Page 341 - Biomedical Engineering and Design Handbook Volume 2, Applications

NUCLEAR MEDICINE IMAGING INSTRUMENTATION 319

although other sizes are available. The first scintillation camera was developed by Hal O Anger in
1
1958. Although this system was very crude, it contained the fundamental components of all future
designs: thallium-activated sodium iodide [NaI(T1)] as the primary detector, weighted signals from
an array of photomultiplier tubes to determine the location of detected events, and lead collimation
as the imaging aperture.
The gamma ray detector of the scintillation camera is a large, thin piece of NaI(T1). Although the
crystals originally had a circular cross section, most scintillation cameras now use a rectangular crys-
tal with dimensions as large as 40 × 55 cm. The thickness of NaI(Tl) in most conventional cameras is
9.5 mm, but in systems that are used for coincidence detection, the crystal may be twice as thick.
Sodium iodide is a scintillator that converts gamma ray energy into visible light with a relatively high
efficiency. The amount of light generated is directly proportional to the absorbed energy and the
absorption of one 140-keV gamma ray will yield approximately 5000 visible light photons. There are
a number of advantages associated with NaI(Tl) in addition to its high light output. It efficiently
absorbs the 140-keV gamma rays of 99m Tc with a photopeak efficiency of 85 percent, and it has a
moderate energy resolution. Energy resolution is an important property since it provides the means to
discriminate against scattered radiation. Gamma rays that undergo scattering within the patient
degrade the quality of images. However, scattered gamma rays necessarily have less energy than
unscattered gamma rays and can be selectively eliminated on that basis. Another positive feature of
NaI(Tl) is that it can be manufactured in many shapes and sizes. There are disadvantages though.
Sodium iodide is hygroscopic and actively absorbs water vapor from the air, resulting in a loss of
transparency of the crystal to the scintillation. The detector must be hermetically sealed, and loss of
this seal results in irreparable damage. Another disadvantage is that the persistence of the scintillation
is long enough that it limits the count rate that the detector can accurately handle. Because conven-
tional nuclear medicine imaging requires very low sensitivity collimators, high count rate is rarely an
issue. However, count rate limitations are a very real problem for PET imaging systems where coin-
cidence detection replaces collimation.
Converting the gamma ray energy to visible light is only the first step in the detection process. In
order for the information from the scintillation to be useful it has to be converted into an electronic sig-
nal with a photomultiplier tube (PMT). The PMT is a vacuum tube with a photoemissive surface called
the photocathode. Visible light hitting this surface knocks off electrons. These electrons are accelerated
to an electric terminal called a dynode. The first dynode has a potential of approximately 100 V higher
than the photocathode, and the electrons hit it with enough force to knock off about 4 more new elec-
trons. The next dynode is another 100 V higher so the process is repeated. The same process occurs
over a total series of 9 dynodes, resulting in a signal amplification of 1,000,000. Proportionality is main-
tained throughout this amplification so that the size of the resulting electronic pulse is directly propor-
tional to the energy deposited by the gamma ray.
A scintillation camera needs to be able to record the location of gamma ray events over the large
area of the NaI(Tl) crystal. This requires uniform sampling by an array of photomultiplier tubes. The
PMTs are arranged in a close-packed array that covers the entire surface of the NaI(Tl) crystal
(Fig. 11.1). The PMTs used in scintillation cameras are usually 2 or 3 in across, so that as many as
120 PMTs may be used. PMTs have been manufactured in variety of different cross sections in order
to maximize their areal coverage. Circular, hexagonal, and square tubes have all been used for this
purpose. The signals obtained from the PMTs are used to determine two important properties about
the gamma ray interaction: the event location and the energy deposited in the detector from the inter-
action. The energy information is crucial for normalizing the position signals and for discriminating
against scattered radiation.
While the energy of the interaction can be inferred by directly summing the PMT signals in the
proximity of the event, the estimation of the event location requires that the PMT signals have spa-
tial weighting factors that are determined by the physical location of each PMT within the array.
Separate weighting factors are used for the x and y coordinates, and the location of the event is
determined by summing the weighted PMT signals. This process is referred to as Anger logic, since
it is the scheme developed by Hal Anger in the first scintillation camera. In the initial designs, all
the PMTs in the array participated in the energy and position signal summations. It was subse-
quently found that the signals from PMTs located far from the event had a poor signal-to-noise