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Si-PIN photodiode readout for a scintillating optical fiber dosimeter
Radiation Measurements 47 (2012) 947e950
Contents lists available at SciVerse ScienceDirect
Radiation Measurements journal homepage: www.elsevier.com/locate/radmeas
Si-PIN photodiode readout for a scintillating optical fiber dosimeter Florbela Rêgo a, Luis Peralta a, b, * a b
Laboratório de Instrumentação e Física Experimental de Partículas, Av. Elias Garcia 14-1, 1000-149 Lisboa, Portugal Faculdade de Ciências da Universidade de Lisboa, FCUL, Edificio C8, Campo Grande, 1749-016 Lisboa, Portugal
h i g h l i g h t s < Photodiode S9195 from Hamamatsu is used as a readout device for an optical fiber dosimeter. < Doses in the mGy range were measured with the fiber dosimeter prototype. < The S9195 photodiode signal temperature dependence was measured in the 0e50 C range. < A 0.2% C1 temperature coefficient was obtained for the 15e25 C range and visible region.
a r t i c l e i n f o
a b s t r a c t
Article history: Received 23 March 2012 Received in revised form 28 June 2012 Accepted 26 July 2012
For more than a decade plastic optical fiber based dosimeters have been developed for medical applications. PMT’s have been widely used as the readout photodetector but for a number of applications, SiPIN photodiodes are a good alternative. They are robust, present good stability over time and have enough sensitivity so that an electrometer can be used as a measuring device. This work describes the tests made with the S9195 Si-PIN photodiode from Hamamatsu used as the readout photodetector for a scintillating optical fiber dosimeter using a BCF-60 polystyrene optical fiber. Ó 2012 Elsevier Ltd. All rights reserved.
Keywords: Scintillating fibers Photodiode Dosimetry
1. Introduction A scintillation dosimeter has three main components (Fontbonne et al., 2002); a scintillator, a clear light-guide and a photodetector. Plastic scintillators can be used as the first component Beddar et al., 1992b), while PMMA fibers have been widely used as the second component (Andersen, 2011). Photomultipliers (PMTs) have been a popular choice as a readout photodetector (Beddar et al., 1992a,b; Beddar, 2007; Williamson et al., 1999) since they have high gain (106 or more) and a dark current in the range of 1e10 nA for typical high voltage (HV) bias values (Hamamatsu, 2006). Si-PIN photodiodes have also been used as photodetectors in scintillation dosimetry (Létourneau et al., 1999; Lee et al., 2007; Rego and Peralta, 2008) although they present smaller gain. On a scintillation dosimeter light is emitted by the dopant of the plastic bulk material when irradiated by ionizing radiation. The amount of light emitted is proportional to the energy
* Corresponding author. Faculdade de Ciências da Universidade de Lisboa, FCUL, Edificio C8, Campo Grande, 1749-016 Lisboa, Portugal. Tel.: þ351 217500945. E-mail address: [email protected] (L. Peralta). 1350-4487/$ e see front matter Ó 2012 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.radmeas.2012.07.019
deposited, at least for low ionizing density radiation, like X or gamma photons (Knoll, 2000). Unfortunately, fluorescent light is also produced at the scintillator and clear fiber bulk material. As shown in earlier works (Williamson et al., 1999) it introduces a dependence of the output signal on the incident radiation energy. Another well-known source of background is the Cherenkov light emitted by electrons (Lambert et al., 2008). Cherenkov light is generated in a medium of refraction index n when electrons travel faster than c/n, where c is the speed of light in vacuum. For optical fibers, the signal from Cherenkov light is highly dependent on the fiber-beam relative configuration and particularly the angle between the electron direction and the fiber axis (Beddar et al., 2004). Moreover, the amount of deposited energy needed to produce Cherenkov photons is much less than the amount needed to produce scintillation photons of the same wavelength (Leo, 1994), introducing an energy dependence on the output signal. This subject has been extensively studied in the literature (Beddar et al., 1992a; Clift et al., 2000; Frelin et al., 2005; Law et al., 2006) and several solutions have been proposed to overcome the Cherenkov light influence in the final signal (Arnfield et al., 1996; Frelin et al., 2005; Lambert et al., 2008). These solutions have particular relevance for the high energy
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F. Rêgo, L. Peralta / Radiation Measurements 47 (2012) 947e950
beams used in external radiotherapy. In radiology, the energy beam is below the Cherenkov threshold and the effect is not present. For brachytherapy applications that may use sources with intermediate energies (in the hundred keV range) some Cherenkov light will be produced. Most of the scintillator based dosimeters developed over the years (Becks et al., 2000; Fluhs et al., 1996; Frelin et al., 2005; Justus et al., 2004; Lambert et al., 2008) use complex readout systems to avoid the Cherenkov light influence in the output signal, but a system based on a red-green-blue photodiode has been successfully used to remove the steam effect on a scintillator dosimeter used in Ir-192 HDR brachytherapy (Therriault-Proulx et al., 2011). For medium/low energy the contribution due to Cherenkov light will be small or nonexistent and the readout system can be simplified, provided the fluorescent light contribution is small. Several radiological exams deliver radiation doses in the range of hundreds of mGy to a few mGy (Mettler et al., 2008), which is enough for a plastic scintillator to produce a light signal that can be detected by a Si-PIN photodiode (Rego, 2010). This type of photodetector can in fact be an alternative to the more expensive and mechanically fragile PMT. The PMT needs HV to work and the gain is highly dependent on the applied HV and as a consequence, the calibration