<<

for Medical Imaging Enabling life-saving diagnostics

by Steven J. Duclos

cintillators are high lumi- nescent materials that convert X- rays to visible . These materials have been used since the turn of the century to dramat- S ically enhance the sensitivity of film cartridges used in medical diagnos- tics. Coupling to photodi- odes, charge coupled devices (CCDs), and photomultiplier tubes (PMTs) is an effi- cient way to convert X-rays to electrical signals. This has ushered in an era of diag- nostic imaging which requires advanced scintillators with enhanced properties, such as increased X-ray conversion effi- ciency, faster decay times, lower afterglow, and greater stability in the radiation field. In addition to medical diagnostics, scintillators are used in X-ray security apparatus, industrial inspection, and in high-energy physics calorimetry. Therefore, the field of scintillator mate- rials is quite extensive, and for more com- plete discussions the reader is directed to reviews by Weber (1), Blasse, et al. (2), FIG. 1. Room temperature luminescent spectra for scintillators used for radiography. These are the spectra and Lempicki, et al. (3). excited by 60 kVp X-rays. The spectra have been normalized to their peak output. Although the storage phos- The present review is restricted to phor BaFBr:Eu is not measured when directly excited by X-rays, its output spectrum is essentially the same inorganic scintillator materials that are when a readout is used to excite the stored energy. currently used in the field of medical

Table 1: Scintillators used in mediccal X-ray detection and their properties.

34 The Electrochemical Society Interface • Summer 1998 imaging. Despite this restriction we will holes have given up sufficient energy tron or hole followed by excitation of the encounter a variety of , oxysul- to the lattice they can pair to form exci- luminescent center. Such delayed emis- fide, fluoride, and alkali halide mate- tons. While the theoretical limit of sion is referred to as afterglow, and can rials. Depending on the application the the number of created is be quite deleterious to a scintillator in scintillators are in the form of single n’e-h = Ex-ray/Egap, where Egap is the scin- certain applications. On the other hand, , polycrystalline ceramics, or tillator , the measured storage phosphors used in radiography β β powders. Optimization of a scintillator number (4) is ne-h = n’e-h/ , where is depend on the efficient trapping and for an application requires an under- typically greater than 2. It is the excitons storage of these charges. This second step standing of and hole transport, that release their energy to the lumines- in the scintillation process is the subject electronic defect creation, lumines- cent centers, so the scintillation effi- of most scintillator materials engi- cence processes, and optical transport ciency will be optimized for materials neering. The control of trapping defects β in materials. Following a brief discus- with smaller s and lower Egap. The role through processing and secondary sion of the scintillation process and of in the X-ray stopping can have a profound effect on properties required by scintillators, rep- process clearly indicates that those mate- critical properties of a scintillator. resentative materials will be reviewed rials with higher electron will The third step is the radiative emis- by medical imaging modality: (1) radi- stop X-rays in shorter distances. There- sion of a from the luminescent center. The total probability of a transi- tion occurring from an to the ground state is the sum of the proba- bility of the radiative emission of a photon, Ar, and a non-radiative proba- bility, Anr. The competition between Ar and Anr is a complex function of the strength and symmetry of the field around the luminescent center, the host phonon spectrum, the transition energy, and that energy’s proximity to other energy levels in the material. The probability Ar will be largest when the initial and final states are of opposite parity and the same total spin (5). Theo- retically, the radiative probability is oth- erwise zero, however, mixing of higher energy states into the excited state gener- ally results in a finite radiative transition probability. The decay rate of the scintil- lator once the X-ray excitation has been turned off is determined by e-t/τ, where τ = 1/Ar. For medical applications such as PET, which involve the counting of indi- vidual X-ray , the counting rate and signal can be limited by the scintil- FIG. 2. Room temperature luminescent spectra for scintillators used for computed tomography. These are the spectra excited by 60 kVp X-rays. The spectra have been normalized to their peak output. lator speed. Processes that increase Anr result in lower of the luminescent center and lower scintilla- ography, (2) computed tomography fore scintillator materials typically tion efficiency. Non-radiative processes (CT), and (3) nuclear cameras and employ heavy elements, and have densi- include the energy relaxation via the emission tomography (PET). ties greater than 5 g/cm3. emission of phonons and transfer of the TABLE 1 lists the scintillators that will be The second step in the scintillation excited state energy to other centers covered in this review, as well as some process is the capture of the exciton nearby, including impurities. of their physical properties. energy by the luminescent center or acti- The three steps outlined above vator. As in phosphors, such determine the emission efficiency of The Scintillation Process and Its centers can be intrinsic to the host lattice the scintillator material. There are other Relation to Materials Properties or can be ions deliberately doped into important materials aspects that deter- the lattice. Not every exciton results in mine the efficiency of the scintillator The generation of visible photons an excited luminescent center since elec- when it is incorporated into an X-ray from X-rays is a three step process. The tronic defects also trap electrons and detection system. For example, the par- X-ray must first be stopped and its holes, and often result in release of this ticle size of a scintillator screen used in energy dissipated into the material. energy through non-radiative processes. conjunction with film will affect the Depending on the energy of the X-ray, Such defect trapping can result in color light properties and, if not Ex-ray, its dissipation results from the centers and optical absorption at the optimized, can degrade the spatial reso- , Compton scat- luminescent , which results lution of the device. For crystalline tering, or . In each case in radiation induced damage to the scin- scintillators the light output can be the energy is ultimately dissipated into tillator efficiency. If the defect trap is affected by the efficiency of light reflec- phonons and mobile electrons and shallow, thermal processes can result in tion and refraction at interfaces. A scin holes. When these free electrons and the delayed regeneration of a free elec- (continued on next page)

