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Abstract
Fiber dosimeters have recently drawn much interest for measuring in vivo and in real time the dose of medical radiations. This paper presents the first miniaturized fiber dosimeter integrated at the end of a narrow 125 μm outer diameter optical fiber. Miniaturization is rendered possible by exploiting the concept of a leaky wave optical antenna for interfacing the scintillators and the fiber and by taking advantage of the low propagation loss of narrow silica fibers and high detection yield of single-pixel photon counters. Upon irradiation at 6 MV in air, our fiber probe leads to a linear detection response with a signal-to-noise ratio as high as 195. Although implemented with inorganic scintillators and fiber, our miniaturized fiber probe induces minimum screening effects on ionizing radiations over a negligible area (0.153 mm2). Our nano-optically driven approach may thus result in ultra-compact fiber dosimeters of negligible footprint in the radiotherapeutic processes, even with non-water equivalent fibers and scintillators. This opens new opportunities for a large panel of therapies relying on ionizing radiations (photons or charged particles).
© 2019 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
Recent advances in radiation therapies have prompted the need for tools to accurately probe ionizing radiations at small scales, in confined environments, in real time and with compact invisible devices. Among possible technological strategies, the combination of scintillators and optical fibers has been foreseen to provide end-users with highly flexible dosimeters capable of a real-time in vivo X-ray monitoring [1–17]. In practice though, the low coupling efficiency between scintillators (also called phosphors) and optical fibers has thus far limited their experimental realization onto broad multimode optical fibers (of near-millimeter diameters), which represents limits for the technology. For instance, exigencies of probe invisibility with regards to the therapeutic process has imposed the use of water-equivalent plastic scintillators and fibers [1–7,9,16,18–20]. However, such probes suffer from degraded signal-to-noise ratio due to the existence of a spurious Cerenkov effect within the fiber [21,22]. Filtering out the Cerenkov signal is possible at the expense of a critical broadening of the probe (paired-fiber configuration) [3,21,23] or more-or-less complex signal post-processing [18–20,24].
Scaling down this on-fiber technology by approximately one order of magnitude, to make it integrated at the end of a narrow optical fiber of about the size of a human hair (80-125 $\mu$m), would provide new opportunities in the detection of ionizing radiations. In this paper, we demonstrate the first miniaturized fiber probe integrated at the end of a narrow 125 $\mu$m outer diameter optical fiber. The drop of detection efficiency accompanying the shrinkage of both the fiber and scintillating cell is compensated by implementing a key-interface in-between the luminescent material and the fiber. Based on the concept of optical antenna [25], such a photonic interconnection enables redirecting the radiation-triggered luminescence towards the fiber and efficiently phase-match the optical waves to the fiber modes.
Optical antennas have demonstrated performances in controlling the emission directionality of fluorescent optical sources [26]. Recently, the horn nano-optical antenna has been proposed to transfer the emission of a point-like dipole source into an optical fiber [27]. This antenna has then been successfully used to interface a scintillating cluster to a single-mode optical fiber, thereby leading to a micron size X-ray fiber probe [28]. Such a system however finds limits in probing the high-energy ionizing radiations used in medical therapies, due to the lower absorption of these radiations by scintillators (and matter in general). Here, we overcome limits using the concept of leaky-wave optical antenna [29–33] in a reciprocal optical approach.
Beyond gain of compactness, the probe miniaturization opportunity offered by such a nano-optically driven technology leads to near-invisible fiber detectors even with inorganic non water equivalent fibers and scintillators [34–36]. The use of narrow fibers offers other advantages such as lower propagation loss than plastic fibers and the use of mono-pixel photon counters which develop higher detection yield than the multi-pixel systems used with plastic fibers.
2. Principle
The fiber-mediated coupling process of luminescent material to a photodetector can be accurately described using reciprocity. If the detector is replaced by a light source at $\lambda$ ($\lambda$ is the vacuum wavelength), the intensity distribution at the output surface of the fiber in contact to the scintillators provides a reciprocal map of the coupling efficiency of the scintillators to the detector through the fiber. The regions of higher leakage reciprocally reveal the regions of higher detection efficiency. In this respect, any source of leakage at the fiber end is reciprocally converted in a coupling region of an incoming light into the fiber. Therefore controlling the leakage at the end portion of a fiber reciprocally enables optimizing the detection sensitivity of a fiber dosimeter based on that fiber.