constants are also dependent on the HV when the device is operated in current mode. On the other hand, Si-PIN photodiodes are robust and cheap. They can be operated without any bias voltage, in which case the leakage current is zero. This is an advantage over a bias operated photodiode, where a more complex acquisition circuit is required (El Mourabit et al., 2009). The photodiode can be DC coupled to an electrometer, greatly simplifying the output circuit. Furthermore, typical photo-sensitivities are in the range of a few hundred mA/W and span over 300e1000 nm in wavelength (with maximum values in the near infrared), making them suitable for light detection from a wide selection of scintillators. The aim of this study is to demonstrate the ability to use Si-PIN photodiodes as a readout device of a dosimeter capable of measuring dose in the 0.1 to several tens of mGy range, suitable for many radiological applications. Plastic scintillators have a density close to water, which makes them nearly invisible on a radiological scan. This type of dosimeter can thus have an important role in dose control of frequent exams such as panoramic dental radiography, where the target structures to scan have a density higher than water. 2. The dosimeter prototype In this work the chosen photodiode was the S9195 from Hamamatsu. It is a relatively low-cost photodiode, presenting a large sensitive area (5.0 5.0 mm2) and high photo-sensitivity (0.28 A/W for l ¼ 405 nm). The S9195 spectral response spans from the near ultraviolet to the infrared, allowing for use with several different scintillator types. The dosimeter prototype uses a 20 mm-long, 2 mm in diameter BCF-60 scintillating fiber, connected to a 1.5 m-long, 2 mm in diameter PMMA clear fiber (ESKAÒ acrylic fiber optics developed and manufactured by Mitsubishi). The scintillating fiber is wrapped in a white TeflonÒ film for more efficient light collection. The scintillating fibers and clear fibers are protected from light by a thin black plastic jacket. For the tests, coupling between scintillating and clear fiber is made with optical grease (Saint-Gobain BC-630). For the final prototype fibers could be glued for greater mechanical stability. This clear fiber is connected to the S9195 photodiode while the optical coupling between them is also assured by optical grease. The photodiode is fixed in a PVC holder which ensures the clear fiber immobilization through a plastic screw. A photograph of the dosimeter prototype is displayed in Fig. 1.
Fig. 1. The optical fiber dosimeter prototype. The scintillating fiber is seen outside the protective black-jacket. The S9195 photodiode is inside the PVC holder.
The box is then closed and wrapped in aluminum foil to ensure that is completely light-tight. To avoid the production of signal at the photodiode due to scattered radiation the box was shielded by 5 cm-thick lead bricks. The scintillating fiber was placed at the surface of a 20 15 cm2 PMMA phantom and exposed to the X-ray beam produced by a Philips PW2184/00 tube with a Tungsten anode and a current of 10 mA. The X-ray beam passed through a circular lead collimator producing a radiation field with a diameter of 20 cm at 40 cm from the tube exit window. An additional filtration of 1 mm of Al was used for the tests. The unbiased photodiode was directly connected to a PTW UNIDOS E electrometer to measure the charge produced. To measure the dose at the fiber position a PTW M23342 parallel plate ionization chamber connected to a PTW UNIDOS E electrometer was used. The ionization chamber was placed inside an identical PMMA phantom in such a way the sensitive volume was at the phantom’s surface. The center of the chamber’s sensitive volume was then placed in the exact point where the scintillating fiber center was before. Data were acquired at several distances from the tube’s exit window, so that different doses could be measured for two X-ray tube voltages (30 and 50 kV). At each position the electric charge produced in the photodiode within 10 s was measured. The same procedure was used to obtain the dose measured by the ionization chamber. The obtained charge versus dose plot is presented in Fig. 2, where linear fits of the data are also included. These fits show the strong linear relationship between the signal produced by the photodiode and the delivered dose. The experimental uncertainties on the charge measurements are below 2% while the uncertainties on the dose measurements are below 1%. The difference in the slope between the two lines is due to energy dependence of the scintillation dosimeter. The energy dependence of plastic dosimeters for low energy photon-beams has been assigned to ionization quenching in the scintillator (Frelin et al., 2008; Williamson et al., 1999) and to the variation of the mass coefficient ratio between plastic and water (Oliveira et al., 2011) in that energy range. For applications with a fixed photon energy beam (as for most diagnostic X-ray exams) the effect on the dose measurement should be small, provided a calibration of the device is made for the beam energy. 3. Photodiode temperature dependence To use the photodiode as a photodetector in dosimetry, good temperature stability is an important requirement. The temperature dependence of the photodiode response to visible and infrared light was measured using a special setup. This setup included
F. Rêgo, L. Peralta / Radiation Measurements 47 (2012) 947e950
BCF-60
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S9195 photodiode 1.1
30 25
Normalized Signal
Charge (pC)
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0.95 LED Green LED Blue LED Red LED IV
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0 0
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Dose (mGy) Fig. 2. Charge produced by the photodiode as a function of the dose for two X-ray tube voltages. The uncertainties on the charge measurements are below 2%. The line is a linear fit to the data points.
