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Development of a wavelength-separated type scintillator with optical fiber (SOF) dosimeter to compensate for the Cerenkov radiation effect

Abstract The scintillator with optical fiber (SOF) dosimeter consists of a miniature scintillator mounted on the tip of an optical fiber. The scintillator of the current SOF dosimeter is a 1-mm diameter hemisphere. For a scintillation dosimeter coupled with an optical fiber, measurement accuracy is influenced by signals due to Cerenkov radiation in the optical fiber. We have implemented a spectral filtering technique for compensating for the Cerenkov radiation effect specifically for our plastic scintillator-based dosimeter, using a wavelength-separated counting method. A dichroic mirror was used for separating input light signals. Individual signal counting was performed for high- and low-wavelength light signals. To confirm the accuracy, measurements with various amounts of Cerenkov radiation were performed by changing the incident direction while keeping the Ir-192 source-to-dosimeter distance constant, resulting in a fluctuation of <5%. Optical fiber bending was also addressed; no bending effect was observed for our wavelength-separated SOF dosimeter.


INTRODUCTION
Recently, as CT-based 3D radiation treatment-planning systems (RTPs) for brachytherapy have become more accessible, the importance of sparing organs at risk (OARs) from tolerant dose has been increasing for high-precision radiation therapy. Although in vivo dosimetry is preferable for verifying the delivered dose for tumors and OARs, some tumors and OARs are located deep inside the patient body, making direct measurement difficult. The small Si photo-diode detector and the MOSFET detector appear to be suitable in vivo dosimeters [1,2]. However, the material composition of these semiconductor detectors is not tissue-equivalent. Moreover, metal wires and cables are necessary for supplying electric power and transmitting signals from these detectors, which inevitably affect the measured dose for therapeutic radiation.
The use of a small plastic scintillation dosimeter coupled to an optical fiber was proposed and developed by Beddar  [3][4][5][6][7][8][9]. We have also developed a novel plastic scintillator-based dosimeter with pulse-counting mode [10]. While dosimeters using materials other than plastic coupled to an optical fiber have long been available [11,12], a plastic scintillator dosimeter is practically more suitable for photon dosimetry because the amount of light emissions from the scintillator is proportional to the absorbed dose. The composition of a plastic scintillator is similar to human tissue; therefore, it has little influence on the measured dose. The mean mass-energy absorption coefficient of the plastic scintillator is relatively close to that of water above 200 keV, as is the mass collision stopping power and the mass angular scattering power. Furthermore, light signals from the scintillator to the counting device can be transmitted through a plastic optical fiber, which maintains the tissueequivalent composition of the entire detector.
On the other hand, high-energy radiation produces Cerenkov radiation in materials [13]. This Cerenkov radiation is well known as an additional signal for plastic scintillation dosimeters, especially coupled with an optical fiber. Beddar et al. first described the Cerenkov Effect in 1992 [3], employing a simple subtraction method to remove the Cerenkov radiation effect and quantifying the stem effect [6,14]. Several research projects have been undertaken seeking to minimize the Cerenkov radiation effect. Beddar et al. and Clift et al. tried to compensate for the effect by subtracting two fibers, one with and one without a scintillator [15][16][17]. A spectral filtering technique, or a gating method, has also been investigated [18,19]. Archambault et al. and Guillot et al. tried to separate signals between scintillation and Cerenkov light using a dichroic filter [20,21]. Lambert et al. developed a Cerenkov-free dosimetry system by using an air-core light guide [22].
In this paper, we implement a spectral filtering technique for compensating for the Cerenkov radiation effect specifically for our plastic scintillator-based dosimeter by using a wavelengthseparated counting method and report the results of our investigation for high-dose-rate 192 Ir irradiation. Figure 1 shows the components of our scintillation dosimetry system reported previously [10]. In the succeeding discussion, we refer to this dosimeter as the scintillator with optical fiber (SOF) dosimeter. The scintillator material we used was BC490 plastic manufactured by Saint-Gobain Ltd. BC490 is partially polymerized and hardened by a catalyst, which makes it possible to be tightly attached to the tip of a plastic optical fiber (MH4001 manufactured by Mitsubishi Rayon Co. Ltd, reflective index of 1.49, 1-mm diameter optical fiber with 2.2-mm diameter polyethylene shielding). The plastic scintillator is shaped into a hemisphere with a diameter of 1 mm and an active volume of 0.26 mm 3 . Hemispherical shape is advantageous for better angular response to the incident radiation. The photon signals are relayed through the optical fiber onto the Photon Counting Head (Hamamatsu H7155), then converted into 30-ns-width TTL pulses. The counting unit is made up of a photomultiplier tube, a charge pre-amplifier and a discriminator. The pulse counts are transmitted to a personal computer via a universal serial bus (USB) connection. The measurable range of the SOF dosimeter, which is from 11.4 cGy min −1 to 1.15 × 10 4 cGy min −1 , covers a range of almost 10 3 cGy min −1 without switching measurement range [10].