Leaky-wave antennas belong to the family of non-resonant travelling-wave antennas [29]. They achieve broadband directional radiation in free space from a lossy waveguide. This concept has been recently extended to optics for achieving controlled emission from a tiny fluorescent source [31]. The loss channel was here provided by a high refractive index substrate in a "‘fast-wave"’ process. The leaky mode continuously loses energy upon propagation, which is released in the substrate by virtue of momentum matching [29]. Importantly, such a leakage mechanism is tailored by the leaky mode itself via its intrinsic dispersion properties.
Our concept of a leaky-wave optical antenna is shown in Fig. 1. It consists of a dielectric cone on top of a cleaved optical fiber, surrounded by a high permittivity medium (used as leakage medium). The fiber core and cladding and the surrounding medium are characterized by refractive indices $n_1$, $n_2$ and $n_3$, respectively. The cone refractive index and base diameter match those of the fiber core. Multimode fibers are considered, with core diameters ranging from 25 $\mu$m to 100 $\mu$m. Such fibers are known to propagate four types of modes, namely the $HE_{ij}$, $EH_{ij}$, $TM_{0j}$ and $TE_{0j}$ modes ($i$ and $j$ are integers) [37]. Figure 2 shows the effective indices ($n_{eff}$) of the $HE_{ij}$ and $EH_{ij}$ modes as a function of the core diameter, for core and cladding refractive indices of 1.46 and 1.44, respectively. We see that these modes reach cutoff at higher core diameters for increasing values of $i$ and $j$. According to that property, the fiber modes of higher orders (i.e., of lower refractive indices and larger cutoff diameters) will undergo maximum leakage immediately after entering the cone whereas the modes of lower orders (i.e., of higher refractive indices and smaller cutoff diameters) will show the higher leakage close to the cone apex. The leakage mechanism is thus tailored by the incoming fiber modes themselves via their intrinsic dispersion properties.
Fig. 1. Principle of our fiber probe: reciprocal approach. Our leaky-wave optical antenna is constituted of a cone surrounded by a high refractive index material. The structure is positioned at the end of the fiber. Light injected into the fiber is turned into a leaky wave within the optical antenna (total internal reflection is deliberately lost at the cone interface). Each incoming fiber mode leads to a leakage distribution along the cone surface defined by its dispersion properties. While increasing the orders of the fiber modes, the leakage occurs from the apex and move towards the basis of the cone. By properly designing such a leaky-wave optical antenna, the multimode nature of the fiber ensures a homogeneous leakage all along the cone, even if the structure is sharp and elongated. Reciprocally, the coupling area between a high refractive index luminescent material and the fiber is dramatically enhanced as compared to a cleaved fiber, thus leading to improved detection efficiencies.
Fig. 2. Dispersion properties of a multimode fiber with core and cladding indices equal to 1.46 and 1.44, respectively: effective refractive index $n_{eff}$ of the $HE_{ij}$ and $EH_{ij}$ fiber modes ($i$ and $j$ are integers) as a function of the core diameter ($\lambda$=600 nm). (a) $i=1$, (b) $i=10$, (c) $i=20$ and (d) $i=40$.
We numerically study this antenna concept using finite difference time domain method (FDTD) [38]. We perform simulations with the body-of-revolution 2D FDTD method. The wavelength is 600 nm in all calculations. We address a 50-$\mu$m core diameter multimode fiber with core and cladding refractive indices of 1.46 and 1.44, respectively. We model a 250-$\mu$m long cone surrounded by a continuous medium of refractive index equal to 2.3. The cone basis lies on a 30-$\mu$m long slice of the fiber (i.e., 50 wavelengths) in such a way that its circular basis covers the fiber core. The computation area extends up to 30 $\mu$m in the radial coordinate and spans -30 $\mu$m below the cone basis and terminates 40 $\mu$m beyond the cone apex. A uniform grid resolution of 50 nm is defined across the computation area. A single fiber mode $HE_{ij}$ is launched into the fiber from its free end facet. At the end of the fiber section, the guided mode is turned into a leaky mode which decays upon propagation within the cone and loses power in the surrounding medium. Figure 3(a) represents the leakage accumulated along the cone for impinging $HE_{11}$ and $HE_{47,3}$ modes within the fiber, respectively. We see that $HE_{47,3}$ leads to maximum leakage in the first half of the cone whereas the power coming from $HE_{11}$ leaves the cone in the last 70 $\mu$m of the structure, close to its apex. The large number of fiber modes in between these two extrema ensure a near-homogeneous leakage all along the cone, as shown Fig. 3(b). The cone antenna thus ensures a complete leakage of the fiber mode over a sub-millimeter distance along the fiber axis with a near-constant optical intensity along the structure.
