Key points of the technical report (6.1) entitled A study of the photoreceptors of the indirect medical X-ray digital imaging systems

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1 Title «Medical Image SCIence through Luminescence», MISCIRLU, code 1476 Key researcher: Prof. Ioannis Kandarakis Key oints of the technical reort (6.1) entitled A study of the hotorecetors of the indirect medical X-ray digital imaging systems Authors: 1. Ioannis Valais. George Fountos 3. Nektarios Kalyvas 4. Panagiotis Liaarinos 5. Christos Michail 6. Stratos David

2 Contents I. Introduction II. Materials and Methods a. Sectral matching factor b. Signal and Noise transfer c. Photon Transfer Curve (PTC).7 d. Resonse time III. Results and Discussion..,, 8 a. Sectral matching factor.,,..8 b. Signal and Noise transfer...9 c. Photon Transfer Curve (PTC) d. Resonse Time IV. Conclusion V. References VI. Aendix...16

3 INTRODUCTION In recent years CMOS (Comlementary Metal Oxide Semiconductors) hotorecetors are widely used at indirect digital detectors, due to their higher resolution and their fabrication method (system on a chi) which allows cost minimization. CMOS detectors are divided into CMOS Passive Pixel Sensors- PPS and CMOS Active Pixel Sensors-APS. CMOS_PPS have only one hotodiode and a transistor, which allows for higher active detector surface (higher fill factor). When a reset voltage is alied the circuit is closed and the charge collected by the sensor is further distributed in the detector circuit. Ideally the charge is roortional to the number of otical hotons iminge on the CMOS surface. The CMOS- APS recetors are discretized. Each unit consists by a hotodiode and three MOSFET transistors. Every unit (i.e. ixel) is considered a discrete detector. Signal recording consists by a reset hase, an integration hase (electrons counting) and a readout hase (electron current is turned into voltage for further rocessing) [1]. The arameters which can be checked in the hotorecetors are (a) the linear range, (b) full well caacity (c) dynamic range and (d) quantum efficiency. Several hotorecetors have two mode of oeration by aroriately take into consideration the signal by several ixels so as to increase the full wel caacity and the dynamic range [-6]. When light incidents on a semiconductor electron-hole airs are roduced. When a -tye semiconductor roduces electrons these may be diffused in adjacent ixels and increase image blur. [1]. MATERIALS AND METHODS Sectral matching factor The otical hotons that are incident on a hotorecetor may not all detected due to a sectral mismatch between the otical hoton sectrum and the hotorecetor sensitivity sectrum. Sectral comatibility is characterized through a sectral matching factor. a s is zero for zero comatibility and 1 for erfect comatibility. It can be calculated from the following equation [7-9] : a s S ( ) S S D ( ) d ( ) d [1] Where S (λ) is the otical hoton sectrum and S D (λ) is the hotorecetor sensitivity sectrum for wavelength λ. 3

4 As sensitivity is considered the hotocurrent roduced when otical hotons incident at the hotorecetor surface. Quantum efficiency is a related arameter defined as the number of electron-hole airs roduced er incident otical hoton. In this reort sectral matching factor will be calculated for different hoshors hotorecetors combinations. S(λ) and S D (λ) were found from literature [7-9], while for Gd O S:Pr,Ce,F S(λ) was measured with a method already described in literature [7-9]. Signal and Noise Transfer In this study the Linear Cascaded Systems Theory was used to redict the signal and noise in the hotorecetor outut. The theory assumes a series of successive stages were signal and noise are affected by system roerties. The outut of a stage is the inut of the next one. These stages describe various hysical roerties of the detector like otical hoton cature, electron-hole generation, electron diffusion etc. The stages may be gain stages were the signal is changed or transformed and blur stages were the signal is dislocated (i.e. electron diffusion). Every stage has (i) an inut S in (u ), where u is the satial frequency (ii) a mean signal inut value x in, a satial frequency distribution outut S out (u ) and a mean signal outut. x out. Every gain stage described a statistical rocess with mean value q and variance q, while every blur stage is characterized by an MTF (u) [1-9]. The stages may be stochastic or deterministic. [,7, 8]. The mean signal outout and the corresonding satial distribution can be calculated by the following equations (i) S (u ) q S (u ) xin and xout qxin, for gain stages, (ii) S out and in in in out (u ) ( S (u ) x )MTF (u ) x and xout xin for stochastic blur stages and (iii) S (u ) S (u )MTF (u ) xout xin for deterministic blur stages [10-1]. in q out Let F otical hotons following a Poisson distribution to be incident at a CMOS ixel with dimension a ix. A fraction of them will be detected. This is a gain stage characterized by the CMOS fill factor ff and its in active area, a d, where a d = ix ff a. This rocess follows a binomial distribution with mean value ff and variance ff (1- ff ). Due to sectral matching only a fraction of these hotons will be further considered. This is gain stage characterized by the sectral matching factor. The rocess mean value equals to a s and its variance equals to a s (1- a s ). The absorbed otical hotons roduce electron hole airs e-h. This is a binomial distribution with mean robability value Q and variance Q (1- Q ).Only a fraction of these electron-hole airs will reach the CMOS outut. This is also assumed a binomial distribution with mean robability (NPS I (u)) equals to: Q e and variance Q e (1- Q e ). By alying LCS theory the CMOS NPS 4