Description of the scintillation dosimeter
The absorbed dose in the water D W can be described as where C Sc is the pulse count corresponding to the number of energy deposition events in the scintillator from the radiation with energy greater than E 0, which was denoted by the symbol C total in our previous paper [10], Δt is the duration of the measurement, α and β are the correction coefficients for converting count to dose, which can be expressed using calibration coefficients a [counts s −1 J −1 ] and b [J −1 ] by where, m is the mass of the scintillator, (m en =r) w and (m en =r) sc are the mass energy absorption coefficient for water and scintillator, respectively, and E 0 is the threshold energy for pulse counting. The parameter a is nuclei dependent, whereas parameter b is almost energy independent. C Sc can also be expressed as By using Eq. (3), βΔt can be expressed as Finally, the absorbed dose in the water D w is simplified as Signal separation by a dichroic mirror Figure 2 shows photographs of the wavelength-separated type SOF (WS-SOF) dosimeter used in this study. A dichroic mirror was placed facing at 45°from the optical fiber axis. As shown in Fig. 2, signals reflected by the dichroic mirror were defined as Ch1, while signals transparent through the dichroic mirror were defined as Ch2. Ch1 and Ch2 consist of the sum of scintillation signals C Sc and Cerenkov signals C Cere with some scale factors. A pair of the H7155 photon counting heads detect separated light signals. The scintillation spectrum of the BC490 plastic scintillator and Cerenkov radiation generated by 192 Ir is shown in Fig. 3. The plastic scintillator has a peak wavelength at 420 nm, whereas the spectrum of Cerenkov radiation varies over a wide range in the visible spectrum. Since these two spectra are quite different, we tried to distinguish contributions from them separately by using a dichroic mirror, setting an adequate threshold wavelength (λ th ). The spectrum of the BC490 plastic scintillator is distributed relatively lower compared with that of Cerenkov radiation. The threshold wavelength was set at 465 nm to preserve as much signal as possible with adequate removal of the Cerenkov radiation effect. The dichroic mirror was ordered with a 465-nm threshold wavelength at 45°respect to the incident light path. As shown in Fig. 3, we defined I L_BC490 as the sum of intensities of BC490, which were lower than the threshold, I L_Cere , the sum of intensities of Cerenkov radiation, which were lower than the threshold, I H_BC490 , the sum of intensities of BC490, which were higher than the threshold, and I H_Cere , the sum of intensities of Cerenkov radiation, which were higher than threshold.
Assuming that scintillation signals are counted as C Sc by the single photon-counting unit, signals divided by the dichroic mirror should have smaller pulse counts due to smaller pulse height with an adequate discriminator level for eliminating electronic noise. Count reduction factors due to separating signals by the dichroic mirror for the scintillation and Cerenkov radiation are defined a Sc and a Cere , respectively. To simplify the equations, efficiency ratios are defined as k Sc and k Cere due to the efficiency difference between the photomultiplier tubes for Ch1 and Ch2. Here, k Sc and k Cere not Fig. 3. Measured spectra of BC490 plastic scintillator and Cerenkov radiation. I L_BC490 , I L_Cere , I H_BC490 , I H_Cere represent sum of intensities of BC490 lower than threshold, sum of intensities of Cerenkov radiation lower than threshold, sum of intensities of BC490 higher than threshold, sum of intensities of Cerenkov radiation higher than threshold, respectively.

