Evaluation of the Red Emitting ${\rm Gd}_{2}{\rm O}_{2}{\rm S}\!\!:\!\!{\rm Eu}$ Powder Scintillator for Use in Indirect X-Ray Digital Mammography Detectors

$Page 1: Evaluation of the Red Emitting ${\rm Gd}_{2}{\rm O}_{2}{\rm S}\!\!:\!\!{\rm Eu}$ Powder Scintillator for Use in Indirect X-Ray Digital Mammography Detectors$
IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 58, NO. 5, OCTOBER 2011 2503

Evaluation of the Red EmittingPowder Scintillator for Use in Indirect X-Ray

Digital Mammography DetectorsChristos M. Michail, George P. Fountos, Ioannis G. Valais, Nektarios I. Kalyvas, Panagiotis F. Liaparinos,

Ioannis S. Kandarakis, and George S. Panayiotakis

Abstract—The aim of the present study was to investigatethe imaging transfer characteristics and the luminescenceefficiency (XLE) of powder scintillator for usein X-ray mammography detectors. emits in thered part of the visible spectrum, having very good spectralcompatibility with optical sensors employed in digital imagingsystems. Three powder scintillating screens, withcoating thicknesses 33.1, 46.4 and 63.1 , were preparedin our laboratory. The imaging performance of these screenswas assessed by experimental determination of the modulationtransfer function (MTF), the noise transfer function (NTF) andthe detective quantum efficiency (DQE) as well as a single indeximage quality parameter such as the information capacity (IC).A theoretical model, describing radiation and light transfer,was used to fit experimental MTF data. This has allowed theestimation of optical attenuation coefficients of the scintillator.In addition, a previously validated Monte Carlo code, based onthe X-ray attenuation properties and on the Mie light scatteringtheory, was used to estimate the X-ray detection efficiency, theSwank factor and the zero frequency DQE of thescintillator. Results showed that exhibits high MTFand DQE values, which are comparable to those of a commerciallyemployed screen. In addition showshigh compatibility (effective gain) to CCDs and to recently in-troduced CMOS based detectors. Considering our image qualityparameters and luminescence efficiency results, this material canpotentially be considered for use in digital X-ray mammographydetectors.

Index Terms—Digital mammography, scintillator, x-raydetectors.

I. INTRODUCTION

I NDIRECT radiation detectors, employed in X-ray medicalimaging, consist of a scintillating (fluorescent) screen emit-

ting visible light coupled to optical sensors. Screens are fabri-cated by specially developed scintillators (phosphors), of which

Manuscript received May 06, 2010; revised January 24, 2011 and May 17,2011; accepted June 27, 2011. Date of publication August 18, 2011; date ofcurrent version October 12, 2011. This work was supported by the Greek StateScholarships Foundation (I.K.Y.).C. M. Michail and G. S. Panayiotakis are with the Department of Medical

Physics, Medical School, University of Patras, 265 00 Patras, Greece (e-mail:[email protected]; [email protected]).G. P. Fountos, I. G. Valais, N. I. Kalyvas, P. F. Liaparinos, and I. S.

Kandarakis are with the Department of Medical Instruments Technology, Tech-nological Educational Institution of Athens, Egaleo, 122 10 Athens, Greece(e-mail: [email protected]; [email protected]; [email protected]; [email protected]; [email protected]).Digital Object Identifier 10.1109/TNS.2011.2162002

terbium-doped (GOS:Tb), emitting green light, isone of the most efficient and popular [1]–[5]. This is principallydue to its high X-ray absorption and light conversion efficiency.However since most recent silicon (Si) based optical sensors(i.e. CCD and CMOS) are more sensitive to longer wavelengthranges, and particularly in the red wavelength range, it would beof interest to investigate the imaging performance of red emit-ting phosphors [2], [3], [6]. is a red emitting phos-phor which has been previously found comparable to terbium-activated phosphors in terms of optical output [7] and have adecay time of the order of one millisecond (slightly higher than

