Nov 20, 2002 - Cassette-based computed radiography (CR) systems have ... in parallel with integrated, instant readout digital radiography (DR) systems.
INSTITUTE OF PHYSICS PUBLISHING
PHYSICS IN MEDICINE AND BIOLOGY
Phys. Med. Biol. 47 (2002) R123–R166
PII: S0031-9155(02)24746-6
TOPICAL REVIEW
The physics of computed radiography J A Rowlands Sunnybrook and Women’s College Health Sciences Centre, University of Toronto, 2075 Bayview Avenue, Toronto, Ontario, Canada M4N 3M5
Received 23 July 2002 Published 20 November 2002 Online at stacks.iop.org/PMB/47/R123 Abstract Cassette-based computed radiography (CR) systems have continued to evolve in parallel with integrated, instant readout digital radiography (DR) systems. The image quality of present day CR systems is approaching its theoretical limits but is significantly inferior to DR. Further improvements in CR image quality require improved concepts. The aim of this review is to identify the fundamental limitations in CR performance. This will provide a background for the development of new approaches to improve photostimulable phosphor CR systems. It will also guide research in designing better CR systems to possibly compete with DR systems in terms of image quality parameters such as detective quantum efficiency and yet maintain CR convenience in being portable and more economical.
1. Introduction Present day computed radiography (CR) is based on the use of photostimulable phosphors, which are also known as storage phosphors (Sonoda et al 1983). They are commercially the most successful detectors for digital radiography. The phosphors used are most often in the barium fluorohalide family (Barnes 1993) in powder form and deposited onto a substrate to form an imaging plate or screen. X-ray absorption mechanisms are identical to those of conventional phosphor screens used with film. They differ in that the useful optical signal is not derived from the light emitted in prompt response to the incident radiation, but rather from subsequent emission when the latent image, consisting of trapped charge, is optically stimulated and released from metastable traps. This triggers a process called photostimulated luminescence (PSL) resulting in the emission of shorter wavelength (blue) light in an amount proportional to the original x-ray irradiation. In CR, an imaging plate (IP) containing the storage phosphor is positioned in a light-tight enclosure, exposed to the x-ray image and then read out by raster scanning with a laser to release the PSL. The blue PSL light is collected with a light guide and detected with a photomultiplier tube (PMT). The PMT signal is digitized to form the image on a point-by-point basis (Fujita et al 1989). The broad acceptance of CR has been due to its large dynamic range, digital nature, easy portability and uniqueness rather than its intrinsic image quality. CR systems have improved in the almost 20 years that they have been available. They are now technological and engineering 0031-9155/02/230123+44$30.00
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marvels, but have been unable to transcend their inherent image quality weaknesses. CR based on the use of storage phosphor screens in a cassette is seen to be complementary to rather than directly competitive with integrated readout digital radiography (DR) systems. There are numerous possible approaches to DR (Yaffe and Rowlands 1997) but to date the most successful are flat panel systems based on active matrix arrays (Rowlands and Yorkston 2000). Flat panels are very close to being ‘perfect’ in terms of quantitative measures of the efficiency in which the x-ray aerial image is transformed to a digital image. In this regard DR has far outstripped the present ability of CR. Thus even to equal the image quality obtainable with screen–film and to match DR, change is needed in the basic approach to CR. Several very promising new approaches suggest that CR can rise to the challenge presented by flat panel based DR systems. 1.1. Historical background The fundamental innovation in the development of CR was by Kodak (Luckey 1975) who conceived the storage of an x-ray image in a phosphor screen. It required significant technical steps and conceptualization of the application by Fuji (Kotera et al 1980) to produce the first medical x-ray images. Fuji, the main developer of CR in the eighties, used BaFBr:Eu2+ phosphor and a cassette-based approach. During this time, Agfa and Kodak performed research and development on the same method but were constrained from commercialization by patent issues and ambivalence due to the fear of damaging their installed base of screen–film, respectively. In this era the storage effect was also being observed in screen–film applications where it caused the unwanted effect of print through, i.e. a ghost image of a prior exposure to the screen that appears on a subsequent film exposed in the same cassette. The storage effect is related to the phenomenon of thermally induced luminescence of irradiated materials, i.e. thermoluminescence. Both photoluminescence and thermoluminescence have a long history that can be traced back to 1603 (McKeever 1985, Kato 1994, Seibert 1997) and forward to present day applications in medicine (e.g. radiation dosimetry), biology (e.g. readout of radioactively tagged electrophoresis gels) and elsewhere (e.g. archaeological dating). 1.2. Outline of review There have been two previous comprehensive reviews of CR science and technology (Kato 1994, Seibert 1997). A recent text (Beutel et al 2000) though not specifically targeting CR is highly relevant. A new review is timely because the science and technology of existing systems have plateaued and new concepts are being actively investigated. In the present review the fundamental operation of photostimulable phosphors is first outlined. Then the two components of present day CR systems, the screens (or imaging plates) and the most commonly used reader type (flying spot), are described. The combination of a CR plate and reader forming a fully functioning CR system follows. From this the limitations of present day CR systems are extracted and novel, new approaches are identified. The current capabilities of clinical systems are described and potential areas of improvement are identified. 2. Photostimulable phosphors 2.1. Operation of photostimulable phosphors The following sections will outline the types of photostimulable phosphors used in CR. The physical phenomena arising in conventional phosphors and the new phenomena unique to photostimulable phosphors will be discussed.
