Surgery – Diagnostic testing – Detecting nuclear – electromagnetic – or ultrasonic radiation
Reexamination Certificate
2001-05-15
2002-11-19
Manuel, George (Department: 3737)
Surgery
Diagnostic testing
Detecting nuclear, electromagnetic, or ultrasonic radiation
C250S370090, C250S370080, C250S363030
Reexamination Certificate
active
06484051
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates to gamma ray imaging, and in particular, to nuclear medical imaging using Compton scattering of gamma rays simultaneously emitted by a radio-nuclide.
2. Background of the Invention
Nuclear medical imaging is an important research and clinical tool. The two common techniques used are Single Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET). These are extensively used to investigate the function of various organs and to determine the location and morphology of malignant tissue. Both SPECT and PET techniques use radio-nuclides which are administered to a patient. In SPECT a variety of radio-pharmaceuticals are used to study various tissue functions such as heart and kidney function and to locate cancer tissue. An exemplary application of PET imaging includes the use of positron emitting isotopes such as
15
O and
18
F to investigate brain function.
In SPECT applications, a radio-pharmaceutical is administered to a patient. Nuclei of the radio-pharmaceutical decay with the emission of gamma rays. The location of the gamma ray emission is imaged using collimated, position-sensitive gamma ray detectors. Each detected gamma ray is assumed to have arrived at the detector through the collimator. A single SPECT detector can provide a two-dimensional image of the distribution of the radio-nuclides. By moving the detector to obtain many views of a region-of-interest from different directions or using several detectors to simultaneously view the region-of-interest from different directions, coupled with the use of computer algorithms, it is possible to reconstruct a three-dimensional structure of the radioactivity.
SPECT is most frequently used with relatively low-energy gamma rays because it is difficult to fabricate collimators that are effective for gamma rays above several hundred keV. A very widely-used radio-nuclide for SPECT is
99m
Tc which emits a 140 keV gamma ray with a half-life of approximately six hours. Other commonly used radio-nuclides used with SPECT include
67
Ga (93 keV),
111
In (172 and 245 keV),
131
I (364 keV), and
201
TI (70-80 keV X-rays).
A major limitation with SPECT is the small field-of-view provided by the collimators and the associated very low fraction of the gamma rays that pass through the collimator to the detector. Typically only 0.001%-0.01% of the gamma rays pass through the collimator to the detector. This results in a trade-off between sensitivity and resolution of SPECT imaging. Coarse collimators transmit a higher percentage of the gamma rays but with a poor image resolution. For example, an average image resolution for many clinical applications is around one cm. Conversely, fine collimators provide improved imaging resolution, e.g. around three to five mm, but with lower sensitivity and/or higher radiation doses to the patient. Because of the low detection efficiency, rather large doses of tens of milli-curies must be administered to the patient to achieve adequate images.
In PET, a positron-emitting radio-nuclide is administered to a patient. Commonly used PET radio-nuclides include
11
C(&tgr;
1/2
=10 m),
15
O(&tgr;
1/2
=2 m), and
18
F(&tgr;
1/2
=110 m). Positrons interact with electrons in the surrounding tissue of the patient and emit two 511 keV gamma rays in opposite directions, at an angle of almost exactly 180 degrees. These gamma rays interact in position-sensitive gamma ray detectors, and restrict the origin of the gamma rays to the line joining the two detection positions. Accumulation of a large number of such events along many differing lines of position enables the reconstruction of a three-dimensional image of the radioactivity. The position resolution in PET is typically three to five mm, and is limited by the position resolution in the detectors and the range of the positrons in the body of the patient.
An advantage of PET over SPECT is that higher efficiency is achieved since collimators are not required. However, a disadvantage of PET is a higher cost of PET systems relative to SPECT systems.
Both SPECT and PET have the disadvantage that for each nuclear decay event detected (i.e., a single gamma ray in SPECT and coincident 511 keV gamma rays in PET), the location of the radio-nuclide is only determined to a line within the region-of-interest. As a result, a large number of events (e.g., thousands to millions) must be recorded and processed by computers to generate a three-dimensional image of the region-of-interest. Due to the requirement of recording a large number of events, coupled with the relatively low detection efficiencies for each radio-nuclide decay, the overall capabilities of both SPECT and PET are limited.
One method to overcome the very low gamma ray throughput of the collimators and the associated increased radiation dose required in SPECT includes electronic collimation. In electronic collimation, Compton scattering is used to reconstruct three-dimensional images from the processing of Compton direction cones of individual gamma rays detected.
An example of implementing electronic collimation for imaging includes a Compton imager that detects multiple Compton scattering interactions in arrays of 1-2 mm thick silicon strip detectors. Specifically, the energies and positions derived from the multiple Compton scattering interactions in the detectors are analyzed for consistency with Compton kinematics to select the correct interaction sequence and thereby determine the most probable incoming direction cone for the gamma ray. Because of the low efficiency of silicon arrays, an alternate electronic collimation imaging concept includes a calorimeter where scattered gamma rays leaving the silicon array are totally absorbed.
Previous Compton imaging concepts are based on operating in a single photon mode. That is, the three-dimensional image reconstruction is achieved by back-projecting Compton direction cones into the presumed source volume and using computer algorithms to reconstruct the morphology of the radioactivity. In this approach, each decay leads to a large conical-shaped annular volume within which the decay occurred.
An alternative Compton imaging concept uses detectors formed of a plurality of thin position-sensitive silicon detector layers. A Compton scattered electron is tracked through two or more layers of the detectors. With this information, the direction of the incident gamma ray can be reduced from a cone to a segment of a cone, with the advantage of improved image reconstruction.
One disadvantage of this alternative Compton imaging system includes a compromise in energy determination of the scattered electron at a first interaction site, and therefore, an associated degradation in the imaging resolution, due to compromised energy resolution in the multiple, thin detector layers. Another significant limitation of this imaging system is the difficulty of tracking electrons with energies below about one MeV due to multiple scattering events. As a result, the most commonly used nuclear medical radio-nuclides cannot be employed by this alternative concept.
A further Compton imager concept uses noble gas detectors. This imager assumes a first interaction that is a Compton scatter event followed by a photoelectric, full energy absorption, at the second interaction site. Advantages of this approach include the use of large volume, position-sensitive detectors, and the moderately good energy resolution. A significant disadvantage is the rather low efficiencies achievable with gas detectors. Another disadvantage is a poorer imaging resolution resulting from the moderate energy resolution of gas detectors.
A limitation relating to most presently available Compton imager concepts is that these devices require that the incident gamma ray energy is totally absorbed. This is necessary to properly determine the direction cone from the Compton scatter formula at the first interaction site. Requiring that the full energy of the incident gamma rays be absorbed will usually
Larson & Taylor PLC
Manuel George
Shah Devaang
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