), which is acceptable for applications that do notinvolve high framing rates [8], [9]. These include stationarydigital and conventional radiography, i.e. general radiographyand mammography. has already been employedin some applications, such as single pulse dual energy radiog-raphy [10], CCD based digital mammography and diffractionenhanced breast imaging [11]–[14]. However its imaging per-formance has not yet been systematically examined in the mam-mography energy range. In a recent study by our group [15], dataon the scintillator were reportedmainly concerningintrinsic properties such as X-ray absorption, intrinsic X-ray tolight conversion, emitted spectrum and spectral compatibility tovarious optical sensors. In the present study, a systematic experi-mental and theoretical investigation of the imaging performanceof powder scintillating screens was performedunder X-ray mammography conditions. Quality metrics such asQuantum Gain (DQG), Modulation Transfer Function (MTF),Noise Power Spectrum (NPS) and spatial-frequency dependentDetective Quantum Efficiency (DQE) were estimated [16]. Inaddition, a custom-developed and validated Monte Carlo (MC)simulation code [17] was used to simulate light spread in orderto predict the zero-frequency DQEwithin the framework ofMielight scattering theory [17].

II. MATERIALS AND METHODS

A. Theoretical and Monte Carlo Modelling

1) Emitted Light and Signal Transfer Efficiency: The lightemitted by an irradiated phosphor screen can be expressed by

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2504 IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 58, NO. 5, OCTOBER 2011

the emitted light energy fluence- (light energy per unit ofarea):

(1a)

where denotes the incident X-ray energy fluence spec-tral distribution (i.e. energy fluence per energy interval). isthe X-ray photon energy and is the maximum energy of theX-ray spectrum. is the energy absorption efficiency(EAE), which is the fraction of incident X-ray energy absorbedlocally at the points of X-rays interaction within the phosphor.is the intrinsic X-ray to light conversion efficiency expressing

the fraction of absorbed X-ray energy converted into light en-ergy within the screen [18], [21], [24]. is the light trans-mission efficiency [18]–[22], which represents the fraction oflight escaping the scintillator. denotes the depth of each thinlayer from the screen surface. The parameters and are op-tical attenuation coefficients [18], [20]–[23] representing the re-ciprocal light diffusion length and the inverse relaxation lengthrespectively. The emitted light can be also expressed by theemitted light photon fluence- (light photons per unit of area)[18]–[23]:

(1b)where denotes the incident X-ray photon fluence spectraldistribution (i.e. photon fluence per energy interval).is the quantum detection efficiency (QDE), being the fractionof the total number of incident X-ray quanta interacting in thephosphor [4]. is the intrinsic quantum conversion gain, i.e.the number of light quanta generated within the phosphor perabsorbed X-ray of energy . (1a) and (1b) may be expressedin the spatial frequency domain. Details on the definition of op-tical parameters and their spatial frequency dependence may befound in previous works [21], [22], [25].In the present work the light energy fluence and the light

photon fluence were modeled following previously publishedtheoretical models [18], [19], [21]. If (1a) and (1b) are dividedby the corresponding X-ray energy and X-ray photon fluence,two parameters expressing the light emission efficiency of phos-phor screens may be defined. These two parameters are: (a) theX-ray luminescence efficiency-XLE , defined as the ratioof the emitted light energy fluence over the incident X-ray en-ergy fluence : and (b) the detector quantumgain- defined as the ratio of the emitted light photonfluence over the incident X-ray photon fluence:[15], [18], [19], [21].Within the framework of the present model, detector quantum

gain may be expressed by a gain transfer function (GTF), de-fined as follows:

(2)

where denotes spatial frequency and is the spa-tial frequency-dependent emitted light photon fluence.The spatial resolution properties of a phosphor screen,

independently of the emission efficiency, are often expressedthrough the modulation transfer function (MTF) [16], [20],[23], [25]–[29] defined by the spatial frequency-dependentlight photon fluence normalized to zero frequency, as follows:

(3)

If is dividedby the incident X-ray photon fluence the gain transfer func-tion may be expressed as follows [15]:

(4)

In medical imaging, where fluorescent screens are used incombination with optical detectors (films, photocathodes, pho-todiodes), the spectral matching between the emitted phosphorlight and the optical detector sensitivity must be taken into ac-count. This is because the degree of spectral matching affectsthe amount of light utilized to form the final image. The spec-tral matching is given by a factor , expressing the fractionof emitted light that can be detected by the optical detector, asgiven below:

