be detected in the second scintillation layer 1920. In this embodiment, the two scintillators selected for the two scintillation layers differ not only in their gamma radiation stopping power but also in their scintillation characteristics, such as light output measured in light photons per unit of energy of the absorbed radiation, and their scintillation light emission temporal characteristics. It is desirable that both scintillators have similar and high light outputs, while their temporal characteristics, expressed in time duration of the scintillation signal, are sufficiently different so as to enable temporal analysis of the shape of each scintillation pulse to be performed by dedicated electronic signal analyzing circuitry incoφorated in the probe system of the present invention. This circuitry and the proper software algorithm installed in the probe system's computer enable differentiation of the pulses produced from lower energy gamma photons interacting in the first scintillator layer 1910 from the signals produced in the second scintillator layer 1920, which mostly originate from higher energy gamma rays. In this embodiment, Nal(Tl) and GSO constitute a very well balanced pair of practical scintillators possessing optimal characteristics for fulfilling the above discussed requirements for a proper operation of the dual-energy probe. In this dual energy detection system embodiment of the present invention, interference between the two radiation components coming from two or more radiopharmaceuticals is minimized. Independent dynamic curves for mixture of at least two radiopharmaceuticals can be obtained. While complicating the probe design, this added analytical power of the present embodiment of the screening probe system can become of prime importance in the future, due to the increasing role of multi-radiopharmaceutical nuclear medicine imaging and multi-modality diagnosis in general. An exemplary photo-multiplier solution of this embodiment is a commercially available single 7.5 cm square model R2487 (or a similar model R3941) position sensitive photo-multiplier (PSPMT) from Hamamatsu Photonics. Its useful photocathode size is about 5 cm square, which is entirely adequate for the second example of a screening probe embodiment of the present invention. With respect to the previously discussed embodiment of the screening probe system with two different probe designs for the first (simpler, no pixellation) and the second (pixellated) probe, the above design is applicable to both probes of the screening system. While the present invention is illustrated with particular examples, other system and methods have been described and are within the scope of the hereto appended claims. The invention and the claims are to be construed as including modifications, alterations, and equivalents that will occur to those of ordinary skill in the art upon reading and understanding this specification. INDUSTRIAL APPLICABILITY The apparatus and associated methods may be used in the field of medicine. The apparatus and methods may be used to image the heart, evaluate pulmonary function, and screen for an early stage of CAD.

Claims: What is claimed is: 1. A method of detecting partial occlusion of a coronary circulation system in a patient having a bloodstream, comprising: injecting the bloodstream with a bolus comprising a radioactive tracer; obtaining measurements related to a time variation in concentration of the radioactive tracer for a plurality of adjacent regions in the patient that encompass an area greater than a left ventricle region; determining a subset of the plurality of regions that correspond to the left ventricle region; and assessing whether the coronary circulation system is partially occluded based at least in part on those of the measurements that correspond to the left ventricle region. 2. The method of claim 1 , wherein the measurements are obtained with one or more gamma particle detectors comprising a scintillator array and one or more position-sensitive photo- multipliers. 3. The method of claim 2, wherein the measurements are obtained with electronics that process output from the one or more photo-multipliers in parallel. 4. A method of obtaining data relating to occlusive coronary artery disease in a patient having a bloodstream and a heart having a left ventricle, comprising: determining the location of the heart; using the determined location of the heart, positioning a gamma camera to detect gamma particles emitted from a left ventricle region of the patient; injecting the bloodstream with a bolus comprising a first radioactive tracer; and using the gamma camera to obtain measurements related to a time variation in concentration of the first radioactive tracer in the left ventricle region. 5. The method of claim 4, wherein the position of the heart is determined by ultrasound, MRI, x-ray, computed tomography, planar nuclear medicine, positron emission spectroscopy, single photon emission computed tomography, or a second radioactive tracer. 6. The method of claim 4, wherein the position of the heart is determined by a process involving the detection of a second radioactive tracer, which is different from the first radioactive tracer. 7. A method of screening a patient for an early stage of coronary artery disease, comprising: injecting the patient with a bolus comprising a radioactive tracer; measuring a first time variation in a concentration of the radioactive tracer in a left ventricle region of the patient; employing the first time variation to screen the patient for the early stage of coronary artery disease 8. The method of claim 7, further comprising: measuring a second time variation in a concentration of the radioactive tracer in a second portion of the patient located downstream of the left ventricle region; and comparing the first time variation and the second time variation in screening the patient for the early stage of coronary artery disease. 