Fig. 2.
Photographs of the wavelength-separated type SOF dosimeter. A dichroic mirror was placed just after the SMA connector to split the spectrum of light signals by threshold wavelength. only include efficiency difference but also the reflection/ transparent ratio due to the dichroic mirror.
The basic concept is very similar to the spectral filter technique proposed by Fontbonne et al. [18]. From equations (6) and (7), C Sc and C Cere are deduced as follows: Scintillation signals and Cerenkov signals can be estimated by the number of counts of Ch1 and Ch2 with correction coefficients k Sc and k Cere .

Experiments
Calibration for correction coefficient k Sc and k Cere Generally, low-energy photons (e.g. under 100 keV) do not generate Cerenkov radiation in the plastic optical fiber. However, generation of a fluorescence from the optical fiber itself might be observed when irradiated with even low-energy photons (10 keV), and thus the optical fiber was shielded with 2-mm-thick lead sheet for this calibration process. A 10-kV X-ray source generated by ECLIPSE III (manufactured by AmpTek Co. Ltd, Ag target) was used to evaluate k Sc . The influence of Cerenkov radiation can be ignored by using lowenergy photons. Eqs. (6) and (7) can be simplified as follows: Then k Sc can be calculated by Calibration for the coefficients of k Cere was performed by two measurements for which the dose at the scintillator was the same but the amount of Cerenkov radiation was different.
Assuming that the calibration geometry of A and B have the same gamma-ray intensity as the scintillators but have different Cerenkov intensities to the optical fiber, each channel count can be expressed as follows: a Sc C Sc;A a Sc C Sc;B a Cere C Cere;A a Cere C Cere;B ð13Þ From Eqs. (13) and (14), k Cere can be deduced as Similarly, the water equivalent dose can be expressed by the following equation: In this study, microSelectron HDR (manufactured by Nucletron Co. Ltd) was used for the 192 Ir source. Source calibration details are shown in Table 1. For the probe calibration, the Afterloading calibration phantom (type 9193 manufactured by Nucletron Co. Ltd) was used. As shown in Fig. 4, the phantom is made of 20 cmφ × 12 cm Polymethylmethacrylate (PMMA) with four holes with diameter of 2 cmφ at 8 cm distance from the 192 Ir source position (phantom center). At the distance of 8 cm, the absorbed dose rate is adjusted tõ 1 cGy/min when the 37 GBq 192 Ir source is used (type 9193 phantom user's manual) [23]. The water-equivalent dose is calculated as follows: Here, _ D w;zp is the reference absorbed dose rate to water in the cylindrical phantom at 8 cm distance from the source center, _ K a ) a À is the air kerma rate in air at 100 cm distance, g w is the relative Bremsstrahlung loss in water for 192 Ir, and (m en =r) w and (m en =r) a are the mass energy absorption coefficient for water and air, respectively. k r and k zp are correction factors for the inverse square of the distance and geometry for the cylindrical phantom, respectively; here these values are defined as k r = (8/100) 2 = 0.0064 and k zp = 1.187 ± 0.012. The absorbed dose rate at the calibration point was confirmed by measurement using a Farmer type ionization chamber (PTW30004). The difference between calculation and measurement was 0.13%.

Scintillation-Cerenkov signal separation
In this study, two probes with differing sensitivities were used. Probe #1 is a normal probe with an estimated volume of 0.234 mm 3 , while probe #2 has a smaller scintillator with a volume of 0.106 mm 3 to enhance the Cerenkov radiation effect. In order to assess the accuracy of the scintillation-Cerenkov signal separation method, measurements were performed with the geometry shown in Fig. 5. A 30 × 30 × 2 cm 3 PMMA phantom with a guide sleeve was placed at a 5-cm distance from the phantom center, which is the location of the SOF dosimeter probe. For gathering scattering gamma rays, 5-cm-thick Mix-DP phantoms were placed on both sides of the PMMA phantom as shown in Fig. 5. Zero degrees is defined toward the tip of the SOF dosimeter.
Measurements were acquired at 0, 30, 60, 90, 120, 150, 160, 170 and 180°. The 180°position was the closest to the 192 Ir source; the distance was~2 mm. Acquisition times for each measurement point were 20 s, repeated five times. Calibration measurements were acquired for 60 s, repeated 10 times.