a constant flux of dry air passing through a sample-box containing the photodiode and maintaining a constant temperature inside (Fig. 3). The sample-box was made of aluminum. The wall was 10 mmthick and was wrapped in a plastic insulating sheet. The box had two nozzles, one for air entry and another for air exit on the opposite wall. Inside the box the temperature was measured with a LM135 integrated temperature sensor with an accuracy of 0.5 C when calibrated at 25 C. Dry air at room temperature was introduced into a cooling-heating circuit to get it at the desired temperature inside the sample-box. In a first stage, the air passed through an ice-cooled copper pipe coil and its temperature lowered. Then, the air passed through a heating box where an electrical resistance boosted the air temperature to higher values before entering the sample-box. A LED driven by a current source produced the light signal that was conducted by an optical fiber to the photodiode inside the sample-box. In this, the LED was always kept at constant room temperature. The signal produced in the photodiode was measured by a Standard Imaging MAX 4000 electrometer. Four different LEDs were used in the measurements and the peak wavelength of each LED was: 470 nm for blue, 565 nm for green, 660 nm for red and 930 nm for infrared. The LEDs current was measured in order to assess their stability, and variations were less than 0.5%. The photodiodes output signal was measured for temperatures between 0 and 50 C. The output values were normalized to the room temperature (20 C) value for each LED (Fig. 4). For this photodiode a variation no greater than 4% in the
Fig. 3. Experimental setup to study the temperature dependence of the photodiode’s response.
0
10
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30
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50
Fig. 4. Normalized photodiode response as a function of temperature when irradiated with light from different LEDs.
output signal is observed in the temperature range between 0 and 50 C and less than 2% in the 15e25 C range. The very low temperature dependence is a positive factor relative to the use of a PMT as photodetector. For instance, a typical PMT used for medical dosimetry applications (Williamson et al., 1999; Fluhs et al., 1996), R647 from Hamamatsu has a decrease in sensitivity of about 20% in the 0e40 C range. The PMT temperature dependence when operated in DC mode, introduces the need for larger calibration factor corrections relative to those used by photodiodes. 4. Discussion and conclusions Si-PIN photodiodes have photo-sensitivities in the range of a few hundred of A/W, which are enough for several applications in scintillator dosimetry. In this work the Hamamatsu S9195 photodiode was used with effective results. The S9195 photo-sensitivity increases from blue to infrared, enhancing the long wavelength detection. Doses in the mGy range were measured with this prototype and the measurement of lower dose values can be envisaged. In order to reduce the influence of electronic noise during the measurements the electrometer should be as close as possible to the photodiode. Also care must be advised on the placement of the photodiode box near the X-ray beam. The photodiode is sensitive to the radiation and a direct signal can be generated in the photodiode due to scattered radiation in the room walls. In the preset tests the photodiode box was shielded with lead bricks. A more manageable solution is to use a longer clear fiber placing the photodiode far away from the radiation source. The photodiode response dependence on the temperature has been studied. Under normal clinical conditions (15e25 C) a 2% variation was found within that range. This variation can be translated into a temperature coefficient of 0.2% C1 in the 15e 25 C range, compatible with the very low temperature coefficient claimed by the photodiode’s manufacturer (Hamamatsu, 2004) in the visible and near infrared region. Although yield of scintillation light has a small dependence on temperature (Beddar et al., 1992) a variation of a few percent has been reported for unloaded and doped polystyrene (Nowotny and Taubeck, 2009). Depending on which dye is used (Nowotny and Taubeck, 2009), polystyrene scintillator can exhibit 1e2% change in light yield within the 15e25 C range. In some cases variation in the scintillator light yield is opposite to the variation of the photodiode response. This leads to a compensation of the two effects with a positive outcome in the dosimeter temperature stability.
