Development of a Gaseous Electron-Tracking
Compton Imaging Camera
Using LaBr3:Ce Scintillator
and Its Application to Proton Therapy

Shunsuke Kurosawa
Department of Physics, Faculty of Science, Kyoto University
Kitashirakawa Oiwake-cho, Sakyo-ku, Kyoto, 606-8502, Japankurosawa@cr.scphys.kyoto-u.ac.jp

Abstract
MeV and sub-MeV gamma-rays imaging has been important for several scientiﬁc ﬁelds suchas astronomy and medical imaging. However, present imaging detector works with a lower
sensitivity than those in other energy range, because it is diﬃcult to reconstruct Compton
scattering which is a main interaction of photons with matter.
We have developed an electron-tracking Compton camera (ETCC) which consists of gaseoustime projection chamber (TPC) and a position-sensitive scintillator. Since the gaseous TPC
measures the electron-tracking, the ETCC is able to reconstruct Compton scattering event-byevent.
This camera is expected to be applied to above ﬁelds. Actually, we have already succeeded in
imaging the cosmic and atmospheric gamma rays using our small size camera onboard a balloon
in 2006. Then we have developed a larger camera in order to detect the bright celestial objects
such as Cyg X-1.
In addition, we have developed the ETCC for the medical imaging for a diagnosis of a tumor,
and the ETCC already have imaged small animals with agents including radio isotopes. Our
ETCC is also expected to use in proton therapy, which is the one of the radiation therapies.
Since charged particles have the sharp peak of the energy deposit (Bragg peak) near the endof the range, the tumor is killed eﬃciently. On the other hand, the normal cells around thetumor are less damaged than X-ray radiation therapy. Here, optimization of the proton range isdecided by only the calculation before the therapy (planning step), while no monitoring system
usually veriﬁes the proton dose during the therapy.
During the proton therapy, prompt gamma rays are generated, and those prompt gamma
rays are reported to trace the Bragg peak well.
Here, we demonstrated the imaging ability of the ETCC for those prompt gamma rays using
water phantom irradiated with 140-MeV proton beam at Research Center for Nuclear Physics(RCNP) of Osaka University in 2010. The aim of this experiment is to detect the diﬀerence of the
distributions between the annihilation and prompt gamma rays. As a consequence, we succeededin obtaining diﬀerent images between annihilation and prompt gamma-ray distribution, which isconsistent with simulations. Thus we found that the ETCC has an imaging ability for continuum
gamma rays, and this ability is useful for both medical imaging and astronomy.
The angular resolution of the ETCC depends on the energy resolution of absorber of the
ETCC, and imaging performance for gamma rays strongly depends on the angular resolutionof the ETCC. At present, a GSO:Ce array camera is used as an absorber, which has an energy
resolution of 10.8 ± 0.3 % at 662 keV (FWHM). LaBr3 :Ce scintillator has an excellent FWHM
energy resolution of ∼ 3 % at 662 keV for a monolithic crystal. Although this scintillator has ahygroscopic nature, we developed LaBr3 :Ce arrays consisting of 8 × 8 pixels with a size of 5.8

mm ×5.8 mm ×15.0 mm. The FWHM energy resolution of all 64 pixels reached 5.7 ± 0.5 % at
662 keV. Using the LaBr3 :Ce arrays, an ETCC has an FWHM angular resolution of 4.2 ± 0.3◦ at
662 keV. In addition, we estimated an FWHM angular resolution of the array would be ∼ 3.5◦
at 662 keV, if depth of interaction technique is used in the pixel scintillator.

2

Contents
1 Introduction

1

2 Science of MeV Gamma Ray

7

2.1

Interactions of Photons with Matter . . . . . . . . . . . . . . . . . . . . . . . . .

7

2.2Interactions of Charged Particle with Matter . . . . . . . . . . . . . . . . . . . .

10

2.3

Mechanism of Gamma-ray Radiation . . . . . . . . . . . . . . . . . . . . . . . . .

12

2.4

Sub-MeV and MeV Gamma-ray astronomy . . . . . . . . . . . . . . . . . . . . .

15

2.5

Medical Imaging and Proton Therapy . . . . . . . . . . . . . . . . . . . . . . . .

