Nuclear Instruments and Methods in Physics Research A 417 (1998) 86 94

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1 Nuclear Instruments and Methods in Physics Research A 417 (1998) Experimental determination of detector gain, zero frequency detective quantum efficiency, and spectral compatibility of phosphor screens: comparison of CsI : Na and Gd S:Tb for medical imaging applications I. Kandarakis, D. Cavouras*, E. Kanellopoulos, C.D. Nomicos, G.S. Panayiotakis Department of Medical Instrumentation Technology, Technological Educational Institution of Athens, Ag. Spyridonos Street, Aigaleo, Athens, Greece Department of Electronics, Technological Educational Institution of Athens, Ag. Spyridonos Street, Aigaleo, Athens, Greece Department of Medical Physics, Medical School, University of Patras, Patras, Greece Received 14 January 1998 Abstract Gd O S : Tb and CsI : Na phosphors mostly used in X-ray medical imaging detectors were compared at various X-ray tube voltages between 50 and 150 kvp. Laboratory prepared test screens of 80, 110, and 150 mg/cm coating weight were evaluated on the basis of the number of emitted photons per incident X-ray (NEP) and DQE, both expressing the signal-and-noise transfer properties of image receptors. NEP and DQE were determined by a method employing luminescence and emission spectrum measurements. The spectral compatibility of the two phosphors with films, photocathodes and the Si photodiode was also examined. CsI : Na exhibited high performance in the kvp X-ray tube voltage range while Gd O S : Tb was found better at higher X-ray energies. The CsI : Na film combination could provide an excellent X-ray radiographic detector, if effectively protected from humidity. Gd O S : Tb is more suitable for high tube voltage digital imaging but also adequate for conventional imaging Elsevier Science B.V. All rights reserved. Keywords: Scintillators; Phosphor screens; DQE; Spectral matching 1. Introduction Among the various phosphors used in X-ray detectors of medical imaging systems, Gd S:Tb *Correspondence address: Esperidon Street, Athens, Kallithea, Greece. Tel: # ; fax: # ; cavouras@hol.gr. or cavouras@medisp.teiath.gr. and CsI : Na are widely accepted as the highest performing X-ray to light converters suitable for most X-ray imaging applications. Gd S : Tb is a high density and effective atomic number material with K-absorption edge at 50.2 kev, which is well within the X-ray spectra often employed in medical imaging. Gd S : Tb also exhibits one of the highest intrinsic X-ray to light conversion efficiencies (15 20%) [1,2]. These properties /98/$ Elsevier Science B.V. All rights reserved. PII: S ( 9 8 ) X

2 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) augment X-ray detection efficiency and light output both allowing for patient dose and image noise reduction. On the other hand, CsI : Na exhibits lower K-absorption energies (36 and 33.2 kev) and lower X-ray detection efficiency and X-ray to light conversion efficiency (10%). However, its light output is high enough due to the CsI : Na non-granular intrinsic crystal structure that forms needle-like columns [3,4], giving highly directional light propagation and minimized optical scattering. These properties reduce light spread and optical loses within the phosphor material, resulting in increased light output and excellent spatial resolution. Nevertheless, a major drawback of CsI : Na is its hygroscopic properties demanding suitable light transparent protective covers to prevent damage from humidity. To our knowledge a systematic comparative investigation of Gd S : Tb and CsI : Na suitability to various X-ray imaging modalities has not been carried out. In this study a new experimental method for determining phosphor detector performance was developed in order to compare Gd S : Tb and CsI : Na phosphors for medical imaging applications. The method was based on X-ray luminescence efficiency (XLE) and emission spectrum measurements. Phosphor screens of both materials and of various coating weights were prepared in the laboratory and were evaluated at x-ray tube voltages from 50 to 150 kvp. Performance parameters determined were: 1. The detector gain which is defined as the number of emitted optical photons (NEP) per incident X-ray. 2. The zero frequency detective quantum efficiency (DQE) expressing the signal-to-noise ratio (SNR) squared. 3. The spectral matching between the phosphor light emission spectra and the spectral sensitivities of various optical photon detectors used in X-ray imaging. 2. Materials and methods Gd S : Tb and CsI : Na phosphors were used in the form of screens prepared in the laboratory. The coating weight of the screens was approximately 80, 110, 150 mg/cm. Gd S : Tb was supplied in powder consisting of 7 μm average size grains. Granular-type screens were formed by sedimentation of the Gd S : Tb powder on silica substrates [5]. Non-granular CsI : Na screens were developed by evaporation techniques using pure CsI and appropriate amount of NaI for the Na activator NEP or detector gain The number of optical photons m emitted by a phosphor screen following the absorption of an X-ray quantum of energy E is given by: m " [E/E ]G, (1) where is the intrinsic X-ray to light conversion efficiency expressing the fraction of absorbed X-ray energy that is converted into optical energy, E is the mean energy of the emitted optical photons and G is the fraction of optical photons produced within the screen material escaping from the screen output surface. The detector gain or NEP (number of optical photons emitted per incident X-ray quantum) is given by NEP" m " [E/E ]G, (2) where is the X-ray quantum detection efficiency expressing the probability of an incident X-ray photon to interact with the detector material. The product G gives the ratio of the emitted optical energy flux over the incident X-ray energy flux, defined as the X-ray luminescence efficiency or XLE of a phosphor detector [6]. Thus, a phosphor detector s gain can be determined by measuring the XLE and E according to the relation NEP"η Φ E/E, (3) where η Φ denotes the XLE The detective quantum efficiency DQE is defined [7] by DQE"[SNR /SNR ] (4)