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Plastic scintillator dosimeters can play an important role in radiological applications where an image-invisible dosimeter is required, such as the dental panoramic radiography. The present work demonstrates that charges in excess of a pC can be obtained for doses in the mGy range using the S9195 photodiode coupled a PMMA optical cable and a plastic scintillator. For these tests an UNIDOS E electrometer was used. The measured current is within the input range of the available low cost monolithic electrometer operational amplifier AD549 (Analog Devices, 2008), which will be used in further developments of the project. As in earlier works (Létourneau et al., 1999; Williamson et al., 1999), a second channel will be present for the subtraction of the fluorescent light produced in the PMMA cable. Acknowledgments We want to thank to Laboratório de Instrumentação e Física Experimental de Partículas (LIP) for the financial support of this work and in particular the LIP’s mechanical workshops for the construction of the water phantom container. We thank Prof. Maria Luisa Carvalho for the use of the X-ray facility at Lisbon University. We are grateful to Ms. Ashley Merrill for corrections to the text. References Analog Devices, 2008. AD549: Ultralow Input-Bias Current Operational Amplifier. http://www.analog.com/static/imported-files/data_sheets/AD549.pdf (accessed 2012). Andersen, C.E., 2011. Fiber-coupled luminescence dosimetry in therapeutic and diagnostic radiology. AIP Conf. Proc. 1345, 100e119. http://dx.doi.org/10.1063/1. 3576161. Arnfield, M.R., Gaballa, H.E., Zwicker, R.D., Islam, Q., Schmidt-Ullrich, R., 1996. Radiation-Induced light in optical fibers and plastic scintillators: application to brachytherapy dosimetry. IEEE Trans. Nucl. Sci. 43, 2077e2084. Becks, K.-H., Drees, J., Goldmann, K., Gregor, I.M., Heintz, M., 2000. A multichannel dosimeter based on scintillating fibers for medical applications. Nucl. Instrum. Methods. Phys. Res. A. 454, 147e151. Beddar, A.S., 2007. Plastic scintillation dosimetry and its application to radiotherapy. Radiat. Meas. 41, S124eS133. Beddar, A.S., Mackie, T.R., Attix, F.H., 1992a. Cerenkov light generated in optical fibres and other light pipes irradiated by electrons. Phys. Med. Biol. 37, 925e 935. Beddar, A.S., Mackie, T.R., Attix, F.H., 1992b. Water-equivalent plastic scintillation detectors for high-energy beam dosimetry: I. Physical characteristics and theoretical consideration. Phys. Med. Biol. 37, 1883e1900. Beddar, A.S., Suchowerska, N., Law, S., 2004. Plastic scintillation dosimetry for radiation therapy: minimizing capture of Cerenkov radiation noise. Phys. Med. Biol. 49, 783e790. Clift, M.A., Sutton, R.A., Webb, D.V., 2000. Dealing with Cerenkov radiation generated in organic scintillator dosimeters by bremsstrahlung beams. Phys. Med. Biol. 45, 1165e1182.
El Mourabit, A., Lu, G.-N., Pittet, P., 2009. On the design of a low noise readout circuit for in-vivo dosimeter, IEEE 2009 International Conference on Microelectronics, http://dx.doi.org/10.1109/ICM.2009.5418587. Fluhs, D., Heintz, M., Indenkampen, F., Wieczorek, C., 1996. Direct reading measurements of absorbed dose with plastic scintillator e the general concept and applications to ophthalmic plaque dosimetry. Med. Phys. 23, 427e434. Fontbonne, J.M., Iltis, G., Ban, G., Battala, A., Vernhes, J.C., Tillier, J., Bellaize, N., Le Brun, C., Tamain, B., Mercier, K., Motin, J.C., 2002. Scintillating fiber dosimeter for radiation therapy accelerator. IEEE Trans. Nucl. Sci. 49, 2223e2227. Frelin, A.M., Fontbonne, J.M., Ban, G., Colin, J., Labalme, M., Batalla, A., Isambert, A., Vela, A., Leroux, T., 2005. Spectral discrimination of Cerenkov radiation in scintillating dosimeters. Med. Phys. 32, 3000e3006. Frelin, A.M., Fontbonne, J.M., Ban, G., Colin, J., Labalme, M., 2008. Comparative study of plastic scintillators for dosimetric applications, for dosimetric applications. IEEE Trans. Nucl. Sci. 55, 2749e2756. Hamamatsu, 2004. http://sales.hamamatsu.com/assets/pdf/parts_S/S9195.pdf (accessed 2012). Hamamatsu, 2006. Photomultiplier Tubes, Basics and Applications. Hamamatsu Photonics