21

3 Compton and Other Gamma-Ray Imaging

28

3.1

MeV Gamma-Ray Imaging Detector in Astronomy . . . . . . . . . . . . . . . . .

28

3.2

MeV Gamma-Ray Imaging Detector in Medical Imaging . . . . . . . . . . . . . .

34

3.3

Electron-Tracking Compton Imaging . . . . . . . . . . . . . . . . . . . . . . . . .

37

4 Gaseous TPC and Scintillator

41

4.1

µ-PIC (Micro Pixel Chamber) . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

41

4.2

Gas Electron Multiplier . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

45

4.3

Time Projection Chamber with µ-PIC (µ-TPC) . . . . . . . . . . . . . . . . . . .

46

4.4

Scintillation Camera . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

50

5 Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays inProton Therapy

56

5.1

Veriﬁcation of Dose in Proton therapy . . . . . . . . . . . . . . . . . . . . . . . .

56

5.2

Setup of the Beam Experiment Using the Compton Camera . . . . . . . . . . . .

60

5.3

Results of the Beam Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

69

5.4

Analysis of Imaging . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

73

5.5

Future Works . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

82

6 Improvement of Scintillator

86

6.1

Position-Sensitive Scintillation Camera . . . . . . . . . . . . . . . . . . . . . . . .

86

6.2

Optimizing the Scintillator and the Pixel Size . . . . . . . . . . . . . . . . . . . .

88

6.3

ARM Calculation for the Selection of a Scintillator and the Pixel Size . . . . . .

88i

CONTENTS
6.4

LaBr3 :Ce Camera . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

91

6.5

Materials for LaBr3 :Ce Array Camera . . . . . . . . . . . . . . . . . . . . . . . .

93

6.6

Assembly of LaBr3 :Ce and GSO:Ce arrays . . . . . . . . . . . . . . . . . . . . . .

96

6.7

Performance of Array Camera . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 100

6.8

Improvement of a LaBr3 :Ce Array . . . . . . . . . . . . . . . . . . . . . . . . . . 104

6.9

Comparison with Other Position-Sensitive LaBr3 :Ce Camera . . . . . . . . . . . 111

6.10 Performance of the Compton Camera with a LaBr3 Array . . . . . . . . . . . . . 112
6.11 Possibility of Improvement in Angular Resolution . . . . . . . . . . . . . . . . . . 115
7 Conclusion

117

ii

Chapter 1

Introduction
Recently sub-MeV and MeV gamma-ray imaging detector has been an important device in theseveral scientiﬁc ﬁelds such as astronomy, and nuclear medicine (e.g. [1, 2]). As for astronomy,
an imaging detector in this energy band is expected to reveal a lot of interesting celestial objects
such as black holes, gamma ray bursts (GRBs), and active galactic nuclei (AGNs). As for
medical ﬁeld, gamma-ray imaging is widely used in the diagnosis for searching a tumor with anagent.
However the quality of MeV gamma-ray imaging looks certainly less satisﬁed for its applications due to its low sensitivity, a poor angular or position resolution than imaging detectors
used in other energy ranges. The main reason of the low sensitivity is that a main interactionof a photon in the matter is Compton scattering, which is a smallest cross section in the energy
range of a few hundred keV to a few MeV. In order to obtain the direction of MeV photons, wemust reconstruct the Compton scattering event-by-event, for which a Compton point, directions
and energies of the scattered gamma ray, and those of the recoil electron have to be measured.
However it is diﬃcult to measure those all parameters, especially the direction of the recoil
electron.
For MeV gamma-ray astronomy, several groups have so far developed imaging detectors
using the reconstruction of Compton scattering, named a Compton camera. The Compton
camera consists of a converter where the Compton scattering occurs, and an absorber where thescattered gamma-ray photon is measured. However such Compton cameras do not measure thedirection of electrons in the parameters required for the reconstruction of Compton scattering,and thus localize the incident direction of a gamma ray only to an event circle as shown in Fig.
1.1 (a-1). Therefore they, written as ”classical Compton cameras”, must use three gamma-ray
event circles at least in order to obtain the real direction of the point source as shown in Fig. 1.1(a-2). In an actual observation, a lot of events should be accumulated to acquire some imaging
quality. In addition, classical Compton cameras have little ability to reject the backgrounds
like those coming from out of the ﬁeld of the detector. In particular, in space, cosmic rays
interact with the vessel of the satellite or the atmosphere, and then a lot of radioisotopes (RIs)
are generated. Those RIs emit huge MeV gamma rays as much backgrounds for the detector.
1Chapter 1. Introduction
For example, a classical Compton camera named COMPTEL detector onboard the Compton
Gamma Ray Observatory (CGRO) satellite had provided various scientiﬁc results. However thelarge backgrounds deteriorated dramatically its imaging quality. As a consequence, COMPTEL
had a less than half sensitivity (∼ 100 mCrab) of that designed before a launching [3], andonly 32 objects were detected with the COMPEL in the MeV region [4], while 271 objects
were detected with a GeV gamma-ray imaging detector [5], called the Energetic Gamma Ray
Experiment Telescope (EGRET) on the CGRO.