(5)

where is the spectrum of the light emitted by the phos-phor and is the spectral sensitivity of the optical detectorcoupled to the phosphor [15], [28]. By taking into account ,we may define the effective gain transfer function as follows:

(6)

2) Monte Carlo Simulation (MC): The Monte Carlo modeluses only physical (complex refractive index, light wavelength)and structural (grain size, packing density) characteristics of thephosphor as input data for the description of light propagation[17], [30]. The simulation codewas based on: (a) the basic X-rayinteractions within the phosphor mass and (b) the light interac-tions modelled by the Mie scattering algorithm [17]. The vali-dation as well as the limitations of the model have been givenin previous reports [17], [30]. According to the Monte Carlosimulation model, the Swank factor was predicted from thefluctuations in the number, , of light photons emitted fromthe screen surface per absorbed X-ray photon, and is defined asfollows [17]:

(7)

where is the nth moment of the light pulse height statisticaldistribution (statistical distribution of values of ) [30].This statistical distribution expresses the fluctuations in thenumber of light photons emitted by the screen, per detected

MICHAIL et al.: EVALUATION OF THE RED EMITTING POWDER SCINTILLATOR 2505

TABLE IPHYSICAL CHARACTERISTICS OF SCINTILLATORS FOR

MAMMOGRAPHIC RADIATION DETECTORS

X-ray photon. Taking into account the Swank factor as well asthe quantum detection efficiency of the screen, the zerofrequency DQE can be estimated according to the following(8) [30]:

(8)

Zero frequency DQE is a very useful metric to characterizeimportant signal and noise transfer properties, as well as themaximum achievable performance of scintillator materials.

B. Experiments and Calculations

was purchased in powder form (Phosphor Tech-nology Ltd, England, code: UKL63/N-R1) with mean grain sizeof approximately 8 and a density of 7.3 . Particle sizeand morphology parameters of the powder phos-phor were verified via scanning electron microscope (SEM) mi-crographs using the JEOL JSM 5310 Scanning Electron Micro-scope (SEM) employing INCA software. Gold was used to ob-tain figures from sites of interest of the specimen.For the elementary particle analysis, the carbon thread evapo-ration process was used. Carbon was flash evaporated, undervacuum conditions, to produce a film suited for theSEM specimen, in a BAL-TEC CED 030 carbon evaporator

.The phosphor was used in the form of thin layers (test

screens) to simulate the intensifying screens employed inindirect X-ray mammography detectors. For the purposes ofthe present study, three relatively thin screens with thick-nesses (coating density) of 33.1, 46.4 and 63.1 , wereprepared on fused silica substrate (spectrosil B). During thesedimentation process, sodium orthosilicate wasused as binding material between the powder grains [15].The properties of , compared with commercially

available phosphors, are given in Table I [5], [31].Experiments were performed on a General Electric

Senographe DMR Plus X-ray mammography unit. For thedetermination of the XLE, X-ray tube voltages ranging from22 to 40 kVp with molybdenum (Mo) anode target and molyb-denum filter was used. The filter changed automatically frommolybdenum to rhodium (Rh) at 35 kVp, and from rhodiumto aluminum (Al) at 42 kVp respectively. Tube voltage andincident exposure rate was checked using a Mammographic ionchamber (Victoreen Model 6000-529, chamber volume of 3.3

) combined with a Victoreen Model 4000M+ dosimeter.Direct diagnostic X-ray spectra measurements were per-

formed with a portable Amptek XR-100T-CdTe spectrometer,

based on a CdTe solid-state detector [33]. The detector wascalibrated for energy scales, linearity checks and energy resolu-tion, by using and -ray calibration sources. X-rayspectroscopy was performed in the mammographic energyrange under clinical conditions. The Mo spectra were producedby the mammography X-ray unit and then transmitted througha 0.03 mm Mo filter at 28 kVp. The detector was placed on thecassette holder with a 57 cm focus to detector distance, and thecompression paddle was removed during the measurements.In order to minimize pile-up distortions, we used a dedicatedtungsten (W) collimation system (W collimator, 400 diam-eter and 1 mm thick). The exposure X (mR) was calculated byusing (9) [33]:

