The Gamma Camera: Basic Principles

A

General Concepts of Radionuclide Imaging

The purpose of radionuclide imaging is to obtain a picture of the distribution of a radioactively labeled substance within the body after it has been administered (e.g., by intravenous injection) to a patient. This is accomplished by recording the emissions from the radioactivity with external radiation detectors placed at different locations outside the patient. The preferred emissions for this application are γ rays in the approximate energy range of 80 to 500 keV (or annihilation photons, 511 keV). Gamma rays of these energies are sufficiently penetrating in body tissues to be detected from deep-lying organs, can be stopped efficiently by dense scintillators, and are shielded adequately with reasonable thicknesses of lead (see Fig. 6-17 —soft tissue has attenuation properties similar to water). Alpha particles and electrons (β particles, Auger and conversion electrons) are of little use for imaging because they cannot penetrate more than a few millimeters of tissue. Therefore they cannot escape from within the body and reach an external radiation detector, except from very superficial tissues. Bremsstrahlung (see Fig. 6-1 ) generated by electron emissions is more penetrating, but the intensity of this radiation generally is very weak.

Imaging system detectors must therefore have good detection efficiency for γ rays. It is also desirable that they have energy discrimination capability, so that γ rays that have lost positional information by Compton scattering within the body can be rejected based on their reduced energy (see Chapter 6 , Section C.3). A sodium iodide [NaI(Tl)] scintillation detector (see Chapter 7 , Section C) provides both of these features at a reasonable cost; for this reason it is currently the detector of choice for radionuclides with γ-ray emissions in the range of 80-300 keV.

The first attempts at radionuclide “imaging” occurred in the late 1940s. An array of radiation detectors was positioned on a matrix of measuring points around the head. Alternatively, a single detector was positioned manually for separate measurements at each point in the matrix. These devices were tedious to use and provided only very crude mappings of the distribution of radioactivity in the head (e.g., left-side versus right-side asymmetries).

A significant advance occurred in the early 1950s with the introduction of the rectilinear scanner by Benedict Cassen (see Fig. 1-2 ). With this instrument, the detector was scanned mechanically in a raster-like pattern over the area of interest. The image was a pattern of dots imprinted on a sheet of paper by a mechanical printer that followed the scanning motion of the detector, printing the dots as the γ rays were detected.

The principal disadvantage of the rectilinear scanner was its long imaging time (typically many minutes) because the image was formed by sequential measurements at many individual points within the imaged area. The first gamma-ray “camera” capable of recording at all points in the image at one time was described by Hal Anger in 1953. He used a pinhole aperture in a sheet of lead to project a γ-ray image of the radionuclide distribution onto a radiation detector composed of a NaI(Tl) screen and a sheet of x-ray film. The film was exposed by the scintillation light flashes generated by the γ rays in the NaI(Tl) screen. Unfortunately, this detection system (especially the film component) was so inefficient that hour-long exposures and therapeutic levels of administered radioactivity were needed to obtain satisfactory images.

In the late 1950s, Anger replaced the film-screen combination with a single, large-area, NaI(Tl) crystal and a photomultiplier (PM) tube assembly to greatly increase the detection efficiency of his “camera” concept. This instrument, the Anger scintillation camera, or gamma camera, has been substantially refined and improved since that time. Although other ideas for nuclear-imaging instruments have come along since then, none, with the exception of modern positron emission tomography systems (see Chapter 18 ), has matched the gamma camera for a balance of image quality, detection efficiency, and ease of use in a hospital environment. The gamma camera has thus become the most widely used nuclear-imaging instrument for clinical applications.

B

Basic Principles of the Gamma Camera

1

System Components

Figure 13-1 illustrates the basic principles of image formation with the gamma camera. The major components are a collimator, a large-area NaI(Tl) scintillation crystal, a light guide, and an array of PM tubes. Two features that differ from the conventional NaI(Tl) counting detectors described in Chapter 12 are crucial to image formation. The first is that an imaging collimator is used to define the direction of the detected γ rays. The collimator most commonly consists of a lead plate containing a large number of holes. By controlling which γ rays are accepted, the collimator forms a projected image of the γ-ray distribution on the surface of the NaI(Tl) crystal (see Section B.3). The second is that the NaI(Tl) crystal is viewed by an array of PM tubes, rather than a single PM tube. Signals from the PM tubes are fed to electronic or digital position logic circuits, which determine the X-Y location of each scintillation event, as it occurs, by using the weighted average of the PM tube signals (see Section B.2).

Individual events also are analyzed for energy, E , by summing the signals from all PM tubes. When the pulse amplitude of an event falls within the selected energy window, it is accepted and the X and Y values are binned into a discrete two-dimensional array of image elements, or pixels. An image is formed from a histogram of the number of events at each possible X-Y location. Large numbers of events are required to form an interpretable image because each pixel must have a sufficient number of counts to achieve an acceptable signal-to-noise level. Because images often are formed in 64- × 64-pixel or 128- × 128-pixel arrays, the counting requirements are some 10 ³ to 10 ⁴ times higher than for a simple counting detector.

Images are displayed on a computer monitor, where image brightness and contrast may be manipulated and different color tables may be employed. More sophisticated digital image processing is discussed in Chapter 20 .

Most modern gamma cameras are completely digital, in the sense that the output of each PM tube is directly digitized by an analog-to-digital converter (ADC). The calculation of X-Y position and pulse-height are performed in software based on the digitized PM tube signals, and errors in energy and positioning caused by noise and pulse distortions caused by the analog positioning circuitry are eliminated. This approach also permits improved handling of pulse pile-up at high counting rates, as described in Section B.2.

The gamma camera can be used for static imaging studies, in which an image of an unchanging radionuclide distribution can be recorded over an extended imaging time (e.g., minutes). Single contiguous images of the whole body can be obtained by scanning the gamma camera across the entire length of the patient. This can be achieved by moving either the bed or the gamma camera while adjusting the event positioning computation to account for this movement. Clinically important whole-body studies include bone scans of the skeleton, and the localization of tumors or their metastases in the body.

The gamma camera also can be used for dynamic imaging studies, in which changes in the radionuclide distribution can be observed, as rapidly as several images per second. This allows physiologic information to be obtained, such as the rate of tracer uptake or clearance from an organ of interest. Images also can be synchronized to electrocardiogram signals, permitting images of the heart in different phases of the cardiac cycle to be formed. These gated images can provide important information on cardiac function.