The Electrochemical Society Interface • Summer 1998 35 tillation photon, which is emitted from the luminescent center isotropically, can undergo many interactions with the scintillator wall before encoun- tering the crystal face coupled to the . Optical transparency is especially important in geometries where the photon path length is long. Internal scattering will increase the path length and make the material more sensitive to radiation induced optical absorption. For example, scintil- lator transparency and efficient wall reflectivity is critical to detector pixels used in computed tomography where spatial resolution and protection of the photodiode from direct hits by trans- mitted X-rays requires a high thickness to width aspect ratio.

Radiography

Radiography refers to the technique where the patient is placed between an X-ray source and an X-ray detector, typ- FIG. 3. Room temperature luminescent spectra for scintillators used for nuclear cameras and PET. These are ically a film cartridge. Anatomic fea- the spectra excited by 60 kVp X-rays. The spectra have been normalized to their peak output. tures, such as bones and tumors, differ in X-ray absorption from their sur- roundings and their shadows are pro- with the resulting loss of spatial resolu- tured on CCDs, generating digital data jected onto the detector. X-ray film tion due to light scattering. Synthesis which can be transmitted and analyzed itself is not sufficiently dense to stop techniques are used that minimize the remotely. CsI:Tl has a broad emission the 30-100 keV X-rays used, so scintil- creation of electronic defects since (see FIG. 1) from 350 nm to 700 nm lator screens are typically employed to excessive afterglow will result in a ghost with a peak at 550 nm. The emission is j 2 increase the stopping power. The film image of one patient onto the film of from parity allowed (sp s ) transi- 3 j 1 then registers the light from the scintil- the next patient, which is a situation tions P1 S0 and has a lifetime of lator rather than the X-rays themselves. that must be avoided. 980 nsec, which is fast enough to allow More recently film has been replaced In 1983, the storage these devices to be used for fluo- with electronic systems which also BaFBr:Eu was introduced (8) which roscopy. The allowed nature of this depend on scintillators for efficient cap- stores the X-ray energy in electronic transition contributes to its high light ture of the transmitted X-rays. One defects. These defects are emptied at a output. Its application in image intensi- device is the storage screen which, after later time by scanning with a laser fiers is aided by the blue component of 2+ exposure to X-rays, is read out with a which releases energy to the Eu acti- its scintillation light, which is particu- 2+ laser system and digitized. The scintilla- vator. The Eu blue emission (FIG. 1) is larly efficient in generating electrons in tion process is critical to all of these proportional to the X-ray flux stored on the . Undoped CsI also devices and the materials used will be the panel and is detected with a blue luminesces from states that trap the described below. sensitive (PMT). free excitons generated by X-rays. These The sensitivity and X-ray capture of An image is generated by recording the self-trapped excitons recombine and a film screen is enhanced by sand- PMT output versus the position of the release that energy radiatively. How- wiching the film between layers of a excitation laser. The scintillator has a ever, CsI without doping has consider- 3 2+ powdered scintillator, typically gado- density of 5.20 g/cm , and the Eu ably lower light output due to linium oxysulfide doped with excited state has a lifetime of 800 nsec, competition by non-radiative defects which determines the rate at which