Fig. 3. (a) Simulation by FDTD of the accumulated leakage (in fraction of the incoming power) along the cone axis $(0z)$ for the incoming $HE_{11}$ and $HE_{47,3}$ modes into the fiber. These two guided modes of the fiber are turned into leaky modes within the cone. The position $z$ along the cone axis given in the abscissa is defined in the figure inset. The dashed line represents constant leakage along the cone. (b) Positions along the cone axis $(0z)$ where 50 % (lower panel) and 90 % (upper panel) of the power of the incoming fiber modes $HE_{ij}$ has leaved out the cone, for $i$=1, 10, 20, 30 and 40. Integers $i$ and $j$ identify the incoming guided modes of the fiber. $\lambda$=600 nm.
Reciprocally, part of the light isotropically generated within the highly refractive medium will be incoupled into the fiber in a mirror-symmetric mechanism of the above-described mode leakage. As a result, the overall surface of the elongated cone will participate with near constant efficiency to the in-fiber light collection process. The light collection area of the fiber is thus extended by a factor of $h/R$ regarding the cleaved fiber, where h is the cone height and R is the core radius. In the configuration under consideration, the collection area is enhanced by a factor of 10. The fiber modes with the lower (higher, respectively) $i$ and $j$ indices will propagate the light mainly collected at the cone apex (at the cone basis, respectively).
The concept of a miniaturized fiber dosimeter emerges from considering the high refractive index luminescent medium of the antenna as being a scintillating material under irradiation (i.e., a high density ensemble of luminescent point-like sources embedded in a high refractive index semiconductor matrix). Given the scattering phenomenon within scintillation materials, we limit our analysis of the optical leakage/collection efficiency right at the cone surface.
3. Demonstration
3.1 Fabrication
The miniaturized fiber dosimeters are engineered onto narrow 125-$\mu$m outer diameter multimode fiber (core diameter of 100 $\mu$m). The fiber end facets are first etched in a buffered fluorhydric acid to obtain elongated cones. The cones are about 500 $\mu$m long (see Fig. 4(a)), leading to a cone surface about 10 fold larger than the surface of the fiber core. Next, scintillating powder is mixed with a home-made photopolymer which combines a multifunctional acrylate monomer and a dye sensitizer. The powder clusters are homogeneously embedded with a high concentration into the viscous photopolymer, leading to a final mixture with minimum polymer. This mixture is attached to the fiber cone by a photopolymerization involving an illumination via the fiber (i.e., light is injected from the other end facet of the fiber and hardens locally the polymer right at the fiber output). That way, the cone is selectively covered with a thin homogeneous layer of scintillating emulsion. We choose an inorganic scintillating material, europium-doped gadolinium oxysulfide (Gd$_2$O$_2$S:Eu) from Phosphor Technology, which generates visible light upon exposure to ionizing radiations (photons or charged particles). Such type of scintillators demonstrated good performances and linearity in the local probing of ionizing radiations in fibered architectures [10,14,39]. They are usually characterized by a refractive index of about 2.3. Figure 4(b) displays the optical image of a resulting miniaturized fiber probe. The scintillating material takes the form of a quasi-uniform 85-$\mu$m thick layer that selectively covers the cone. The maximum width of the probe does not exceed 250 $\mu$m, i.e., the full diameter of the optical fiber with its plastic cladding. As an option to increase detection efficiency, the probe can be coated with a thin aluminum layer (200 nm thick) aimed at back-reflecting the initially wasted light leaving the probe outside the fiber. To that end, a conventional metal deposition can be performed with a standard evaporation or sputtering technique. A few nanometer thick chromium layer should be deposited prior to aluminum coating to improve aluminum adhesion onto the probe. Aluminum is known for its good transparency to ionizing radiations (owing to its low atomic number Z=13) and high reflectivity to light. Therefore, spurious photoelectric effect and other energy down-conversion phenomena triggered by such thin aluminum layers can be neglected [40].
