SSPM based radiation sensing: Preliminary laboratory and clinical results

Radiation Measurements 46 (2011) 76e87

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SSPM based radiation sensing: Preliminary laboratory and clinical results Daniel C. Konnoff a, *, Thomas K. Plant b, Elizabeth Shiner c a

Radiation Health Physics Department, Oregon State University 116 Radiation Center, Corvallis, OR 97331, USA School of Electrical Engineering and Computer Science, Oregon State University, 1148 Kelley Engineering Center, Corvallis, OR 97331, USA c Good Samaritan Regional Medical Center: Cancer Center, 501 NW Elks Dr., Corvallis, OR 97330, USA b

a r t i c l e i n f o

a b s t r a c t

Article history: Received 29 January 2010 Accepted 3 August 2010

Recent Solid State Photomultiplier (SSPM) technology has matured, reaching a performance level that is suitable for replacement of the ubiquitous photomultiplier tube in selected applications for environmental radiation monitoring, clinical dosimetry, and medical imaging purposes. The objective of this work is low signal level laboratory and high signal level clinical testing of the Hamamatsu MPPC (S10362-11-050C), Photonique SSPM (0810G1), and Voxtel SiPM (SQBF-EKAA/SQBF-EIOA) SSPMs coupled to different inorganic scintillator crystals (Prelude 420, BGO), inorganic doped glass scintillator material SiO2:Cu2þ and organic BCF-12 plastic scintillating fibers, used as detector elements. Plastic Optical Fibers (POFs) and Glass Optical Fibers (GOFs) are used as signal conduits for laboratory and clinical testing. Further, reduction of electron-beam-generated Cerenkov light in optical fibers is facilitated by the inclusion of metalized air-core capillary tubing between the BCF-12 plastic scintillating fiber and the POF. In a clinical setting dose linearity, percent depth dose, and angular measurements for 6 MV/18 MV photon beams and 9 MeV electron beams are compared with and without the use of the air-core capillary tubing for BCF-12 plastic scintillating fiber. These same measurements are repeated for SiO2:Cu2þ scintillator material without air-core capillary tubing. Ó 2010 Elsevier Ltd. All rights reserved.

Keywords: Solid state photomultiplier Medical imaging Clinical radiotherapy Diagnostic dosimetry Environmental radiation sensing

1. Introduction Optical fiber (OF) dosimetry using plastic (POF) and glass (GOF) fibers coupled to a variety of scintillators has been reported for clinical (Akselrod et al., 2007; Anderson et al., 2009; Beddar, 1994; Beddar et al., 2001), diagnostic radiography (Hyer et al., 2009), and environmental radiation sensing purposes (Klein et al., 2005). Copper-doped glass, SiO2:Cu2þ, may be used as an Optically Stimulated Luminescence (OSL) material or as a real-time Radio Luminescence (RL) scintillator (Huston et al., 2001; Justus et al., 2004), while BCF-12 plastic scintillator may be used in the RL mode (photomultiplier tube as photosensor) (Beierholm et al., 2008). Water equivalence and energy independence of plastic scintillators as a detector material is well documented, and considerations of detector size and S/N ratio have been analyzed (Beddar et al., 2005; Clift et al., 2000; Archambault et al., 2005). BCF-12 plastic scintillator represents a good choice for a dosimeter, as variants of this material have produced good results in clinical studies (Bartesaghi et al., 2007). Recent work by groups in Japan and Russia have advanced Solid State Photomultiplier (SSPMs, SiPMs, MPPCs) technology (Golovin * Corresponding author. Tel.: þ1 541 737 2984; fax: þ1 541 737 1300. E-mail address: [email protected] (D.C. Konnoff). 1350-4487/$ e see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.radmeas.2010.08.002

and Saveliev, 2004; Gomi et al., 2007; Heckathorne et al., 2006), making it a viable alternative to photomultiplier tubes (PMTs) for laboratory, environmental, and clinical dosimetry; medical imaging applications (PET, CT, SPECT); dosimetry in diagnostic radiology; and in-vivo applications (catheters and brachytherapy (Anderson et al., 2009; Suchowerska et al., 2007)). Their small size, high gain, low bias voltage, and non-magnetic characteristics are distinct advantages when compared with the PMT (Saveliev and Golovin, 2000). Previous clinical studies using optical fibers and different inorganic/organic scintillators have used standard photodiodes as an optical photon detector. The large signals generated by Linac (linear accelerator) beams coupled with external amplification proved sufficient for analyzable signals (Beddar, 1994; Lee et al., 2006). However, the added advantages of built-in gain together with low noise make an SSPM a more suitable photon detector than standard photodiodes in a wider variety of high- or low-energy applications. Overall SSPM efficiency, sometimes referred to as Photon Detection Efficiency (PDE), given by

PDE ¼ QEðlÞ þ 3Geiger þ 3Geometry

(1)

Here QE(l) is the SSPM quantum efficiency, 3Geiger is the probability of an avalanche breakdown, and 3Geometry is the SSPM

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photosensitive area packing factor (Buzhan et al., 2006; Gomi et al., 2007). SSPM pixel capacitance (Cpixel) at the breakdown voltage Vbreakdown is used to calculate gain using (Buzhan et al., 2006)

or, more recently, the use of metalized air-core capillary tubing (Matsuura and Miyagi, 2004). In this work the air-core capillary tube was chosen to reduce the Cerenkov light generated within OFs (Lambert et al., 2008).

Gain ¼ Vov Cpixel =e ¼ Qpixel =e

2. Materials and methods

(2)

where gain is proportional to the overvoltage, Vov ¼ Vbias Vbreakdown above the SSPM breakdown voltage. The total amount of charge produced per incident photon is

Qtotal

¼ 3Numphotons Qpixel ½C

(3)

here 3 ¼ 3Geometry QEðlÞ and Numphotons is the number of incident photons (Pavlov et al., 2005; Renker and Lorenz, 2009). SSPM die capacitance is used to estimate Dark Count Rate (DCR) using (Pavlov et al., 2005)