(9)where is the measured X-ray spectrum

at energy and isthe X-ray mass energy absorption coefficient of air at energyobtained from the literature [34].The X-ray luminescence efficiency (XLE) was determined

by performing X-ray photon flux and emitted light energy fluxmeasurements. The latter were performed using an experimentalsetup comprising a light integration sphere (Oriel 70451) cou-pled to a photomultiplier (EMI 9798B), connected to a Cary 401vibrating reed electrometer [15]. For the calculation of XLE,the X-ray energy flux, in (1a), was determined directly fromthe CdTe spectrometer measurements, by multiplying the X-rayphoton fluence with the corresponding X-ray energy.MTF was experimentally determined by the edge spread

function (ESF) method according to the IEC 61267-2 report[35]. A PTW tungsten edge test device (L659136) was used toobtain the slanted edge images, placed at a slight angle to avoidaliasing. The W edge test device consists of a 1 mm thick Wedge plate (100 75 mm) fixed on a 3 mm thick lead plate. Thescreen was brought in close contact with a radiographic film(Agfa LT 2B) enclosed in a light tight cassette. The exposureconditions employed for the MTF measurements were: 28 kVat 31.9 entrance surface air Kerma-ESAK (i.e. KineticEnergy Released in Matter). The experiment was carried out inreflection mode (in contrast to the XLE measurement) for com-parison with the gold standard powder phosphor ,employed in the commercially available Kodak Min-R 2000screen-film system. In this mode, the film was placed behindthe test pattern (in front of the screen) and the light emittedby the irradiated screen side was measured. After irradiation,films were developed in an Agfa Classic EOS film processor,operated at 360 and at 90 s processing time. Pattern images,obtained on the films, were digitized in an Agfa Duoscanscanner with scanning parameters 1000 dpi, 8 bit. Prior todigitisation it was verified that the optical density values of thefilm were within the linear part of the H&D characteristic curveand then digital images were expressed in units of air kerma. Aregion of interest (ROI) was used to extract a 5 5 cm sectionof the image with the edge roughly at the centre. The angleof the edge was then determined using a simple linear least


squares fit and the 2D image data were re-projected around theangled edge to form an ESF. The ESF was smoothed with amedian filter of five bins and then was differentiated to obtainthe line spread function (LSF). Then, the Fourier transform(FT) was employed to obtain the MTF [35], [36]. The latterwas divided by the MTF of the scanner and the MTF of thefilm to extract the MTF of the screen. Finally, the gain transferfunction was estimated from MTF and DQG values accordingto (4).The noise power spectrum of the scintillating screens is ex-

pressed as the FT of the autocorrelation function [37].

(10)

where the terms inside the brackets stand for ensemble av-erage. is the difference between the average image signaland the signal at points x, y of the spatial coordinates, sampledat regular intervals . Then the NTF was obtained ac-cording to (11).The noise transfer function (NTF) is expressed as [22]:

(11)

where is the zero frequency NPS.The detective quantum efficiency of a phosphor screen has

been defined by the relation [16].In this relation and are the output and inputsignal to noise ratios respectively. The spatial frequency de-pendent DQE was estimated using MTF, NPS, anddata, according to [20]:

(12)where is the Lubberts fraction defined as the ratio of

which is a measure of the differentbehaviour in the transfer efficiency of the signal and the noise,as they pass through the screen [20], [22]. is the meanlight photon energy of phosphor, taken frompreviously published measurements [15]. The experimentallydetermined quantity denotes photon fluence whichis typically employed in the definition of DQE [38]. Since thespatial frequency sampling steps of MTF and NPS are generallynot the same, NPS was linearly interpolated at the frequencysampling points of MTF and then DQE was calculated at thesepoints.If the MTF describes the signal response of a system, and the

NPS describes the amplitude variance in the spatial frequencyspace, then the ratio of these quantities, properly normalized,gives information about the maximum available signal-to-noiseratio as a function of frequency. This concept has been describedas the noise equivalent quanta (NEQ), assuming the measured