(b)

(a-1)
Incident γ
converter

Incident γe-absorber

Scattered γ(a-2)

Circle 1

convertere-absorber

Scattered γ

Circle 2
Circle 3

Figure 1.1: Schematic view of classical and advanced Compton camera as shown in (a-1) and(b), respectively. To determine a source direction with the classic one, three events must be
measured and a cross point of obtained three circles is regard as the solution indicted as an X
in Fig. (a-2).

Other types of imaging detectors using a collimator or a coded mask were developed andwidely used in the energy regions of sub-MeV and MeV [6, 7, 8, 9, 10, 11]. Several groups
succeeded in the observations in space using these detectors. However both detectors achieved
a less sensitivity than the classical Compton camera in the sub-MeV and MeV region, because
they do not select the signal photon or reject the noise one event-by-event by the measurementof the gamma-photon direction. As a result, the imaging by those detectors needs a lot of
gamma photons and consequently degrades its sensitivity. Also the dynamic ranges of thesedetector are limited under a few hundred keV owing to i) diﬃculty in detecting events with
suﬃcient statistics, and ii) Compton scattering or transparency for high energy gamma rays2

in the collimator or coded mask. In addition, in order to remove backgrounds coming fromthe outside of the ﬁeld of view, a veto counter surrounding the detector such as a scintillator
is indispensable. The veto counters which are usually heavier than the detector itself, are adisadvantage of the observation in space.
For medical imaging, the collimator technique is widely used in single photon emission computed tomography (SPECT) [12, 13]. For the SPECT imaging, an agent including gamma-ray
emitting RI is used as a tracer which is injected into the body. Usually the agent concentratingin the active tissues such a tumor are used. In order to obtain a ﬁne image, the ﬁne collimationis necessary. However its available energy range is limited to approximately 80 - 300 keV due to
its transparency for higher energy gamma rays.
An imaging detector specialized to the 511-keV annihilation gamma rays named a positronemission tomography (PET) has been recently popular on biological or medical imaging [14, 15].
By surrounding the target, the PET measures the back-to-back gamma rays, from an agent
including a positron emitting RI in the body. Although the PET has a good spatial resolutionof 0.5 cm for human imaging, the available RIs are limited to only positron emitting RIs.
In the energy range of sub-MeV gamma ray region, only the PET and SPECT are used as
the medical imaging instruments, while in the MeV region, no instrument is used. Thus anavailable energy ranges for the medical imaging are restricted. However, there are a lot of RIs
out of those energy ranges, and the RIs may have unknown potential to use as a new type of
the agent.
For the treatment, gamma-ray imaging has not been used usually, but expected to be quite
useful in the radiation therapy. Bennett et al. suggested the use of the PET imaging for the
proton therapy [16]. Proton therapy is one of the charged beam radiation therapy using the
Bragg peak, under which tumor cells are killed [17]. Here the proton energy is adjusted so thatits Bragg peak position coincides with the tumor. This therapy is able to reduce unwanted side
eﬀects for the normal cells under the beam line. Although the optimization of the Bragg peak
position is calculated before the treatment (planning), no detector measures the proton dose
directly during the beam irradiation. In the planning step, the eﬀect of the irradiation is usuallyveriﬁed by an X-ray computed tomography (CT) [18, 19]. However, the X-ray CT measuresthe electron density in the body. Then the optimizing calculation sometimes does not lead the
Bragg position accurately when the treatment area includes complicated bone structures. At
present, there is no method to measure the energy deposit of the proton beam around the tumor.
Therefore it is necessary to monitor the proton beam range in the patient body during proton
beam therapy.
Recently several groups have shown that the PET may be used for the determination of theradiation dose distribution in the patient after the proton therapy using the positron emitter inthe body generated by the proton beam [20, 21]. However, the distribution of generated positron
emitters in the patient is not considered to match well with the distribution of the radiationdose by the beam. The reason is that the positron emitters are more eﬃciently generated by3