  DCR Vop dark ¼ I Vop dark =Vop Cdie ½Hz

(4)

where I(Vop) is the SSPM current measured in the dark at the manufacturer recommended operating voltage Vop, and Cdie is the total measured device capacitance (Cpixel  number of pixels) (Renker and Lorenz, 2009). For our laboratory and clinical measurements PDE at the peak scintillator wavelength was the best indicator of overall SSPM system performance, followed by SSPM-scintillator/SSPM-Optical Fiber coupling considerations, SSPM DCR, then by SSPM gain. For low-level environmental radiation sensing or when used together with OFs, SSPM DCR must remain small to detect signals attenuated by long POFs (10s of meters), where a high system S/N ratio is necessary for detection of low-level scintillator signals (Pavlov et al., 2005; Beddar, 2007). In clinical applications the generation of Cerenkov light within OFs when using photon/electron beams (Archambault et al., 2006; Clift et al., 2002) for radiotherapy degrades the S/N ratio significantly, introducing additional noise into the system. (Beddar et al., 1992; Elsey et al., 2007) Cerenkov generated light varies with the speed of the particle in the medium and the index of refraction of the material, and is a function of impinging electron beam angle on the OF. (Law et al., 2007) However, it is an important noise source only for electron beams above 200 keV, and can be neglected for diagnostic dosimetry purposes. The incident electron beam angle at which this light is predicted to be a maximum (for relativistic particles) is given by

qCerenkovMax ¼ cos1

1 ncore

(5)

where ncore is the refractive index of the fiber core. (Jelly, 1958) The approximate theoretical intensity curve for the Cerenkov radiation in OF is given by

8 >

9 >

= pncore ; : sin g n2core  1 >

2.1. SSPMs Custom electronic charge and transimpedance (MAR8ASM þ RF) amplifiers were constructed for the three SSPMs tested: MPPC S10362-11-050C, Photonique 0810G1, and Voxtel SQBF-EKAA/SQBF-EIOA. (Hamamatsu, 2008a; Photonique, 2008; Voxtel, 2008) Two versions of Voxtel SSPMs were evaluated, one a room temperature device (SQBF-EIOA) in a T0-8 package and another using the same semiconductor die mounted to a 3-stage thermoelectric cooler in a modified T0-8 package (SQBF-EKAA) to minimize DCR. (Dhulla et al., 2007) Dedicated circuitry (Maxim MAX1978) was used for the temperature controller of the SQBFEKAA for laboratory tests. The amplified SSPM pulse or charge signal was coupled to a custom-designed discriminator circuit feeding an ORTEC 996 counter/timer or host computer running analysis software, as shown in Fig. 1. The photocurrent for SSPM tests was generated using blue and red light from high intensity LEDs using a custom-designed pulser circuit with extensive modifications based upon the Kapustinsky design (Kapustinsky et al., 1985). A reference photomultiplier tube (Hamamatsu PMT H5783P) was used as a benchmark photon detector coupled to the same photon counting electronics used with the SSPMs. A Hamamatsu S1223 Si PIN photodiode was used as a reference photodetector for calibration purposes due to its known responsivity. A Tiffen P Series ND.6 Grad SE neutral density (ND) filter mounted to an optical xyz stage was used to attenuate the photon flux from the LEDs and LEDs coupled to test fibers from the pulser. The LED pulser/ND filter combination allows attenuation of light photons in a controlled manner. A Newport digital power meter Model 815 was used to measure the optical power. Special light-tight mating jigs were used to hold in place each scintillation crystal or POF to each SSPM device or the PMT for laboratory testing purposes. Signal acquisition was also done using LabVIEWÓ software controlling Tektronix oscilloscopes. CurrenteVoltage (IeV) curves for each SSPM were measured with an Agilent 4156 semiconductor parameter analyzer, and pixel capacitance was measured with an HP 4263B LCR meter e both measurements with the DUT SSPM in total darkness. Thirty engineering samples of the SQBF-EIOA were tested for dark current using the 4156 semiconductor parameter analyzer. The lowest dark current sample was chosen for further testing. The operating temperature for all SSPM tests was ambient, 25  C, with the exception of the SQBF-EKAA, which was tested at an operating temperature of 20  C.

(6)

where r is the OF core radius, Dn is the difference between the OF core and cladding indices of refraction, and g is the angle between the electron beam and the OF axis in the direction of the detector (Law et al., 2007). Various Cerenkov light reduction schemes proposed by different groups typically use one of several methodologies: a background fiber in parallel with the signal fiber (Lee et al., 2007a), optical filters to attenuate the Cerenkov light which is centered at a different frequency than the signal of interest (spectral discrimination) (Frelin et al., 2006), temporal methods (Justus et al., 2006)

2.2. Scintillators Two different inorganic scintillation materials and one organic material were used during SSPM laboratory testing: 1 mm  5 mm square cylindrical Prelude 420 (Lu1.8Y.2SiO5:Ce), 8 mm  10 mm cylindrical BGO (Bi4Ge3O12) crystals, and 1 mm  3 mm cylindrical lengths of BCF-12 plastic scintillating fiber (Saint-Gobain Crystals, Hiram OH.). Scintillators were coupled to each SSPM active area using optical grease (Saint-Gobain Crystals, Hiram OH.) (Globus et al., 2002). BCF-12 scintillators were polished (both ends), painted with BaSO4 reflective paint except for one polished end, and finally covered with black epoxy before coupling to each

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Fig. 1. Laboratory and clinical measurement system diagram.

SSPM active area using optical grease. The count rate from each SSPM-scintillator combination was measured when excited by low activity radiation sources: (1.842 mCi 60Co, 0.661 mGy/s and 1 mCi 137 Cs, 0.201 mGy/s) were used to stimulate the BGO and BCF-12 scintillators in the laboratory along with a 1 mCi 137Cs check source for detector efficiency. The Prelude 420 crystals are self-poisoned with radioactive 176Lu, providing scintillations that can be detected without using the low activity external radiation sources. The photon count rate (cpm) with 0.5 p.e. (photo electron equivalent) threshold was measured in the laboratory (5 samples each; averaged over 1 min count times) for each SSPM with Prelude 420, BGO, and BCF-12 scintillators directly attached to the transparent window surface of the package. These same scintillators were then attached to 2.5 m lengths of Mitsubishi Eska GH4001 POF, coupled to each SSPM, and the photon counting measurements repeated. 2.3. Linac testing Clinical measurements with the GOF-coupled SiO2:Cu2þ and POF-coupled BCF-12 scintillators were done using the MPPC S10362-11-050C SSPM together with a USB-based photon counter and PC software supplied by Hamamatsu using 3.5 p.e. threshold. The MPPC device was used clinically because it held a slight measurable edge in gain, PDE, and DCR after laboratory tests. A Varian Clinac 2100CD Linear Accelerator was used for percent depth dose, dose linearity, and angular dependence measurements. Electronics were in the Clinac vault during tests. All clinical measurements were verified with a benchmark calibration using a Wellhofer CC13 ion chamber with 0.13 cm3 detector volume (hereafter the expected value). Unless otherwise noted, 10 cm  10 cm radiation fields (photon, electron beam) were used during measurements. The central beam axis was at 90 to the POF axis with a source-to-surface distance (SSD) of 1.0 m. A water phantom/tank from Med-Tec, Inc., model MT-DDA with a hand-cranked, gear-driven depth adjustment (in increments of 0.1 mm) was used. Positional accuracy error for percent depth dose was 2 mm from the effective point of measurement in the water bath. Sensitive ends of the dosimeter cables were completely submerged in the water bath during tests. Dose linearity measurements were done on top of an acrylic 20 cm  30 cm  5 cm