5 NPS Ia ( u) ( ff a s Q Q ) F( u) F( u) ff a Q e s Q (1 e ff a Q Q ) s e () where ff a s Q Qe qe is the CMOS quantum efficiency. The total number of electrons in the outut ( X e ) is calculated equal to: Xe F ff a s Q Q e (3) The electrons are diffused towards the exit. This is a blur stage characterized by the MTF of the CMOS, MTF inh,,. The final NPS equals to: I NPS Ia ( u) XeI MTF ( u XeI NPS ( u) inh ) (4) If the electronic samling, the finite size of the ixel and the system electronics were not considered then the ideal detective quantum efficiency DQE ideal would be equal to [11]: X emtfinh ( u) MTFix ( u) DQEideal (5) FNPS ( u) I When the signal is finelly detected the electrons are intergrated due to the finite ixel size. This is a deterministic blur stage deended uon a ix. The related MTF ix (u) equals to [1] MTF ix ( u) sinc( a u) (6) ix Thus: a 4 NPS ( u) MTF ( ) NPS ix( u) ix I ix u (7) and X ix a Xe (8) d 5

6 Due to samling the aliasing should be considered. In this case the outout NPS, NPS out, equals to [1]: NPS out 1 ( u) NPS ix( u) Wadd( u) ff (9) where, W add is the additive electronic noise The final modulation transfer funciton, MTF, is calculated as the roduct of the MTF of each blur stage. Thus MTF( u) MTF ( u) MTF ( u) inh ix (10) And the DQE alias is calculated by X ix MTF( u) DQEalias FNPS ( u) (11) out Where it has been assumed that [1] MTF inh ( u) sinc( a u). ad In this work the effect of ixel size, fill factor and the energy ga of the hotorecetor on MTF, DQE ideal and DQE alias will be examined. For calculation uroses it as been assumed that a 1000 otical hotons of energy Ε λ =.46eV, emitted by a light source are incident on a CMOS detector. The fill factor was taken equal to 0.9 (RadEye CMOS), while Q was assumed equal to Q E E g [13], where Ε g is the silicon energy ga. It is customary assumed that Q 1 [10]. In this reort however a value of α Q 0.49, E 5, has also been considered.. Finally Q 0. 55(a tyical value for RadEye CMOS, g e which is calculated by solving ff a s Q Q qe, for ff=0.8 and α =0.9). s e 6

7 In this study the charge sharing effect was not considered in MTF inh. In medical imaging the hotorecetors are in conduct with scintillators. Due to the Comton effect in the scintillator, electrons may iminge on the hotorecetor and ionize it. This may rovide image blur. Therefore this tye of electron interactions in silicon was studied with the Monte Carlo software ypenelope [14]. Photon Transfer Curve-PTC PTC is a owerful tool for measuring linear range, full well caacity dynamic range anf quantum efficiencu of hotorecetors. This technique assumes non correlated noise sources which divides them into readout noise (corresonds to the first art of the curve where low signal levels are recorded), shot noise (statistics of X-ray hotons), fixed attern noise (FPN), due to atterns associated with the detector (i.e. image of CMOS electronics) and the saturation, where variance is zero [-7]. For the urose of this reort the PTC curve of a Gd O S:Tb/CMOS RadEye combination in real mammograhic imaging conditions was examined [10]. The eculiarities of clinical imaging conditions are the following: (a) The smallest available signal may lead to variances higher than readout noise. (b) Shot noise does not have a Poisson distribution [15] (c) FPN also includes the structure noise of the scintillator [15]. (d) Clinical imaging conditions usually do not reach saturation region, for atient radiation rotection uroses All measurements erformed with ImageJ [16] Resonse Time The hotorecetor electronics resonse time is a crucial arameter when erforming digital examinations in ulse mode (angiograhy, coronograhy), or a lot of X-ray shots are erformed (comuted tomograhy). In angiograhy and coronograhy, the frame rate could be as high as 5 frames/s in fluoro mode or 40 ms er X-ray image in CINE mode. In addition in Comuted Tomograhy during a gantry rotation the X-ray tube maybe oerated 180 times. If the rotation time is 1s then er 6 ms the detector receives new signal. Furthermore the scintillator has a decay time of 3 μs for Gd O S:Pr or u to 1 ms for Gd O S:Tb and Gd O S:Eu. If F(t) is the X-ray ulse, S(t) is the resonse of the scintillator and Φ(t) is the total resonse time of the hotorecetor and its electronics then the total resonse time Τ(t) can be calculated by the following convolution: 7

