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PERFORMANCE LIMITS OF PLASTIC SCINTILLATION DETECTOR CCD-BASED DOSIMETERS USING CURRENT TECHNOLOGY
D OSE
MEASUREMENT : SMALL AND COMPOSITE FIELDS AND FLATTENING FREE
S 151
100 Hz Varian Clinac linac pulse repetition rate. Potentially then, a PSD system could be designed to resolve a single dose pulse. The graph presents the precision as a function of integration time for this case.III.B MAXIMUM SENSITIVITY model indicates that a 0.1 cGy dose will suffice to provide a precision of ±1% for a 1x3 mm detector.III.C MAXIMUM SPATIAL RESOLUTION OR MINIMUM DETECTOR VOLUMEWhat is the minimum scintillator volume that can be achieved while providing a precision of ±1% at a dose rate of 400 cGy/min for 1 second of integration time? The model predicts that a scintillator volume of only 0.02 mm3 would be sufficient to obey the above criterias.
Conclusions: The sensitivity of the D-I curves varied according to the bandwidth. Specifically, band-pass filters with narrower bandwidths reduced the quantity of light travelling to the detector. To improve the sensitivity of radiochromic film for dosimetry, it is necessary to select an effective combination of bandwidth for the band-pass filters and intensity for the scanner light source.
Dose measurement: Novel detectors 405 poster (Physics Track) PERFORMANCE LIMITS OF PLASTIC SCINTILLATION DETECTOR CCD-BASED DOSIMETERS USING CURRENT TECHNOLOGY F. Lacroix1 , S. Beddar2 , L. Gingras3 , L. Beaulieu3 1 CHUM - H ÔPITAL N OTRE -DAME, Département de Radio-Oncologie, Montréal, Canada 2 U.T. M.D. A NDERSON C ANCER C ENTER, Radiation Physics, Houston, TX, USA 3 CHUQ-H OTEL -D IEU DE Q UEBEC, Quebec city, Canada
Purpose: A numerical model was developed that allows the calculation of the image signal-to-noise ratio (SNR) as a function of the system parameters for a plastic scintillation detector (PSD). The reason that SNR is interesting is that it is inversely related to the standard deviation in the spot Analog Digital Unit (ADU) counts (i.e. image pixel values) and therefore characterizes the measurement precision of the dosimeter as a function of the system parameters (scintillator volume, deposited dose, coupling efficiency to the CCD, etc.). The SNR has therefore the potential to become an important parameter in the design of PSDs. The SNR calculation model is used to explore the fundamental precision and spatial resolution limits in scintillation dosimetry. Delimiting the performance limits of PSDs will help position this technology amongst the other dosimeter technologies available. Delimiting the optimal application areas of PSDs might help drive the development of novel commercial PSD systems and speed up the adoption of this technology in the clinic. Materials: The SNR in CCD based optoelectronic systems is given by equation 1, appropriately called the "CCD equation" [8]. Where is the photon fluence (photons pixel-1 s-1 ), is the quantum efficiency of the photodetector, T is the integration time for one image in seconds, D is the variance of the dark noise (electrons) and is the variance of the readout noise (electrons). In general, the imaged object (the fiber) subtends a finite number of pixels (n). To obtain the SNR for a collection of pixels, a summation of the signal and noise for each pixel over n pixels must be performed. The number of pixels included in one optical spot will depend on the optical magnification of the objective lens used. The scintillation photon production rate is known for each Gy of dose deposited for most plastic scintillators [9]. The photon fluence rate on the CCD can be calculated by taking into account all the losses suffered by the scintillation photons throughout the optical train (scintillator to fiber coupling loss, fiber attenuation loss, objective lens coupling loss). These losses are either provided by the component manufacturer, can be extracted from the scientific literature or can be calculated from geometrical optics considerations. Calculating the SNR for a specific PSD system is therefore fairly straightforward. SNR=Φp ηq T 2 Φp ηp T +DT +Nr
Results: An "optimal" or close to optimal PSD system employing current technology can be modeled based to derive the ultimate performance limits of scintillation dosimeters using current off-the-shelf technology. The PSD system selected is based on the one presented in [5]. III.A MAXIMUM PRECISIONA precision of ±1% on the dose is attained in less than 10 msec of integration time. Using 10 seconds on integration time allows to reach a precision of ± 0.02%. Note that 10 msec of integration time is on par with the
Conclusions: We have shown that PSD systems using fiber tapers to couple between the optical fibers and the photodetector could achieve a precision of ±1%:1. In less than 10 msec integration time using a 2.35 mm3 volume scintillator (1x3 mm3 ).2. With an absorbed dose of 0.1 cGy using a 2.35 mm3 volume scintillator (1x3 mm3 ).3. With a 0.02 mm3 volume for 1 second of integration time.The SNR is one of the most important parameters to optimize when designing PSD systems.