Chapter 1. Introduction
the higher-energy protons than ∼ 10 MeV. Moreover, the positron emitters such as 45 Ca tend tobe accumulated in the bone, and the distribution of annihilation gamma rays from 45 Ca might
trace the bone. In addition, the PET imaging cannot be operated during the beam-on time
under the high counting rate due to the backgrounds.
During the proton beam radiation, prompt gamma rays are also generated by nuclear reaction. The simulation indicates that distribution of prompt gamma rays near the proton range
is correlated with the proton dose, and the imaging of prompt gamma rays is expected to indicate the accurate dose distribution in the body [22]. Especially MeV gamma rays tend to becorrelated with the proton dose than sub-MeV gamma rays. Thus, a novel detector is required
to have ability of the imaging of continuum gamma rays above 1 MeV under the beam-on.
As mentioned above, MeV gamma-ray imaging is eagerly expected in several scientiﬁc ﬁelds.
To obtain both a higher sensitivity and the ability of a good background rejection, we havedeveloped an electron-tracking Compton camera (ETCC) which consists of a gaseous time projection chamber (TPC) and a scintillator array as shown in Fig. 1.2 [24, 25, 26, 27]. The TPC
has a sensitive volume of 10 cm ×10 cm ×10 cm or 30 cm ×10 cm ×30 cm [28, 29, 30], whichis equipped with a gaseous avalanche detector consisting of a gas electron multiplier (GEM)[31, 32] and a micro pixel chamber (µ-PIC) [33, 34]. When an incident gamma ray is scattered
via Compton process in the TPC, both the three-dimensional track and energy of a Compton
recoil electron are measured [28, 29]. At the same time, the scintillator array measures theenergy and position of the scattered gamma ray. By combining these measurements, the ETCC
is able to reconstruct a full process of Compton scattering event-by-event, which enables us torestrict the incident direction of a gamma ray on the arc (not circle) region as shown in Fig. 1.1(b). In addition, the ETCC has a wide ﬁeld of view (∼ 3 str) and wide energy dynamic range(∼ 0.1 -a few MeV). Moreover the ETCC can reject backgrounds eﬃciently using a geometricalﬁtting for an angle between a recoil electron and a scattered gamma ray.
Thus the ETCC has several unique features, and hence we have continued a developmentof the ETCC. Our goal for astronomy is to explore the all sky with a higher sensitivity than
COMPTEL by an order of magnitude. As a ﬁrst step, we had developed a small ETCC including
a 10-cm squared µ-PIC successfully, and observed cosmic and atmospheric gamma rays using aballoon in 2006 [35, 36]. Then we are improving a larger ETCC with a higher sensitivity in orderto detect the bright celestial objects such as Crab and Cyg X-1 by a balloon-born experimentplanned in 2012.
In addition, I analyzed which component of the ETCC has a dominant eﬀect to improve theangular resolution. Then based on its results, I improved the energy and angular resolution of
the ETCC by developing a new scintillator array, and the angular resolution was expected tobe better than COMPTEL. This novel scintillator camera was successfully assembled into thearray with my own technique.
Since the ETCC is also surely expected to be useful for medical imaging, and we havecarried out the performance test of the ETCC such as spatial resolution, and the energy range
4

error region(arc)absorberconverter

PMT (photo multiplier tube)

Figure 1.2: Schematic view of the ETCC (advanced Compton camera).