phantom, repeated for both types of BCF-12 dosimeter, and the SiO2:Cu2þ dosimeter. Dose linearity measures the ability of the OF probes to accurately determine dose between scaled monitor units (MUs) or activity from ionizing radiation sources. The signal ratio to MU from the linear accelerator should be a constant value for each probe tested, but will not be the same value between probes using different scintillators. For example, ideally each raw signal value should be 2 the previous value 3% (20 MU ¼ 2  10 MU) but also (4  5 MU), etc. The variation between the signal ratios within the same data set was analyzed using the slope of the linear best-fit line. The slope of the best-fit line is the average of the signal/MU ratio. The y-intercept is not zero because the accelerator is not linear when delivering dose at the low end of the range. Accelerators typically under-deliver in the low MU (<4 MU) ranges which is why a lower limit of 5 MUs per segment is used for IntensityModulated Radiation Therapy (IMRT) radiation fields. A typical clinical specification requires that the linearity results are between 3%. Percent depth dose, equation (7), measures the dose deposited at a particular depth in water, normalized to the depth of maximum dose.

PDD ¼

Dd  100 Dd max

(7)

The depth dose distribution is a function of the energy and the type of radiation used in a treatment setting. Each of the depth dose figures shows the results for the OF dosimeters measured in a water tank, using different electron and photon energies, and comparing against the expected value from the ion chamber. Each depth dose data set was normalized to its own dmax depth: 1.5 cm for 6 MV photons, 3.0 cm for 18 MV photons, and 2.0 cm for 9 MeV electrons. 2.4. SiO2:Cu2þ dosimeter design A 400 mm  5 mm cylindrical length of SiO2:Cu2þ scintillator, supplied by the Naval Research Laboratory, Washington DC., was butt-coupled (Norland UV cement, Polyacrylate glue) to a 2.5 m length of GOF using an FC fiber connector to mate with the photon counting electronics. The scintillator head and fiber junction were

D.C. Konnoff et al. / Radiation Measurements 46 (2011) 76e87

covered with black epoxy to minimize scattered light and then coated with Dupont TefzelÔ to shield the scintillator from light. Due to the stimulation energy needed to generate an RL signal using this material, it was not laboratory tested with the low activity gamma radiation sources, but tested clinically using the linear accelerator as noted above. Fig. 2 shows the construction of the SiO2:Cu2þ dosimeter cable. 2.5. BCF-12 dosimeter design A 1 mm  3 mm cylindrical length of BCF-12 scintillator was butt-coupled (Norland UV cement, Polyacrylate glue) to a 2.5 m length of Mitsubishi Eska Premier GH4001 jacketed POF using an FC fiber connector to mate with the photon counting electronics. BCF-98 clear light guide was also evaluated but rejected due to permanent radiation damage and RL considerations within the fiber itself (Wick and Zoufal, 2001; Nowotny, 2007). Each scintillator was covered with BaSO4 reflective paint, then further covered with Dupont TefzelÔ to shield light. POF and BCF-12 scintillator ends were prepared following the general procedures outlined in (Lee et al., 2004; Ayotte et al., 2006). Fig. 2 shows the construction of the BCF-12 dosimeter cable and Fig. 3 shows the completed BCF-12 dosimeter cable with FC connector and dosimeter end. Laboratory testing of count rates (cpm) were measured for each of the MPPC, Photonique, and Voxtel SSPMs with the BCF-12 scintillator stimulated using 1.842 mCi 60Co and 1 mCi 137Cs gamma radioactive sources. Clinical testing was done using the linear accelerator as noted above. 2.6. Angular measurements To minimize the effect of Cerenkov radiation generated in POFs when using electron beams during clinical radiotherapy, a 20 cm length of silver-coated air-core capillary tubing (PolyMicro Technologies LLC, Scottsdale AZ.) was inserted between the BCF-12 scintillator and Mitsubishi Eska GH4001 POF (Norland UV cement, Polyacrylate glue), as shown in Fig. 4 (Lambert et al., 2008).

Fig. 3. Eska GH4001 POF and BCF-12 scintillator dosimeter cable.

Fig. 2. BCF-12 and SiO2:Cu2þ dosimeter cables.

79

80

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Fig. 4. BCF-12 Air-core capillary tube dosimeter.

Clinical measurement of signal magnitude was recorded as a function of electron beam angle for BCF-12 and BCF-12/capillary tube dosimeters. The measurement setup shown in Fig. 5 was used. Two 25 cm  25 cm styrofoam blocks 30 cm apart were used to position the dosimeter at the isocenter of the electron beam. The dosimeter was fixed in position between the blocks using Micropore tape and small cylindrical paper tubes to keep scattered radiation to a minimum. A constant 4 cm  20 cm field size was used throughout the angular dependence measurements. Five readings each at dose rate 400 MU min1 with 160 MU total dose delivered were averaged for each type dosimeter measurement. The dosimeter was irradiated in air with a 9 MeV electron beam using angles from 0 to 100 in 10 steps. 3. SSPM results and discussion Fig. 6 shows the measured IeV characteristics for each SSPM, and Fig. 7 shows the measured pixel capacitance up to the breakdown voltage for each SSPM, both are needed to estimate device performance. Table 1 summarizes the number of SSPM pixels (Npixel), measured dark current, measured pixel capacitance (Cdie/ Npixel), calculated gain (using equation (2)), calculated DCR (using

Fig. 5. Photo of angular dependence measurement setup.