K.K. http://sales.hamamatsu.com/assets/applications/ETD/pmt_ handbook_complete.pdf (accessed 2012). Justus, B.L., Falkenstein, P., Huston, A.L., Plazas, M.C., Ning, H., Miller, R.W., 2004. Gated fiber-optic-coupled detector for in vivo real-time radiation dosimetry. Appl. Opt. 43, 1663e1668. Knoll, G., 2000. Radiation Detection and Measurement. John Wiley &Sons, New York. Lambert, J., Yin, Y., McKenzie, D.R., Law, S., Suchowerska, N., 2008. Cerenkov-free scintillation dosimetry in external beam radiotherapy with an air core light guide. Phys. Med. Biol. 53, 3071e3080. Law, S.H., Fleming, S.C., Suchowerska, N., McKenzie, D.R., 2006. Optical fiber design and the trapping of Cerenkov radiation. Appl. Opt. 45, 9151e9159. Lee, B., Jang, K.W., Cho, D.H., Yoo, W.J., Kim, H.S., Chung, S.C., Yi, J.H., 2007. Development of one-dimensional fiber-optic radiation sensor for measuring dose distributions of high energy photon beams. Opt. Rev. 14, 351e354. Leo, W.R., 1994. Techniques for Nuclear and Particle Physics Experiments: a How-To Approach. Springer-Verlag Berlin, Heidelberg. Létourneau, D., Pouliot, J., Roy, R., 1999. Miniature scintillating detector for small field radiation therapy. Med. Phys. 26, 2555e2561. Mettler, F.A., Huda, W., Yoshizumi, T.T., Mahesh, M., 2008. Effective doses in radiology and diagnostic nuclear medicine: a catalog. Radiology 248, 254e263. Nowotny, R., Taubeck, A., 2009. A method for the production of composite scintillators for dosimetry in diagnostic radiology. Phys. Med. Biol. 54, 1457e1468. Oliveira, C.H., Rego, F., Peralta, L., Abreu, M.C., 2011. 2011 IEEE Nuclear Science Symposium Conference Record, 1612e1615, http://dx.doi.org/10.1109/NSSMIC. 2011.6154644. Rego, F., 2010, Desenvolvimento de dosímetros de estado sólido para dosimetria em radiologia e braquiterapia, PhD thesis (in Portuguese), University of Lisbon, http:// repositorio.ul.pt/bitstream/10451/2414/1/ulsd058995_td_Florbela_Rego.pdf, (accessed 2011). Rego, F., Peralta, L., 2008. Scintillating optical fiber dosimetry with photodiode readout. In: Nuclear Science Symposium Conference Record, NSS ’08. IEEE, pp. 1035e1037. http://dx.doi.org/10.1109/NSSMIC.2008.4774574. Therriault-Proulx, F., Beddar, S., Briere, T.M., Archambault, L., Beaulieu, L., 2011. Removing the stem effect when performing Ir-192 HDR brachytherapy in vivo dosimetry using plastic scintillation detectors: a relevant and necessary step. Med. Phys. 38, 2176e2179. Williamson, J.F., Dempsey, J.F., Kirov, A.S., Monroe, J.I., Binns, W.R., Hedtjarn, H., 1999. Plastic scintillator response to low-energy photons. Phys. Med. Biol. 44, 857e871.
Contents lists available at SciVerse ScienceDirect
Radiation Measurements journal homepage: www.elsevier.com/locate/radmeas
Si-PIN photodiode readout for a scintillating optical fiber dosimeter Florbela Rêgo a, Luis Peralta a, b, * a b
Laboratório de Instrumentação e Física Experimental de Partículas, Av. Elias Garcia 14-1, 1000-149 Lisboa, Portugal Faculdade de Ciências da Universidade de Lisboa, FCUL, Edificio C8, Campo Grande, 1749-016 Lisboa, Portugal
h i g h l i g h t s < Photodiode S9195 from Hamamatsu is used as a readout device for an optical fiber dosimeter. < Doses in the mGy range were measured with the fiber dosimeter prototype. < The S9195 photodiode signal temperature dependence was measured in the 0e50 C range. < A 0.2% C1 temperature coefficient was obtained for the 15e25 C range and visible region.
a r t i c l e i n f o
a b s t r a c t
Article history: Received 23 March 2012 Received in revised form 28 June 2012 Accepted 26 July 2012
For more than a decade plastic optical fiber based dosimeters have been developed for medical applications. PMT’s have been widely used as the readout photodetector but for a number of applications, SiPIN photodiodes are a good alternative. They are robust, present good stability over time and have enough sensitivity so that an electrometer can be used as a measuring device. This work describes the tests made with the S9195 Si-PIN photodiode from Hamamatsu used as the readout photodetector for a scintillating optical fiber dosimeter using a BCF-60 polystyrene optical fiber. Ó 2012 Elsevier Ltd. All rights reserved.