8 T( t) F( t) S( t) ( t) [1] In this reort different resonse times of the scintillator and the hotorecetor will be tested. It has been assumed that the resonse of the X-ray and the hotorecetor are square, while the resonse of the scintillator is exonential. The calculations were erformed with Octave 3.8. [17] RESULTS AND DISCUSSION Sectral matching factor Table 1. Sectral matching factor of different scintillators hotorecetors combinations Photorecetor Lu O 3 :Eu Gd O S:Pr,Ce,F Gd O S:Eu CsI:Tl Gd O S:Tb CMOS 0.5μm Pgate 0,1 0,7 0,9 0,86 0,79 Hybrid CMOS with NIR AR Coatong 0,81 0,81 0,98 0,86 0,83 Hybrid CMOS with BLUE AR Coating 0,81 0,96 0,99 0,98 0,98 Monolithic 0.5μ CMOS- Image sensor 0,46 0,88 0,96 0,93 0,9 CCD BI With IR AR coating 0,65 0,74 0,98 0,8 0,94 CCD with ITO gates 0,41 0,78 0,9 0,8 0,79 CCD with oly gates 0,8 0,55 0,84 0,66 0,6 CCD with ITO gates&μlens 0,58 0,78 0,94 0,87 0,86 It is observed from table 1 that for every scintillator examined, a suitable hotodetector which yields a sectral matching factors above 0.9 can be found. 8

9 Signal and Noise Transfer Figure1: MTF and DQE of hotorecetor with a ixel size of 5μm. In figure 1 the MTF and DQE curves of a hotorecetor with ixel size 5μm are demonstrated. It may be observed that MTF is slightly affected (inversely roortional) by the fill factor. On the contrary the DQE curves are roortionally affected by the fill factor as well as Q. In every case DQE alias is lower that he ideal DQE due to the higher noise as it is calculated by equation 11 9

10 Figure1: MTF and DQE of hotorecetor with a ixel size of 100μm. In figure and DQE curves of a hotorecetor with ixel size 100 μm are demonstrated. The results are similar to that of Figure 1. However the MTF values for the 100 μm ixel is well degraded with resect to the ones with the 5μm ixel size. In addition the DQE curve dros significantly after 50 cycles/cm due to aliasing. In order to test the effect of charge sharing the electron sread in Si, which was obtained from literature [18], is resented in Figure 3. From the figure it can be observed that charge sharing may be of imortance in small ixel detectors. The Gausian curve which is fitted into the data has a statistical deviation σ=11.5 μm 10

11 Figure 3: Electron satial distribution due to charge sharing, for X-ray energies 0keV to 40keV In figure 4 the effect of the Comton electrons by the scintillator on the hotorecetor is demonstrated, for 80keV electron energy. It can be seen by figure 4 that the maximum disersion is about 5μm, so there effect is less imortant than charge sharing Figure 4: Electron sread in silica. The electron energy was 80 kev 11

12 Photon Transfer Curve-PTC Figure 5: PTC curves for X-ray mammograhy irradiation conditions In figure 5 it can be observed (ι) the PTC curve for 8 kv Mo/Mo filter target combination and tube current ranging from 4 mas to 100 mas. The variance measured at the 4 mas exosure corresonds to read out noise,. That is The RadEye CMOS manual reorts that the read out noise is less than 1 [19]. The exosure between 4 mas and 40mAs are related to shot noise, while the others corresonds to FPN. FPN incororated also the Gd Ο S:Tb structure noise. In this clinical setu the maximum signal (S max ) corresonded to a ixel value S max 80. For this case the clinical dynamic range (DR) can be calculated as DR 0log [-5]. By substitution, DR=4 db. However according to RadEye CMOS secifications the dynamic range is 4000:1 or 7 db. The measured difference is due the fact that exosure was ket as low as ossible. If we used a setu of 00 mas, which corresonds to a ixel value of 19 (lease see Figure 6), the new dynamic range would be 53dB. R R R 1