equation (4)), and s$DCR for each SSPM. s$DCR shows the expected number of dark counts added to a detected signal when using a s ¼ 1 ms sampling time. To decrease the DCR contribution to any signal, the sampling time s must be shorter or the device DCR must be smaller. When using fast scintillators such as BCF-12 or Prelude 420 (decay times 3.2 ns and 41 ns, respectively), s can be shortened (Pavlov et al., 2005). Mainly this can be done by lowering I(Vop)dark holding the other parameters constant. For the SSPMs tested in this study, I(Vop)dark was indeed decreased with each new generation of device. Vop, Cpixel, and Npixel often change with each device improvement, topology, and application (PET, SPECT, environmental monitoring, etc.) (McNally and Golovin, 2009). Table 2 shows the measured photon count rate (cpm) for the reference PMT (H5783) and each SSPM type with Prelude 420, BGO, and BCF-12 scintillators directly attached and coupled using 2.5 m Mitsubishi Eska GH4001 POF. Laboratory measurements confirm that all three room temperature SSPM types may be used with inorganic scintillators coupled directly to their photosensitive die areas for photon counting, gamma spectroscopy functions, and remote environmental radiation monitoring. The use of long length OFs coupled to scintillators for remote monitoring purposes depends upon several additional parameters, among them: coupling efficiency between scintillator and OF, signal attenuation vs. length in the OF itself, coupling

Fig. 6. IV-characteristics for tested SSPMs.

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81

Table 2 Laboratory photon count rates (cpm) 60Co: direct scintillator-SSPM attachment and using POF. Photon Detector

BGO

P450

BCF12

BGO-POF

P450-POF

BCF12-POF

H5783PMT-ref PH0810G1 S10362-11-050C SQBF-EIOA SQBF-EKAA

5713 927 1082 358

3672 3109 3281 2613 138

2378 1127 1682 173

732 92 182

1466 385 526

1039 302 413

a

a

a

a

a

a

a

a

SSPM Vop: 30.5 V, 70 V, 39.5 V, 40.5 V respectively. a Undetectable.

Fig. 7. Die capacitance for tested SSPMs.

efficiency between OF and SSPM, and noise considerations (manmade or natural). When testing SSPM/scintillator combinations in the laboratory using low activity gamma radiation sources, FC fiber connectors were observed to cause excessive signal loss. It was necessary to place the fiber core in direct contact with the SSPM photosensitive area, eliminating any air gap to maximize the acquired signal. If used as a low-level environmental radiation sensor the additional signal loss imposed by standard FC connectors is high. The Voxtel (SQBF-EIOA) and Photonique (0810G1) devices use clear epoxy windows as a protective material covering their photosensitive areas. This material is hard, yet scratchable when butting sharp crystal edges or OF cores against its surface; however, shallow scratches may be polished out. The MPPC (S10362-11-050C) uses a clear gelatin-like polymer as a protective material covering its photosensitive area which easily scratches whereby sharp crystal edges or OF cores can penetrate to the photosensitive area, potentially damaging the die or breaking the bonding wires. When using the S10362-11-050C it is necessary to design a coupling solution that prevents sharp edges from crystals or OF cores from actually touching the material e adding to system cost. SSPMescintillator coupling was an issue for the cooled TO-8 packaged Voxtel (SQBF-EKAA) with its transparent glass window, where the photosensitive die is set back approximately 2.7 mm from the coupling surface. A custom lens design (GRIN or convex system) is needed to take full advantage of this cooled SSPM. Indeed even when using a GRIN lens for 1 mm POFeSSPM coupling, it was not possible to transmit usable amounts of light to the photosensitive die for POF testing purposes during laboratory tests.

When using the uncooled TO-8 packaged Voxtel (SQBF-EIOA) with 1 mm core POF centered directly over the photosensitive die (butted against the protective epoxy), the lower PDE of this device when compared with the MPPC (S10362-11-050C) and Photonique (0810G1) during laboratory tests (as shown by lower photon counts), rendered it unusable for our (G,P)OFeSSPM clinical system. However the uncooled Voxtel SSPM (SQBF-EIOA) could be used as part of a remote radiation monitoring system with scintillators directly attached to the photosensitive die. A simple package redesign, which places the cooled (SQBFEKAA) photosensitive die nearer (<0.05 mm) to the OF core or scintillator crystal surface, while eliminating the need for custom lenses, takes advantage of the superior dark noise performance of the SQBF-EKAA making it an ideal candidate for low-noise (G,P) OFeSSPM coupled systems or direct scintillator-coupled environmental monitoring systems. Semiconductor device theory predicts (Sze, 2006) that thermally generated free carriers contributing to SSPM dark count noise are reduced by approximately a factor of 2 for every 8  C drop in temperature. (Renker and Lorenz, 2009) A similar dependency was noted for the SQBF-EKAA (see I/V curve Fig. 6). This metric alone and others that depend upon temperature (I/V, gain) are not sufficient by themselves to choose one SSPM over another for a given fiber-based application as noted above. The MPPC (S1036211-050C) is a lower noise, higher gain device at small overvoltage (Vov) with better PDE at the scintillator emission frequencies tested here, than the Photonique (0810G1) and the Voxtel SQBF-EIOA devices (Fig. 6, Table 1) at room temperature. However, the Photonique (0810G1) has greater dynamic latitude for overvoltage (Vov) than the MPPC (S10362-11-050C), which can been seen from the slope of their respective I/V curves in Fig. 6. This characteristic is important, for example, when trading off noise against gain (bright scintillator light source, long fiber length, remote monitoring of isotopes) or signal acquisition without a preamplifier where high gain is a requirement. Test results indicated these two devices are closely matched. The room temperature Voxtel SQBF-EIOA, despite a large dynamic overvoltage (Vov) range, exhibited lower gain with higher noise than the MPPC and Photonique counterparts (Table 1). Optical signal coupling with 1 mm POF/scintillator combinations was poor for this device. Based upon results: (including I/V and

Table 1 Summary of characteristics for Photonique, Hamamatsu, and Voxtel SSPMs. SSPM PH0810G1 S10362-11-050C

Npixel 556 400

Pixel Cap. (fF)

Gain (105)

DCR

2.5  10

7

80.0

10.0

1.0  10

7

92.75

8.69

12.3

2.30

2585.4 Hz/pixel 1.437 MHza 1796.9 Hz/pixel 718.77 kHza 3969.8 Hz/pixel 4.065 MHza 392.76 Hz/pixel 402.188 kHza

I(Vov)dark(A)

SQBF-EIOA

1024

1.5  107

SQBF-EKAA

1024

2.5  108

a

At Vov: 2 V, 1.5 V, 3 V, 7 V respectively.