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2.1.1. Types of photostimulable phosphors. The photostimulable phosphor first used for CR was BaFBr:Eu2+. Its crystal structure is non-cubic, i.e. a layered structure that gives rise to phosphor grains with a plate-like rather than the more desirable cubic morphology (Blasse and Grabmaier 1994). BaFBr:Eu2+ is a good storage phosphor in that it can store a latent image for a long time, e.g. the latent image 8 h after irradiation will still be ∼75% of its original size (Kato 1994). The family of phosphors BaFX:Eu2+ where X can be any of the halogens Cl, Br or I (or an arbitrary mixture of them) have been studied extensively. The decay time after photostimulation of all these phosphors is now known to be approximately the same (∼0.7 µs) and so they can all be used in CR. In earlier literature there was a long decay noted for BaFCl:Eu2+ which can now be eliminated. In recent years most manufacturers have used BaFBr0.85I0.15:Eu2+ not for the marginal increase in x-ray absorption compared to BaFBr:Eu2+, but rather for the better optical match of the wavelength of maximum stimulation of the phosphor to diode lasers. Recently Konica has utilized pure BaFI:Eu2+ in commercial systems (Nakano et al 2002) where the change in absorption is significant. RbBr:Tl+ is cubic and has the advantage that it can be made into a needle-structured layer. By guiding the light to the surface, even a thick layer can achieve high resolution. However it has the disadvantage of a rapid (tens of seconds) loss of the latent image that makes it unsuitable for cassette-based systems (Nakazawa et al 1990). Konica has used this material in integrated readers where the CR plate can be rapidly read out in situ immediately after the termination of the exposure. CsBr:Eu2+ is also cubic, can be made in the needle structure, has a stable latent image and can be photostimulated. Agfa has proposed using this material in both cassette-based and integrated readers (Leblans et al 2001). The spectrum of light emitted by an efficient phosphor is controlled by a dilute ( Eg. In fact W ≈ 3Eg .
(1)
This behaviour has been shown to apply to practically all semiconductors, photoconductors and—at sufficiently high fields—insulators (Klein 1968). It therefore represents the fundamental limiting energy requirement for ehp production in the solid state. However, in conventional phosphors there is one further stage before light can be emitted, the ehps must be allowed to recombine with the emission of radiation. This process can be made to approach 100% efficiency by incorporating appropriate activators, which form luminescent recombination centres. Thus equation (1) also applies to activated phosphors and can be used to establish an approximate value for the lower limit of W on any phosphor once Eg is known. 2.1.3. Mechanisms of trapping and photostimulation. In photostimulated phosphors there are several more stages in the conversion of incident energy to light than in a conventional phosphor. Each stage causes further inefficiency and therefore an increase in W. In table 1 the physical properties of several important conventional and photostimulable phosphors including W and wavelength λ of the light emitted by the activator have been listed and represent an update of the classic table given by Arnold (1979). A notable feature is that the
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Phosphor
Z
EK (keV)
Gd2O2S:Tb3+ BaFCl:Eu2+
64 56 56 56/53 56/53 55/53 55 37/35
50.2 37.4 37.4 37.4/33.2 37.4/33.2 36/33.2 36 15.2/13.4
BaFBr:Eu2+ BaFBr0.85I0.15:Eu2+ BaFI:Eu2+ CsI:Tl+ CsBr:Eu2+ RbBr:Tl+
G Light Spectrum for Eg Density W (photons/ Decay time emission stimulation (eV) (g cm−3) (eV) 50 keV) (µs) peak (nm) (nm) (∼8) 8.3 (∼8) (∼8) 6.2 7.3
7.34 4.56 5.1 (5.1) (∼5.6) 4.52 4.45 3.35
20 2500 25 2000 360∗ 140∗ 360∗ 140∗ 20 2500 250∗ 200∗
∼3 0.7 0.7 0.7 0.6 0.98 0.7 0.35
550 390 390 390 405 550 440 433
na 500–600 500–650 550–700 550–700 na 685 735
Conventional phosphors, i.e. those used in screen–film combinations, are shown in normal type and the photostimulable phosphors in bold. Entries that are unknown are left blank or if an entry is inappropriate ‘na’ is used. In some cases the value has been estimated by comparison with similar known materials, in this case the value is enclosed in parentheses. Z is the atomic number of the radiologically important, i.e. heavy element or elements, EK is K-absorption edge energy of the heavy elements, Eg is the energy gap of the host crystal in electron volts (eV), density is quoted for a single crystal, i.e. 100% packing factor, G = conversion gain (emitted light photons per 50 keV of absorbed x-ray energy). For the conventional phosphors this is the direct light emitted and for the photostimulable phosphors is the PSL light obtained when the trapped energy is fully stimulated and shown with an asterisk. Decay time constants are usually the same whether the light is given promptly on stimulation by x-rays or after stimulation of trapped charge by a laser since they are limited by characteristics of the same luminescence centre and not the trapping, and release mechanisms. Data were compiled from Arnold (1979), Blasse and Grabmaier (1994), Holl et al (1988), Leblans et al (2001), Schweizer (2001) and Sonoda et al (1983).
conversion gain for the photostimulable phosphors is an order of magnitude smaller than that for conventional phosphors. In a storage phosphor, excitons can be trapped without the emission of light. It is believed that if photostimulation is to occur later, the trapping must occur on sites spatially correlated with the activator. This is called the PSL complex shown in figure 1(b). The energy levels in the crystal are critical to effective storage phosphor operation. The energy difference between the electron traps and the conduction band edge must be small enough to allow stimulation with laser light, yet sufficiently large to prevent significant random thermal release of the charge carriers from the traps. In BaFBr:Eu2+ the image storage is due to: (i) electron trapping at positive ion (Br or F) vacancies, forming an F-centre, or (ii) hole trapping at an unidentified site (Blasse and Grabmaier 1994). An activated photostimulated luminescent site or PSL centre is therefore thought to be an arrangement of three spatially correlated components: an electron trap, a hole trap and the luminescent activation centre. The PSL emission spectrum has been correlated with an internal transition within the activator, Eu2+. The stimulation spectrum has been correlated with the absorption spectrum of the F-centre showing that the first step in the stimulation process is excitation of the trapped electron. However, it is believed that a great inefficiency arises because ∼80% of the electrons are trapped at F sites and ∼20% at Br sites but only the latter contribute to PSL (von Seggern et al 1988, Thoms et al 1991). It was first thought that the hole was trapped on the activator site itself (Eu2+) thereby increasing the valency to Eu3+ (Takahashi et al 1984, 1985). However as there is no change in the electron spin resonance spectrum following x-ray irradiation, this trapping mechanism cannot be operative (Schweizer 2001). Thus the nature of the hole trap is in doubt and the details of the entire process are only known approximately. In the needle phosphor RbBr:Tl+ the luminescent centre is the Tl+ ion; the electron is trapped at a Br vacancy; and the hole is assumed to be trapped at a Tl+ ion (Blasse