13 MPV Ερευνητικό Πρόγραμμα ΓΓΕΤ «ΑΡΙΣΤΕΙΑ» RadEye CMOS y = 0,769x - 0,157 R² = 0,9937 mas mean Γραμμική (mean) Figure 6: Linearity resonse curve of RadEye CMOS. A oint worth mentioning is that the resented digital image in grey level reresentation incororates all signal maniulation erformed. In addition the otical hotons incident on RadEye CMOS do not have a Poisson distribution but the distribution of a X-ray triggered scintillator. Finally in clinical imaging conditions we are mainly interested in low exosures therefore the shot noise is of interest. Therefore in clinical conditions only a range of the hotorecetor dynamic range is utilized. Resonse Time In Figure 7 the total resonse curves for different resonse times of the X-ray tube and the hotorecetor are resented. The scintillator decay tiem was ket constant at 5μs. We may deduce that a resonse time ratio of the hotorecetor resonse time over the X-ray tube resonse time which is equal to 1/10, does not significantly alters the total resonse time.. 13

14 Figure 7: Total resonse time for CT (6 ms/rojection) και angiograhy (40ms/X-ray). CONCLUSIONS We may conclude that there are hotorecetors sensitive to various wavelengths. The DQE of the hotorecetor is roortional to the fill factor and inversely roortional to the energy required to create an electron-hole air. In addition MTF is ugraded when the ixel size is decreased. If the charge sharing effects are considered, the electrons have a disersion of 60μm which is imortant esecially for small ixel detectors. The effect of Comton electrons originating by the scintillator is smaller due to the smaller disersion of the electrons (6 μm). In clinical conditions it is suggested that the study of the hotorecetor should include the scintillator. It was found that only ortion of the hotorecetors dynamic range is clinically utilized esecially the art corresonding to low atients exosures. Finally a ration of the hotorecetor time resonse over the X-ray tube trigger time less that 1/10 does not affects image acquisition 14

15 REFERENCES 1. Jun Ohta «Smart CMOS Image Sensors and Alication» New York CRC Press Taylor and Francis Grou, LLC Bohndiek, Sarah Elisabeth Active ixel sensors for the breast biosy analysis using X-ray diffraction PhD Thesis, UCL, Janesick J and Putnam G (003) Develoments and alications of higherformance 4. CCD and CMOS imaging arrays. Annu Rev Nucl Part S Janesick J R (001) Scientic Charge Couled Devices. SPIE Press. 6. Janesick J R, Andrews J T and Elliott T (006) Fundamental erformance differences between CMOS and CCD imagers; Part I. In Proc. SPIE, volume Kalyvas N., Valais I., Costaridou L., Kandarakis I., Cavouras D., Nomicos C.D., and Panayiotakis G.: "Evaluating otical sectra matching of hoshor-hotodetector combinations" JINST P07003, July Michail C., Fountos G., Liaarinos P., Kalyvas N et al.: "Light emission efficiency and imaging erformance of GdOS:Eu owder screens under X-ray radiograhy conditions" Med. Phys., 37(7), , P. Magnan, Detection of visible hotons in CCD and CMOS: A Comarative view Nucl. Instrum. and Meth. A. 504, 199-1, (003) 10. Michail C.M., Syrooulou V.A., Fountos G.P., Kalyvas N.I., Valais I.G., Kandarakis I.S. and Panayiotakis G.S.: "Exerimental and Theoretical Evaluation of a High Resolution CMOS Based Detector under X-ray Imaging Conditions" IEEE Transactions on Nuclear Science 58(1), 314-3, Liaarinos P., Kalyvas N., Kandarakis I., Cavouras D. "Analysis of the imaging erformance in indirect digital mammograhy detectors by linear systems and signal detection models" Nuclear Instruments and Methods in Physics Research A 697, 87-98, H K Kim, S.M. Yun, J.S. Ko, G.Cho and T. Graeve: Cascade Modeling of ixelated cintillator Detectors for X-ray Imaging IEEE TNS, 55, , Kalivas N., Costaridou L, Kandarakis I., Cavouras D., Nomicos C.D. Panayiotakis G.: "Effect of intrinsic gain fluctuations on quantum noise of hoshors used in medical x-ray imaging detectors". Alied Physics A 69, , htt://yeneloe.sourceforge.net/ 15. Kalivas N., Costaridou L., Kandarakis I., Cavouras D., Nomicos C.D. and Panayiotakis G.: Modeling quantum and structure noise of hoshors used in medical x-ray imaging detectors. Nuclear Instruments and Methods in Physics Research A 490, , htt://imagej.nih.gov/ij/ 17. htt:// 18. K. Mathieson, M.S. Passmore, P. Seller, et. al. Charge sharing in silicon ixel detectors Nucl. Instrum. Method. Phys. Res. A 48, 113-1, Remote RadEye X-Ray Camera Datasheet. 15