3 88 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) SNR is the output signal-to-noise ratio, associated with the image produced by the detector, and SNR is the input signal-to-noise ratio, related to the incident X-ray beam. The output signal of a phosphor detector is described by the mean number N of optical photons emitted by a unit of detector area during the time of observation, when N X-ray quanta hit the detector per unit of detector area during the time of observation. N can be calculated by N "N [E/E ]G "N η Φ [E/E ], (5) where N,,, G and η Φ express mean values of the corresponding parameters averaged over the detector area. The output quantum noise is represented by the variance in the number N given [8] by var[n ]"N ( [E/E ]G )#N [E/E ]G (6a) or using relations (1), (5), and (6a) var[n ]"N [m ]#N. (6b) Alternatively, the output quantum noise can be expressed using relations (3), (5), and (6) in terms of XLE (η Φ ) and E var[n ]"N Φ [E/E ]m #N Φ [E/E ]. (7) Eq. (6a) was derived assuming that 1/the quantities N,, [E/E ], and G are four stochastic variables independent of each other, 2/N and [E/E λ ] follow Poisson statistics while the processes and G have binomial probability distributions. Thus, the variances of N and [E/E ] are equal to the corresponding mean values while the variances of and G are equal to (1! ) and G (1!G ) respectively. Additionally, it was considered that there were N events of the process, N events of [E/E ] photons production and N [E/E ] events of the G process. Using relations (5) and (7), the output signal-tonoise ratio at zero spatial frequency may be given by SNR " N [var(n )] " N η Φ[E/E ] m #1. (8a) Eq. (8a) can also be written in the form SNR " N m m #1 (8b) which indicates that for large m the output SNR is governed by the Poisson statistics of the absorbed X-ray photons. The input signal-to-noise ratio may be expressed in terms of the number N of incident X-rays, which follow Poisson statistics (var[n ]"N ) [4,8,9]: SNR " N [var(n )] "[N ]. (9) From relations (4) (9) we obtain DQE" η Φ[E/E ] m #1. (10) Thus, the zero frequency detective quantum efficiency of a phosphor detector can be determined by measuring η Φ, E and calculating m. The latter is given by m "NEP/η, where η can be determined by the relation η "1!e, (11) where μ(e) is the X-ray attenuation coefficient of the phosphor material and w is the coating weight of the phosphor. μ(e) may be calculated from the data on chemical elements [10]. Relation (10) can also be written in a form similar to Eq. (8b) indicating the importance of η in determining DQE The spectral matching factor The spectral compatibility between a phosphor s emission spectrum S and the spectral sensitivity S of an optical detector was evaluated by determining the spectral matching factor a [11,12] by the relation a " S (λ)s (λ)dλ, (12) S (λ)dλ where λ, λ are the wavelength limits of the phosphor spectrum.