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and Grabmaier 1994). Optical stimulation excites the electron, which recombines with the hole on the Tl+ yielding the PSL light. There are many other possible photostimulable phosphors but little is published relevant to their application in CR. Significant studies have been made into large classes of materials (Shimada et al 1991). It is probable that there are more to be discovered and improvements to be made to already known phosphors. It is important to continue to better understand the operation of the storage phosphors so as to be in a position to improve their imaging properties. 2.2. X-ray properties of photostimulable phosphors The number of x-rays used limits the image quality since x-ray image noise arises from the random interactions of x-rays with the detector. The square of signal-to-noise ratio (SNR) at the input of a detector SNR2in, i.e. the ultimate SNR which a perfect detector could achieve, is equal to the number of x-rays incident on the detector NI. At the output of the detector SNR2out = Nd (the number of x-rays detected) for a detector that counts x-rays. The ratio of detected to incident x-rays is called the quantum efficiency AQ. Thus for a photon counting detector SNR2out Nd = = AQ . NI SNR2in
(2)
From equation (2) it is evident that AQ is the single most important determinant of the ultimate image quality possible from an x-ray detector. In figure 2, AQ is plotted against incident x-ray energy E for some common photostimulable phosphors. In figure 2(a) the AQ for BaFBr0.85I0.15:Eu2+ is plotted for phosphor loadings of 40 and 70 mg cm−2 representing the typical values for a high resolution and standard IPs, respectively, used in CR (Kato 2002). In figure 2(b) the attenuation of BaFI:Eu2+ is shown for the same phosphor mass loadings and in figure 2(c) a more direct comparison of BaFBr0.85I0.15:Eu2+ and BaFI:Eu2+ layers of the same physical thickness of 200 µm. It is seen that the K-edge is very pronounced and the total absorption is never very large but can be significant in the energies just above the K-edge of Ba and/or I. The very large variation of AQ with energy, shown in the relatively thin layer of the powder phosphors, gives rise to an undesirably large absorption of scattered radiation relative to the primary (unscattered) radiation in these screens. This is a disadvantage (Tucker et al 1993 Shaw et al 1994 Yip et al 1996) compared to for example, Gd2O2S screens with a higher K-edge (Bogucki et al 1995, Floyd et al 1991). There is controversy in the literature as to whether the increase in the scatter to primary ratio in CR is due to the position of the K-edge or is correlated to the relatively poor absorption of CR screens (McLean and Gray 1996). Thus, in general, the SNR is increased in CR by reduction of scatter (Miettunen and Korhola 1991). In figure 2(d) AQ is shown for the needle-structured phosphors. In this case the absorption is (with the exception of RbBr) larger than in the powder phosphors and the increase in attenuation seen at the K-edge is less pronounced. 3. Screens and CR imaging plates (IPs) The design and physics of IPs are very similar in concept to conventional phosphor screens used with film. Thus a brief review of the design and operation of conventional screens follows.
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(b) BaFBr0.85I0.15
1.0
0.8
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AQ
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BaFI
1.0
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Figure 2. X-ray attenuation curves for phosphors. (a) BaFBr0.85 I0.15:Eu2+ for 40 and 70 mg cm−2. (b)BaFI:Eu2+ for 40 and 70 mg cm−2. (c) Comparison of BaFBr:Eu2+ with BaFI:Eu2+ at the same phosphor layer thickness. (d) Comparison of the needle-structured phosphor RbBr, CsBr and CsI for the same 400 µm thick layer (graphs courtesy of G DeCrescenzo).
3.1. Conventional phosphor screens The phosphor grains are highly scattering particles due to the high refractive index of phosphors compared to the plastic binder and air pockets within the screen. The scattering is sufficient in that the flow of photons can be considered diffusive and the layer turbid. This scattering limits the spreading of light from its point of origin to a lateral distance comparable to the phosphor layer thickness. Consequently, if high resolution is desired, a thin screen must be used also reducing AQ. While travelling within the phosphor, light will be spread by scattering–a random walk process—the amount of lateral diffusion being proportional to the path length required to escape the phosphor. The fraction of light escaping from the screen depends on: the bulk absorption of the screen, generally negligible for most screens unless a dye is intentionally added; and the boundary condition on the screen, primarily the nature of the backing, whether it is reflective or absorptive. The boundary condition is used to adjust the light escape efficiency as a function of depth of x-ray absorption. The blurring for x-rays in a given screen also depends on the depth of x-ray absorption. Thus the optical boundary conditions (and bulk absorption due to dye) can be optimized to appropriately weight the contribution from x-rays absorbed at different depths. Analytical models of light transport in phosphor screens have been developed to permit the calculation of screen resolution from their
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Figure 3. Manufacturing methods and structure of powder phosphor screens. (a) Doctor blade method for depositing a thick uniform layer of phosphor slurry on the moving web of backing material from a supply roll. (b) Overall structure of screen.
basic properties (Swank 1973). The models use parameters to represent the scattering length (usually the phosphor grain size but this can be increased by using a binder with a similar refractive index to the phosphor grains (Kato 1994)), the absorption length and the boundary conditions. The phosphor grain size is a critical design parameter for several reasons. First, there are factors related to the intrinsic luminescence properties. If the grains are too small, then the defects created on the outside of the grain, e.g. non-photostimulable electron or hole traps, may dominate the bulk properties of the material. Secondly, there are factors related to the diffusion of light through the screen. If the grains are too large, then non-uniformity of light output or structural noise will be more pronounced. Thirdly, a larger scattering length will reduce the resolution of the screen. Thus screens designed for a low-resolution task, e.g. chest, can use larger phosphor grains than a high-resolution, e.g. mammographic, task. Screens are usually designed to be ∼10–20 phosphor grains thick. 3.2. Manufacture of IPs CR IPs are powder phosphor screens made with photostimulable phosphor. Phosphor grains are combined with a polymer binder and deposited on a substrate. The ratio of binder volume to phosphor volume in the mixture controls the fractional volume of the phosphor layer finally occupied by air pockets or voids. A solvent is used to liquefy the phosphor/binder mixture into slurry in preparation for deposition onto the flexible but strong backing layer of the screen. The doctor blade controls the phosphor layer thickness deposited on the backing as shown in figure 3(a). The backing material is transported as a web below the doctor blade at a carefully controlled spacing. The doctor blade deposits the phosphor slurry onto the backing, and its edge establishes the thickness of the phosphor layer as shown in figure 3(b). The final screen thickness will be less than the height of the blade due to shrinkage as the solvent is driven off during the subsequent drying process. Typically the binder is nitrocellulose, polyester, acrylic or polyurethane and the backing material is also a polymer, e.g. polyethylene terephthalate 200–400 µm thick (Kato 1994). The use of a black or white backing permits adjustments of the reflectivity and absorptivity at the phosphor interface. In BaFBr0.85I0.15:Eu2+ IPs the typical phosphor grain size is 4 or 5 µm (Kengyelics et al 1998a, Matsuda et al 1993) with a trend to yet smaller sizes as the ability to make smaller grains with good PSL properties improves. A recently introduced capability in the design of IPs is the possibility of having different optical boundary conditions for the laser (absorptive) and PSL light (reflective) by use of
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analyser
Figure 4. Pulse height spectra in phosphor screens. Top row: measurement method. Bottom row: example pulse height spectra obtained under the conditions shown.