9. A method of imaging a beating heart within a patient, comprising: injecting the patient with a bolus comprising a radioactive tracer; using high temporal resolution gamma particle detectors, measuring the concentration of the radioactive tracer at a plurality of positions within the patient comprising the heart; and forming an image of the heart from the measured concentrations. 10. The method of claim 9, wherein the high temporal resolution gamma particle detectors are used to obtain image data from at least to different angles simultaneously. 11. The method of claim 9, wherein the gamma particle detectors have a spatial resolution no greater than about 1 cm. 12. The method of claim 9, wherein the gamma particle detectors have a temporal resolution of at least about 0.03 seconds. 13. The method of claim 9, wherein the gamma particle detectors have a temporal and spatial resolution that provides a measurement error no greater than about 10% within a region of interest when the bolus contains 20 milliCurie of Tc99m. 14. A method of assessing pulmonary health, comprising: injecting the bloodstream with a bolus comprising a radioactive tracer; employing a first gamma particle detector to determine the bolus shape as it passes a point upstream of a heart; determining whether the bolus shape is satisfactory for a bolus transit time measurement; and if the bolus shape is satisfactory, measuring the bolus transit time with a second gamma particle detector and assessing pulmonary health from the bolus transit time. 15. A dedicated fast rate heart gamma imager, comprising: an array of scintillators producing a scintillator output; an optical coupling for transmitting the scintillator output; one or more position-sensitive photo-multiplier tubes forming a single detector head for receiving the transmitted scintillator output, the position-sensitive photo-multiplier tubes producing a photo-multiplier output; and electronic circuitry for receiving the output from the position-sensitive photo-multiplier tubes (740). 16. The imager of Claim 15, wherein the one or position-sensitive photo-multiplier tubes form an array. 17. The imager of Claim 15, wherein the one or position-sensitive photo-multiplier tubes are flat. 18. A dual energy fast rate heart gamma detector, comprising: a housing having an open end; a collimator located at the open end of the housing; a first scintillator layer located within the housing for receiving radiation via the collimator, the first scintillator layer producing a first scintillator output; a second scintillator layer located within the housing for receiving radiation via the collimator and via the first scintillator layer, the second scintillator layer producing a second scintillator output; a photo-multiplier located within the housing for receiving the first scintillator output and the second scintillator output; and electronic circuitry coupled to the photo-multiplier. 19. The detector of Claim 18, further comprising a first optical coupling located between the first scintillator layer and the second scintillator layer. 20. The detector of Claim 18, wherein the first scintillator layer is pixellated. 21. A system for fast rate heart imaging and detection of cardiovascular hemodynamics, the system comprising: a first gamma probe; and a second gamma probe; wherein the first gamma probe and the second gamma probe each include: a housing having an open end; a collimator located at the open end of the housing; a scintillator array located within the housing, the collimator being located between the scintillator array and the open end of the housing, wherein the scintillator array produces an output; at least one photomultiplier tube module located within the housing for receiving the scintillator array output; and electronic circuitry coupled to the at least one photomultiplier tube module. 22. The system of Claim 21 , further comprising a support gantry/harness for supporting the first gamma probe and the second gamma probe, wherein the support gantry/harness provides co- registration of the first gamma probe and the second gamma probe. 23. The system of Claim 21, further comprising a central processing unit that processes a first set of heart information received from the first gamma probe and a second set of heart information received from the second gamma probe. 24. The system of Claim 23, further comprising: a third gamma probe, wherein the third gamma probe receives a third set of heart information; wherein the third gamma probe includes a housing having an open end; a collimator located at the open end of the housing;a scintillator array located within the housing, the collimator being located between the scintillator array and the open end of the housing, wherein the scintillator array produces an output; at least one photomultiplier tube module located within the housing for receiving the scintillator array output; and electronic circuitry coupled to the at least one photomultiplier tube module; wherein the central processing unit processes a third set of heart information received from the third gamma probe.

Inventors: BISHOP HARRY, MAJEWSKI STANISLAW, UMENO MARC M, BISHOP, HARRY, MAJEWSKI, STANISLAW, UMENO, MARC, M.
Applicants: NEOMED TECHNOLOGIES INC, NEOMED TECHNOLOGIES, INC.
Application year: 2001
Main classification: A61B
IPC codes: A61B5, G01T1