16 APPENDIX Aendix 1. Scrit of ypenelope as it is automatically created. TITLE Si80. >>>>>>>> Electron beam definition. SKPAR 1 [Primary articles: 1=electron, =hoton, 3=ositron] SENERG [Energy of the electron beam, in ev] SPOSIT [Coordinates of the electron source] SDIREC [Direction angles of the beam axis, in deg] SAPERT 0.0 [Beam aerture, in deg] SDIAM 1e-06 [Beam diameter, in cm]. >>>>>>>> Material data and simulation arameters. MFNAME S100.mat [Material file, u to 0 chars] MSIMPA [EABS(1:3),C1,C,WCC,WCR]. >>>>>>>> Geometry of the samle. GEOMFN substrate.geo [Geometry definition file, 0 chars]. >>>>>>>> Job roerties. TRJSC 1 [Track secondary electrons?] NTRJM 1000 [Number of electrons] 16

17 XXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXXX XXXXXXXXXX Substrate SURFACE ( 1) Cylinder of radius 1.50 along z-axis INDICES=( 1, 1, 0, 0,-1) X-SCALE=( E+00, 0) (DEFAULT=1.0) Y-SCALE=( E+00, 0) (DEFAULT=1.0) Z-SCALE=( E+00, 0) (DEFAULT=1.0) OMEGA=( E+00, 0) DEG THETA=( E+00, 0) DEG PHI=( E+00, 0) DEG X-SHIFT=( E+00, 0) Y-SHIFT=( E+00, 0) Z-SHIFT=( E+00, 0) SURFACE ( ) Plane Z=0.00 INDICES=( 0, 0, 0, 1, 0) X-SCALE=( E+00, 0) (DEFAULT=1.0) Y-SCALE=( E+00, 0) (DEFAULT=1.0) Z-SCALE=( E+00, 0) (DEFAULT=1.0) OMEGA=( E+00, 0) DEG THETA=( E+00, 0) DEG PHI=( E+00, 0) DEG X-SHIFT=( E+00, 0) Y-SHIFT=( E+00, 0) Z-SHIFT=( E+00, 0) SURFACE ( 3) Plane Z=-3.00 INDICES=( 0, 0, 0, 1, 0) X-SCALE=( E+00, 0) (DEFAULT=1.0) Y-SCALE=( E+00, 0) (DEFAULT=1.0) Z-SCALE=( E+00, 0) (DEFAULT=1.0) OMEGA=( E+00, 0) DEG 17

18 THETA=( E+00, 0) DEG PHI=( E+00, 0) DEG X-SHIFT=( E+00, 0) Y-SHIFT=( E+00, 0) Z-SHIFT=( E+00, 0) MODULE ( 1) substrate MATERIAL( 1) SURFACE ( 1), SIDE POINTER=(-1) SURFACE ( ), SIDE POINTER=(-1) SURFACE ( 3), SIDE POINTER=( 1) OMEGA=( E+00, 0) DEG THETA=( E+00, 0) DEG PHI=( E+00, 0) DEG X-SHIFT=( E+00, 0) Y-SHIFT=( E+00, 0) Z-SHIFT=( E+00, 0) END Aendix. Calculating convolution with Octave Txray=6000; % irradiation time μs Tdecay=5; % scintillation decay time Thoto=1000; % hotorecetor resonse for i=(1:txray); a(i)=1; %irradiation ulse end Tzero=5*Tdecay; for i=(1:tzero); end b(i)=ex(-i/tdecay); % scintillation decay ulse 18

19 for i=(1:thoto); c(i)=1; %recetor ulse end d=conv(a,b); e=conv(d,c); 19