anti-halation backing layers (Gingold and Schaetzing 2001). The exposed phosphor surface and the back of the screen have additional protective layers. The purpose of both these layers is to protect the optical surface of the phosphor layer because during stacking and transport within the reader, the bottom surface rubs against the top of other IPs. 3.3. Gain-fluctuation noise in photostimulable phosphor screens Ideally in a phosphor screen, the same amount of light would be given off from the absorption of every x-ray. However, in practice there are variations in the light emitted per x-ray, which gives rise to gain-fluctuation noise. Such effects can be visualized by considering a pulse height spectrum PHS (Ginzburg and Dick 1993). This is a histogram of the number of events in which a specific number of light photons are given off after the absorption of a single monoenergetic x-ray (figure 4). Such a PHS may be measured experimentally for conventional phosphor screens (Drangova and Rowlands 1986) as shown in figure 4(a) or may be modelled (Fahrig et al 1995). The physics of the mechanisms giving rise to changes in the PHS is usually clear (Trauernicht and Van Metter 1988) and illustrative examples are given in figure 4(b)-(d). Furthermore, the effect of gain-fluctuation noise on image quality can be calculated by use of the Swank factor, which is derived directly from the PHS (Swank 1973a). The underlying idea is that the gain-fluctuations sensed by an energy-integrating detector will decrease SNRout compared to a counting detector. The generalization of the quantum efficiency concept is recognized in the terminology as detective quantum efficiency, which is used to understand the noise properties of an integrating detector. Thus the following relationship can be derived: SNR2out = DQE. (3) SNR2in This concept has been described fully by Swank (1973a) and can be further generalized to include effects that occur as a function of spatial frequency f. If we look first at the large area effect, that is for f = 0, Swank showed that the added gain-fluctuation noise could be expressed as a multiplicative factor now called the Swank factor, AS: (4) DQE(0) = AQ AS .
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Thus the correction term AS converts the DQE of a device used as a counting detector to that for an integrating detector. The Swank factor can, to a useful approximation, be subdivided into further multiplicative components with the common feature that they are all ideally unity but in practice are usually less than unity by only a small factor. We will now examine these components of AS. When an x-ray spectrum is used, the screen reacts to x-rays with a different energy by giving off a different amount of light. The correction term for a broad x-ray spectrum is called AXED (x-ray energy distribution). Its value depends both on the spectrum and the absorption of the screen and is modest with AXED ∼ 0.9. There are ongoing discussions to redefine DQE(0) for radiographic and fluoroscopic detectors so as to exclude this effect. The current literature still includes it. Further variations of light output from the screen from x-ray to x-ray arise from two effects shown in figure 4(d): (i) K-fluorescence occurs when an x-ray of energy E above the K-edge of a material in the screen may give off a K-fluorescent x-ray of energy EK and the K-photon may escape (resulting in a PHS output with a total energy per x-ray of E–EK) or be reabsorbed (the PHS output has a total energy per x-ray of E ). The effect on gainfluctuation noise arising from K-escape and reabsorption is called AAED (absorbed energy distribution). Values for commonly used radiographic phosphors have been tabulated (Swank 1973a, Rowlands and Yorkston 2000) and calculated by Monte Carlo methods (Chan and Doi 1984, Boone et al 1999). The AAED including K-reabsorption is unity for E below the K-edges, i.e. there is no effect on noise, but AAED drops considerably for E > K-edge. The smaller the total absorption in the screen, the greater K-escape above the K-edge and so the smaller is AAED. Due to the relatively low AQ in powder phosphor screens AAED ∼ 0.7 is reached at energies just above the K-edge. It then slowly increases back to unity with increase in energy. Averaged over the whole x-ray spectrum AAED is ∼0.75–0.85 for powder phosphor screens and in needle phosphor screens, with a smaller K-escape fraction, AAED ∼0.9–0.95. (ii) The effects due to the interactions of the depth of absorption of the x-ray and the optical boundary conditions (e.g. for a screen with an absorptive backing the amount of light emitted for an x-ray depends on whether it is absorbed near the output face of the screen {full light emitted} or the back of the screen near the absorber {no light emitted}). This gives rise to the correction factor AOPD (optical pulse distribution). For screens with a reflective backing the PHS is quite narrow showing that the amount of light collected from each x-ray is independent of the depth of x-ray absorption. Thus AOPD ∼ 1. However for a black backing or a bulk dye the light collected is highly dependent on the depth of x-ray absorption resulting in an exponential PHS and AOPD ∼ 0.5. For a practical system it is also necessary to consider contributions to DQE(0) from structural noise Astruct. This understanding of conventional screens can be applied to CR IPs. There are two additional important effects: (i) the fraction of the trapped charge latent image released by the simulating light, i.e. the discharge fraction F, depends nonlinearly, on the stimulating light intensity (Lubinsky et al 1987). Therefore AOPD depends on F. For example, in IPs with absorptive (black) backing AOPD ∼ 0.5 at low F, i.e. F < 0.1, but AOPD ∼ 1 as F → 1 where all depths of the IP are completely discharged. (ii) The number of photoelectrons emitted from the photocathode of the PMT per x-ray absorbed in the IP is relatively small so that secondary quantum noise (specifically called luminescence noise in the context of CR) represented by the parameter ASQ also plays a role in the total DQE(0). Other sources of noise, which will be ignored here because they can be generally reduced to insignificant values, include stimulating laser noise, multiplicative gain noise in the electronic amplifier and quantization noise. Especially note that amplifier dark noise, which can be a dominant factor degrading DQE in DR systems (e.g. flat panel detectors, Rowlands and
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Figure 5. Screen optics—the path of laser light entering and scattering within the phosphor layer and the same for the resulting PSL emission depending on the optical boundary conditions defined by the backing layer. The small graphs represent the depth distribution of laser light IL and escape probability IPSL for stimulated PSL generated at a given depth d in the IP. First for transport of laser light into the IP for (a) reflective, (b) absorptive backing layers and for transport of the PSL out of the IP, (c) reflective and (d) absorptive backing layers. Full lines represent the situation where there is no bulk absorption of light in the phosphor layer and the dashed lines are for the situation with a significant amount of bulk light absorption.