the for the exciton energy. as an , Gd2O2S:Tb (6). The emission spectrum of Gd O S:Tb under panel can be read by the laser. Consider- 2 2 able effort has gone into determining Computed Tomography X-ray excitation is shown in FIG. 1. It has a strong peak at 540 nm due to the and optimizing the defect responsible 5 j 7 for the energy storage (9, 10). Computed tomography (CT) mea- D4 FJ transitions within the f man- ifold of the Tb ion. This wavelength is Cesium iodide activated with thal- sures the attenuation of X-rays through well matched to green sensitive medical lium, CsI:Tl, is used as a scintillator in the body as an X-ray source and film. The morphology and size of the image intensifiers. A of CsI:Tl opposed detector rotate 360° in a plane particles in the scintillator film is crit- is placed in front of a photocathode around the patient. Images of internal ical to avoid excessive light scattering which generates electrons from the organs are generated by back projecting which degrades the resolution of the scintillation light. These electrons are detector readings at the collected system (7). The thickness of the scintil- then electrostatically amplified and angles. Typically fan beam X-ray lator layer must also be carefully con- converted back to light by a cathodolu- sources and arc detectors are employed trolled, and it is designed to balance the minescent material for viewing of the to increase the information gathered added stopping power of a thicker layer image. These images can also be cap- and improve image quality. The scintil-

36 The Electrochemical Society Interface • Summer 1998 lators used in CT must be capable of that are sintered to transparency (12). Li2GeF6 is used as a HIP aid, which measuring attenuation differences of 1 The structure of (YxGd(1-x))2O3 is cubic improves the bulk transparency and scin- part in 1000, show little afterglow, and at its preparation temperatures for x < tillator light output (16). be stable in the radiation field over the 0.4. The isotropic optical properties of time of an extended scan (typically 60 the cubic structure allows for sintering to Nuclear Cameras and PET seconds). For example, afterglow will complete transparency. The emission of result in an image artifact if a dense the Eu is from 550 nm to 720 nm (FIG. 2) Nuclear cameras and positron emis- object (such as a bone) rapidly eclipses with a strong peak at 611 nm, which is sion tomography (PET) both work by the X-ray intensity and the signal in well suited to detection by . detecting high energy gamma rays the detector is dominated by the scin- The lifetime of this state is slightly less emitted from a patient that has ingested tillator afterglow rather than the X-ray than 1 ms. At x = 0.3 the material density short-lived radio-pharmaceuticals. photons passing through the anatomy is 5.95 g/cm3, sufficient to stop 99.96% Detecting the spatial distribution of these and striking the detector. Three scintil- of 70 keV X-rays in 3 mm. pharmaceuticals yields information on lators currently used in CT machines The ceramic process allows for the biological activity in the organs of interest, will be reviewed here: homogeneous doping of additives, in although the spatial resolution of these tungstate, a single crystal scintillator addition to the activator, which are cameras is less than for CT. In nuclear with intrinsic emissions, and two useful for the control of critical scintil- cameras a large area detector is comprised ceramic scintillators, yttria-gadolinia lation properties (13). Similar to the of scintillator segments coupled to PMTs. oxide activated by Eu, and well known process of doping semi- Interpolation of the PMT outputs deter- oxysulfide activated by Pr. conductors to control electronic prop- mines the position that the