Comparison with MCNPX calculation
In order to justify the measurements obtained from WS-SOF dosimeter, a comparison with MCNPX 2.4 Monte Carlo code was performed. Measurement and calculation was performed in a 30 × 30 × 30 cm 3 water phantom with a 5-mm-thickness PMMA wall. An 192 Ir source was fixed at the center of the phantom. The SOF dosimeter probe was scanned along an axis 1 cm away from the 192 Ir source. The scan range was −100 mm to +100 mm on the axis with 10-mm intervals. Additional measurements were performed within 10 mm with 1-mm intervals. The measurement at each position was performed four times with a 10-min acquisition time. The 192 Ir source strength at the experiment date was 26.42 mGy m 2 h −1 .

Calibration for correction coefficients k Sc and k Cere
Correction coefficients k Sc and k Cere are shown in Table 2.
For k Cere calibration, the 192 Ir source was placed at the 0°& 160°positions and the 0°& 180°positions in the same geometry of Fig. 5. Conversion factor α′ cannot be deduced in our theory; however, α′/a Sc can be estimated by comparing the value of a Sc C Sc with the absolute dose measurement. As mentioned in the Experiments section, the calibration process requires two different conditions which have different amount of Cerenkov radiation to the fiber with exactly same gamma-ray incidence from Ir-192 source. As shown in Fig. 5, the 0°and 180°positions are the minimum and maximum position influenced by Cerenkov radiation, respectively. On the contrary, a small Cerenkov radiation effect is expected at the 160°position. The 0°& 160°and 0°& 180°positions were selected for the calibration process in order to compare the calibration accuracy due to the amount of the Cerenkov radiation. Figure 6 shows the angular dependency for separated signals of scintillation and Cerenkov radiation as a function of incident angle for two probes for which calibrations were calibrated with 0°and 160°source position values. Figure 7 shows the angular dependency with 0°and 180°source position values. All values in Figs 6 and 7 were calculated by Eq. (6) and Eq. (7), normalized at the scintillation signal of 0°position. As shown in Fig. 6, the estimated scintillation signals for calibration with 0°& 160°positions were influenced by an increase of Cerenkov light. A more accurate calibration can be performed using 0°& 180°data, which contains sufficient Cerenkov signals. Relative estimation errors % s D W for the relative absorbed dose %Dw in Figs 6 and 7 were deduced from error propagation formula as described below:

Experiment for scintillation-Cerenkov signal separation
Here, s Ch1 and s Ch2 are measurement uncertainty for Ch1 and Ch2, and s k Sc and s k Cere are estimation uncertainty for parameters of k Sc and k Cere , respectively. Figure 8a and 8b shows the results of the SOF measurement and the MCNPX 2.4 calculation. The values for the MCNPX calculation were normalized at the −50 mm position with the value of SOF probe #1. The maximum differences between the SOF measurement and the MCNPX calculation for probes #1 and #2 were 5.06% and 4.80%, respectively, whereas for the limited range within 10 mm, the maximum differences were 0.90% and 2.07%, respectively. Therriaut-Proulx et al.

Comparison with MCNPX calculation
[23] performed a phantom study using a plastic scintillator detector (PSD) based on in vivo dosimetry for HDR 192 Ir brachytherapy. The measurement agreed within 6% compared with the calculation of a treatment-planning system. Our results are comparative with these results.
On the other hand, if the compensation for Cerenkov radiation is not performed, the maximum differences between the SOF measurement and the MCNPX calculation for probes #1 and #2 were 15.1% and 25.9%, respectively. In this case, since the calibration was performed without employing any compensation for Cerenkov radiation, significant signal loss was observed near the 192 Ir source position.

