Como parte fundamental de um sistema de detecção composto por um material cintilador, as fibras ópticas são responsáveis pela transmissão do sinal luminoso gerado pelo dosímetro até à fotomultiplicadora ou fotodiodo. Essa transferência de sinal no interior da fibra é possível devido à diferenças de índices de refração entre a casca e o núcleo.
De acordo com a Lei da Refração de Snell, quando um raio luminoso passa de um meio de índice de refração ' para outro de índice de refração (, respectivamente com um ângulo de incidência θ) e de refração θ*, tem-se que
(1,21,27):
+ θθθθ, + θθθθ (equação 3)
• Quando ' é maior que (, o raio refratado se afasta da normal (θ* θ)), como mostra a Figura 9.
Figura 9 - Refração da luz para . Adaptado de (21).
• Aumentando-se o ângulo de incidência θ), o ângulo de refração θ* se aproxima de 90/, de modo que 01 θ* 1 e, portanto:
θ θθ
θ, θθθθ3 + 45
6 (equação 4)
Como mostrado na Figura 10, o ângulo de incidência (θ)) passa a
Figura 10 - Refração da luz quando θθθθ se aproxima de . Adaptado
de (21).
• Na Figura 11, ilustra-se o fenômeno denominado Reflexão Interna Total. Ele ocorre quando o ângulo de incidência θ) é
maior que o ângulo crítico θ7, fazendo com que a luz retorne ao
meio 1. Neste caso, os ângulos de incidência e reflexão são iguais.
Figura 11 - Esquema que ilustra o fenômeno de Reflexão Interna Total. Adaptado de (21).
É através do princípio da Reflexão Interna Total que se explica o efeito pelo qual a luz é guiada no interior de fibras ópticas. O raio luminoso incidente é primeiramente transmitido para dentro do núcleo da fibra que apresenta um índice de refração '. Em seguida, o raio atinge a casca da fibra de índice de refração (, onde (8 ', é retroespalhado e permanece sofrendo reflexão interna total até o final da fibra, onde ele é transmitido para o meio externo, como mostrado na Figura 12.
Figura 12 - Vista longitudinal de uma fibra óptica cuja propagação luminosa é explicada pela Reflexão Interna Total. Adaptado de (21).
O ângulo máximo de incidência () que um raio luminoso, vindo do meio externo, pode formar com o eixo central da fibra, para que se propague no núcleo pelo fenômeno de Reflexão Interna Total, é caracterizado pelo parâmetro denominado Abertura Numérica (9:) (1,21,27). Matematicamente, ela pode ser definida como:
;< + = (equação 5)
Aplicando-se a Lei da Refração de Snell na ;< + = (equação 5), chega-se à relação entre o ângulo de incidência e os índices de refração da casca e do núcleo da fibra, onde:
;<
Assim como o AN, a dimensão do núcleo da fibra são parâmetros que estão fortemente relacionados à eficiência com que o núcleo aceita a luz incidente do meio externo ( #), caracterizando a eficiência de detecção da fibra.
As fibras ópticas podem ser classificadas como (1,21,27):
• Fibra multimodo de índice degrau: fabricada em sílica, vidro ou acrílico é caracterizada por apresentar um núcleo de índice de refração constante.
• Fibra multimodo de índice gradual: caracterizada por apresentar um núcleo com várias camadas de vidro e, consequentemente, com índices de refração que decrescem gradualmente à medida que se afastam do centro em direção à casca.
• Fibra Monomodo: possui um perfil de índice degrau, porém com núcleos extremamente pequenos (cerca de 10 vezes menor que a casca).
Outro fator importante que caracteriza a fibra óptica é a atenuação da intensidade da luz (em decibel) ao percorrer toda sua extensão (1,21,27). Pode ser definida matematicamente como:
@A !10log E""
#F
Onde "# e " são, respectivamente, a intensidade inicial da luz e a intensidade após ter percorrido toda a extensão da fibra.