Yorkston (2000)), has no role in CR due to the essentially zero dark current of the PMT used as the first stage amplifier. In summary, the product of all the individual factors modifies the detective quantum efficiency and thus DQE(0) = AQ AXED AAED AOPD (F )AStruct ASQ .
(5)
3.4. Resolution of photostimulable phosphor screens Using a flying spot scanner, resolution is not dependent on the scattering of the PSL but rather on the scattering of the stimulating laser light. The light transport properties are shown for both PSL and laser light in figure 5. The boundary conditions for the penetration of the stimulating laser light and escape of PSL will be examined separately. In figures 5(a) and (b) the spreading of the laser light is illustrated from its point of incidence on the front surface of the screen to the back. The spreading is less if the backing absorbs rather than reflects the laser light. X-rays absorbed closer to the surface of the IP will have better resolution than those absorbed at greater depths and the resolution is worse for an IP of greater thickness. In order to understand the efficiency with which the latent image is read out as a function of depth d in the screen, it is necessary to know the amount of stimulating light reaching each
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Figure 6. Comparison of the geometry and orientation of screens used in different radiographic imaging systems. The backing layer defines the optical boundary condition and in most cases is highly reflecting but can, in some cases be absorbing. In (a) orientation of mammographic screen with x-rays incident on light emitting surface and (b) illustrating similar geometry and screen thickness for computed radiography are shown. Note the x-ray attenuation curve showing the relatively larger absorption of x-rays at the incidence surface. In (c) the arrangement for flat panel imagers used in DR is shown, where the absorption is at the less favourable back surface of the screen, far from the readout active matrix. (d) Shows that for screen–film with dual screens one of the screens is in the front screen orientation and the other in the back screen.
layer. The quantity IL is defined as the total intensity of the laser light at d independent lateral spread of the laser and is shown in figure 5. If, as is often the case, the bulk absorption of light in the IP can be neglected, then for reflective backing IL is essentially independent of d. (There is, extraordinarily, a smaller intensity of laser light close to the incident face than deeper within the phosphor layer. This arises because the laser light is trapped in the IP by multiple scattering and it can more readily escape close to the surface.) In marked contrast, for an absorptive backing IL drops linearly from the front surface to almost zero at the back. Bulk absorption, which may be intentionally created by the incorporation of dye to absorb the laser light preferentially. The effect of the dye absorber on IL is shown by the dotted lines on the graphs in figure 5. The efficiency of escape of PSL, IPSL, has the same general nature as IL, i.e. constant for a reflective backing and linearly dependent on the distance from the backing for an absorptive backing (figures 5(c) and (d)). The PSL is more blurred from x-rays at the back than the front of the IP. (The PSL blurring has no effect on resolution if the flying spot readout method is used.) Note that the paths of most optical quanta will be the shortest and hence blurring least if the read out is performed at the x-ray entrance side of the phosphor as is normal in CR (figure 6(b)) and single screen–film as used in mammography (figure 6(a)) but not in flat panel DR imagers (figure 6(c)) or conventional dual screen screen–film (figure 6(d)). This provides a much-needed advantage to CR over flat panel imagers. It also is the explanation of the surprising, but little remarked fact that CR is far more universal than other imaging systems. A single type of CR plate (standard resolution) is usable for all imaging tasks except mammography, which is to be compared to three types of screen–film in general use for radiography (high resolution, general purpose and high speed).
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Figure 7. Overall concept of CR readout systems. (a) Cassette-based requiring carrying cassette back and forth between the x-ray and readout systems. (An important component in practical application not shown in the diagram is a stacker needed to buffer the demand on the system.) (b) Integrated readout systems requiring no operator intervention in the exposure readout cycle.
4. CR readers—flying spot scanner Present day CR systems are of two general types: (i) cassette-based systems as shown in figure 7(a) where the IP is enclosed in a light-tight cassette for the x-ray exposure, and subsequently moved by hand to the readout system; (ii) integrated readout systems, shown
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in figure 7(b) where the IPs are captive within the readout system, re-circulated and reused without handling. Both types use a flying spot readout system, i.e. a laser spot is scanned with a mirror over the exposed IP in a point-by-point raster pattern. The flying spot scanner is not the only possible approach but is common in commercially available medical CR systems and will therefore be described in detail. 4.1. Gaussian optics of laser beam For mirror scanning, highly collimated beams produced by lasers are essential. A circular beam with a Gaussian intensity in cross section is desirable and this can be accomplished using helium–neon gas lasers or solid-state laser diodes (Kengyelics et al 1998b). Gas lasers naturally have a circular Gaussian form if operated in their fundamental transverse mode (TEM00). Astigmatic lenses are needed with solid-state lasers due to their highly elliptical beam shape. With either kind of laser, once a circular Gaussian beam has been obtained, its cross-sectional diameter can be modified using a laser beam expander. However, the beam cannot be collimated to a parallel beam of the small size required at the IP as such a beam will have a large inherent divergence caused by diffraction. A larger diameter, less divergent beam must be used and focused to a small spot at the image plane. The diameter of the collimated Gaussian beam Wl where it enters the focusing lens of focal length f is related to the focused Gaussian spot size Wf by the relationship (Siegman 1986) Wl Wf = f λ/π.