4 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) Measurements The X-ray luminescence efficiency was measured by exposing the phosphor screens to X-rays, using a Siemens Stabilipan X-ray unit. X-ray tube voltage varied from 50 to 150 kvp. The emitted light energy flux was measured by a photomultiplier (EMI 9558 QB) coupled to a Cary 401 vibrating reed electrometer. The incident X-ray energy flux was determined from exposure rate measurements, using the appropriate conversion factor α"ψ/xq [13,14], where Ψ is the X-ray energy flux and XQ is the exposure rate. Since both incident X-rays and emitted optical photons exhibit polyenergetic spectra the ratio E/E in relations (2) and (10) was determined by calculating the average energy values using the formulas (13) and (14) S (E)E de E" S (E)dE and E " (13) S (E )E de hc where E " λ. (14) S (E )de S (E) is the X-ray spectrum, calculated as described in previous studies [2,12,15,16] (see appendix), E is the maximum energy of the X-ray spectrum, which is numerically equal to the tube voltage, S is the phosphor s emission spectrum measured with an Oriel 7240 grating monochromator, E and E are the lower and upper limits of the emission spectrum. S was also used in the calculation of the spectral matching factor (relation 12) while the spectral sensitivity distributions S were taken from the corresponding manufacturer s published data. Errors in our results were mainly due to systematic and statistical errors in light flux measurements. Smaller errors may also be attributed to exposure and S (λ) measurements and to and S (E) calculations. The total error was estimated to be within 5%. 3. Results and discussion Fig. 1 shows the results obtained for the variation of detector gain (NEP) with X-ray tube voltage for three CsI : Na screens. For the 80 and 110 mg/cm screens, high gain values were obtained in the range of X-ray tube voltages between 60 and 90 kvp. X-ray exposures in this range, are very often employed in various X-ray imaging applications. For higher X-ray voltages NEP decreases very slightly in a short range and remains practically constant thereafter. This is mainly due to two competing effects: (1) the X-ray quantum efficiency ( ) reduces at high voltages and (2) the X-ray to optical energy ratio E/E increases with tube voltage, since E is always constant depending on the type of activator and on the material quantum structure. Results concerning the 150 mg/cm screen are very similar to those of the 80 and 110 mg/cm screens except for voltages over 90 kvp where NEP is increasing slightly. As shown in Fig. 1 NEP increases with screen coating weight. This behavior is due to the following reasons: (1) the X-ray quantum efficiency ( ) increases with coating weight and (2) the fraction G of light transmitted through the screen material is very high for all CsI : Na screens, since in this a non granular phosphor, light scattering and light attenuation are minimized. Thus, light transmission is not significantly affected by the increase in screen coating weight. Fig. 2 shows a comparison of a CsI : Na screen with a Gd S : Tb screen both having 150 mg/cm coating weights. The CsI : Na screen was found to exhibit higher detector gain in the useful for X-ray imaging range between 50 and 97 kvp. For higher X-ray energies the Gd S:Tb gain was found superior. This comparison shows that Gd S : Tb is more suitable for relatively high tube voltage X-ray imaging. The differences observed for the two phosphor materials can be explained by considering the following: (1) The X- ray quantum efficiency of Gd S : Tb is superior to that of CsI : Na due to the higher density and effective atomic number of Gd S : Tb. CsI : Na cannot efficiently absorb X-ray quanta produced at high voltages, even at the high coating weight of 150 mg/cm. Thus, its detector gain is practically

5 90 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) Fig. 1. Number of emitted optical photons per incident X ray (NEP) versus X-ray tube voltage for three CsI : Na screens of 80, 110, 150 mg/cm coating weight. Fig. 2. NEP comparison between CsI : Na and Gd S : Tb 150 mg/cm screens. constant or slightly reduced after 90 kvp. (2). The fraction G of the light transmitted to the phosphor emitting surface must be significantly higher for CsI : Na due to its needle-like columnar structure and the absence of scattering grains. Thus, optical quanta are easier channeled out of the CsI : Na screens without diffusion. Hence, at tube voltages lower than 90 kvp, where CsI : Na absorbs more X-ray quanta, its detector gain is augmented.