9. Radiação Cerenkov
A radiação Cerenkov é gerada no interior das fibras ópticas que compõem o sistema de detecção de um dosímetro de cintilação. O fenômeno ocorre quando partículas carregadas atravessam um meio dielétrico, como o núcleo da fibra, com uma velocidade maior que a da luz neste local (26,28,29,30). Assim, o sinal da luz Cerenkov se soma ao sinal luminoso produzido pelo cintilador, tornando-se um fator indesejável nas medidas do dosímetro. Somado a ele, há ainda o ruído decorrente da fluorescência gerada pelas impurezas presentes na fibra.
No caso desse trabalho, a radiação Cerenkov é produzida no interior de uma fibra óptica de poly(methyl methacrylate), PMMA, de 1 mm de diâmetro. Essa fibra foi utilizada para fazer a transferência do sinal luminoso gerado pelo dosímetro até o sistema de detecção (fotomultiplicadora e eletrômetro), que permanece fora da sala onde é feita a irradiação.
Sendo G a velocidade da luz no vácuo, a intensidade da luz Cerenkov
" , para um determinado comprimento de onda H , gerada por uma partícula que atravessa um meio de índice de refração com velocidade I, é proporcional a (26,29,31,32,33):
J !3K (equação 7)
Os fótons Cerenkov são emitidos em um cone de luz com relação à direção da velocidade da partícula (34,35,36), como ilustrad na Figura 13.
Figura 13 – Comportamento luminoso da radiação Cerenkov. Os fótons Cerenkov gerados são emitidos através de um ângulo com relação à direção da velocidade da partícula, produzindo um cone de luz.
O ângulo L entre o vetor da direção dos fótons Cerenkov com o vetor da direção da velocidade da partícula incidente é dado por:
MNO P K3 (equação 8)
No limite relativístico onde a velocidade da partícula se aproxima de G, o ângulo é dado por:
P MNO45
6 (equação 9)
Dessa forma, por exemplo, para a fibra de PMMA, cujo índice de refração é de 1,49 e ângulo de aceitação/propagação de 60 , o ângulo L é de aproximadamente 42 . Assim sendo, como L está dentro do ângulo de propagação da fibra, a luz Cerenkov permanecerá confinada no interior da mesma e seu sinal se somará à luz emitida pelo cintilador.
Objetivos
• Construir e caracterizar um dosímetro de cintilação para dosimetria relativa e controle de qualidade de feixes de megavoltagem, a partir do desenvolvimento de uma mistura cintiladora encapsulada por fibra capilar; • Desenvolver um sistema de transmissão e leitura do sinal luminoso,
construído com cabos de fibra óptica tradicional e dispositivos eletrônicos de base;
• Analisar o desempenho do dosímetro em feixe clínico a partir da comparação com um sistema dosimétrico de referência, para avaliação da viabilidade do mesmo no uso em rotina.
Artigo
Development of a scintillation dosimeter for photon
clinical beams
Renata Rodrigues dos Santos Lixandrão, BSc.
Department of Internal Medicine, Faculty of Medical Sciences, University of Campinas, Campinas, São Paulo, Brazil.
Cristiano Monteiro de Barros Cordeiro, Ph.D.
Institute of Physics “Gleb Wataghin”, University of Campinas, Campinas, São Paulo, Brazil.
Arnaldo Luis Lixandrão Filho, BSc.
Three-dimensional Technologies Department, Renato Archer Information Technology Center, Campinas, São Paulo, Brazil.
José Renato Oliveira Rocha, MSc. In memoriam
Vilma Aparecida Ferrari, MSc.
Medical Physics Department, Biomedical Engineering Center, University of Campinas, Campinas, São Paulo, Brazil.
José Barreto Campello Carvalheira, MD, Ph. D.
Department of Internal Medicine, Faculty of Medical Sciences, University of Campinas, Campinas, São Paulo, Brazil.
Corresponding Author
Renata Rodrigues dos Santos Lixandrão Av. Santa Isabel, 310 – Apto. 32
Barão Geraldo Campinas-SP 13084-012
renata.fisicamedica@gmail.com
Abstract
Purpose: In the present work we have developed and validated a scintillator radiation detection system designed for megavoltage photon clinical beams.