Cadmium tungstate, CdWO4, is a erties, these additives are used to affect intercepted the camera. Energy resolution colorless oxide single crystal which is the electronic defects and energy is used to eliminate those gamma rays grown at 1240°C by Czochralski tech- migration between them, which are that have been scattered into the detector. niques (11). Its density is 7.99 g/cm3 responsible for afterglow and radiation Linearity of the scintillator signal with X- and has a broad emission between 400 damage. As an example, the addition of ray energy is critical for proper discrimina- and 600 nm peaking at 480 nm (FIG. 2). either Pr or Tb to the (YxGd(1-x))2O3:Eu tion between scattered and direct gamma The light output is adequate for CT scintillator can reduce afterglow by rays. In PET the radio-pharmaceutical gen- scanners, and its afterglow is among more than an order of magnitude, even erates which are rapidly annihi- the lowest of the materials used in at levels of 100 ppm. The Eu activator is lated, generating two 512 keV gamma these detectors. The emission originates a strong electron trap which can leave rays emitted in opposite directions. A ring from the WO4 tetrahedra in this mono- holes available for trapping at defect of detectors surrounds the patient and clinic structure material. CdWO4 sites. The addition of Pr or Tb, which timing coincidence is used to detect an exhibits a cleavage plane along the are strong hole traps, competes for . An image of annihilation (010) crystallographic plane which is so holes with the afterglow trap and decay events, and therefore density of the radio weak that the material can be cut with non-radiatively when associated with a pharmaceuticals, is generated by sum- a razor blade. Chipping and cracking nearby Eu ion. Although this process ming lines projected between detector ele- during dicing into pixels can result in also decreases the efficiency of the scin- ments that detect coincident gamma rays. undesirable pixel-to-pixel nonuniformi- tillator, the high intrinsic efficiency of For scintillators used in PET very fast ties in light collection efficiency. this material allows for this tradeoff decay times (τ < 300 ns) and high density Defects generated in the material (14). Uniformity of scintillation proper- (ρ> 7 g/cm3) are critical for proper registra- upon exposure to X-rays result in ties demands that such potent additives tion of gamma ray coincidence. absorption into the visible region and be uniformly mixed into the lattice, activated with thal- self-absorption of the scintillation emis- which is accomplished in the ceramic lium, NaI:Tl, is used in nuclear cameras, sion. Therefore, CdWO4 has a radiation process by coprecipitating the additives although it has relatively low density of damage which must be compensated along with the host ions. 3.67 g/cm3. It can be grown in large for during image reconstruction. Recent A second ceramic now being used as a single crystal boules that makes possible developments in the growth techniques scintillator in CT scanners is gadolinium the thick transparent scintillators of CdWO4 have reduced this radiation oxysulfide activated with Pr (15), required for stopping gamma rays. Rapid damage susceptibility. These include the Gd2O2S:Pr. This material has a predomi- growth rates of these crystals also results minimization of impurity defects and nant emission at 513 nm (FIG. 2), which in low cost of manufacture. The emission 3 j 3 3 reduction of Cd vacancy defects by results from the P0 HJ, FJ transitions is a broad band extending from 300 nm replacement of the CdO evaporated within the Pr ion, which is nominally to 500 nm, peaking at 415 nm (FIG. 3), during high temperature growth (11). doped to 1%. This transition is spin- which is well matched to blue sensitive 3 t 1 Extensive doping of the crystal to allowed (although it still is an f-f configu- PMTs. The decay time of the P1 S0 improve properties such as radiation ration transition so is parity disallowed), transition on the Tl+ activator is 230 ns damage have not been accomplished, which makes it reasonably fast with a which is adequate for X-ray photon most likely due to difficulties in uni- decay time of 3 µs. The structure of this counting devices. This material is a pop- formly doping the single crystals. material is hexagonal, which results in ular choice for scintillator applications Solid of yttria and decreased light transmission due to scat- that are not sensitive to the effects of its gadolinia with Eu as an activator, tering at index mismatches at the ran- radiation damage due to color center (Y,Gd)2O3:Eu, represent a dense version domly-oriented grain boundaries. Hot formation and its relatively low density. of the high quantum efficiency isostatic pressing (HIP) is used to improve Also, the material is hygroscopic and Y2O3:Eu red phosphor. Its melting tem- the transparency (16), and doping has must be protected from ambient atmos- perature of 2439°C makes single crystal been used to improve the scintillation phere. Notice that the emission in growth difficult, so these materials are properties. Ce is added at approximately NaI:Tl has a considerably shorter wave- prepared as polycrystalline ceramics 100 ppm for afterglow reduction. (continued on next page)