Fig. 4. (a) Microscope image of the tapered fiber. (b) Microscope image of the fiber probe obtained after the scintillator-to-taper attachment. (c) Schematics of the experimental set-up. (d) Emission spectrum of the europium-doped gadolinium oxysulfide powder (Gd$_2$O$_2$S:Eu) used as the probe scintillating material.
3.2 Experimental set-up
The miniaturized probes are engineered at one of the two ends of 9-meter long fibers covered with 0.9 mm black hytrel cladding, the other end being terminated with a FC-PC connector. Their detection performances are tested under a clinical linear accelerometer used for external radiation beam therapy (Model 2100 CD version 9.01 from Varian; Fig. 4(c)). The fiber probe is positioned on the treatment bed at the center of a $10 \times 10$ cm$^2$ or $5 \times 5$ cm$^2$ ionizing beam at a STD (Source-to-Target Distance) of 100 cm. Upon irradiation, europium-doped gadolinium oxysulfide shows a multi-peak emission spectrum within the bandwidth 580-710 nm (Fig. 4(d)). At the wavelength of maximum emission, propagation losses of the fiber (SEDI-ATI) are lower than 10.1 dB/km. The optical signal leaving the fiber was recorded with an Aurea Technology SPD-A-VIS single pixel photon counter located in the control room. This optical detector ensures a quantum yield larger than 60% within the emission bandwidth of the scintillators. Dose monitoring is realized with a computer connected to the photon counter. Miniaturization thus takes benefit of the better performances of narrow silica fibers and single-pixel photon counters regarding plastic fibers and the multipixel photodiodes used with plastic fibers, respectively.
3.3 Results
The response of a miniaturized fiber dosimeter is determined in air for a field size of $10\times 10$ cm$^2$ and a normal X-ray photon energy of 6 MV. We thus test our probe to the primary irradiation beam emerging from the medical source. Figure 5(a) shows the detection signal upon irradiation at a dose rate of 600 MU/min with a total dose of 600 MU (Monitory Units). In this investigation, 1 MU corresponds to an absorbed dose of 1 cGy in a water-equivalent environment under standard reference conditions. The intrinsic integration time of the photon counter is set to one second (sampling rate of one signal value per second), which is compatible with a real-time monitoring of the irradiation power. At the beginning and at the end of the acquisition, the irradiator is off to estimate a background (dark) noise level of the photon counter smaller than 15 counts/s. When the radiation beam is activated, a signal larger than 429 kcounts/s is measured with a signal-to-noise ratio of 195. Although such a signal-to-noise level is high enough to perform accurate dose monitoring, the observed signal fluctuations remain 2.98 fold higher than the intrinsic detection noise of the photon counter at this photon flux. Such a discrepancy may be explained by high-frequency fluctuations in the irradiation beam. Higher signal-to-noise ratio could be obtained by increasing the integration time of the photodetector to a few seconds (two or three seconds to ensure real-time acquisition). Figure 5(a) also shows a Cerenkov signal (measured with a blank fiber, see inset) of 18.1 kcounts/s, that is 23 fold smaller than the total luminescence signal. Even small, Cerenkov signal cannot be neglected in the probe detection process. It can however be easily filtered out in a parallel paired-fiber configuration [3,21,23]. In this case, the use of narrow optical fibers would solve problems of probe compactness encountered with plastic fibers in intracorporal dosimetry. The resulting two-fiber dosimeters should indeed remain at least two times narrower than the single-fiber probes developed so far [14–16]. More complex signal processing aimed at removing Cerenkov signal in a single-fiber dosimeter architecture are also possible [18–20,24] .
Fig. 5. Probe response in air. (a) Detected optical signal (in kcounts/s) over time for an exposure at a beam energy of 6 MV and at a dose of 600 MU with rate of 600 MU/min (the dose is defined in water under standard reference conditions). Blue curve: detected signal from the fiber probe. Red curve: Cerenkov signal from the unprocessed bare fiber. Figure insets: Diagrams of the two detection configurations. (b) Cumulated optical signal during a 100-second exposure (corresponding to a total delivered dose of 10 Gy in water under standard reference conditions). (c) Cumulated optical signal as a function of the dose rate. Red squares: over 1 minute exposure, blue circles: over the time needed to deliver 1 Gy in water under standard reference conditions. (d) Detected optical signal for 10 exposures of 22 seconds at 600 MU/min (field size: $5 \times 5$ cm$^2$). Lower panel: detected signal over time. Upper panel: Squares: average signal at each acquisition, dashed lines: Average signal over the 10 exposures and standard deviation (0.12 % of the average signal).
