17.3.1 Basic Principles
Scintillators are materials that emit visible light or UV radiation after the interaction of ionizing radiation with the material. Scintillators are the oldest type of radiation detectors; Roentgen discovered x-radiation and the fact that x-rays induce scintillation in barium platinocyanide in the same fortuitous experiment. Scintillators are used in conventional film-screen radiography, many direct digital radiographic image receptors, fluoroscopy, scintillation cameras, CT scanners, and positron emission tomography (PET) scanners.
Although the light emitted from a single interaction can be seen if the viewer’s eyes are dark adapted, most scintillation detectors incorporate a means of signal amplification. In conventional film-screen radiography, photographic film is used to amplify and record the signal. In other applications, electronic devices such as photomultiplier tubes (PMTs), photodiodes, or image-intensifier tubes convert the light into electrical signals. PMTs and image-intensifier tubes amplify the signal as well. However, most photodiodes do not provide amplification; if amplification of the signal is required, it must be provided by an electronic amplifier. A scintillation detector consists of a scintillator and a device, such as a PMT, that converts the light into an electrical signal.
When ionizing radiation interacts with a scintillator, electrons are raised to an excited energy level. Ultimately, these electrons fall back to a lower energy state, with the emission of visible light or UV radiation. Most scintillators have more than one mode for the emission of visible light or UV radiation, and each mode has its characteristic decay constant. Luminescence
is the emission of light after excitation. Fluorescence
is the prompt emission of light, whereas phosphorescence
(also called afterglow
is the delayed emission of light. When scintillation detectors are operated in current mode, the prompt signal from an interaction cannot be separated from the phosphorescence caused by previous interactions. When a scintillation detector is operated in pulse mode, afterglow is less important because electronic circuits can separate the rapidly rising and falling components of the prompt signal from the slowly decaying delayed signal resulting from previous interactions.
It is useful, before discussing actual scintillation materials, to consider properties that are desirable in a scintillator.
The conversion efficiency, the fraction of deposited energy that is converted into light or UV radiation, should be high. (Conversion efficiency should not be confused with detection efficiency.)
For many applications, the decay times of excited states should be short. (Light or UV radiation is emitted promptly after an interaction.)
The material should be transparent to its own emissions. (Most emitted light or UV radiation escapes reabsorption.)
The frequency spectrum (color) of emitted light or UV radiation should match the spectral sensitivity of the light receptor (PMT, photodiode, or film).
If used for x-ray and γ-ray detection, the attenuation coefficient (µ) should be large, so that detectors made of the scintillator have high detection efficiencies. Materials with large atomic numbers and high densities have large attenuation coefficients.
The material should be rugged, unaffected by moisture, and inexpensive to manufacture.
In all scintillators, the amount of light emitted after an interaction increases with the energy deposited by the interaction. Therefore, scintillators may be operated in pulse mode as spectrometers. When a scintillator is used for spectroscopy, its energy resolution (ability to distinguish between interactions depositing different energies) is primarily determined by its conversion efficiency. A high conversion efficiency is required for superior energy resolution.
There are several categories of materials that scintillate. Many organic compounds exhibit scintillation. In these materials, the scintillation is a property of the molecular structure. Solid organic scintillators are used for timing experiments in particle physics because of their extremely prompt light emission. Organic scintillators include the liquid scintillation fluids that are used extensively in biomedical research. Samples containing radioactive tracers such as 3H, 14C, and 32P are mixed in vials with liquid scintillators, and the light flashes are detected and counted by PMTs and associated electronic circuits. Organic scintillators are not used for medical imaging because the low atomic numbers of their constituent elements and their low densities make them poor x-ray and γ-ray detectors. When photons in the diagnostic energy range do interact with organic scintillators, it is primarily by Compton scattering.
There are also many inorganic crystalline materials that exhibit scintillation. In these materials, the scintillation is a property of the crystalline structure: if the crystal is dissolved, the scintillation ceases. Many of these materials have much larger average atomic numbers and higher densities than organic scintillators and therefore are excellent photon detectors. They are widely used for radiation measurements and imaging in radiology.
Most inorganic scintillation crystals are deliberately grown with trace amounts of impurity elements called activators. The atoms of these activators form preferred sites in the crystals for the excited electrons to return to the ground state. The activators modify the frequency (color) of the emitted light, the promptness of the light emission, and the proportion of the emitted light that escapes reabsorption in the crystal.