The Electrochemical Society Interface • Summer 1998 37 length than CsI:Tl which is due to the times ranging from 40 to 60 ns in the port and postal security, oil , high sensitivity of the exposed s2 elec- orthosilicates. The crystal structure of and high energy physics experimenta- + trons on the Tl ion to the stronger Gd2SiO5 is monoclinic which makes tion. As pointed out throughout this crystal field of the NaI lattice. the material susceptible to cleavage review, no ideal scintillator exists that has germanate, Bi4Ge3O12, is an during dicing. These materials are the properties required by all applica- intrinsic emission scintillator with a high grown as single crystals using the tions, and it is expected that specialized density of 7.13 g/cm3 and rapid decay Czochralski technique, at rates of about scintillator materials will continued to be ■ time of 300 ns. Although its scintillation 1 mm/h. In the case of Lu2SiO5 the seg- optimized for specific applications. efficiency is considerably lower than regation coefficient of Ce is 0.2, and NaI:Tl, its energy resolution and linearity care must be taken to adjust the melt are excellent. Its density allows for the composition during growth to com- absorption of 95% of 512 keV gamma pensate for the increasing Ce concen- About the Author rays in 3 cm thickness, which is ideal for tration (18). Homogeneous doping is coincidence counting applications such important to ensure linearity of light Dr. Steven J. Duclos is is a physicist in the Ceramic Laboratory of the General as PET. Bi4Ge3O12 has a broad band output with X-ray energy. For example, emission (FIG. 3) from 400 nm to 550 nm a lower concentration of activator on Electric Corporate Research and Devel- peaking at 480 nm, which is well cou- the front volume of the crystal, where opment Center in Schenectady, New pled to the PMT sensitivity. The material lower energy X-rays are preferentially York. His current work focuses on has a relatively low melting temperature captured, would result in a lower than improvements in the properties of the (1044°C) for an oxide material, and expected light output for these X-rays. ceramic scintillator used in General single crystals are typically grown by the Electric’s Computed Tomography Czochralski method (17). Summary imaging machines. A third material used in high energy X-ray detection is the family of We have covered here a few of the doped gadolinium and lutetium scintillator materials used in the med- References orthosilicates, (GdxLu(1-x))2SiO5:Ce. ical field. The list is not exhaustive, and These materials make use of the parity advancements in the understanding of 1. M. J. Weber, Calorim. High Energy Phys., Proc. 5 t 4 5th Int. Conf., H. A. Gordon and D. Rueger, allowed d f transitions within the luminescence physics and material pro- eds., World Scientific: Singapore, p. 17 (1995). Ce ion. The emission for Gd2SiO5:Ce cessing is expected to result in scintilla- 2. G. Blasse and B. C. Grabmaier, Luminescent (see FIG. 2) is broad, extending from tors with improved properties. These Materials (Springer-Verlag, New York 1994). 