Basic Physics of Nuclear Medicine/Print version

Gamma Camera

The basic design of the most common type of gamma camera used today was developed by an American physicist, Hal Anger and is therefore sometimes called the Anger Camera. It consists of a large diameter NaI(Tl) scintillation crystal which is viewed by a large number of photomultiplier tubes.

A block diagram of the basic components of a gamma camera is shown below:

The crystal and PM Tubes are housed in a cylindrical shaped housing commonly called the camera head and a cross-sectional view of this is shown in the figure. The crystal can be between about 25 cm and 40 cm in diameter and about 1 cm thick. The diameter is dependent on the application of the device. For example a 25 cm diameter crystal might be used for a camera designed for cardiac applications while a larger 40 cm crystal would be used for producing images of the lungs. The thickness of the crystal is chosen so that it provides good detection for the 140 keV gamma-rays emitted from ^99mTc - which is the most common radioisotope used today.

Scintillations produced in the crystal are detected by a large number of PM tubes which are arranged in a two-dimensional array. There is typically between 37 and 91 PM tubes in modern gamma cameras. The output voltages generated by these PM tubes are fed to a position circuit which produces four output signals called ±X and ±Y. These position signals contain information about where the scintillations were produced within the crystal. In the most basic gamma camera design they are fed to a cathode ray oscilloscope (CRO). We will describe the operation of the CRO in more detail below.

Before we do so we should note that the position signals also contain information about the intensity of each scintillation. This intensity information can be derived from the position signals by feeding them to a summation circuit (marked ? in the figure) which adds up the four position signals to generate a voltage pulse which represents the intensity of a scintillation. This voltage pulse is commonly called the Z-pulse (or zee-pulse in American English!) which following pulse height analysis (PHA) is fed as the unblank pulse to the CRO.

So we end up with four position signals and an unblank pulse sent to the CRO. Let us briefly review the operation of a CRO before we continue. The core of a CRO consists of an evacuated tube with an electron gun at one end and a phosphor-coated screen at the other end. The electon gun generates an electron beam which is directed at the screen and the screen emits light at those points struck by the electron beam. The position of the electron beam can be controlled by vertical and horizontal deflection plates and with the appropriate voltages fed to these plates the electron beam can be positioned at any point on the screen. The normal mode of operation of an oscilloscope is for the electron beam to remain switched on. In the case of the gamma camera the electron beam of the CRO is normally switched off - it is said to be blanked.

When an unblank pulse is generated by the PHA circuit the electron beam of the CRO is switched on for a brief period of time so as to display a flash of light on the screen. In other words the voltage pulse from the PHA circuit is used to unblank the electron beam of the CRO.

So where does this flash of light occur on the screen of the CRO? The position of the flash of light is dictated by the ±X and ±Y signals generated by the position circuit. These signals as you might have guessed are fed to the deflection plates of the CRO so as to cause the unblanked electron beam to strike the screen at a point related to where the scintillation was originally produced in the NaI(Tl) crystal. Simple!

The gamma camera can therefore be considered to be a sophisticated arrangement of electronic circuits used to translate the position of a flash of light in a scintillation crystal to a flash of light at a related point on the screen of an oscilloscope. In addition the use of a pulse height analyser in the circuitry allows us to translate the scintillations related only to photoelectric events in the crystal by rejecting all voltage pulses except those occurring within the photopeak of the gamma-ray energy spectrum.

Let us summarise where we have got to before we proceed. A radiopharmaceutical is administered to the patient and it accumulates in the organ of interest. Gamma-rays are emitted in all directions from the organ and those heading in the direction of the gamma camera enter the crystal and produce scintillations (note that there is a device in front of the crystal called a collimator which we will discuss later). The scintillations are detected by an array of PM tubes whose outputs are fed to a position circuit which generates four voltage pulses related to the position of a scintillation within the crystal. These voltage pulses are fed to the deflection circuitry of the CRO. They are also fed to a summation circuit whose ouput (the Z-pulse) is fed to the PHA and the output of the PHA is used to switch on (that is, unblank) the electron beam of the CRO. A flash of light appears on the screen of the CRO at a point related to where the scintillation occurred within the NaI(Tl) crystal. An image of the distribution of the radiopharmaceutical within the organ is therefore formed on the screen of the CRO when the gamma-rays emitted from the organ are detected by the crystal.

What we have described above is the operation of a fairly traditional gamma camera. Modern designs are a good deal more complex but the basic design has remained much the same as has been described. One area where major design improvements have occurred is the area of image formation and display. The most basic approach to image formation is to photograph the screen of the CRO over a period of time to allow integration of the light flashes to form an image on photographic film. A stage up from this is to use a storage oscilloscope which allows each flash of light to remain on the screen for a reasonable period of time.