8.88

3.88

s$DCR 1.44 0.71 4.06 z0.4

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DCR, gain at Vov), PDE at scintillator wavelength, and system analysis considerations (fiber coupling issues, and availability of portable electronics), the MPPC (S10362-11-050C) was chosen for the initial clinical portion of this work.

Table 3 Photon and electron dose linearity differences (range extrema) from reference ion chamber (in %). Dosimeter SiO2:Cu2þ

4. Clinical results and discussion 4.1. Angular dependence Results for the 9 MeV electron beam angular dependence tests are shown in Fig. 8. Using equation (5) the Cerenkov generated maximum peak at approximately 47.91 is clearly visible for the standard BCF-12 dosimeter, as is a 40% reduction when using the BCF-12 capillary tube POF dosimeter. The theoretical Cerenkov intensity curve is also plotted for comparison using equation (6). The spectrum of Cerenkov light generated is independent of the angle between the beam and the fiber axis (Lambert et al., 2009). Due to a scarcity of material to meet the clinical testing schedule, the length of air-core capillary tube used for this measurement was not of sufficient length to completely avoid Cerenkov effects (as described in Lambert et al., 2008) from the beam. We believe this greater variation in result is due to three factors: the shorter length of air-core capillary tubing used during the angular measurements e as the 2100CD Clinac rotates to extreme angles more of the POF is in the beam, coupling issues between the POF-air-core capillary tubing-scintillator junctions, and manufacturing differences between the two dosimeter cables. There remains a measurable Cerenkov effect when irradiating the BCF-12 capillary tube POF dosimeter.

6 MV

5.5 þ0.5 BCF-12 þ9.0 þ0.6 BCF-12 21.2 Capillary Tube 0.8 a

18 MV 6 MeV 9 MeV 12 MeV 16 MeV 20 MeV a

a

3.0 þ6.4 19.2 þ22.4

0.8 þ7.1 17.8 þ1.6

4.4 þ0.2 16.0 þ8.2 15.7 þ7.1

a

a

a

6.5 þ0.8 20.3 þ14.0

5.1 þ9.4 15.4 þ8.0

6.2 þ5.0 18.2 þ6.6

Not tested.

Fig. 9(b) shows the measured photon dose linearity for the capillary tube POF dosimeter: BCF-12 scintillator. When using SSPM-OFs as photon detectors, variation in dose linearity for the SiO2:Cu2þ (6 MV only) and standard BCF-12 dosimeters (6 MV, 18 MV) confirms previous studies on the accuracy of OF dosimetry (Beddar, 1994; Beddar et al., 2001; Huston et al., 2001). Here the variation in dose linearity ranged from 5.5, þ0.5% for SiO2:Cu2þ dosimeters, and ranged from þ9.0, þ0.6%, 3.0, þ6.4% for BCF-12 dosimeters, respectively, when compared against the expected values from the ion chamber. For the BCF-12 air-core capillary tube dosimeter (6 MV, 18 MV),

4.2. Dose linearity The linear dose responses of the OF dosimeters in the RL scintillating mode are shown in the following figures. Five readings each was integrated and averaged, then compared against the expected value from the ion chamber over a dose range of 5 MUe160 MU at 400 MU min1 dose rate. The dose ranges discussed below and shown in Table 3 are range extremas. Fig. 9(a) shows the measured photon dose linearity for the standard POF dosimeter and GOF dosimeter: BCF-12, and SiO2:Cu2þ scintillators.

Fig. 8. Angular dependence of standard and capillary tube POF dosimeters, BCF-12 scintillator: measured and theoretical results.

Fig. 9. Photon dose linearity: standard and capillary tube dosimeters. (a) BCF-12 and SiO2: Cu2þ Dosimeters. (b) BCF-12 Capillary Tube Dosimeter.

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the variation in dose linearity ranged from 21.2, 0.8% and 19.2, þ22.4% when compared against the expected value. Fig. 10(a) shows the measured electron dose linearity (9 MeVe20 MeV) for the standard POF dosimeter, GOF dosimeter: BCF-12, and SiO2:Cu2þ scintillators. Fig. 10(b) shows the measured electron dose linearity for the capillary tube POF dosimeter (6 MeV, 9 MeV): BCF-12 scintillator. Fig. 10(c) shows the measured electron dose linearity for the capillary tube POF dosimeter (12 MeV, 16 MeV, 20 MeV): BCF-12 scintillator. When using SSPM-OFs as electron detectors variation in dose linearity for the SiO2:Cu2þ dosimeter (9 MeV only) ranged from 4.4, þ0.2%, and for the standard BCF-12 dosimeter (6 MeV, 9 MeV, 12 MeV, 16 MeV, 20 MeV) ranged from 0.8, þ7.1%, 16.0, þ8.2%, 6.5, þ0.8%, 5.1, þ9.4%, and 6.2, þ5.0% respectively, when compared against the expected value. For the BCF-12 air-core capillary tube dosimeter, the variation in dose linearity ranged from 17.8, þ1.6%, 15.7, þ7.1%, 20.3, þ14.0%, 15.4, þ8.0%, and 18.2, þ6.6% when compared against the expected value. The figures show a fairly linear response over the dose range of interest for radiation oncology treatments. Overall the figures show

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the agreement is good for the SiO2:Cu2þ and fair for standard BCF-12 dosimeters. However, in a clinical setting no more than 3%e 5% error is permitted. The error is too large for actual clinical measurements using the currently manufactured OF dosimeters. Table 3 compares the measured photon and electron dose linearity’s for the GOF dosimeter, standard POF dosimeter, and capillary tube dosimeters: BCF-12, and SiO2:Cu2þ scintillators. 4.3. Depth dose measurements: photon and electron beam Depth dose ranges discussed below represent range extremas. Fig. 11 shows the result of the 6 MV photon depth dose measurements for the standard POF dosimeter, using both BCF-12 and SiO2:Cu2þ scintillators as well as the capillary tube dosimeter using a BCF-12 scintillator. Fig. 12 shows the result of the 18 MV photon depth dose measurements for the standard and capillary tube dosimeters: BCF12 and SiO2:Cu2þ scintillators. When using SSPM-OFs as a photon detector for depth dose (6 MV, 18 MV), variation in accuracy for SiO2:Cu2þ, BCF-12, and BCF12 capillary tube dosimeters ranged from: 7.4,1.2%, 9.1,1.3%,

Fig. 10. Electron dose linearity: standard and capillary tube dosimeters. (a) BCF-12 and SiO2: Cu2þ Dosimeters. (b) BCF-12 Capillary Tube Dosimeter: 6MeV, 9 MeV. (c) BCF-12 Capillary Tube Dosimeter: 12 MeV, 16 MeV and 20 MeV.