Keywords: Scintillating fibers Photodiode Dosimetry
1. Introduction A scintillation dosimeter has three main components (Fontbonne et al., 2002); a scintillator, a clear light-guide and a photodetector. Plastic scintillators can be used as the first component Beddar et al., 1992b), while PMMA fibers have been widely used as the second component (Andersen, 2011). Photomultipliers (PMTs) have been a popular choice as a readout photodetector (Beddar et al., 1992a,b; Beddar, 2007; Williamson et al., 1999) since they have high gain (106 or more) and a dark current in the range of 1e10 nA for typical high voltage (HV) bias values (Hamamatsu, 2006). Si-PIN photodiodes have also been used as photodetectors in scintillation dosimetry (Létourneau et al., 1999; Lee et al., 2007; Rego and Peralta, 2008) although they present smaller gain. On a scintillation dosimeter light is emitted by the dopant of the plastic bulk material when irradiated by ionizing radiation. The amount of light emitted is proportional to the energy
* Corresponding author. Faculdade de Ciências da Universidade de Lisboa, FCUL, Edificio C8, Campo Grande, 1749-016 Lisboa, Portugal. Tel.: þ351 217500945. E-mail address: [email protected] (L. Peralta). 1350-4487/$ e see front matter Ó 2012 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.radmeas.2012.07.019
deposited, at least for low ionizing density radiation, like X or gamma photons (Knoll, 2000). Unfortunately, fluorescent light is also produced at the scintillator and clear fiber bulk material. As shown in earlier works (Williamson et al., 1999) it introduces a dependence of the output signal on the incident radiation energy. Another well-known source of background is the Cherenkov light emitted by electrons (Lambert et al., 2008). Cherenkov light is generated in a medium of refraction index n when electrons travel faster than c/n, where c is the speed of light in vacuum. For optical fibers, the signal from Cherenkov light is highly dependent on the fiber-beam relative configuration and particularly the angle between the electron direction and the fiber axis (Beddar et al., 2004). Moreover, the amount of deposited energy needed to produce Cherenkov photons is much less than the amount needed to produce scintillation photons of the same wavelength (Leo, 1994), introducing an energy dependence on the output signal. This subject has been extensively studied in the literature (Beddar et al., 1992a; Clift et al., 2000; Frelin et al., 2005; Law et al., 2006) and several solutions have been proposed to overcome the Cherenkov light influence in the final signal (Arnfield et al., 1996; Frelin et al., 2005; Lambert et al., 2008). These solutions have particular relevance for the high energy
948
F. Rêgo, L. Peralta / Radiation Measurements 47 (2012) 947e950
beams used in external radiotherapy. In radiology, the energy beam is below the Cherenkov threshold and the effect is not present. For brachytherapy applications that may use sources with intermediate energies (in the hundred keV range) some Cherenkov light will be produced. Most of the scintillator based dosimeters developed over the years (Becks et al., 2000; Fluhs et al., 1996; Frelin et al., 2005; Justus et al., 2004; Lambert et al., 2008) use complex readout systems to avoid the Cherenkov light influence in the output signal, but a system based on a red-green-blue photodiode has been successfully used to remove the steam effect on a scintillator dosimeter used in Ir-192 HDR brachytherapy (Therriault-Proulx et al., 2011). For medium/low energy the contribution due to Cherenkov light will be small or nonexistent and the readout system can be simplified, provided the fluorescent light contribution is small. Several radiological exams deliver radiation doses in the range of hundreds of mGy to a few mGy (Mettler et al., 2008), which is enough for a plastic scintillator to produce a light signal that can be detected by a Si-PIN photodiode (Rego, 2010). This type of photodetector can in fact be an alternative to the more expensive and mechanically fragile PMT. The PMT needs HV to work and the gain is highly dependent on the applied HV and as a consequence, the calibration constants are also dependent on the HV when the device is operated in current mode. On the other hand, Si-PIN photodiodes are robust and cheap. They can be operated without any bias voltage, in which case the leakage current is zero. This is an advantage over a bias operated photodiode, where a more complex acquisition circuit is required (El Mourabit et al., 2009). The photodiode can be DC coupled to an electrometer, greatly simplifying the output circuit. Furthermore, typical photo-sensitivities are in the range of a few hundred mA/W and span over 300e1000 nm in wavelength (with maximum values in the near infrared), making them suitable for light detection from a wide selection of scintillators. The aim of this study is to demonstrate the ability to use Si-PIN photodiodes as a readout device of a dosimeter capable of measuring dose in the 0.1 to several tens of mGy range, suitable for many radiological applications. Plastic scintillators have a density close to water, which makes them nearly invisible on a radiological scan. This type of dosimeter can thus have an important role in dose control of frequent exams such as panoramic dental radiography, where the target structures to scan have a density higher than water. 2. The dosimeter prototype In this work the chosen photodiode was the S9195 from Hamamatsu. It is a relatively low-cost photodiode, presenting a large sensitive area (5.0 5.0 mm2) and high photo-sensitivity (0.28 A/W for l ¼ 405 nm). The S9195 spectral response spans from the near ultraviolet to the infrared, allowing for use with several different scintillator types. The dosimeter prototype uses a 20 mm-long, 2 mm in diameter BCF-60 scintillating fiber, connected to a 1.5 m-long, 2 mm in diameter PMMA clear fiber (ESKAÒ acrylic fiber optics developed and manufactured by Mitsubishi). The scintillating fiber is wrapped in a white TeflonÒ film for more efficient light collection. The scintillating fibers and clear fibers are protected from light by a thin black plastic jacket. For the tests, coupling between scintillating and clear fiber is made with optical grease (Saint-Gobain BC-630). For the final prototype fibers could be glued for greater mechanical stability. This clear fiber is connected to the S9195 photodiode while the optical coupling between them is also assured by optical grease. The photodiode is fixed in a PVC holder which ensures the clear fiber immobilization through a plastic screw. A photograph of the dosimeter prototype is displayed in Fig. 1.
