Gamma camera

Principle

A gamma camera is an medical imaging device, most commonly used as a imaging device in nuclear medicine. It produces images of the distribution of gamma rays emitted by metastable radionuclides (isotopes), also called metastable nuclear isomers. A gamma ray comprises gamma photons, which are high energetic photons (at least 5000 times those of visible light). They are produced from sub-atomic particle interaction, such as electron-positron annihilation and radioactive decay. (Annihilation is the collision of a positron with an electron, followed by vanishing of both. Two (sometimes more) (gamma) photons are produced moving in almost opposite directions.) A radionuclide can also produce subatomic particles (which give ionization). Excited metastable isomers de-excite with sending a gamma photon mostly within much les than one picosecond, but some isomers are far much slower. These are for example the Technetium isomers ^99mTc (here indicated without atom number; m indicates metastable; half-life 6.01 hours) and ^95mTc (half-life of 61 days) are used in medical and industrial applications.

Fig. 1 Diagrammatic cross section of a gamma camera detector comprising the scintillating crystal layer and the photomultipliers (red circle).

A gamma camera is a complex device consisting of one or more detectors mounted on a gantry. It is connected to an acquisition system for operating the camera and for storing the images. The system counts gamma photons that are absorbed by a crystal in the camera, usually a large flat crystal of NaI with thallium doping in a light-sealed housing. The crystal scintillates in response to incident gamma radiation: when a gamma photon knocks an electron loose from an iodine atom in the crystal, a faint flash of light is produced when the electron again finds a minimal energy state. The initial phenomenon of the excited electron is similar to the photoelectric effect (an electron hitting an atom, with as a result the emission of another electron and back-scatter of the electron with a lower speed) and (particularly with gamma rays) the Compton effect. This is generally an electron hitting an atom, with as a result the emission of a photon and back-scatter of the electron with a lower speed. So actual it is fluorescence, but here for impinging gamma rays. The flash of light must be detected. Photomultiplier tubes (extremely sensitive detectors of UV, near-IR and visible light) behind the crystal detect the fluorescent flashes and a computer sums the fluorescent counts. The computer in turn constructs and displays a two dimensional image of the relative spatial count density on a monitor. This image then reflects the distribution and relative concentration of radioactive tracer elements present in the organs and tissues imaged.

Application

SPECT (Single photon emission computed tomography) machines are based at gamma cameras. Multi-headed gamma cameras can also be used for PET (Positron emission tomography), provided that their hardware and software can be configured to detect 'coincidences' (near simultaneous events on 2 different heads). Gamma camera PET is markedly inferior to PET imaging with a purpose designed PET scanner, as the scintillator crystal has poor sensitivity for the high-energy annihilation photons, and the detector area is significantly smaller. However, given the low cost of a gamma camera and its additional flexibility compared to a dedicated PET scanner, this technique is useful where the expense and resource implications of a PET scanner cannot be justified.

In order to obtain spatial information about the gamma emissions from an imaging subject (e.g. a person's heart muscle cells which have absorbed an intravenous injected radioactive, usually thallium-201 or technetium-99m, medicinal imaging agent) a method of correlating the detected photons with their point of origin is required.

The spatial resolution is a major limitation for heart muscle imaging systems. The thickest normal heart muscle in the left ventricle is about 1.2 cm and most of the left ventricle muscle is about 0.8 cm, always moving and much of it beyond 5 cm from the collimator face. To help compensate, better imaging systems limit scintillation counting to a portion of the heart contraction cycle, called gating, however this further limits system sensitivity (see also PET).

More Info

Generally, each multiplier tube has an exposed face of about 3 inches in diameter and the tubes are arranged in hexagon configurations, behind the absorbing crystal. The electronic circuit connecting the photodetectors is wired so as to reflect the relative coincidence of light fluorescence as sensed by the members of the hexagon detector array. All the photomultiplier tubes which simultaneously detect the (presumed) same flash of light. Thus the spatial location of each single flash of fluorescence is reflected as a pattern of voltages within the interconnecting circuit array. The location of the interaction between the gamma ray and the crystal can be determined by processing the voltage signals from the photomultipliers; in simple terms, the location can be found by weighting the position of each photomultiplier tube by the strength of its signal, and then calculating a mean position from the weighted positions. The total sum of the voltages from each photomultiplier is proportional to the energy of the gamma ray interaction, thus allowing discrimination between different isotopes or between scattered and direct photons.

The conventional method is to place a collimator (device that selects the parallel rays to go through) over the detection crystal/ photomultiplier tubes array. The collimator essentially consists of a thick sheet of lead, typically 1-3 inches thick, with thousands of adjacent holes through it. The individual holes limit photons which can be detected by the crystal to a cone; the point of the cone is at the midline center of any given hole and extends from the collimator surface outward. However, the collimator is also one of the sources of blurring within the image; lead does not totally attenuate incident gamma photons, there can be some crosstalk between holes.

Unlike a lens, as used in visible light cameras, the collimator attenuates most (>99%) of incident photons and thus greatly limits the sensitivity of the camera system. Large amounts of radiation must be present so as to provide enough exposure for the camera system to detect sufficient scintillation dots to form a picture.

Other methods of image localization (pinhole, rotating slat collimator with CdZnTe and others) have been proposed and tested; however, none have entered widespread routine clinical use.

The best current camera system designs can differentiate two separate point sources of gamma photons located a minimum of 1.8 cm apart, at 5 cm away from the camera face. Spatial resolution decreases rapidly at increasing distances from the camera face. This limits the spatial accuracy of the computer image: it is a fuzzy image made up of many dots of detected but not precisely located scintillation.