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Fig. 11. 10 cm  10 cm 6 MV photon depth dose in water tank.

11.0,1.4% and 12.9,þ1.8%, 1.8,þ7.4%, 2.7,þ15.3% respectively, when compared against the expected values. Fig. 13 shows the result of the 9 MeV electron depth dose measurements for the standard BCF-12 POF dosimeter, the GOF SiO2:Cu2þ dosimeter and the BCF-12 capillary tube dosimeter. When using SSPM-OFs as an electron detector for depth dose (9 MeV), variation in accuracy for SiO2:Cu2þ, BCF-12, and BCF-12 capillary tube dosimeters ranged from: 7.3,þ4.9%, 11.3,þ6.4%, and 21.2,þ3.9% respectively, when compared against the expected values. Overall the figures show the agreement is good for the SiO2:Cu2þ and fair for the standard BCF-12 dosimeters; however, as noted previously, in a clinical setting no more than 3%e5% error is

permitted. The error is too large for actual clinical measurements using the current manufactured OF dosimeters. Table 4 compares the measured photon and electron depth doses for the GOF dosimeter, standard POF dosimeter, and capillary tube dosimeters: BCF-12, and SiO2:Cu2þ scintillators. 5. Uncertainty analysis SiO2:Cu2þ and BCF-12 material dose linearities have been well characterized (Huston et al., 2001; Justus et al., 2004, 2006; Beierholm et al., 2008) while a capillary tube dosimeter using a PMT optical photon sensor showed much better linearity than measured here in a previous study using a 60 cm capillary tube

Fig. 12. 10 cm  10 cm 18 MV photon depth dose in water tank.

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Fig. 13. 10 cm  10 cm 9 MeV electron depth dose in water tank.

(Lambert et al., 2008). In a POF-coupled radiation sensing system, sources of signal degradation are numerous: three optical junctions in a standard cable (scintillator-POF, POF-FC, FC-SSPM), and four optical junctions in a capillary tube cable (scintillator-capillary tube, capillary tube POF, POF-FC, FC-SSPM). Each of these junctions has associated with it a signal coupling attenuation factor which, when multiplied together, greatly reduce the number of optical photons reaching the SSPM (Beddar, 2007). In this work scintillators of varying diameters (1 mm2 square, 1 mm2 circular, and 4 mm2 circular) were used to gain insight into coupling with 1 mm2 die area SSPMs and 1 mm diameter POFs for remote monitoring applications. Photon counting results indicate that significant light is lost when using a larger diameter scintillator without an optical concentrator system. Using a smaller scintillator (smaller than die area) introduces excess noise into the system due to DCR considerations. ESKA GH4001 plastic fiber cable attenuates light at 0.19 dB/m (650 nm) and 0.22 dB/m (400 nm). For example, for 15 m of fiber at 650 nm, this is 2.85 db with a signal attenuation factor of 0.518 (2.85 db ¼ 10 logT). For remote radiation monitoring applications even this short POF length degrades the signal substantially (fewer optical photons reaching the SSPM), underscoring the need for a low DCR SSPM with high gain and high PDE at scintillator emission wavelengths. Similarly, in clinical applications it is desirable to have electronics physically separated from noise sources (removed from Linac vaults, x-ray generator rooms, etc.). At long POF cable lengths needed to achieve this separation, signal attenuation in non-coupled fiber systems is significant, multiple junctions further Table 4 Photon and electron percent depth dose differences (range extrema) from reference ion chamber (in %). Dosimeter

6 MV

18 MV

9 MeV

SiO2:Cu2þ

7.4 1.2 9.1 1.3 11.0 1.4

12.9 þ1.8 1.8 þ7.4 2.7 þ15.3

7.3 þ4.9 11.3 þ6.4 21.2 þ3.9

BCF-12 BCF-12 Capillary tube

degrade the S/N ratio. Multiple junction POF systems necessitate that precision manufacturing techniques be used on each component to avoid optical photon attenuation and loss. Current state of the art SSPMs have greater DCRs than PMTs (200 nA vs. 2 nA). (Hamamatsu, 2008b,a) Replacing PMTs in radiation monitoring applications using POFs as signal conduits requires that system S/N ratio budget be carefully considered. For direct-coupled, low-level radiation detection systems we recommend that: the size of SSPM die and scintillator diameter be closely matched or an optical concentrator system be used, that the scintillator end surfaces must be polished to optical industry standards, and that an index matching fluid be used at the interface. Finally, the PDE of the SSPM must be high near the emission wavelength of the scintillator. Once POF is introduced into the system the importance of these constraints is magnified as noted above. During laboratory tests length extensions of POF beyond 2.5 m with low PDE SSPMs resulted in undetectable signals when using low activity radiation sources. 6. Summary and conclusions Use of SSPMs as photon detectors for clinical and environmental dosimetry has many advantages over a PMT: rugged solid-state construction, small size, high gain, low bias voltage, magnetically insensitive, high quantum efficiency, and (potentially) lower cost. This study has shown that, overall, the performance of the SSPM-OF dosimeters tested indicates that SSPMs can be used to replace PMTs for clinical and environmental dosimetry applications. The introduction of air-core capillary tubing in the optical signal path essentially functions as an attenuator and Cerenkov noise reducer. Under clinical conditions the SSPM-air-core capillary tube-OF dosimeter provided a usable signal for analysis. Angular measurements confirm the value of the air-core capillary tube for reducing Cerenkov radiation generated within POFs. Additional lengths of this material (>30 cm) can eliminate measurable Cerenkov noise generated in the POF, whilst providing a simpler noise reduction methodology when compared with other noise reduction schemes discussed previously (Lee et al., 2007b; Frelin et al., 2006).