for medical use [37, 38]. We have obtained gamma-ray images with a high energy RI up to
approximately 1.3 MeV (e.g., 1.116 MeV from 65 Zn), and multi-RI tracers for small animals.
In proton therapy, the monitor of the Bragg peak position and the dose in situ are required.
The ETCC looks a suitable imaging detector measuring prompt gamma rays, and one of the
purposes of this thesis is to demonstrate the feasibility of the ETCC for it. We carried out anexperiment on the gamma-ray imaging of the ETCC using continuous gamma rays including
prompt gamma rays from a water tank irradiated with 140-MeV proton beam at ResearchCenter for Nuclear Physics (RCNP) of Osaka University in 2010. This beam test was the ﬁrst
experiment as a continuous gamma-ray imaging with a Compton camera, and it provided two
important suggestions. One was that the ETCC is able to use as a real-time monitoring systemof proton therapy. The other was that the ETCC has an advantage for MeV gamma astronomy
because most objects in the universe emit a continuum gamma ray spectrum.
The outline of this thesis is as follows. In Chapter 2, I describe the interactions of photons
with matter, and sciences of astronomy and medical imaging. In particular, I show the detail ofproton therapy in the section of medical imaging.
Chapter 3 and 4 are expositions about imaging detectors. In Chapter 3, I illustrate a history
and mechanism of the MeV gamma-ray imaging detectors including Compton camera and otherdevices. In Chapter 4, the structure of the ETCC including a gaseous TPC and scintillator aredescribed.
In Chapter 5, the development of the real-time monitor for proton therapy is shown. In
5Chapter 1. Introduction
addition, because of the wide energy dynamic range, our ETCC can obtain an image with
continuum spectral gamma rays. Then I demonstrate an image from a water tank irradiatedwith proton beam.
I have developed a new position-sensitive scintillation camera using a LaBr3 :Ce crystal inorder to improve the spatial resolution of an ETCC. In Chapter 6, I explain why I have developed
the new scintillation camera with a higher energy resolution, and the detail of development andperformance of the LaBr3 :Ce camera. In addition, the performance of the ETCC using the
LaBr3 :Ce is shown.

6

Chapter 2

Science of MeV Gamma Ray
2.1

Interactions of Photons with MatterIn this and next sections, I describe the interaction of photons and charged particles with thematter, respectively. The main interactions of high energy photons from X-rays to gamma rays
with the matter are photo-absorption, Compton scattering and pair production as its energy
increasing. The cross sections of these interactions depend on the energy of photons and theatomic number Z of the matter. Figures 2.1 and 2.2 show each cross-section of the interaction
with Argon, and the dominant interaction as a function of the photon energy.
In less than 100 keV region, the photoelectric absorption is a dominant interaction. This
process provides the energy of a photon into an electron in the atom, and ejects the electron,
called photoelectron, from the atom When a photon with an energy of E0 makes a photoelectron,
the kinetic energy of the photoelectron is as follows,
Ke = E0 − Ebind ,(2.1)

where Ebind is the binding energy of the photoelectron in the atom. If E0 is higher than thebinding energy of the K shell, the probability of the interaction with an electron in K shell is
maximum. Then the cross section of the photoelectric absorption of a K shell electron is:√σK = 4 2Z 5

(

8π 2
r
3 e

)(

e2
4πε0 ~c

)4 (

me c2
E0

)7/2
,

(2.2)

where re is the classical electron radius:
re =

e2
.
4πϵ0 me c2(2.3)

In the energy band from a few hundred keV to 10 MeV, Compton scattering is the dominantinteraction. this scattering is an elastic scattering of a gamma ray and an electron. Figure 2.3
shows Compton process, in which an incident gamma-ray photon provides a part of the initialenergy to a recoil electron, and simultaneously the gamma-ray photon is scattered as its energy
becoming lower.
7

Chapter 2. Science of MeV Gamma Ray

cross section [cm2/g]

104
3

10

102

1

ic
ctr
ele
oto ionPh sorpt
ab

10In
sc coheatt reeri ntng

To
Att tal
en
ua

10-1

tio
n

Coherescatte nt
ring

10-2
-3

10

tion
Pair crea ld)
fie(Nuclearn
atio
cre field)
Pair ctron
(Ele

-4

10

-5

10 -3
10

10-2

10-1

1

3

102
10
10
Photon Energy [MeV]

Figure 2.1: The cross-section of the interaction between gamma rays and matter (Argon) [39].