Fig. 1. The optical fiber dosimeter prototype. The scintillating fiber is seen outside the protective black-jacket. The S9195 photodiode is inside the PVC holder.
The box is then closed and wrapped in aluminum foil to ensure that is completely light-tight. To avoid the production of signal at the photodiode due to scattered radiation the box was shielded by 5 cm-thick lead bricks. The scintillating fiber was placed at the surface of a 20 15 cm2 PMMA phantom and exposed to the X-ray beam produced by a Philips PW2184/00 tube with a Tungsten anode and a current of 10 mA. The X-ray beam passed through a circular lead collimator producing a radiation field with a diameter of 20 cm at 40 cm from the tube exit window. An additional filtration of 1 mm of Al was used for the tests. The unbiased photodiode was directly connected to a PTW UNIDOS E electrometer to measure the charge produced. To measure the dose at the fiber position a PTW M23342 parallel plate ionization chamber connected to a PTW UNIDOS E electrometer was used. The ionization chamber was placed inside an identical PMMA phantom in such a way the sensitive volume was at the phantom’s surface. The center of the chamber’s sensitive volume was then placed in the exact point where the scintillating fiber center was before. Data were acquired at several distances from the tube’s exit window, so that different doses could be measured for two X-ray tube voltages (30 and 50 kV). At each position the electric charge produced in the photodiode within 10 s was measured. The same procedure was used to obtain the dose measured by the ionization chamber. The obtained charge versus dose plot is presented in Fig. 2, where linear fits of the data are also included. These fits show the strong linear relationship between the signal produced by the photodiode and the delivered dose. The experimental uncertainties on the charge measurements are below 2% while the uncertainties on the dose measurements are below 1%. The difference in the slope between the two lines is due to energy dependence of the scintillation dosimeter. The energy dependence of plastic dosimeters for low energy photon-beams has been assigned to ionization quenching in the scintillator (Frelin et al., 2008; Williamson et al., 1999) and to the variation of the mass coefficient ratio between plastic and water (Oliveira et al., 2011) in that energy range. For applications with a fixed photon energy beam (as for most diagnostic X-ray exams) the effect on the dose measurement should be small, provided a calibration of the device is made for the beam energy. 3. Photodiode temperature dependence To use the photodiode as a photodetector in dosimetry, good temperature stability is an important requirement. The temperature dependence of the photodiode response to visible and infrared light was measured using a special setup. This setup included
F. Rêgo, L. Peralta / Radiation Measurements 47 (2012) 947e950
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a constant flux of dry air passing through a sample-box containing the photodiode and maintaining a constant temperature inside (Fig. 3). The sample-box was made of aluminum. The wall was 10 mmthick and was wrapped in a plastic insulating sheet. The box had two nozzles, one for air entry and another for air exit on the opposite wall. Inside the box the temperature was measured with a LM135 integrated temperature sensor with an accuracy of 0.5 C when calibrated at 25 C. Dry air at room temperature was introduced into a cooling-heating circuit to get it at the desired temperature inside the sample-box. In a first stage, the air passed through an ice-cooled copper pipe coil and its temperature lowered. Then, the air passed through a heating box where an electrical resistance boosted the air temperature to higher values before entering the sample-box. A LED driven by a current source produced the light signal that was conducted by an optical fiber to the photodiode inside the sample-box. In this, the LED was always kept at constant room temperature. The signal produced in the photodiode was measured by a Standard Imaging MAX 4000 electrometer. Four different LEDs were used in the measurements and the peak wavelength of each LED was: 470 nm for blue, 565 nm for green, 660 nm for red and 930 nm for infrared. The LEDs current was measured in order to assess their stability, and variations were less than 0.5%. The photodiodes output signal was measured for temperatures between 0 and 50 C. The output values were normalized to the room temperature (20 C) value for each LED (Fig. 4). For this photodiode a variation no greater than 4% in the
Fig. 3. Experimental setup to study the temperature dependence of the photodiode’s response.