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The SiO2:Cu2þ dosimeter had the most accurate response in this study, when compared against the expected value with the exception of its 18 MV photon response. Both BCF-12 dosimeter depth dose measurements were consistently below the reference ion chamber measurements, indicating a S/N ratio, coupling, or scintillator efficiency issue when compared to the SiO2:Cu2þ dosimeter. The SiO2:Cu2þ dosimeter, standard BCF-12 dosimeter, and BCF-12 capillary tube dosimeter can measure depth dose profiles for photon and electron beams. The standard SSPM-OF dosimeters exhibit linear response characteristics to dose and energy independence when using both plastic and SiO2:Cu2þ materials. The SSPM-air-core capillary tube-OF dosimeter, while not as accurate, can be improved with better manufacturing processes to nearly match the performance of the standard SSPM-OF dosimeters. The S/N ratio of the SSPM-OF system is adequate for both clinical and environmental dosimetry. Though no long-term environmental or clinical study of the SSPM-OF system has yet been performed, the first results presented here are encouraging. For future prototypes the air-core capillary tube-dosimetereOF coupling must be optimized by cutting/polishing the OFs, scintillators, and capillary tubes with tighter manufacturing tolerances to improve performance and reduce signal variability. Moreover, computer simulation of the optical photon signal path will assist in identifying coupling problems and in improving the S/N ratio. Further laboratory and clinical studies with new OFeSSPM prototypes, using the Photonique (0810G1) and Voxtel (SBQF-EKAA) with modified package (outside the LINAC vault, with and without a preamplifier), will include long-term measurements, low-energy dose rate studies, ion chamber OF dosimeter calibration measurements, and improved software to operate the system. Acknowledgments The author would like to thank Mr. Roshan Patel from Hamamatsu Corp., San Jose, CA. for MPPC samples, Dr. David McNally from Photonique SA, Geneva, Switzerland for SSPM samples, and Dr. Vinit Dhulla from Voxtel Corp., Beaverton, OR. for SSPM samples. In addition the author thanks Mr. Don Doize from PolyMicro Technologies LLC for samples of their air-core coated capillary tubing, Dr. Alan Huston from the U.S. Naval Research Laboratory for samples of SiO2:Cu2þ scintillating fiber material, and Mr. Michael Mayhugh from Saint-Gobain Crystals for samples of their inorganic scintillator material. References Akselrod, M., Botter-Jenson, L., McKeever, S., 2007. Optically stimulated luminescence and its use in medical dosimetry. Radiat. Meas. 41, S78eS99. Anderson, C., N, S.K., Greilich, S., Helt-Hansen, J., Lindegaard, J., Tanderup, K., 2009. Characterization of a fiber-coupled Al2O3:C luminescence dosimetry system for online in vivo dose verification during 192Ir brachytherapy. Med. Phys. 36 (3), 708e718. Archambault, L., Arsenault, J., Beddar, A., Roy, R., Beaulieu, L., 2005. Plastic scintillation dosimetry: optimal selection of scintillating fibers and scintillators. Med. Phys. 32 (7), 2271e2278. Archambault, L., Beddar, A., Gingras, L., Roy, R., Beaulieu, L., 2006. Measurement accuracy and Cerenkov removal for high performance, high spatial resolution scintillation dosimetry. Med. Phys. 33 (1), 128e135. Ayotte, G., Archambault, L., Gingras, L., Lacroix, F., Beddar, A., Beaulieu, L., 2006. Surface preparation and coupling in plastic scintillator dosimetry. Med. Phys. 33 (9), 3519e3525. Bartesaghi, G., Conti, V., Bolognini, D., Grigioni, S., Mascagna, V., Prest, M., Scazzi, S., Mozzanica, A., Cappelletti, P., Frigerio, M., Gelosa, S., Monti, A., Ostinelli, A., Giannini, G., Vallazza, E., 2007. A scintillating fiber dosimeter for radiotherapy. Nucl. Instrum. Methods Phys. Res. A 581 (1e2), 80e83. Beddar, A., 1994. A new scintillator detector system for the quality assurance of 60Co and high-energy therapy machines. Phys. Med. Biol. 39, 253e263. Beddar, A., 2007. Plastic scintillation dosimetry and its application to radiotherapy. Radiat. Meas. 41, S124eS133.