Figure 2.2: The signiﬁcant interaction between gamma rays and matter [40].

8Interactions of Photons with Matter

Recoil electron
Incident gamma ray
Scattering angle

φScattered gamma ray

Figure 2.3: Schematic view of Compton scattering.

When the gamma ray photon with an energy of E0 scatters through Compton process witha free electron, the energy of the scattered gamma ray is
E′ =

E0
,
E0
1 + me c2 (1 − cos ϕ)(2.4)

and the energy of the recoil electron is
Ke = E0 − E′

E0(1 − cos ϕ)
me c2
=
E0 ,
1 + mEe0c2 (1 − cos ϕ)(2.5)

where ϕ is a scattering angle. The angular distribution of the scattered gamma rays is represented
by the diﬀerential scattering cross section as follows,
(
)2 (
)(
)
dσ
1
1 + cos2 ϕ
k 2 (1 − cos ϕ)2
2
= Zre
1+
, (2.6)dΩ
1 + k(1 − cos ϕ)
2(1 + cos2 ϕ) [1 + k(1 − cos ϕ)]where σ and Ω are the cross section, and the solid angle, respectively, and k = mEe0c2 . The angular
distribution is shown in Figure 2.4.
In a real atom, electrons do not rest and they have the ﬁnite momentum of the orbit.
Therefore, the gamma-ray scattered at the ﬁxed angle from a monoenergetic source has some
intrinsic ﬂuctuation in its energy (the “Doppler broadening”).
When the gamma-ray energy exceeds twice of the rest mass of the electron, a pair creation
occurs, especially it is dominant interaction above 10 MeV. This interaction is to create a pair ofan electron and a positron with the interaction between the photon and an electric ﬁeld aroundthe nucleus. The energies of an electron and a positron emitted from the photon of the energy
E0 must be satisﬁed with the following equation,
E0 = E− + E+ + 2me c2 .

(2.7)

where E− and E+ are kinematic energy of an electron and a positron, respectively. The crosssection is proportional to Z 2 .
9Chapter 2. Science of MeV Gamma Ray

Figure 2.4: The angular distribution of Compton scattering [40].

One of the diﬀerent properties between photons and charged particles is range. Althoughcharged particles deﬁne the range, gamma rays have no well-deﬁned range. X/gamma rays areattenuated exponentially based on the probability function as they pass through matter. The
attenuation of gamma rays is given by:
I(x) = I0 e−µx .

(2.8)

where I(x) is the gamma ray intensity as a function of the distance x into the material, I0 is
constant (intensity at x = 0), and µ is the mass attenuation coeﬃcient. The mass attenuationcoeﬃcient depends on gamma-ray energy and the material with which the gamma rays interact.

2.2

Interactions of Charged Particle with MatterIn the matter, charged particles lose the energy by following processes:
1. Coulomb interaction with an electron or atomic nucleus
2. Bremsstrahlung
3. Interaction with atomic nucleus
4. Cherenkov radiation
I mention on the lower energy particles with a less than a few hundred MeV, and I ignore
the interaction with atomic nucleus in this thesis. Also, Cherenkov radiation is not described inthis section, because the energy loss is small by Cherenkov for our experiment.
In Coulomb interaction, ionization occurs in a target atom of the matter by a charged
particle. The energy of an emitted electron Ee is follows:
Ee = Eint − I,