Methods and Materials: The developed dosimeter contains a 6mm3 radiation sensitive volume of milled anthracene mixed with an optical polymer inside a SiO2 capillary. The scintillation light was guided through optical fibers to a
photomultiplier outside of the irradiation room connected to an electrometer. The electrometer measured signal is related with dosimeter exposed dose. The Cherenkov signal was measured using a reference fiber for later background subtraction. Characteristics and performance tests were done using a 6 MV linear accelerator photon beam. Dosimeter performance and some beam parameters were obtained, analyzed and compared with an FC65 Farmer Ionization Chamber.
Results: The optical polymer that was mixed with anthracene reduced the scattering effects of the dose-sensitive volume and increased the luminescence capture by the guide fiber. The dose-sensitive volume is around 100 times smaller than the FC65 Chamber. It was detected that there is a time of 6.2s before stabilization of the dose rate readings and that the amount of light emitted by scintillator and hence the readings obtained by the detection system, increases
linearly with dose rate. Reproducibility measurements were evaluated using ANOVA. The dosimeter Percentage Depth Dose (PDD) and dose profile curves are in close agreement with reference readings.
Conclusions: Real time measurements with the developed dosimeter have high accuracy after the stabilization time making it reliable for quality assurance. Due to its reduced size has great potential to be used when a high spatial resolution is needed, i.e., in situations that presents high dose gradient regions, like build-up regions and small radiation fields dosimetry.
Keywords
Radiation therapy, Quality Assurance, Scintillator Dosimeter, Optical Fiber, Anthracene.
Introduction
Quality assurance in radiation therapy includes those procedures that ensure a consistent and safe fulfillment of the dose prescription to the target volume with minimal dose to normal tissues and minimal exposure to personnel. A comprehensive quality assurance program is necessary because of the importance of accuracy in dose delivery [1], mainly for Intensity- Modulated Radiotherapy (IMRT) due to its small fields and regions with high dose gradient [2, 3]. Nevertheless, it is difficult to ensure that the radiotherapy planning calculated dose is being in fact delivered to the target volume in the patient, since there are many setup uncertainties [4, 5]. To overcome these problems, and for ensure an effective and safe treatment for the patient, it is necessary the development of more feasible and accurate dosimetric devices and techniques to improve the quality assurance.
Ionization chambers are very efficient dosimeters, being conventionally used in dosimetry. The profile of the beam radiation (flatness), the dose rate (UM/min) and the determination of the Percentage Depth Dose (PDD) curves are, however, difficulty to obtain, especially for small fields, due to the relatively large size of these detectors and because it integrates over dose gradients [2]. It is also necessary to make some corrections in their readings due the dependence of energy mass absorption coefficient, temperature and pressure that may lead to errors in dose calculations [6].
Among other kind of dosimeters, the scintillators are interesting mainly because they have small size, are water-equivalent, have linear response, have low energy dependence, are waterproof and do not require voltage to operate. A comparative of desired properties for radiotherapy dosimeters are shown in Table 1.
Chloride polyvinyltoluene (PVT), the most studied scintillator in radiotherapy [7,8,9], has several characteristics that makes it a good radioluminescent dosimeter. Its sensibility, however, is just about 65% of anthracene, a traditional organic scintillator used for detection of elementary particles. Anthracene (C6H4CH)2, a polycyclic aromatic hydrocarbon has a high
scintillation efficiency with fluorescence in the wavelength range of blue/violet. It was studied as a radiation detector [10, 11] and, in 1954, Robinson and Jentschke showed that his fluorescence has a linear response to X-ray doses of energies greater than 130 keV [12]. Furthermore, according to Pereira Neto anthracene does not presents any carcinogenic, genotoxic or mutagenic effects [13].
Detector systems using plastic scintillators attached to conventional optical fiber has been utilized since 1990 [7] in radiotherapy as they can provide real-time measurements of small fields (implying in higher spatial resolution) and areas with high dose gradients [14]. They also have radiological properties that are similar to the human body tissue. A full detection system is composed by a small amount of scintillation material coupled to a conventional optical fiber that is used
to guide the generated light till a photomultiplier or photodiode. The photomultiplier converts the optical to an electrical signal to be measured with an electrometer [8, 9]. Using fiber optics allows the electrical equipment to operate outside the irradiation room, minimizing radiation and radio frequency noise. For example, linear accelerators generate microwaves that strongly influence measuring instruments.