3. A. Lempicki, A. J. Wojtowicz, and C. 380 nm to 500 nm peaking at 430 nm. advances will also be driven by other Brecher, Wide-Gap Luminescent Materials, These transitions are among the fastest scintillator uses, not discussed here, Stanley R. Rotman, ed., Kluwer Academic in the rare-earth series, with the decay such as industrial X-ray inspection, air- Publishers, Boston, p. 235 (1997). 4. P. A. Rodnyi, P. Dorenbos, and C. W. E. Van Eijk, Scintillator and Phosphor Materials, M. J. Weber, P. Lecoq, R. C. Ruchti, C. Woody, W. M. Yen, R. Zhu, Eds., Materials Research Society Symposium Proceedings 348, p. 379 (1994). The Total for Fuel Cell 5. G. F. Imbusch, Luminescence , M. D. Lumb, ed., Academic, p. 1 (1978). 6. K. A. Wickersheim, R. V. Alves, and R. A. Buchanan, IEEE Trans. Nucl. Sci. 17, 57 (1970). Test and Development 7. J. D. Kingsley, Real-Time Radiologic Imaging: Medical and Industrial Applications, ASTM Electronic Loads with Windows Software STP 716, D. A. Garrett and D. A. Bracher, Eds., American Society for Testing and Control: Measure: Materials, p. 98 (1980). 8. N. Kotera, S. Eguchi, J. Miyahara, S. Mat- Current Voltage/Time sumoto, H. Kato, United States Patent No. 4,239,968 (1980). Voltage Current Density 9. H. H. Ruter, H. von Seggern, R. Reiniger, and Power iR Losses V. Saile, Phys. Rev. Lett., 65, 2438 (1990). 10. F. K. Koschnick, J. M. Spaeth, R. S. Eachus, Mass Flow Tafel Data W. G. McDugle, R. H. D. Nuttal, Phys. Rev. Temperature MEA Resistance Lett., 67, 3571 (1991). Long Term 11. D. S. Robertson, I. M. Young, and J. R. Perfor- Telfer, J. Mat. Sci., 14, 2967 (1979). mance 12. C. Greskovich, D. A. Cusano, R. J. Riedner, and D. Hoffman, Am. Ceram. Soc. Bull., 71, Standard loads from 100 to 5000 watts, 1120 (1992). cell currents of 10 to 500 amperes, NEW! 13. Greskovich and S. Duclos, Annu. Rev. Mater. Sci., 27, 69 (1997). other sizes available. Model 891 14. R. J. Riedner, R. J. Lyon, D. A. Cusano, C. D. All systems under the control of Reformate Greskovich, United States Patent No Simulator 5,521,387. our FuelCell software for Windows. 15. H. Yamada, A. Suzuki, Y. Uchida, M. Yoshida, H. Yamamoto, and Y. Tsukuda, J. Electrochem. Soc., 136, 2713 (1989). Call for more information or to arrange a demonstration: 16. Y. Ito, H. Yamada, M. Yoshida, H. Fujii, G. Toda, H. Takeuchi, and Y. Tsukuda, Jap. J. Scribner Associates, Inc. Phone (910)695-8884 Appl. Phys., 27, L1371 (1988) 150 East Connecticut Ave. Fax (910)695-8886 17. M. J. Weber and R. R. Monchamp, J. Appl. Southern Pines, North Carolina Web Site: http://www.scribner.com Phys., 44, 5495 (1973). 28387, USA E-mail:[email protected] 18. C. L. Melcher, R. A. Manente, C. A. FC198 Peterson, and J. S. Schweitzer, J. Cryst. Growth, 128, 1001 (1993).

38 The Electrochemical Society Interface • Summer 1998