6 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) Fig. 3 shows the results obtained for the CsI : Na DQE variation with X-ray tube voltage. DQE values peaked at kvp and decreased slowly thereafter following a behavior slightly different to that observed for detector gain values (Fig. 1). DQE also increased with coating weight whereas in other materials DQE decreases after a peak coating weight value [2,12]. This difference is readily explained by the G of CsI : Na, which is maintained high even at thick CsI : Na screens. Fig. 4 compares the DQEs of the 150 mg/cm CsI : Na and Gd S : Tb screens. CsI : Na was Fig. 3. Zero frequency detective quantum efficiency (DQE) versus X-ray tube voltage for three CsI : Na screens of 80, 110, 150 mg/cm coating weight. Fig. 4. DQE comparison between CsI : Na and Gd S : Tb 150 mg/cm screens.

7 92 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) Fig. 5. Normalized tungsten X-ray spectrum at 70 kvp filtered by an additional 20 mm aluminum to simulate X-ray attenuation by the patient s body. found to exhibit better DQE than Gd S : Tb in the range from 50 to 90 kvp. At higher voltages Gd S : Tb was superior due to its increased quantum efficiency and detector gain at the corresponding X-ray energies. Differences in both detector gain and DQE observed between the two phosphors are due to differences in their X-ray quantum efficiencies and their capabilities to efficiently transmit the light created within their mass towards the emitting surface. The high quantum efficiency of Gd S : Tb increases its capability to capture primary diagnostic information (signal) contained in the incident X-ray beam. However, the optical scattering caused by the presence of powder grains decreases the output optical signal, thus, contributing to SNR and DQE reduction. On the other hand the high light transmission efficiency of CsI : Na reduces output noise, thus, ameliorating SNR. At high diagnostic tube voltages the primary X-ray capture efficiency of CsI is reduced resulting in its inferior SNR performance for energies corresponding to voltages higher than 90 or 100 kvp. The role of X-ray to optical energy ratio E/E, giving the number of optical photons (NEP) for a perfect phosphor ( "1, "1, G "1), is significant in determining both NEP and DQE. However, this ratio slightly differs between the two phosphors and, thus, it does not seriously affect the results. Results concerning the spectral matching factor of the two phosphors with various optical detectors are shown in Table 1. The optical detectors considered were three films currently in use in radiographic cassettes (Agfa Curix Ortho GS, Kodak X-omatic GR, Fuji UM-MH. Only the matching factor of one film is shown in Table 1, since the matching factors of the other two films were found approximately equal), three photocathodes (ES/20, S9, GaAs) used in fluoroscopic Table 1 Matching factors of Gd S : Tb and CsI : Na with optical detectors Phosphors Optical detectors Gd S : Tb CsI : Na GaAs Si S Film E/S