According to Beddar [14], however, a drawback for using scintillation detector systems in dosimetry is the low signal-to-noise (S/N) ratio. One of the main noise contributions is expected to arise from radiation-induced light produced inside the optical fiber due to a combination of Cerenkov emission, fluorescence or luminescence, depending on the fiber material. For pure fused silica optical fibers, this spurious light emission is predominantly due to Cerenkov [2,15]. Its spectrum spans over the shorter wavelength region of the visible spectrum, overlapping, unfortunately, with the typical emission spectra of most scintillating materials like anthracene. A solution is to remove its influence via a background subtraction [14].
In this paper it is presented a radiotherapy dosimeter that uses anthracene as the scintillator material. It is encapsulated by a specially processed SiO2 capillary and the generated light guided by a multimode standard polymeric
optical fiber. It is shown that its reduced size, good linearity, reproducibility, sensitivity and response makes it suitable to be used for relative dosimetry in megavoltage clinical linear accelerators photon beams.
Methods and Materials
The built detection system (Fig. 1) is composed by the scintillator sample (anthracene) inserted in a capillary fiber. This sensing head is attached to a conventional optical fiber that guides the light signal to the photomultiplier. At last, the electrical signal is measured by an electrometer.
For radiotherapy beams (0.1 to 15 MV), the predominant interaction in low-Z materials is via Compton scattering [14]. Some reports in literature shows anthracene water equivalence [16, 17] and Table 2 summarized some physical properties of scintillator and water regarding to water equivalence.
To study the optical properties of anthracene we measured its transmission using a Perkin Elmer Model Lambda 9 spectrophotometer. Then, we milled and mixed the crystals with an optically clear liquid polymer NOA73 (Norland Optical Adhesive 73). The adhesive cures quickly when exposed to long ultraviolet light and shows an refractive index of 1.56 [18]. Pure anthracene and mixture samples luminescence spectra were obtained by an Optical Spectrum Analyzer (OSA) through exposure to UV light.
To encapsulate the scintillator mixture (anthracene plus optical polymer), a 5 mm long SiO2 capillary was built. Fig. 2 shows optical coupling
details between the capillary and the guide fiber. It should be noted that the detector length and the diameter of the guiding fiber were chosen based on a geometry optical optimization approach introduced by Elsey [20].
As a waveguide, two step index multimode optical fibers were used. These are made of Poly(methyl methacrylate) (PMMA, acrylic) with 1 mm diameter thickness, refractive index of 1.49 and a fluorinated polymer shell. It has numerical aperture (NA) of 0.50, acceptance angle of 60°, att enuation of 0.70 dB/m at 650 micrometer and an operating temperature from -40 °C to 700 °C [19].
Apart the sensing head attached to the guiding fiber, a second head, without radiation-sensitive material, was prepared in order to quantify the background noise (mainly Cerenkov light produced during exposure). The same procedure is widely used by other researchers [2,15] in order to enhance the system accuracy.
The coupling region of the guiding fiber and the dosimeter head was also exposed to ultraviolet (UV) for curing the optical polymer and fixing the fiber in
the capillary. In order to minimize environment spurious optical signal, all guiding fiber exposed to the clinical beams was isolated from the external environment using a vinyl black cover tape [21].
An EMI photomultiplier (model 9829B, from Thorn-EMI GENCOM Inc.) with a spectral range from 300 to 650 nm was used. The tube diameter is 52 mm (active diameter 46 mm) with a sequence of 12 dinodes (stage) linearly focused with typical noise (dark current maximum) of 3 x 10-9 A. For measuring the current supplied by the PMT we used an electrometer (Model 6514, Keithley Instruments).