X-ray fluence is linearly related to the input signal [39].NEQ is typically given as [26], [39]:

(13)

Beyond the spatial frequency dependent parameters, de-scribed above, image quality can be expressed by single indexparameters, expressed through integration over the usefulspatial frequency range. Such a parameter is the informationcapacity (IC). Calculation of this parameter was based on theexperimentally measured values of the MTF and the NPS.The concept of information capacity has been introduced

within the context of Shannon’s information theory, in orderto assess image information content. However, little workrelevant to medical imaging has been published up to now[16], [40]–[45]. Previous published work was based on thedetermination of the MTF and the NPS for radiographicscreen-film combinations, computed tomography (CT) andnuclear medicine systems [16]. According to Shannon’stheory, the IC, per unit of image area, is given by the product

, where is the number of image elements(pixels) per unit of area and is the number of distinguish-able signal intensity levels that can be registered in an imageelement. However since IC can be expressed as a function ofthe output SNR squared [16], [45], it can be expressed as [21],[45]:

(14)It is important to note that, apart from phosphor properties,

information capacity depends on the level of the incident X-rayquantum fluence (input signal). The practical value of IC isthat it defines an imaging performance index that evaluates theimage information quantity by a single numerical value. IC isnot expressed for specific frequency values since it is the out-come of integrating over the spatial frequency bandwidth. ICmay be considered as being roughly proportional to the areaunder the curve of the frequency-dependent image SNR. Thus,IC mainly reflects the quantity and not the quality of image in-formation. Image quality is better described via the frequency-dependent signal parameters. IC strongly depends on the typeof the phosphor material, and to a lesser extent on the layerthickness.

III. RESULTS AND DISCUSSION

A fragment of a scanning electron microscope (SEM) image,for the powder phosphor, is presented in Fig. 1. ByEnergy Dispersive X-ray (EDX), it was shown that the powderhas no admixtures.The measured spectrum of the X-ray mammographic tube

(Mo target) is shown in Fig. 2. The tube settings were: 28 kVpwith 4 cm of additional Perspex filtration. The system repro-ducibility was verified by measuring the X-ray spectrum severaltimes. The measured X-ray spectrum was corrected for the effi-ciency of the CdTe detector [46]. ESAK, estimated via (9), was


Fig. 1. SEM image from a site of interest of the phosphor.

Fig. 2. Mo X-ray spectrum measured with the portable CdTe spectrometer.

Fig. 3. Variation of the XLE, with X-ray tube voltage.

31.9 . The number of incident X-ray photons, per ESAK,per unit area, was found 3641.045 .Fig. 3 shows the experimental data for the X-ray

luminescence efficiency of the phosphor screensfor various tube voltages. The behavior of the XLE is affectedby the X-ray absorption and the light transmission throughthe screen. Low energy X-rays do not penetrate deeply withinthe phosphor mass. Hence, light photons, which are mainlycreated close to the input screen surface, are forced to travellong trajectories to escape the rear surface of the screen andthus light attenuation increases. The significance of this effect

Fig. 4. Comparison of the experimentally determined MTFs of thepowder phosphor screens with a commercial Kodak Min-R 2000 screen-

film.

depends on phosphor thickness, i.e. in the thinner 33.1screen, light photon trajectories are shorter and hence, thecorresponding XLE values were approximately equal or evenslightly higher than those of the 46.4 screen (e.g. at22 and 25 kVp). For the same reason XLE is higher at 46.4

than at 63.1 . As X-rays penetrate deeper inthe phosphor mass at higher voltages light photons are createdcloser to the screen output and escape the phosphor mass.Light is thus more easily transmitted through the phosphorgrains, reaching maximum XLE values. These values appearat different voltages, which are shifted in accordance with thescreen coating thickness (i.e. higher voltage at higher coatingthickness). At higher X-ray tube voltages the XLE values forall screens were found to decrease. In this case, although lightis easily transmitted through phosphor mass, the quantity of thetotal light created within the phosphor is relatively low. This isdue to a corresponding decrease in the X-ray absorption, whichis more significant at this energy range.Fig. 4 shows experimentally obtained MTF curves, of the