Dose measurement: small and composite fields and flattening free 406 poster (Physics Track) AN ANALYTICAL FUNCTION TO ESTIMATE BACKSCATTER CONTRIBUTION FROM THE JAWS FOR SMALL PHOTON BEAMS. J. Puxeu Vaqué1 , M. C. Lizuain1 , I. Modolell i Farré1 , R. De Blas Pinol1 , C. Picon1 , F. Pino Sorroche1 , I. Sancho1 , J. Vilar Palop1 , D. Navarro Gimenez1 I NSTITUT C ATALÀ D ’O NCOLOGIA, Medical Physics, L’Hospitalet de Llobregat, Spain 1
Purpose: The purpose of this work is to present a simple analytical approach that has been developed to model monitor chamber backscatter for clinical photon beams used in radio-surgery. Many other investigators have previously studied the effect of modeling monitor chamber backscatter. In order to obtain the contribution of the collimator factor, authors habitually use in air output factors. Some authors assume that backscattered radiation to the monitor chamber has a linear relationship with the irradiated area on the jaw’s upper surface. In this paper, total scatter factors measured for a fixed cone beam collimator have been used. Materials: Measurements have been performed on a Clinac 600D (VARIAN) linear accelerator with a photon beam of nominal energy 6MV and dose rate 240MU/min. The detection system used was a small volume ionization chamber (PTW-31016) in a Solid Phantom (Scanditronix-Welhfer). Total scatter factors were determined using two different cones from BrainLab (4 cm and 2.5 cm diameter). For each cone, the backup jaws were increased in 0.5 cm intervals ranging from 4.5 cm to 7 cm for the 4 cm cone and from 3 cm to 7 cm for the 2.5 cm cone.The factors were normalized to the 7 cm backup jaw value, thus avoiding the contribution of the Sp, since the irradiated volume in the phantom was not changed for a given cone. As other authors have previously reported, it is a good approximation to assume that the backscattered radiation to the monitor chamber has a linear relationship with the area of jaws irradiated, and that the upper jaws will have a greater influence. Using this relationship, and assuming that beam irradiating the jaws is not flat, we assumed a polynomial function (Z) where upper Y jaws have a linear and a quadratic term and the lower X jaws a linear term only. Such a function is
MEASUREMENT : SMALL AND COMPOSITE FIELDS AND FLATTENING FREE
S 151
100 Hz Varian Clinac linac pulse repetition rate. Potentially then, a PSD system could be designed to resolve a single dose pulse. The graph presents the precision as a function of integration time for this case.III.B MAXIMUM SENSITIVITY model indicates that a 0.1 cGy dose will suffice to provide a precision of ±1% for a 1x3 mm detector.III.C MAXIMUM SPATIAL RESOLUTION OR MINIMUM DETECTOR VOLUMEWhat is the minimum scintillator volume that can be achieved while providing a precision of ±1% at a dose rate of 400 cGy/min for 1 second of integration time? The model predicts that a scintillator volume of only 0.02 mm3 would be sufficient to obey the above criterias.
Conclusions: The sensitivity of the D-I curves varied according to the bandwidth. Specifically, band-pass filters with narrower bandwidths reduced the quantity of light travelling to the detector. To improve the sensitivity of radiochromic film for dosimetry, it is necessary to select an effective combination of bandwidth for the band-pass filters and intensity for the scanner light source.