The dosimeter characteristics and performance were analyzed using a 6 MV CLINAC 2100C Linear Accelerator photon beam from Varian Inc. Measurements of reproducibility and linearity were obtained using a field size of 10x10 cm2, a source-skin distance (SSD) of 100 cm and depth of 5.0 cm. Some parameters related to the beam, such as PDD and crossplane dose profile were also analyzed using a field size of 10x10 cm2 and 20x20 cm2, respectively.These data were compared with reference measurements provided by the Blue Phantom dosimetry system of Wellhöfer Dosimetrie GmbH - Scanditronix Medical AB. This system uses an Ionization Chamber FC65-G (Farmer type chambers) of 6.2 mm inner diameter, 0.65 cm3 active volume and wall material such as graphite [22].
Results
The transmittance and the luminescence spectra of the scintillator mixture is show in Fig. 3a and Fig. 3b, respectively. As can be observed by Fig. 3a the material strongly absorbs in wavelengths shorter than 400 nm and, from there, transmits with high efficiency as confirmed by Fig. 3b and by previous work of Bushuk [23]. In Fig. 3b it is also possible to observe that the addition of optical polymer in anthracene particles causes an enlargement of luminescence spectra as an increase of its intensity. These results confirm the choice of a PMT with spectral response between 300 and 650 nm.
The next characterization step was to expose the anthracene dosimeter to clinical beams (CLINAC). Reproducibility tests were performed once a week, during 10 weeks, with set of 30 readings. Analysis of variance (ANOVA) performed was not able to demonstrate that the groups of measurements do not obey the same phenomenon. This indicates that such dosimeter is stable over long periods of time.
Fig. 4a shows the characteristics curves of dose rate (80, 160, 240, 320 and 400 MU/min) for several Monitor Units (MU). It was observed that for each curve there is a time before stabilization of the measurement system. After that, it is possible to measure the exactly dose rate of the CLINAC.
In Fig. 4a, for each dose rate curve it is possible to approximate the readings into two linear fittings (Current=A+B*MU), before and after the stabilization of the measurement system. From the intersection point was calculated that the mean time for stabilization is 6.2 s with a relative standard error of 1.6 %. The mean values after the stabilization of the measurement system is shown in Fig. 4b. Its behavior confirms the linearity of the dosimeter readings, i.e., the amount of scintillating light increases linearly as MU and dose rate increases [24]. From the linear coefficient it is possible to obtain a dark current of 2 nA for the PMT. This value agrees with the expected value for a PMT.
Fig. 5a shows a PDD curve performance comparative of our scintillator and a reference ionization chamber. Both curves are in agreement, indicating that the dosimeter was capable of properly characterize the beam behavior in the build- up region (peak at 1.5 cm). The PDD curves were normalized to their doses at 1.5 cm from the water surface.
The crossplane dose profile of the CLINAC beam is shown in Fig. 5b. The dosimeter performance agrees with the reference chamber and both curves were normalized to their doses from the central axis.
Discussion
Quality assurance complexity in radiotherapy treatment is due to present-day needs of treatments. Most recent treatment techniques such as IMRT and radiosurgery require dosimeters with small dose-sensitive volume to prevent dose averaging in high-dose-gradient regions. Ionization chambers, like the FC65, have a dose-sensitive volume of about 600 mm3 that is too large for such measurements. To overcome this was developed a dosimeter with only 6 mm3 of dose sensitive volume.
Our results have shown that the developed scintillator mixture increases the signal detection efficiency in comparison with pure anthracene. Note that the optical polymer used and the anthracene presents close refractive index (1.56 and 1.62, respectively) resulting in a more homogeneous and transparent mixture, decreasing the luminescence scattering and increasing the detection efficiency when compared to a medium composed of air (refractive index equal to 1) and anthracene. For reducing the background the measurement system (PMT plus electrometer) was kept outside of the irradiation room. However, the Cherenkov radiation produced inside the fiber lowered the signal noise ratio. There are several reports in the literature of some techniques to remove this noise [25,26,27,28]. One possibility is to subtract the Cherenkov radiation using a reference fiber. All these efforts resulted in a very low dark current (2 nA) for the PMT when the system is in standby.
A set of procedures to ensure that the system is reliable enough for clinical radiation detection is necessary. To accomplish this, the dosimeter reproducibility was evaluated with ANOVA test. Dose rate and stabilization time were also characterized in order to determine the sensor linearity and, consequently, the dark current of the PMT. The percentage depth dose and the