(6)
From equation (6) it can be seen that the smaller the focal spot required, the larger the beam diameter before entering the lens. For CR, f = 50 cm for a field-of-view of ∼35 cm. Applying equation (6) for the desired Wf = 100 µm (diameter measured at the 1/e2 intensity points (Siebert 1997)), then Wl ∼ 1 mm. This is a useful beam size, often encountered in laboratory lasers, as a 1 mm beam does not diverge significantly over a distance of a few metres. 4.2. Laser types The advantages of solid-state laser diodes compared to continuous gas lasers are: (i) the output intensity of a solid-state laser can be controlled electrically; Note that gas lasers, e.g. HeNe with wavelength λ = 633 nm, need an additional external device such as an electro-optical or electro-acoustical modulator; (ii) The solid-state diodes used in CR (λ = 680 nm) are more compact, energy efficient, and have a longer operational lifetime than gas lasers. They do not, however, have as good a match to the optical stimulation requirements of BaFBr:Eu2+. This necessitated using both a greater laser power and a redesign of the phosphor (replacing Br by Br0.85I0.15, Matsuda et al (1993)). 4.3. Readout rate limits In designing a readout system one must know the required readout rate. The plate throughput of medical scanners is ∼30–110 plates per hour (Seibert1997), which is adequate for the workload of a typical clinic. The engineering limit on the transport and stacking of IPs is a component of the total readout time. There is also a fundamental limit related to the characteristics of the photostimulable phosphor, i.e. the intrinsic decay time of the luminescence centre (Eu2+ in the case of BaFX). PSL from a previously stimulated region will continue to glow. This decays with a time constant, which is characteristic of the activator and the host lattice. For BaFBr0.85I0.15:Eu2+, the time constant is 0.7 µs (table 1). If scanning were performed too
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1.0
1.0
0.8
0.8
PSL output from pixel
PSL output from complete plate
(a)
0.6
0.4
0.2
0.6
0.4
0.2
0.0
0.0 0
1
2
3 2
Laser Irradiance (J/m )
4
0
10
20
30
40
50
Laser power (mW)
Figure 8. The nonlinear, stimulation energy dependence of PSL (a) measured by uniform irradiation of complete imaging plate (Kato 1994) and (b) measured for a single pixel in scanning mode (Arakawa et al 2000). In both graphs the full line is an exponential fit to the discharge curve, which can be seen not to be quite exact. It can, under certain circumstances, be represented by a power law.
quickly, the PSL signal from one pixel would not be completely decayed before the PSL from the next was initiated. Consequently, it would bleed into the next pixel and cause spatial blurring. To avoid this, several time constants should elapse between the readout of one pixel and the next, i.e. usually 5 time constants (1/e5 < 1% lag) resulting in ∼4 µs per pixel. Thus with 2000 × 2000 pixels the shortest readout time would be 16 s. Fuji AC-3 readers take ∼30 s to readout an array of this size (Fetterly and Hangiandreou 2000). Another practical limitation on readout time is the laser power available. The PSL output with stimulating laser irradiance is shown in figure 8 for BaFBr0.85I0.15:Eu2+ (a) for the complete plate irradiated uniformly and (b) for an individual pixel. Note that since the PSL output saturates, i.e. the IP is within 10% of complete discharge, with laser input of ∼2 J m−2, there is no need to use more light. The energy deposited by a ∼30 mW laser spread uniformly over the whole area of a typical IP (0.33 m × 0.33 m ∼ 0.1 m2) is 0.3 J m−2 s−1. Thus it will take 2/0.3 ∼ 7 s to discharge the IP by 90%, i.e. F = 90%. In practice, therefore, the characteristic phosphor decay time provides the limit on readout time in flying spot scanners. 4.4. Beam scanning The diode laser beam is sent through several subsystems before reaching the CR plate as shown in figure 9(a). The laser beam is divided (not necessarily equally) into two with a beam splitter such as a partially silvered mirror. The main beam passes to the scanning system; the side beam is sent to a photodiode used to monitor, and with feedback, stabilizes the laser output intensity. The laser focusing lens is generally of the F/θ design. It has three further functions: (i) to make the focal plane flat so that focus is uniform across the IP, (ii) to convert the constant angular motion of the scan mirror into a constant linear speed at the image plane so that the pixel spacing is constant and (iii) to move the entrance aperture of the lens, i.e. locus defining region where any ray within the angle of acceptance will be imaged by the lens, significantly out from the body of the lens so that the rapidly scanning mirror has space to
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Figure 9. Flying spot CR readout scanner. (a) Scanner components also showing the graph of direct output of PMT response to a step in intensity on the IP, (b) processing stages, (c) response curves of processing stages and (d) typical form of output signal as it passes through the processing stages.
operate. Either a rotating polygonal mirror driven by a synchronous motor or an oscillating flat mirror driven by a galvanometer can perform the scanning function. The advantage of the rotating polygon mirror is that the transition from one facet to the next performs the flyback, i.e. retrace which occurs at the end of one line as the beam rapidly returns to the start of the next line. This maintains a high laser duty cycle, i.e. the fraction of the time the laser is actually reading out the IP. In contrast, a galvanometer has to be driven in an oscillatory manner by a sawtooth signal and after flyback takes a fixed time to return to stable operation. Its duty cycle is therefore lower than a polygon and at high line rates, the duty cycle decreases even further. Thus the polygon is used for faster scanning speeds but it has two disadvantages which lead to periodic errors that appear as banding artefacts in the subscan direction. (i) The reflective
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properties of its facets may each differ slightly, which requires a further correction to the laser output (Matsuda et al 1993). (ii) Unintentional scanning of the beam perpendicular to the scan direction results in an effect called cross-scan error, caused by slight angular shifts between facets. This can be corrected passively using a pair of cylindrical optics (Matsuda et al 1993). CR, as in any destructive readout, is very sensitive to cross-scan error. If the beam moves one way, there will be reduced signal as it rereads already partially discharged regions of the IP and in the other way, it will steal signal from the next line. Control of cross-scan error to 100,000 10000
2000 Flat panel 1800
1300
1000
2
3
Output electrons
0.1
5.5 Electrons released
Photostimulable centres
1
1
Collected photons
Electron-hole pairs
10
Emitted photons
100
Absorbed x-ray
Number of quanta
1000 130 Line scan CR 65 133 33 67 Flying spot CR 22
4
5
6
7
0.01
Stage
Figure 11. Quantum accounting diagram for several different radiographic imaging systems namely the flying spot CR system, line scanner CR system and flat panel DR system.