17.3.2 Inorganic Crystalline Scintillators in Radiology
No one scintillation material is best for all applications in radiology. Sodium iodide activated with thallium [NaI(Tl)] is used for most nuclear medicine applications. It is coupled to PMTs and operated in pulse mode in scintillation cameras, thyroid probes, and γ-well counters. Its high content of iodine (Z = 53) and high density provide a high photoelectric absorption probability for x-rays and γ-rays emitted by common nuclear medicine radiopharmaceuticals (70 to 365 keV). It has a very high conversion efficiency; approximately 13% of deposited energy is converted into light. Because a light photon has an energy of about 3 eV, approximately one light photon is emitted for every 23 eV absorbed by the crystal. This high conversion efficiency gives it a very good energy resolution. It emits light very promptly (decay constant, 250 ns), permitting it to be used in pulse mode at interaction rates greater than 100,000/s. Very large crystals can be manufactured; for example, the rectangular crystals of one modern scintillation camera are 59 cm (23 inches) long, 44.5 cm (17.5 inches) wide, and 0.95 cm thick. Unfortunately, NaI(Tl) crystals are fragile; they crack easily if struck or subjected to rapid temperature change. Also, they are hygroscopic (i.e., they absorb water from the atmosphere) and therefore must be hermetically sealed.
PET, discussed in Chapter 19
, requires high detection efficiency for 511-keV annihilation photons and a prompt signal from each interaction because the signals must be processed in pulse mode at high interaction rates. PET detectors are thick crystals of high-density, high atomic number scintillators optically coupled to PMTs. For many years, bismuth germanate (Bi4
, often abbreviated as “BGO”) was the preferred scintillator. The high atomic number of bismuth (Z
= 83) and the high density of the crystal yield a high intrinsic efficiency for the 511-keV positron annihilation photons. The primary component of the light emission is sufficiently prompt (decay constant, 300 ns) for PET. NaI(Tl) was used in early and some less-expensive PET scanners. Today, lutetium oxyorthosilicate (Lu2
O, abbreviated LSO), lutetium yttrium oxyorthosilicate (Lux
O, abbreviated LYSO), and gadolinium oxyorthosilicate (Gd2
O, abbreviated GSO), all activated with cerium, are used in newer PET scanners. Their densities and effective atomic numbers are similar to those of BGO, but their conversion efficiencies are much larger and they emit light much more promptly.
Calcium tungstate (CaWO4
) was used for many years in intensifying screens in film-screen radiography. It was largely replaced by rare-earth phosphors, such as gadolinium oxysulfide activated with terbium. The intensifying screen is an application of scintillators that does not require very prompt light emission, because the film usually remains in contact with the screen for at least several seconds after exposure. Cesium iodide activated with thallium is used as the phosphor layer of many indirect-detection thin-film transistor radiographic and fluoroscopic image receptors, described in Chapters 7
. Cesium iodide activated with sodium is used as the input phosphor and zinc cadmium sulfide activated with silver is used as the output phosphor of image-intensifier tubes in fluoroscopes.
Scintillators coupled to photodiodes are used as the detectors in CT scanners, as described in Chapter 10
. The extremely high x-ray flux experienced by the detectors necessitates current mode operation to avoid dead-time effects. With the rotational speed of CT scanners as high as three rotations per second, the scintillators used in CT must have very little afterglow. Cadmium tungstate and gadolinium ceramics are scintillators used in CT. Table 17-1
lists the properties of several inorganic crystalline scintillators of importance in radiology and nuclear medicine.
TABLE 17-1 INORGANIC SCINTILLATORS USED IN MEDICAL IMAGING
WAVELENGTH OF MAXIMAL EMISSION (nm)
CONVERSION EFFICIENCYa (%)
DECAY CONSTANT (µS)
AFTER GLOW (%)
0.3-5 at 6 ms
83, 32, 8
0.005 at 3 ms
71, 14, 8
Input phosphor of image-intensifier tubes
0.5-5 at 6 ms
Thin-film transistor radiographic and fluoroscopic image receptors
30, 48, 16
Output phosphor of image-intensifier tubes
48, 74, 8
0.1 at 3 ms
Computed tomographic (CT) scanners
20, 74, 8
64, 8, 16
a Relative to Nal(Tl), using a PMT to measure light.
b The light emitted by CsI(Tl) does not match the spectral sensitivity of PMTs very well; its conversion efficiency is much larger if measured with a photodiode.
Data on Nal(Tl), BGO, Csl(Na), Csl(Tl), and CdWO4 courtesy of Saint-Gobain Crystals, Hiram, OH. Data on LSO from Ficke DC, Hood JT, Ter-Pogossian MM. A spheroid positron emission tomograph for brain imaging: a feasibility study. J Nucl Med. 1996;37:1222.
Only gold members can continue reading. Log In