As seen in Fig. 5(b), the time-integrated optical signal shows a constant enhancement of the applied dose during irradiation, which agrees with the constant applied dose rate (600 MU/min in water under standard reference conditions). Figure 5(c) reports the cumulated optical signal as a function of the dose rate. Two cases are considered: a constant exposure time of one minute for all dose rates and an exposure time chosen for each dose rate so that 1Gy (100 MU) is deposited in a water-equivalent medium (under standard reference conditions). This result reveals a linear response of the dosimeter regarding the power of the incoming irradiation beam.
The repeatability of the dose measurement with our fiber dosimeter is confirmed in Fig. 5(d) for a field size of $5\times 5$ cm$^2$. The repeatability was tested over ten exposures of 22 seconds at a dose rate of 600 MU/min with a radiation energy of 6 MV. The detected signal reveals ten plateaus of almost constant signal described by an average value of 364.88 kcounts/s and a relative deviation of 0.12 % of the average value.
6MV radiation is thus detected with a large signal-to-noise ratio and a linear response to both the dose rate and accumulated dose. As a comparison, our level of detected signal is only 1.37 fold smaller than that a former Gd$_2$O$_2$S:Eu based probe built at the end a 0.5-mm core diameter plastic fiber [14]. Such results are prerequisite to the use of our probe for medical applications. Our probe will however need to be specified at least in a water-equivalent environment. Water and materials of equivalent density are good models of biological media. As already evidenced in Ref. [14], the presence of a water-equivalent medium enhances the signal-to-noise ratio of fiber-integrated optical detectors.
4. The benefit of miniaturization: ultra-low footprint fiber probes
Beyond biocompatibility, in vivo X-ray dosimeters must be of negligible footprint for the patient and the therapeutic process, i.e., ultra-compact and invisible to the therapeutic process. It has been reported by Papanikolaou et al that a 5% change in dose may result in a 10 % to 20 % change in tumor control probability at a TCP of 50 %, as well as in a 20 % to 30 % impact on complication rates in normal tissues [41]. Although these results refer to changes caused by homogeneous dose distributions covering the whole tumor, they stress the need to develop detection tools with a minimum absorption of the incoming radiations, especially in the treatment of small tumors.
Figure 6 reports an estimation of the radiation absorption by our fiber dosimeter, when envisaged as a medical probe. To this end, a model of the probe is elaborated from its optical images (see Figs. 4(a) and (b)), the Beer-Lambert law and the X-ray mass-attenuation coefficients of the probe’s constitutive materials [40]. The geometry of our probe model is that of a tapered silica cylinder with a diameter of 125 $\mu$m. The taper angle is 12.4$^{\circ }$ consistent with the angle of the fabricated fiber tip of Fig. 4(a). The cone and the very end of the cylinder are covered with an axis symmetrical layer of gadolimium oxysulfide whose thickness distribution leads to the probe shape of Fig. 4(b). The probe is considered being irradiated from the side, leading to an attenuation area which does not exceed 0.153 mm$^2$.
Fig. 6. (a) and (b) Mappings of the probe-induced intensity contrast of an incoming radiation beam (a) in air and (b) in water. The radiations impinge from the probe side. (c) Solid lines: spectra of the beam intensity contrast calculated at a single point of the fiber probe positioned in water (red curve) and in air (blue curve). This point is identified with the letter A in (a) and (b). The beam energy ranges from 0.4 MV to 6 MV. Dashed lines: spectra of the maximum beam intensity contrast induced by a couple of narrow 125-$\mu$m outer diameter optical fibers.
Figures 6(a) and (b) show mappings of the probe-induced attenuation of a monochromatic radiation at 1MeV, both in air (Fig. 6(a)) and in water (Fig. 6(b)). These two figures represent intensity contrast after the probe, defined as $T(x,y)=(I_0(x,y)-I(x,y))/I_0(x,y)$ where $I(x,y)$ and $I_0(x,y)$ are the 2D radiation intensity distributions with and without the fiber probe, respectively. In air, the probe-induced intensity contrast peaks at 1.08 % and shows values exceeding 1% over an area of only 0.005 mm$^2$ (the total attenuation area is 0.153 mm$^2$). The narrow optical fiber alone leads to an intensity contrast lower than 0.18%. In a water-equivalent (biological) medium, the maximum attenuation of our fiber probe and of the fiber alone falls down to 0.9 % and 0.09 %, respectively. Therefore, our fiber probe can be considered of negligible footprint at 1 MeV, both in air and in a biological environment.