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output signal is observed in the temperature range between 0 and 50 C and less than 2% in the 15e25 C range. The very low temperature dependence is a positive factor relative to the use of a PMT as photodetector. For instance, a typical PMT used for medical dosimetry applications (Williamson et al., 1999; Fluhs et al., 1996), R647 from Hamamatsu has a decrease in sensitivity of about 20% in the 0e40 C range. The PMT temperature dependence when operated in DC mode, introduces the need for larger calibration factor corrections relative to those used by photodiodes. 4. Discussion and conclusions Si-PIN photodiodes have photo-sensitivities in the range of a few hundred of A/W, which are enough for several applications in scintillator dosimetry. In this work the Hamamatsu S9195 photodiode was used with effective results. The S9195 photo-sensitivity increases from blue to infrared, enhancing the long wavelength detection. Doses in the mGy range were measured with this prototype and the measurement of lower dose values can be envisaged. In order to reduce the influence of electronic noise during the measurements the electrometer should be as close as possible to the photodiode. Also care must be advised on the placement of the photodiode box near the X-ray beam. The photodiode is sensitive to the radiation and a direct signal can be generated in the photodiode due to scattered radiation in the room walls. In the preset tests the photodiode box was shielded with lead bricks. A more manageable solution is to use a longer clear fiber placing the photodiode far away from the radiation source. The photodiode response dependence on the temperature has been studied. Under normal clinical conditions (15e25 C) a 2% variation was found within that range. This variation can be translated into a temperature coefficient of 0.2% C1 in the 15e 25 C range, compatible with the very low temperature coefficient claimed by the photodiode’s manufacturer (Hamamatsu, 2004) in the visible and near infrared region. Although yield of scintillation light has a small dependence on temperature (Beddar et al., 1992) a variation of a few percent has been reported for unloaded and doped polystyrene (Nowotny and Taubeck, 2009). Depending on which dye is used (Nowotny and Taubeck, 2009), polystyrene scintillator can exhibit 1e2% change in light yield within the 15e25 C range. In some cases variation in the scintillator light yield is opposite to the variation of the photodiode response. This leads to a compensation of the two effects with a positive outcome in the dosimeter temperature stability.
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Plastic scintillator dosimeters can play an important role in radiological applications where an image-invisible dosimeter is required, such as the dental panoramic radiography. The present work demonstrates that charges in excess of a pC can be obtained for doses in the mGy range using the S9195 photodiode coupled a PMMA optical cable and a plastic scintillator. For these tests an UNIDOS E electrometer was used. The measured current is within the input range of the available low cost monolithic electrometer operational amplifier AD549 (Analog Devices, 2008), which will be used in further developments of the project. As in earlier works (Létourneau et al., 1999; Williamson et al., 1999), a second channel will be present for the subtraction of the fluorescent light produced in the PMMA cable. Acknowledgments We want to thank to Laboratório de Instrumentação e Física Experimental de Partículas (LIP) for the financial support of this work and in particular the LIP’s mechanical workshops for the construction of the water phantom container. We thank Prof. Maria Luisa Carvalho for the use of the X-ray facility at Lisbon University. We are grateful to Ms. Ashley Merrill for corrections to the text. References Analog Devices, 2008. AD549: Ultralow Input-Bias Current Operational Amplifier. http://www.analog.com/static/imported-files/data_sheets/AD549.pdf (accessed 2012). Andersen, C.E., 2011. Fiber-coupled luminescence dosimetry in therapeutic and diagnostic radiology. AIP Conf. Proc. 1345, 100e119. http://dx.doi.org/10.1063/1. 3576161. Arnfield, M.R., Gaballa, H.E., Zwicker, R.D., Islam, Q., Schmidt-Ullrich, R., 1996. Radiation-Induced light in optical fibers and plastic scintillators: application to brachytherapy dosimetry. IEEE Trans. Nucl. Sci. 43, 2077e2084. Becks, K.-H., Drees, J., Goldmann, K., Gregor, I.M., Heintz, M., 2000. A multichannel dosimeter based on scintillating fibers for medical applications. Nucl. Instrum. Methods. Phys. Res. A. 454, 147e151. Beddar, A.S., 2007. Plastic scintillation dosimetry and its application to radiotherapy. Radiat. Meas. 41, S124eS133. Beddar, A.S., Mackie, T.R., Attix, F.H., 1992a. Cerenkov light generated in optical fibres and other light pipes irradiated by electrons. Phys. Med. Biol. 37, 925e 935. Beddar, A.S., Mackie, T.R., Attix, F.H., 1992b. Water-equivalent plastic scintillation detectors for high-energy beam dosimetry: I. Physical characteristics and theoretical consideration. Phys. Med. Biol. 37, 1883e1900. Beddar, A.S., Suchowerska, N., Law, S., 2004. Plastic scintillation dosimetry for radiation therapy: minimizing capture of Cerenkov radiation noise. Phys. Med. Biol. 49, 783e790. Clift, M.A., Sutton, R.A., Webb, D.V., 2000. Dealing with Cerenkov radiation generated in organic scintillator dosimeters by bremsstrahlung beams. Phys. Med. Biol. 45, 1165e1182.
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