Signal attenuation through optical fiber
As shown in Fig. 7, Cerenkov signal removal was adequately performed, even if the Cerenkov signals were nine times higher than the scintillation signals. For probe #1, the efficiency fluctuation was less than 1% for all directions; that is, the angular dependency was quite small. The absorbed doses estimated by the two probes were 18.63 ± 0.16 Gy (#1) and 18.58 ± 0.18 Gy (#2) at the measured position in Fig. 5, respectively; measured doses were quite identical. On the other hand, for the spectrum filtering technique, the possibility exists of efficiency change due to spectrum change that may occur by bending the optical fiber. To confirm the effect of optical fiber bending, wavelength spectra was measured at 180°, as shown in Fig. 5, with the optical fiber lopped five times (diameters 10, 15 and 20 cmφ) using probe #1 and the USB-2000 spectrometer (manufactured by Ocean Optics Ltd). The relative efficiency for the 10, 15 and 20 cmφ loops were 1.000 ± 0.0027, 1.005 ± 0.0026 and 1.007 ± 0.0025, respectively. Here, relative efficiency refers to the ratio of the signal value with and without bending of the optical fiber. The efficiency change due to bending an optical fiber seems thus not to be observed. Figure 9 shows the measured wavelength spectra of the scintillation and Cerenkov signals with and without bending using another scintillator-less probe. As shown in Fig. 9, both the scintillation and Cerenkov light spectrum are almost identical, even if the optical fiber was bent with 10 cmφ × 5 times. For practical usage, looping optical fiber with less than 10 cmφ diameter is rare; the efficiency due to bending will thus be negligible. The Cerenkov spectrum was different from that reported by Frelin et al. (2005) [19], indicating that attenuation of short wavelengths in the optical fiber might occur. However, calibration was performed including such an attenuation effect; influence to the dose evaluation seems therefore to be negligible for light attenuation in the fiber.

Consideration for in vivo dosimetry
For 192 Ir HDR brachytherapy, the source position frequently changes during treatment, and the Cerenkov signals change   accordingly. The compensation method should therefore be effective for not only constant conditions but also for such transient situations. Here, the mixture of two situations (Situation A and Situation B), which have the different scintillation signals of S A and S B and Cerenkov signals of C A and C B , are considered. Note that the count reduction factors a Sc and a Cere are included in each signal in order to simplify the calculation. Signal counts for each channel can be expressed by using Eq. (13) as follows.
where Ch1 A and Ch1 B stand for the signal counts on Ch1 under Situations A and B, respectively, and Ch2 A and Ch2 B stand for signal counts on Ch2 under Situations A and B, respectively. The scintillation signals S A and S B are calculated by Eq. (19), and the total signal is expressed as As shown in Eq. (22), the sum of each scintillation signal is expressed with the sums of each channel. This means that the wavelength-separated SOF dosimeter will be effective in actual measurement, and accurate integral dose can be obtained by using integral counts for each channel.
In practical use, a small-sized probe is better for in vivo dosimetry because the dosimeter passes through a small cavity inside the body. The presented dosimeter can compensate for the Cerenkov radiation effect without the reference fiber as necessary in the subtraction method reported previously [10], which facilitates smooth insertion into the patient body. Moreover, since the 192 Ir source goes through several catheters, angular dependency is important for accurate in vivo dosimetry, and this has been achieved by the hemispherical shape of the scintillator.

CONCLUSIONS
In order to compensate for the Cerenkov radiation effect, input light signals were separated using a dichroic mirror. Signals for high and low wavelengths were counted separately. Measurements with various amounts of Cerenkov radiation were performed by changing the incident direction while keeping the 192 Ir source-to-dosimeter distance constant, resulting in a fluctuation of <5%, even if the 192 Ir source was placed 2 mm from the optical fiber. Accurate measurement is therefore possible at any source position. For in vivo dosimetry, the probe should consist of non-toxic materials and be sterilized. Further investigation will be necessary for applying the proposed dosimeter.