converter to process the signal from the PMT before it entered the a/d as shown in figure 9. This further reduced the effective dynamic range of the signal (Dolazza and Poulo 1984) and the number of bits required. Currently, analogue logarithmic processing or square root processing (Dolazza and Poulo 1984) is used to reduce the dynamic range of the signal before digitization by a 12-bit a/d. Alternatively if logarithmic processing is performed digitally, i.e. after the a/d, using look-up tables, then a 16-bit a/d will be necessary. The prescan approach has been eliminated due to the ready availability of high bit depth converters and the ability of handling large amounts of digital data, which was far from trivial in the early eighties when CR systems were introduced. Finally a shading correction is applied to allow for the varying light collection efficiency of the light guide as a function of the laser position along the line. This is a one-dimensional correction as every line is the same. 4.7. Isolation of PSL from laser light To permit the PSL signal to be isolated from the laser light, the phosphors chosen for CR can be stimulated with laser light with a different wavelength from the PSL light. Is there a problem in isolating these two sources of light, both of which will enter the light guide and reach the PMT? Laser power used in CR is ∼30 mW (Matsuda et al 1993, Seibert 1997, Leblans et al 2001) or ∼2 ×1017 red light photons/s (assuming 2 eV per red light photon). At this power and the 4 µs dwell time per pixel derived earlier, there are ∼8 × 1011 red light photons incident on each pixel. The stimulated light output from the IP for a single fully absorbed 50 keV x-ray is ∼70 blue PSL photons (figure 11). Thus the ability to detect 120 x-rays, i.e. lowest exposure level defined above at which signal needs to be measured—1/10 mean exposure, will require the detection of 8000 blue photons in the presence of 8 × 1011 red photons. Therefore the
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Figure 12. Propagation of noise in CR systems illustrating the contribution of noise from secondary quantum statistics. The effect on image quality by changing g is investigated. In the upper row a representation of the image of a single x-ray photon is illustrated. With g = 1000 each x-ray has enough secondary quanta that its image is smooth and representative of the characteristic blurring of the systems. As g decreases, each x-ray has fewer light photons representing it resulting in a rougher appearance. The corresponding Wiener noise power spectra for the same g values are shown in the lower part of the figure.
ratio of stimulating red to PSL blue photons at the surface of the IP at the worst case of the minimum x-ray exposure is ∼108 or eight orders of magnitude! The first approach used to separate these light photons depends on the PMT. A typical transparent bialkali photocathode has a quantum efficiency of ∼25% in the blue and ∼0.1% in the red. The second approach is the filter, which still needs to selectively remove five orders of magnitude brighter light in the red while efficiently passing light in the blue if interference with the output of the PMT from the laser light is to be avoided. 4.8. Quantum accounting diagram The maximum signal to noise ratio (SNR) of any imaging system occurs where the x-rays are absorbed. If the SNR of the imaging system is essentially determined here, the system is said to be x-ray quantum noise limited in its performance. Inevitably the SNR is reduced as the signal passes through the system. It is possible to make high image quality x-ray detectors because of the large intrinsic conversion gain as the x-ray energy, e.g. 50 keV, is converted to many secondary particles, e.g. 3 eV photons. It is important that the detector maintains a large number of quanta representing each x-ray if secondary quantum noise is to be minimized. A quantum accounting diagram aids in locating the secondary quantum sink. Figure 11 illustrates the propagation of quanta through the CR conversion stages. Each 50 keV x-ray interacting in the detector produces ∼2000 ehp. However, only a small fraction (∼1/15) of these ehps are trapped in a manner that permits them to be laser photostimulable and produce PSL. Of these, only half are actually stimulated, i.e. discharge fraction F = 0.5, and PSL photons emitted. For an IP with a reflective backing we assume all of these PSL photons escape from the IP. However only ∼1/3 can be collected by the light guide and brought to the face of the PMT where ∼1/4 release photoelectrons. Taking into account all these factors an absorbed x-ray is represented on average by ∼5.5 e, i.e. the system gain g = 5.5 electrons per 50 keV x-ray (2000 × 1/15 × 12 × 1/3 × 1/4). Fortunately, the PMT can efficiently and adequately amplify 5.5 electrons such that no significant further noise is added. The noise inherent in the few electrons representing a single x-ray is the weakest noise link in the CR system and is the secondary quantum sink. For comparison, in the quantum accounting diagram for flat panel DR (figure 11) g equals 1000. In figure 12 the effect on image quality by changing
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g is investigated. With g = 1000 each x-ray has enough secondary quanta that its image is smooth and representative of the characteristic blurring of the systems. As g decreases, each x-ray has fewer photons resulting in a rougher appearance. The resulting noise power spectrum is shown in the lower part of figure 12. As g decreases the noise of the secondary quanta NPSSQ approaches, and in the case where g = 1, equals the noise from the x-rays NPSX. This has a deleterious effect on both DQE(0) and as f increases, DQE( f ). Secondary quantum noise is evident in optically coupled x-ray image devices (Yaffe and Rowlands 1997), especially where optical demagnification is significant. It has been investigated in the context of phosphor screens coupled to CCDs using tapered fibre optics (Maidment and Yaffe 1994). The situation is also analogous to the use of film in screen–film (Nishikawa and Yaffe 1990). Film has an optical DQE(0) ∼ 1% (Dainty and Shaw 1974) leading to g ∼ 20, which is comparable to g for CR. What options are available to increase g in CR? 4.9. Possible sources of improvement in conversion gain in flying spot CR systems In figure 11, the sources of loss in g are highlighted. BaFX:Eu2+ has been developed to produce the maximum possible PSL output and practically may have reached a limit. However, it does not appear to violate any conservation laws to expect that a significant fraction of losses seen within the phosphor (failure to trap ehps, non-photostimulable sites) could be eliminated. Using previously established data leads to a theoretically possible increase in g by 15. By careful analysis it may be possible to elucidate the exact mechanism of operation of the phosphor, reduce the losses and improve the gain. The next stage is the light collection system. The theoretical limit would collect all the emitted light resulting in an increase in g ∼ 3. Means for accomplishing this could include reading from the side of the IP opposite to that of the laser and increasing the light guide efficiency by placing it in contact with the IP (Arakawa et al 1999, 2000). The PMT has an optical quantum efficiency of only ∼25% since it relies on the external photoelectric effect, i.e. an electron is given enough energy to eject it from a solid and into the vacuum. In theory, solid-state devices can approach an optical quantum efficiency of 100% because of the internal photoelectric effect where an electron is transferred from the valence band to the empty conduction band within the solid. This potential four times increase in optical quantum efficiency would help to reduce the secondary quantum noise characteristic of CR. However, if a large area photodiode replaced the PMT, then both the electronic noise from the capacitance of the diode (14 cm2 × 50 pF cm−2) and the fluctuations in the dark current (>14 × 5 nA cm−2) would each yield an electronic noise Ne > 5000 e rms. For this to be equal to the x-ray noise from NX x-rays, we have the relationship Ne = g (NX)0.5. Our example is then 5000 = 5.5 (NX)0.5, i.e. NX ∼ 106 x-rays/pixel, which corresponds to 4 mGy incident air kerma. Thus a flying spot CR system using a silicon photodiode would not be quantum noise limited below 4 mGy (three orders of magnitude larger than the typical incident air kerma of 3 µGy used in CR) incident on the IP! Thus, PMTs continue to be used because: (i) the gain achieved by impact ionization using a cascaded series of ∼10 dynodes, each with an adjustable gain up to ∼4, yields an overall gain of a million. This gain is essentially noiseless. Unlike an electronic amplifier, there are no parasitic capacitances and the inherent gain fluctuations are insignificant. (ii) The dark current Id is negligible even for a very large active area (14 cm2) required in mapping the input of the light guide (35 cm × 4 mm). Id < 1000 e cm−2 s−1 or ∼0.06 electrons from the 14 cm2 area of the photocathode during the pixel integration time of 4 µs. Thus the electronic noise and dark current fluctuations, from a practical-sized PMT, are insignificant compared to the ∼60 e (11 x-rays × g = 5.5) noise that corresponds to the lowest signal encountered in CR. Thus for a flying spot scanner, g cannot be expected to be increased significantly in the near future.