Dose measurement: Novel detectors 405 poster (Physics Track) PERFORMANCE LIMITS OF PLASTIC SCINTILLATION DETECTOR CCD-BASED DOSIMETERS USING CURRENT TECHNOLOGY F. Lacroix1 , S. Beddar2 , L. Gingras3 , L. Beaulieu3 1 CHUM - H ÔPITAL N OTRE -DAME, Département de Radio-Oncologie, Montréal, Canada 2 U.T. M.D. A NDERSON C ANCER C ENTER, Radiation Physics, Houston, TX, USA 3 CHUQ-H OTEL -D IEU DE Q UEBEC, Quebec city, Canada
Purpose: A numerical model was developed that allows the calculation of the image signal-to-noise ratio (SNR) as a function of the system parameters for a plastic scintillation detector (PSD). The reason that SNR is interesting is that it is inversely related to the standard deviation in the spot Analog Digital Unit (ADU) counts (i.e. image pixel values) and therefore characterizes the measurement precision of the dosimeter as a function of the system parameters (scintillator volume, deposited dose, coupling efficiency to the CCD, etc.). The SNR has therefore the potential to become an important parameter in the design of PSDs. The SNR calculation model is used to explore the fundamental precision and spatial resolution limits in scintillation dosimetry. Delimiting the performance limits of PSDs will help position this technology amongst the other dosimeter technologies available. Delimiting the optimal application areas of PSDs might help drive the development of novel commercial PSD systems and speed up the adoption of this technology in the clinic. Materials: The SNR in CCD based optoelectronic systems is given by equation 1, appropriately called the "CCD equation" [8]. Where is the photon fluence (photons pixel-1 s-1 ), is the quantum efficiency of the photodetector, T is the integration time for one image in seconds, D is the variance of the dark noise (electrons) and is the variance of the readout noise (electrons). In general, the imaged object (the fiber) subtends a finite number of pixels (n). To obtain the SNR for a collection of pixels, a summation of the signal and noise for each pixel over n pixels must be performed. The number of pixels included in one optical spot will depend on the optical magnification of the objective lens used. The scintillation photon production rate is known for each Gy of dose deposited for most plastic scintillators [9]. The photon fluence rate on the CCD can be calculated by taking into account all the losses suffered by the scintillation photons throughout the optical train (scintillator to fiber coupling loss, fiber attenuation loss, objective lens coupling loss). These losses are either provided by the component manufacturer, can be extracted from the scientific literature or can be calculated from geometrical optics considerations. Calculating the SNR for a specific PSD system is therefore fairly straightforward. SNR=Φp ηq T 2 Φp ηp T +DT +Nr
Results: An "optimal" or close to optimal PSD system employing current technology can be modeled based to derive the ultimate performance limits of scintillation dosimeters using current off-the-shelf technology. The PSD system selected is based on the one presented in [5]. III.A MAXIMUM PRECISIONA precision of ±1% on the dose is attained in less than 10 msec of integration time. Using 10 seconds on integration time allows to reach a precision of ± 0.02%. Note that 10 msec of integration time is on par with the
Conclusions: We have shown that PSD systems using fiber tapers to couple between the optical fibers and the photodetector could achieve a precision of ±1%:1. In less than 10 msec integration time using a 2.35 mm3 volume scintillator (1x3 mm3 ).2. With an absorbed dose of 0.1 cGy using a 2.35 mm3 volume scintillator (1x3 mm3 ).3. With a 0.02 mm3 volume for 1 second of integration time.The SNR is one of the most important parameters to optimize when designing PSD systems.
Dose measurement: small and composite fields and flattening free 406 poster (Physics Track) AN ANALYTICAL FUNCTION TO ESTIMATE BACKSCATTER CONTRIBUTION FROM THE JAWS FOR SMALL PHOTON BEAMS. J. Puxeu Vaqué1 , M. C. Lizuain1 , I. Modolell i Farré1 , R. De Blas Pinol1 , C. Picon1 , F. Pino Sorroche1 , I. Sancho1 , J. Vilar Palop1 , D. Navarro Gimenez1 I NSTITUT C ATALÀ D ’O NCOLOGIA, Medical Physics, L’Hospitalet de Llobregat, Spain 1
Purpose: The purpose of this work is to present a simple analytical approach that has been developed to model monitor chamber backscatter for clinical photon beams used in radio-surgery. Many other investigators have previously studied the effect of modeling monitor chamber backscatter. In order to obtain the contribution of the collimator factor, authors habitually use in air output factors. Some authors assume that backscattered radiation to the monitor chamber has a linear relationship with the irradiated area on the jaw’s upper surface. In this paper, total scatter factors measured for a fixed cone beam collimator have been used. Materials: Measurements have been performed on a Clinac 600D (VARIAN) linear accelerator with a photon beam of nominal energy 6MV and dose rate 240MU/min. The detection system used was a small volume ionization chamber (PTW-31016) in a Solid Phantom (Scanditronix-Welhfer). Total scatter factors were determined using two different cones from BrainLab (4 cm and 2.5 cm diameter). For each cone, the backup jaws were increased in 0.5 cm intervals ranging from 4.5 cm to 7 cm for the 4 cm cone and from 3 cm to 7 cm for the 2.5 cm cone.The factors were normalized to the 7 cm backup jaw value, thus avoiding the contribution of the Sp, since the irradiated volume in the phantom was not changed for a given cone. As other authors have previously reported, it is a good approximation to assume that the backscattered radiation to the monitor chamber has a linear relationship with the area of jaws irradiated, and that the upper jaws will have a greater influence. Using this relationship, and assuming that beam irradiating the jaws is not flat, we assumed a polynomial function (Z) where upper Y jaws have a linear and a quadratic term and the lower X jaws a linear term only. Such a function is