screens, measured at 28 kVp and compared with acommercial Kodak Min-R 2000 screen-film system [38]. MTFdecreases with increasing coating thickness. The MTF of the33.1 screen was found very close to thatof the Kodak Min-R 2000 screen-film system, which has almostthe same screen coating thickness (33.91 ). In granularphosphors,MTF is principally affected by the directivity of lightgeneration and the light attenuation effects (scattering and ab-sorption), e.g. the fraction of laterally directed optical photonsthat arrive at the emissive surface of the screen. These photonsspread out and cause image quality degradation. The amountof this light depends on the thickness of the screen and on thecorresponding light attenuation (light absorption and light scat-tering) properties of the phosphor material. Thicker screens (e.g.the 63.1 ), from phosphors of low light attenuation co-efficients, although exhibiting high emission efficiency, showlower MTF due to the long lateral trajectories of non-attenuatedlight photons, which increase light spreading.Fig. 5 shows a comparison between theoretical and experi-

mental MTF curves. Theoretical curves were determined by fit-ting the model (3) to the experimental data. Calculations were


Fig. 5. Comparison of the experimental and theoretical MTFs of thepowder phosphors.

Fig. 6. Variation of the eGTF with spatial frequency for the 33.1phosphor screen combined with various optical detectors.

performed using the experimentally determinedMoX-ray spec-trum, at 28 kVp, and the values of optical and intrinsic parame-ters corresponding to the best curve fitting on the ex-perimental MTF data, achieved by using (1b) and (3). For thethree screens used in our experiments, i.e. 33.1,46.4 and 63.1 , the values of and giving best fit-ting were as follows: , 183.3, 163.3 respec-tively and , 5.5, 4.9 respectively. was setequal to 0.12, found in a previous publication [15]. The theoret-ical model showed better agreement with the experimental MTFvalues for the 46.4 and 63.1 screens, except 2.4 to 3.5cycles/mm range of 63.1 screen. For the thinner 33.1

screen, the calculated MTF showed the best fit in thespatial frequency range from 2.4 to 5.8 cycles/mm.Fig. 6 shows the variation of eGTF with spatial frequency,

measured at 28 kVp, of the 33.1 screenin combination with various optical detectors (CCD and CMOSbased) [47] and a red sensitive film. Differences between eGTFandMTF curves are due to the light emission effects (e.g. DQG),which are included in the definition of eGTF (see (6)). The cal-culated matching factors corresponding to the optical sensorsindicated in Fig. 6 are shown in Table II. The best combina-tion with optical detectors was obtained for the CMOS sensor.

TABLE IISPECTRAL MATCHING FACTORS

Fig. 7. Experimental NPS of the powder phosphors.

TABLE IIIZERO FREQUENCY DQE, QDE & SWANK FACTOR VALUES

This can be very interesting for some newly developed dig-ital radiography detectors [47], [48] based on CMOS sensors.Very high eGTF values were also obtained with the CCD sensor,which is often used in digital mammography detectors [2], [12],[14], [49]. Finally acceptable eGTF values were obtained withthe amorphous silicon (a-Si) sensor, used in modern flat panelimagers.Fig. 7 shows the experimentally obtained NPS curves (see

(10)) of the phosphor screens. NPS decreases withspatial frequency, although at a slower rate than MTF. This isbecause noise is transmitted more efficiently than signal in thehigher spatial frequencies as it has been shown previously [20],[22], [50].Zero frequency DQE, QDE and Swank factor values for the

screens, as determined by the Monte Carlo pro-gram, are shown in Table III.Fig. 8 shows the DQE of the phosphor

screens. DQE decreases with increasing frequency, sincethe shape of the curves is crucially affected by the ratio

. The DQE at zero frequency islimited by the X-ray absorption efficiency of the phosphor [51].Our DQE findings for screens were comparedto published data for a screen (Kodak Min-R2000 screen film with coating thickness 33.91 ), inFig. 8 [38]. The zero frequency DQE of the was


Fig. 8. DQE of screens at 28 kVp and comparisonwith(Kodak Min-R 2000) screen film.