8 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) image intensifiers and in photomultipliers of radiation detectors, and the Si photodiode employed in X-ray detectors of digital imaging systems. CsI : Na shows excellent spectral compatibility (better than 90%) with the radiographic film, which being orthochromatic has been mainly designed for Tb activated phosphors like Gd S : Tb. The spectral compatibility of CsI : Na with the extended S20 photocathode was also excellent. Taking into account the very good performance of CsI : Na at tube voltages lower than 90 kvp (Figs. 2 and 4) and provided the phosphor may be adequately protected from humidity, the result concerning the CsI : Na film combination is very interesting for medical radiography. Gd S : Tb is better matched to the Si photodiode than CsI : Na. Gd S : Tb is also very well matched to the films and photocathodes and considering the results shown in Figs. 2 and 4, it is well suited for rather high tube voltage X-ray imaging, conventional or digital, and probably for computed tomography. 4. Summary In this study phosphor screens prepared from Gd S : Tb and CsI : Na were compared by determining the number of emitted optical photons per incident X-ray, the detective quantum efficiency and the spectral compatibility to optical detectors used in medical imaging. CsI : Na screens had very good performance at a range of tube voltages kvp, which is very often employed in medical imaging. Gd S : Tb was better at higher X- ray energies. CsI : Na was also found to exhibit excellent spectral compatibility with radiographic films and photocathodes of fluoroscopic imaging systems. Gd S : Tb was also very well matched to films and photocathodes. Additionally, it was found better matched than CsI : Na to the Si photodiode used in digital imaging detectors. These results are very interesting since they show that CsI : Na could be used in radiography, provided it is suitably protected from humidity. Also, Gd S : Tb is very well suited to both digital and conventional imaging but could also be considered for use in computed tomography detectors. Acknowledgements This study is dedicated to the memory of Prof. G.E. Giakoumakis, leading member of our team, whose work on phosphor materials has inspired us to continue. Appendix A. Model for X-ray spectra In the model developed by Tucker et al. [15] for X-ray spectra, the number of X-ray quanta produced by the bremsstrahlung process with energy between E and E#dE is given as follows: S (E)dE"[αr Z/A][dE/E] B (E #m c) E de F(E, E, E ) 1 ρ dx de (A.1) where S is the X-ray quantum fluence per kev, α is the fine structure constant, r is the classical electron radius, Z is the atomic number of the target material (X-ray tube anode) usually tungsten or molybdenum, A is the mass of the target atoms, E is the kinetic energy of the incident electron, E is the penetrating electron energy, m is the rest mass of an electron, c is the velocity of light, (1/ρ) [de /dx] is the mass stopping power of the target material, F(E, E, E ) is the fraction of X-ray quanta transmitted by the anode given by F(E, E, E )" exp[!μ(e)(e!e )], (A2) ρ c sin(ϑ#φ) where μ(e) and ρ is the linear attenuation coefficient and density of the anode material respectively, c is the Thomson Whiddington constant [15], θ is the target angle, φ is the angle off the central axis along which an X-ray quantum travels, x is the depth of electron penetration within the target, B is a function of Z and ¹ given by Tucker et al. [15]. In addition to the bremsstrahlung X-ray quanta given by (A.1), target characteristic X-rays are also produced which are part of an X-ray spectrum. However, in a previous study [12] we have found the contribution of target characteristic X-rays to be very low (0.48% at 80 kvp, 0.67% at 100 kvp) in comparison with the total X-ray exposure, under

9 94 I. Kandarakis et al./nucl. Instr. and Meth. in Phys. Res. A 417 (1998) our experimental conditions of using an additional 20 mm Aluminum filter to simulate X-ray attenuation by the human body. Therefore, in our calculations, we have not taken into account the contribution of target characteristic X-rays considering the error involved to be very small. A typical filtered X-ray spectrum at 70 kvp is shown in Fig. 5. References [1] B.A. Arnold, in: A.G. Haus (Ed.), The Physics of Medical Imaging: Recording System, Measurements and Techniques, American Association of Physicists in Medicine, New York, 1979, pp [2] I. Kandarakis, D. Cavouras, G.S. Panayiotakis, C. Nomicos, Phys. Med. Biol. 42 (1997) [3] C.A. Mistretta, in: A.G. Haus (Ed.), The Physics of Medical Imaging: Recording System, Measurements and Techniques, American Association of Physicists in Medicine, New York, 1979, pp [4] M.J. Yaffe, J.A. Rowlands, Phys. Med. Biol. 42 (1997) 1. [5] I. Kandarakis, D. Cavouras, G.S. Panayiotakis, D. Triantis, C.D. Nomicos, Nucl. Instr. and Meth. A 399 (1997) 345. [6] G.W. Ludwig, J. Electrochem. Soc. 118 (1971) [7] E. Dick, J.W. Motz, Med. Phys. 8 (1981) 337. [8] R. Shaw, R. Van Metter, Proc. SPIE 454 (1984) 128. [9] R. Swank, J. Appl. Phys. 45 (1974) [10] E. Storm, H. Israel, Report LA 3753, Los Alamos Scientific Laboratory of the University of California, [11] E. Giakoumakis, Appl. Phys. A 52 (1991) 7. [12] D. Cavouras, I. Kandarakis, G. Panayiotakis, E.K. Evangelou, C.D. Nomicos, Med. Phys. 23 (1996) [13] J.W. Motz, M. Danos, Med. Phys. 5 (1978) 8. [14] R. Hendee, in: Medical Radiation Physics, Year Book Medical Publishers, Chicago, 1970, pp [15] D.M. Tucker, G.T. Barnes, D.B. Chakraborty, Med. Phys. 18 (1991) 211. [16] E. Storm, Phys. Rev. A 5 (1972) 2328.