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1000
10
(b)
(a) 4
CR Reader Units
PMT output (AU)
10
3
10
2
10
500
S=200 L=4 ST-V HR-V
1
10
S=200 L=2 0
10
0.1
1
10
100
air Kerma (µGy)
1000
10000
0 0.1
1
10
100
1000
air Kerma (µGy)
Figure 13. Linear response of IP to x-ray exposure. (a) CR reader output measured at PMT in arbitrary units (AU) plotted against irradiation to IP expressed in air kerma. This curve shows the linear input–output relationship over more than four orders of magnitude (Bootstrapped from three individual plots from Kato 1994). (b) CR reader output in CR Reader unit plotted against the logarithm of x-ray kerma, i.e. exposure, showing a straight line fit which demonstrates a linear relationship between the PSL output and air kerma (Kengyelics et al 1998b) for four combinations of reader and IP parameters (latitude parameter L and IP type ST or HR).
4.10. Readout linearity Figure 13(a) shows the native characteristic curve, i.e. a plot of PMT signal output before logarithmic compression, of a CR reader. The author constructed this composite graph by bootstrapping data from several individual curves (Kato 1994) to show the overall dynamic range of the system. The exposure sensitivity of the CR plate is shown to be linearly proportional to exposure from ∼0.1–1000 µGy (a dynamic range of 104). This curve must be linear (or linearizeable) for MTF, NPS and thus DQE to be defined. However, the linearity of the characteristic curve, although necessary, is not a sufficient condition. The creation of PSL depends nonlinearly on the stimulation light intensity as can be seen by referring back to figure 8. If the system is to satisfy the requirements for linear analysis, three further conditions must be satisfied. The first condition for effective linearity is that the laser light intensity must be kept constant during the readout scanning process to maintain the discharge fraction F of the IP at a constant level. This is because both the spatial resolution of CR systems (figure 14) and the x-ray to charge conversion gain g, and hence the secondary quantum noise, depends on F. The second condition (satisfied automatically for photostimulable phosphors in current use) is that the discharge process must be very inefficient. This ensures that F is independent of the latent image. If the discharge process were more efficient, then an effect similar to adjacency in film could occur leading to a complete breakdown of linear analysis. This is an edge enhancement that arises close to a high-contrast sharp edge during development (Dainty and Shaw 1974). In both film development and IP readout such effects arise from the latent image using up the developer or laser light, respectively. The third condition is that any nonlinear processing applied to the signal from the PMT, e.g. the logarithmic compression, be reversed.
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Figure 14. Phosphor blunting is illustrated by (a) a plot of MTF plotted for three different laser power intensities to the IP and corresponding relative discharge of plate obtained from figure 8 (Arakawa et al 2000) and (b) schematic illustration of origin of phosphor blunting.
Thus although the readout process is intrinsically nonlinear, the linear parameters defining spatial resolution (modulation transfer function (MTF)), spatial noise (Wiener noise power spectrum (NPS)) and the spatial frequency dependent detective quantum efficiency DQE( f ) can nevertheless be defined for CR systems in current use (with a possible exception to be discussed later). Note that the values of MTF, NPS and DQE will be different for the same reader and IP if factors affecting F are changed. Furthermore, the optimization procedure by which the scan parameters are obtained is a nonlinear process. 4.11. Plate reconditioning A residual latent charge image remains on the CR plate after readout. Erasure of the image using a high intensity light source must be performed before the plate is reused. This is accomplished using a high-pressure sodium or fluorescent lamp (Siebert 1997). The erasure time depends on the brightness of the lamp and the level of erasure required. The level of erasure achieved also depends on the prior x-ray exposure and F. The erasure rate is enhanced if the initial erasure is performed with a light spectrum including ultraviolet followed by a spectrum with the UV filtered out (Matsuda et al 1993). This is probably due to the creation of trapped charge by UV irradiation. Erasure times are 10–20 s for the Fuji AC-3 and ∼50 s for the Lumisys ACR-2000 readers (Fetterly and Hangiandreou 2000). In practice, unless exposed to an excessive x-ray irradiation (such as during a quality control procedure), the previous latent image is effectively removed during a single erase cycle. It is possible to accumulate an image on the CR plate due to natural radioactivity and cosmic radiation. Therefore, before clinical use the IPs should pass through a further erasure cycle if they have been in storage for more than a day (Seibert 1997). 5. Configuration and operation of CR systems After the image has been acquired digitally, it has to be processed and manipulated before it can be displayed effectively and a diagnosis made. Processing for CR images is specific for each different application (Vuylsteke et al 1997). The system configuration, how the components parameters are chosen and the CR system performance will now be described.
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IP type
Phosphor layer mass loading (mg cm−2)
Phosphor layer thickness (µm)
ST-Va
70d
230c
HR-Va
40d
140e
Phosphor packing factor
System gain (e/x-ray)
0.60 0.56
16.6b 4.0b
AQ
DQE(0) E
