Radiant energy – Invisible radiant energy responsive electric signalling – With or including a luminophor
Reexamination Certificate
2000-02-15
2002-05-21
Hannaher, Constantine (Department: 2878)
Radiant energy
Invisible radiant energy responsive electric signalling
With or including a luminophor
C250S363070, C250S363100, C250S369000, C250S371000, C378S062000, C378S149000
Reexamination Certificate
active
06392235
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention is related to the general field of gamma-ray imaging. In particular, the invention provides a new method and a new coded-aperture system for detecting gamma radiation emitted from an object under examination and constructing an image corresponding to the spatial distribution of the source of radiation within the object.
2. Description of the Prior Art
Coded-aperture systems were first suggested in the 1960s for use in x-ray astronomy, where the objects of interest are essentially two-dimensional. This idea was exciting at the time because a coded aperture could be far more efficient than conventional pinhole apertures of collimators in collecting radiation such as x-ray photons. It was later understood that more photons were required with a coded aperture than with a pinhole because each photon conveyed less information about the source. Nevertheless, for many kinds of sources, coded apertures offer a considerable net advantage, and they continue to be actively used in x-ray and gamma-ray astronomy today.
The concept of coded apertures was later extended to nuclear medicine, where the objective is a volumetric, or three-dimensional (3D), distribution of a radioactive tracer. This application of coded apertures has been studied extensively during the past 25 years, and its mathematical and engineering aspects are well understood in the art. In particular, it is generally agreed that a single coded image taken with an aperture at a fixed location is an inadequate data set for accurate 3D tomographic reconstruction. The basic reason is that radiation from each point in the object is collected only over a very limited range of angles.
Several approaches have been proposed to overcome this limited-angle problem. One is to use many different apertures, either at the same location or, preferably, at different angular positions around the object to be imaged. As one skilled in the art would readily appreciate, the reconstruction problem then becomes more complicated because the data consist of many different 2D images; nevertheless, it is well within the capability of modern computers and algorithms currently in use in the art. The main drawback is the mechanical complexity of such systems and the fact that the net advantage in collection efficiency is marginal for large three-dimensional objects.
To take full advantage of coded apertures, a detector with high spatial resolution is required, and the scintillation detectors used routinely in nuclear medicine are deficient in this respect. The best scintillation cameras have a spatial resolution of only about 2-3 mm, and there is very little prospect for any substantial improvement.
Accordingly, much effort has been devoted to the development of a new generation of semiconductor gamma-ray detectors for use in nuclear medicine. Such detectors consist of a slab of semiconductor material, such as cadmium zinc telluride, and a multiplexer readout circuit. For example, a 64×64 detector array with a spatial resolution of 0.38 mm, an order of magnitude better than scintillation cameras, has already been tested successfully in the art (see U.S. Pats. No. 5,245,191 and No. 5,825,033, hereby incorporated by reference) and work on developing new imaging systems to take full advantage of this capability is in progress. The ultimate goal is to build a high-resolution tomographic imaging system based on semiconductor detectors and either conventional pinhole apertures or coded apertures.
The considerable interest found in semiconductor gamma-ray detector arrays arises from the improvements they are able to provide over scintillation cameras in a variety of applications. These devices can be used in nuclear medicine, diagnostic radiology, molecular biology, gamma-ray astronomy, particle-physics and nuclear-weapon applications. Therefore, it is expected that a high-quality semiconductor array combined with improved data processing techniques will have a substantial commercial and scientific impact.
The present application concerns a novel approach in the way data are generated by detector arrays and processed for medical-imaging applications. Because of the background that lead to the invention, this disclosure is based on data and experiments related to the medical field, but its application pertains to all disciplines that can utilize semiconductor sensors for gamma-ray detection.
The goal of medical imaging is to provide a spatial mapping of some parameter, feature, or process within a biological entity. Emission imaging (or nuclear medicine) comprises a class of imaging techniques that produce a functional mapping of the object under observation. Generally speaking, the techniques used in nuclear medicine involve the injection of a radioactive substance into a patient's body and the measurement of the emitted radiation (gamma rays) by radiation sensitive detectors through a system of apertures in an impermeable medium. Typically, before injection the radioactive tracer (radionuclide) is combined with a substance that is known to be preferentially concentrated in the organ of interest, so that the preferential concentration of the resulting radiopharmaceutical will correspond to an indication of blood flow, metabolism, or receptor density within the organ. Thus, an image of the resulting radioactive distribution within the organ of interest will yield functional information about the organ. Either a single projection image of the emission distribution may be taken (planar imaging) or many projection images may be acquired from different directions and used to compute the three dimensional emission distribution (single photon emission computed tomography, generally referred to as SPECT).
Since photons in the energy range used in nuclear medicine are not substantially refracted or reflected, data are collected by placing attenuating apertures between the patient and the detector plane, so that each detector has an associated field of view defined by the aperture. Photons that are recorded by a particular detector element in the detector plane are known to have originated in a certain small portion of the object space. The number of photons detected by a given detector is proportional to a weighted integral of the activity contained in the region it sees. By utilizing the information collected by many detector elements or cells, each viewing different but overlapping regions of the object space, an estimate of the original activity distribution can be produced by a reconstruction algorithm according to analytical methods and techniques that are well understood by those skilled in the art.
Different kinds of apertures are commonly used to provide the desired select field of view of an object. For example, parallel-hole collimators, focused collimators, single and multiple pinhole attenuators, and several other apertures that can be used to restrict the path of the gamma rays between the radioactive object and the detector in a tomographic imaging system. In each case, each detector element ideally receives radiation from a single line from the object through the aperture of the system. U.S. Pats. No. 5,245,191 and No. 5,825,033 describe the use of a semiconductor detector array in combination with multiple-pinhole apertures to produce an improved tomographic imaging system.
Not all imaging in nuclear medicine is tomographic, however. Most routine studies are performed with simple parallel-bore collimators, and the resulting image is a 2D projection of the 3D object. Common jargon for this approach is “planar imaging,” implying not that the object is planar but that the resulting image is.
The improved semiconductor detectors could be used for planar imaging, but the overall spatial resolution would still be severely limited by the collimator. The bore size in currently available collimators is 1-3 mm, which establishes an absolute lower limit to the resolution in the image. For a thin object in contact with the face of the collimator, the resolution is essentially the same
Barrett Harrison H.
Clarkson Eric
Wilson Donald W.
Durando Antonio R.
Durando Birdwell & Janke PLC
Gagliardi Albert
Hannaher Constantine
The Arizona Board of Regents on behalf of The University of Ariz
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