TABLE IVZERO FREQUENCY NEQ & IC FOR THE PHOSPHOR SCREENS

approximately 0.40 [38]. This value is lower than the zero fre-quency DQE of the 33.1 , showing thatthe Europium doped phosphor can achieve higher maximumperformance compared to the corresponding Terbium dopedscintillator. Thereafter showed comparable DQEdue to the higher MTF values (Fig. 4).Finally, zero frequency NEQ (13) and the calculated values

of the information capacity (14) for the powderphosphor screens, determined at 28 kVp, are shown in Table IV.NEQ’s amplitude is affected by the number of incident X-rayphotons distribution measured at the surface of the screen

. Information capacity decreases very slightlywith screen thickness. This is because IC is mainly affected bythe combined effects of MTF and DQG variation with screenthickness. MTF decreases with screen thickness. DQG initiallyincreases rapidly with thickness and remains approximatelyunaltered for thick screens. These data show that, for a givenlevel of incident X-ray fluence, IC is mainly determined by theintrinsic phosphor material properties, while the role of screenthickness is of lower significance. Thin screens exhibit lowlevels of brightness (DQG), which is compensated for theirhigher spatial resolution (MTF), whereas thick screens showpoor resolution properties but with high brightness gain.

IV. CONCLUSIONS

In the present study, three powder phosphorscreens with coating thicknesses 33.1, 46.4 and 63.1were prepared and examined under X-ray mammography con-ditions. The spatial frequency dependent MTF, NPS and DQE,as well as the single index image quality parameter examinedin this investigation, reveal a promising material which can bean alternative to the established phosphors, particularly in CCDand CMOS based X-ray mammography detectors. The 33.1

screen was found to have similar MTF

to the commercially used Kodak Min-R 2000, while the zerofrequency DQE appeared to be better, showing that this phos-phor can achieve higher maximum performance. Taking intoaccount that has (i) high absorption efficiency atlow X-ray energies, (ii) comparable MTF and DQE to thecommercially used Kodak Min-R 2000 and (iii) very highcompatibility (effective gain) with CMOS and CCD opticalsensors, it could potentially be considered for applications inboth digital and conventional X-ray mammography detectors.

APPENDIXANALYTICAL MODELING

The scintillator was assumed to be divided into layers of el-ementary thickness . In each elementary layer, the local ab-sorption of X-rays (without taking into account secondary X-rayinteractions or scatter), the production of light quanta and theircorresponding escape to the output was taken into account. Theanalytical model uses as input to the detector the X-ray photonfluence obtained from the X-ray spectra, experimentally deter-mined with the CdTe detector. The X-rays interact in the scin-tillator mass in a layer of thickness , at depth . The meannumber, , of absorbed X-rays, can be calculated as[20], [52]:

(A1)

where is the total linear attenuation coefficient of. The absorbed X-ray energy, , is converted into

optical photons in the phosphor mass. This statistical process isassumed to follow a Poisson distribution with mean value ,where expresses the number of optical photons producedper absorbed X-ray of energy . In addition, ,where is the intrinsic conversion efficiency of the scintillator,which is an intrinsic property of the material and does not di-rectly depend on incident X-ray energy. is the light photonenergy [52], [53]. Therefore, the mean number ofoptical quanta produced inside the scintillator mass equals to:

(A2)

A fraction of these quanta escape the phosphor screen.is affected by the combined effects of optical scatter

and absorption within the phosphor mass. In the spatial fre-quency domain is expressed by . The total numberof the spatially distributed optical photons can then be expressedin the spatial frequency domain as [20], [52], [53]:

(A3)

where [20], [52], [53]:

(A4)


The parameters expressing: light scattering , light absorption, front and back screen surface reflection

[53]. is the total screen thickness. If all the phosphorlayers as well as the X-ray energy distribution are taken intoaccount, then the total signal distribution in the phosphor outputequals to:

(A5)corresponds to the non-normalized MTF of the screen

[20], that is . For the zero frequencycase, where , corresponds to the total meannumber of light quanta escaping the phosphor.

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