The most modern approach is to feed the position and energy signals into the memory circuitry of a computer for storage. The memory contents can therefore be displayed on a computer monitor and can also be manipulated (that is processed) in many ways. For example various colours can be used to represent different concentrations of a radiopharmaceutical within an organ.

The use of digital image processing is now widespread in nuclear medicine in that it can be used to rapidly and conveniently control image acquisition and display as well as to analyse an image or sequences of images, to annotate images with the patient's name and examination details, to store the images for subsequent retrieval and to communicate the image data to other computers over a network.

Essential elements of a modern gamma camera. MCA: Multi-Channel Analyzer

The essential elements of a modern gamma camera are shown in the figure on the left. Gamma rays emitted by the patient pass through the collimator and are detected within the camera head, which generates data related to the location of scintillations in the crystal as well as to the energy of the gamma rays. This data is then processed on-the-fly by electronic hardware which corrects for technical factors such as spatial linearity, PM tube drift and energy response so as to produce an imaging system with a spatially-uniform sensitivity and distortion-free performance. A multichannel analyzer (MCA) is used to display the energy spectrum of gamma rays which interact inside the crystal. Since these gamma rays originate from within the patient, some of them will have an energy lower than the photopeak as a result of being scattered as they travel through the patient's tissues - and by other components such as the patient table and structures of the imaging system. Some of these scattering events may involve just glancing interactions with free electrons, so that the gamma rays loose only a small amount of energy. These gamma rays may have an energy just below that of the photopeak so that their spectrum merges with the photopeak. The photopeak for a gamma camera imaging a patient therefore contains information from spatially-correlated, unattenuated gamma rays (which is the information we want) and from spatially-uncorrelated, scattered gamma rays. The scattered gamma rays act like a variable background within the true photopeak data and the effect is that of a background haze in gamma camera images.

While scatter may not be a significant problem in planar scintigraphy, it has a strong bearing on the fidelity of quantitative information derived from gamma camera images and is a vital consideration for accurate image reconstruction in emission tomography. It is the unattenuated gamma rays (also called the primary radiation) that contain the desired information, because of their direct dependence on radioactivity. The scatter situation is illustrated in more detail in the figure on the right, which shows estimates of the primary and scatter spectra for 99mTc in patient imaging conditions. Such spectral estimates can be generated using Monte Carlo methods. It is seen in the figure that the energy of the scattered radiation forms a broad band, similar to the Compton Smear described previously, which merges into and contributes substantially to the detected photopeak. The detected photopeak is therefore an overestimate of the primary radiation. The extent of this overestimate is likely to be dependent on the specific imaging situation because of the different thicknesses of tissues involved. It is clear however that the scatter contribution within the detected photopeak needs to be accounted for if an accurate measure of radioactivity is required.

Detected gamma ray energy spectrum for 99mTc (green) with estimates of the scatter (blue) and primary (red) components.

Gamma ray energy spectrum for 99mTc, with energy discrimination settings of 92-126 keV for scatter estimation (blue) and of 126-154 keV, centred on 140 keV, for the photopeak (red).

One method of compensating for the scatter contribution is illustrated in the figure on the left and involves using data from a lower energy window as an estimate for subtraction from the photopeak, i.e. Primary Counts = Photopeak Window Counts - k (Scatter Window Counts) where k is a scaling factor to account for the extent of the scatter contribution. This approach to scatter compensation is referred to as the Dual-Energy Window (DEW) method. It can be implemented in practice by acquiring two images, one for each energy window, and subtracting a fraction (k) of the scatter image from the photopeak image. For the spectrum shown above, it can be seen that the scaling factor, k, is about 0.5, but it should be appreciated that its exact value is dependent on the scattering conditions. Gamma cameras which use the DEW method therefore generally provide the capability of adjusting k for different imaging situations. Some systems use a narrower scatter window than that illustrated, e.g. 114-126 keV, with a consequent increase in k to about 1.0, for instance. A host of other methods of scatter compensation have also been developed. These include more complex forms of energy analysis such as the Dual-Photopeak and the Triple-Energy Window techniques, as well as approaches based on deconvolution and models of photon attenuation. An excellent review of these developments is provided in Zaidi & Koral (2004).

Some photographs of gamma cameras and related devices are shown below:

A single-headed gamma camera.	Another single-headed gamma camera.	The NaI crystal of a gamma camera.	The cathode ray oscilloscope (CRO) of a gamma camera.
The image processing system of a gamma camera.	A dual-headed gamma camera.	Another view of a dual-headed gamma camera.	The image acquisition and processing console of a dual-headed gamma camera.

We will continue with our description of the gamma camera by considering the construction and purpose of the collimator.