Medical irradiators are polychromatic sources and decay channels in matter originate a secondary radiation over a broad spectral range. In that context, we also evaluated the probe-induced intensity contrast over a spectral bandwidth ranging from 0.4 MeV to 6 MeV. Figure 6(c) reports spectra in air and in water at a single point A represented in Figs. 6(a) and (b). We see that the probe-induced radiation attenuation remains under 1 % at energies larger than 0.84 MeV and 0.67 MeV in air and in water, respectively. We also show weak beam attenuation through a couple of parallel fibers (possible configuration for filtering Cerenkov signal). At 0.4 MeV, we anticipate an intensity contrast peaking at 0.53 % in air and 0.26 % in water. One can predict from Fig. 6(c) a sharp increase of the probe-induced attenuation at energies smaller than 0.4 MeV. Lower energy ionizing radiations are emitted by water-equivalent materials themselves upon irradiation. With an attenuation area of 0.153 mm$^2$ and a fiber diameter of 0.125 mm, our miniaturized fiber probe ensures a minimum screening cross-section and thereby a minimum perturbation of the therapeutic processes.
These results show that the provided miniaturization opportunity leads to fiber probes of negligible footprint, even with non water-equivalent scintillators and fibers. This greatly relaxes exigencies in the probe design.
5. Conclusion
We demonstrate a miniaturized fiber probe for locally measuring in real-time the dose of medical ionizing radiations (photons and possibly charge particles). Such a probe relies on the concept of a leaky-wave optical antenna developed at the end of a narrow 125-$\mu$m outer diameter optical fiber. The antenna concept is used as a key-connection between the fiber and the scintillators aimed at enhancing the in-fiber coupling of the radiation-triggered luminescence. Miniaturization also takes benefit of the low propagation loss of narrow silica fibers and high detection efficiency of single-pixel photon counters.
The advantage of miniaturization is here threefold. First, the scintillating material being spread along the fiber axis in the form of a thin film (of micrometer thickness), probe-induced attenuation of ionizing radiations remains at a very low level even with non water-equivalent scintillators and fibers. Second, the scintillating area (of maximum absorption) is only a tenth of a square millimeter. Finally, the Cerenkov signal could be easily filtered out in a parallel paired fiber configuration which remains ultracompact and compatible with intracorporeal dosimetry. Miniaturization thus provides a new strategy to reach in vivo fiber dosimeters of negligible footprint for the patient and the therapeutic process, even with non water-equivalent scintillators and fibers.
Our concept of a miniaturized fiber dosimeter opens new perspectives for radiation dose monitoring and may thus enable unprecedented control in up-to-date and future radiation therapies. As an example, ultracompact and ultralow footprint multiprobe fiber dosimeters may emerge from a parallel implementation of our fiber probe onto a bundle of fibers of different lengths. The opportunity of engineering one probe per fiber would avoid cross-talk between parallel acquisition channels as compared to existing multi-probe devices sharing the same fiber (due to compactness limitations inherent to the use of broad plastic fibers) [42]. Such multiprobe dosimeters would be of crucial importance in the in vivo localization and metrology of sharp dose gradients. Note that our preliminary probe characterization realized in air onto the primary irradiation beam will need to be completed by tests in a water-equivalent environment to comply with medical protocole. Preceding works on similar but larger probes [14] provides promising perspectives for our device in intracorporal dosimetry.
Our fiber probe is thus expected to fulfill the increasing requirement of accuracy in medical dose delivery. It is compatible with the current image guidance techniques for target positioning (such as MRI [43,44]). Owing to their unprecedented spatial resolution, they seem to be well adapted to all current radiation therapies as well as to the newborn microbeam radiation therapy.
Funding
CNRS DEFI Instrumentation aux limites 2018; EIPHI Graduate School (ANR-17-EURE-0002); Conseil régional de Bourgogne-Franche-Comté; Agence Nationale de la Recherche (ANR-18-CE42-0016).
Acknowledgments
The authors are indebted to the company SEDI-ATI for helpful discussions. This work was conducted in close connection with this company in the purpose of probe optimization and encapsulation.
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