Beddar, A., Briere, T., Mourtada, F., Vassiliev, O., Liu, H., Mohan, R., 2005. Monte Carlo calculations of the absorbed dose and energy dependence of plastic scintillators. Med. Phys. 32 (5), 1265e1269. Beddar, A., Kinsella, T., Ikhlef, A., Sibata, H., 2001. A miniature “scintillator-fiberoptic-pmt” detector system for the dosimetry of small fields in stereotactic radiosurgery. IEEE Trans. Nucl. Sci. 48 (3), 924e928. Beddar, A., Mackie, T., Attix, F., 1992. Cerenkov light generated in optical fibers and other light pipes irradiated by electron beams. Phys. Med. Biol. 37 (4), 925e935. Beierholm, A., Anderson, C., Lindvold, L., Kjaer-Kristoffersen, F., Medlin, J., 2008. A comparison of bcf-12 organic scintillators and Al2O3:C crystals for real-time medical dosimetry. Radiat. Meas. 43, 898e903. Buzhan, P., Dolgoshein, B., Filatov, L., Ilyin, A., Kaplin, V., Karakash, A., Klemin, S., Mirzoyan, R., Otte, A., Popova, E., Sosnovtsev, V., Teshima, M., 2006. Large area silicon photomultipliers: performance and applications. Nucl. Instrum. Methods Phys. Res. A 567 (1), 78e82. Clift, M., Johnston, P., Webb, D., 2002. A temporal method of avoiding Cerenkov radiation generated in organic scintillator dosimeters by pulsed mega-voltage electron and photon beams. Phys. Med. Biol., 1421e1433. Clift, M., Sutton, R., Webb, D., 2000. Water equivalence of plastic organic scintillators in magavoltage radiotherapy bremsstrahlung beams. Phys. Med. Biol., 1885e1895. Dhulla, V., Gudkov, G., Gavrilov, D., Stepukhovich, A., Tsupryk, A., Kosobokova, O., Borodin, A., Gorbovitski, B., Gorfinkel, V., 2007. Single photon counting module based on large area APD and novel logic circuit for quench and reset pulse generation. IEEE J. Sel. Top. Quantum Electron. 13 (4), 926e933. Elsey, J., McKenzie, D., Lambert, J., Suchowerska, N., Law, S., Fleming, S., 2007. Optimal coupling of light from a cylindrical scintillator into an optical fiber. Appl. Opt. 46 (3), 397e404. Frelin, A.-M., Fontebonne, J.-M., Ban, G., Batalla, A., Colin, J., Isambert, A., Labalme, M., Leroux, T., Vela, A., 2006. A new scintillating fiber dosimeter using a single optical fiber and a CCD camera. IEEE Trans. Nucl. Sci. 53 (3), 1113e1117. Globus, M., Hrynyova, T., Ratner, M., 2002. Optimization of detection system based on inorganic crystal coupled with a long lightguide. Nucl. Instrum. Methods Phys. Res. A 486, 55e59. Golovin, V., Saveliev, V., 2004. Novel type of a photodetector with Geiger mode operation. Nucl. Instrum. Methods Phys. Res. A 518 (1e2), 560e564. Gomi, S., et al., 2007. Development and study of the multi pixel photon counter. Nucl. Instrum. Methods Phys. Res. A 581, 427e432. Hamamatsu, P., 2008a. S10362-11-050C. http://sales.hamamatsu.com/en/products/ solid-state-division/si-photodiode-series/mppc/part-s10362-11-050c.php URL. Hamamatsu, P., 2008b. H5783e03. http://sales.hamamatsu.com/en/products/electrontube-division/detectors/photomultiplier-modules/part-h5783-20.php URL. Heckathorne, E., Silverman, R., Daghighian, F., Dahlbom, M., 2006. Prototype solid state photomultiplier based intra-operative beta camera. Nucl. Sci. Symp. Conf. Rec. 6, 3571e3575. Huston, A., Justus, B., Falkenstein, P., Miller, R., Hing, N., Altemus, R., 2001. Remote optical fiber dosimetry. Nucl. Instrum. Methods Phys. Res. B. 184, 55e67. Hyer, D., Fisher, R., Hintenlang, D., 2009. Characterization of a water-equivalent fiber-optic coupled dosimeter for use in diagnostic radiology. Med. Phys. 36 (5), 1711e1716. Jelly, J., 1958. Cerenkov Radiation and its Applications. Pergamon Press, London UK. Justus, B., Falkenstein, P., Huston, A., Plazas, M., 2004. Gated fiber-optic-coupled detector for in vivo real-time radiation dosimetry. Appl. Opt. 43 (8), 1663e1668. Justus, B., Falkenstein, P., Huston, A., Plazas, M., 2006. Elimination of Cerenkov interference in a fibre-optic-coupled radiation dosemeter. Radiat. Prot. Dosimetry 120 (1e4), 20e23. Kapustinsky, J.S., DeVries, R.M., DiGiacomo, N.J., Sondheim, W.E., Sunier, J.W., 1985. A fast timing light pulser for scintillation detectors. Nucl. Instrum. Methods Phys. Res. A 241, 612e613. Klein, D., Yukihara, E., Bulur, E., Durham, J., Akselrod, M., McKeever, S., 2005. An optical fiber radiation sensor for remote detection of radiological materials. IEEE Sensors J. 5 (4), 581e588. Lambert, J., Yin, Y., McKenzie, D., Law, S., Suchowerska, N., 2008. Cerenkov-free scintillation dosimetry in external beam radiotherapy with an air core light guide. Phys. Med. Biol. 53, 3071e3080. Lambert, J., Yin, Y., McKenzie, D., Law, S., Suchowerska, N., 2009. Cerenkov light spectrum in an optical fiber exposed to a photon or electron radiation therapy beam. Appl. Opt. 48 (18), 3362e3367. Law, S., Suchowerska, N., McKenzie, D., Fleming, S., Lin, T., 2007. Transmission of Cerenkov radiation in optical fibers. Opt. Lett. 32 (10), 1205e1207. Lee, B., Cho, D., Jang, K., Chung, S., Lee, J., Kim, S., Cho, H., 2006. Measurements and characterization of Cerenkov light in fiber-optic radiation sensor irradiated by high-energy electron beam. Jpn. J. Appl. Phys. 45 (10A), 7980e7982. Lee, B., Hwang, Y., Cho, H., Cho, S., 2004. Fabrication of fiber-optic radiation sensor tips with inorganic scintillator for remote sensing of x or g-ray. Nucl. Sci. Symp. Conf. Rec. 7, 865e868. Lee, B., Jang, K., Cho, D., Chung, S., Tack, G., Eom, G., Yi, J., 2007a. Fabrication and characterization of a minature fiber-optic scintillating detector for electronbeam therapy dosimetry. J. Korean Phys. Soc. 51 (5), 1645e1648. Lee, B., Jang, K., Cho, D., Yoo, W., Tack, G., Chung, S., Kim, S., Cho, H., 2007b. Measurements and elimination of Cerenkov light in fiber-optic scintillating detector for electron beam therapy dosimetry. Nucl. Instrum. Methods Phys. Res. A 579, 344e348.

D.C. Konnoff et al. / Radiation Measurements 46 (2011) 76e87 Matsuura, Y., Miyagi, M., 2004. Hollow optical fibers for ultraviolet and vacuum ultraviolet light. IEEE J. Quantum Electron. 10 (6), 1430e1434. McNally, D., Golovin, V., 2009. Review of solidstate photomultiplier developments by CPTA and photonique SA. Nucl. Instrum. Methods Phys. Res. A 610 (1), 150e153. Nowotny, R., 2007. Radioluminescence of some optical fibres. Phys. Med. Biol. 52, N67eN73. Pavlov, N., Maehlum, G., Meier, D., 2005. Gamma spectroscopy using a silicon photomultiplier and a scintillator. Nucl. Sci. Symp. Conf. Rec., 173e180. Photonique, I., 2008. PH0810G1. http://www.photonique.ch URL. Renker, D., Lorenz, E., 2009. Advances in solid state photon detectors. J. Inst. 4, 1e47.

87

Saveliev, V., Golovin, V., 2000. Silicon avalanche photodiodes on the base of metalresistor-semiconductor (MRS) structures. Nucl. Instrum. Methods Phys. Res. A 442, 223e229. Suchowerska, N., Lambert, J., Nakano, T., Law, S., Elsey, J., McKenzie, D., 2007. A fibre optic dosimeter customized for brackytherapy. Radiat. Meas. 42, 929e932. Sze, S.M., 2006. Physics of Semiconductor Devices, third ed. Wiley-Interscience, New York (Chapter 13). Voxtel, I., 2008. SBQF-EKAA. http://www.voxtel-inc.com/index.php/products/apdsproduct-page/ URL. Wick, K., Zoufal, T., 2001. Unexpected behavior of polystyrene-based scintillating fibers during irradiation at low doses and low dose rates. Nucl. Instrum. Methods Phys. Res. B. 185, 341e345.