(2.9)
10

Interactions of Charged Particle with Matter
where Eint , and I are, Coulomb energy given by a charged particle into the atom, and theionization potential of the atom. In addition, by Coulomb interaction, excitation occurs when
the charged particle has a larger energy than the excite energy, and then a photon emits with anenergy of the diﬀerence between the states in the atom. The event with ionization or excitation
is called an inelastic scattering (inelastic collision).
The energy loss of the charged particle by the inelastic scattering is described by Bethe-Blochformula is as follows:dE
me c2= 4πre2 z 2 2 N Z{logdxβ(2me c2 2βI)− log(1 − β) − β 2 },(2.10)

where c, me , e, z, N , Z, and I are velocity of light, rest mass of the electron, charge of theelectron, particle charge number, number of atom per unit volume in the matter, atomic numberof matter, and mean excitation potential of the matter, respectively. In addition, β = v/c (v isvelocity of the charged particle) and re is a classical electron radius deﬁned by Equation 2.3
This formula means that the energy deposit increases dramatically in the end of the range
where the particle velocity is close to zero. However, this formula 2.10 is derived based on anassumption that the charged particles has a constant charge when the particle travels in thematter. Therefore this formula cannot be used when the particles have Z > 1 except alphaparticle.
In the bremsstrahlung, ionizing process deposits a continuous energy in the matter. When acharged particle with the charge number of z and the mass of M passes the matter of an atomic
number Z, the intensity I is described as follows:
I∝

z2Z 2
.
M2

(2.11)

This equation indicates that gamma-ray emission from the bremsstrahlung process is mostly
caused by an electron or a positron, because the mass of the proton or other charged particle
are heavier than the electron by an order of 103 or more.
This energy loss of an electron or a positron by the bremsstrahlung dE
dx b is approximately
described as follows:dE
ZT (M eV ) dE
=
,
dx b
750
dx e

(2.12)

where Z, T , and dEdx e are an atomic number of target atom, kinematic energy of electron or
positron, and the energy loss by the inelastic scattering, respectively.

11

Chapter 2. Science of MeV Gamma Ray

2.3

Mechanism of Gamma-ray Radiation

By various processes listed below, continuous or line gamma rays are emitted.
⋄ continuous gamma rays
· Thermal radiation· Synchrotron [electron]· Bremsstrahlung [electron, nucleus]· (Inverse) Compton scattering [electron, photon]⋄ line gamma rays / broadening line
· nuclear transitions (de-excitation, radioactive decay, neutron capture) [nucleus, (neutron)]· annihilation line [electron, positron]· pion decay [pion]
Here, in the square bracket of the above list denotes the particle related to the process.
Thermal radiation is emission from a large population of electromagnetically interacting
particles in equilibrium with their mean energy characterized by temperature T . The spectrumof radiation intensity follows the ‘black body’ distribution, and the energy density of radiation
at frequency ν isIν =8πhν 31,
c3 ehν/kB T − 1(2.13)

where h is Planck’s constant, and kB is Boltzmann’s constant. For a thermal gamma-ray photonof 1 MeV, the corresponding temperature would be above 109 K. Thus, thermal process is not
dominant, and nonthermal processes are a more dominant source of gamma rays. In the MeV
region, several nonthermal processes exist as follows.
When a relativistic electron moves in the magnetic ﬁeld B, an electron trajectory will bechanged by the Lorentz force. Because any acceleration of the charged particle can be viewed
as the modiﬁcation of the ﬁeld, the kinetic energy of the electron is transformed into the energyof the electromagnetic ﬁeld. For the relativistic electron, such an energy is lost as an emissionof photon, which is called Synchrotron Radiation.
Synchrotron radiation spectrum has a characteristic peak described as
3h eB⊥hνc =4π me c

(

Ee
me c2

)2
,

(2.14)

where, e, me , Ee are the charge, the mass, the energy of the electron, respectively, and B⊥ =B sin θ, where the pitch angle θ is the angle between the particle trajectory and the direction of
12

Mechanism of Gamma-ray Radiation
the magnetic ﬁeld . The total energy loss rate by synchrotron radiation of an electron moving
in the magnetic ﬁeld B is
(
)
dEe
4
B2
−
= σT c γ 2 β 2 ,dt syn 38π(2.15)

where σT is Thomson cross section described followings:
σT =

8πe4
3m2e c4(2.16)

Another important process of gamma-ray production is ‘Bremsstrahlung’. When an electron
passes very near the nucleus, the trajectory of the electron is substantially changed by the strong
electric ﬁeld of the nucleus, and photons are emitted by the same process as that in synchrotron.
The spectrum of bremsstrahlung radiation remains ﬂat up to roughly the electron kinetic
energy, and it drops sharply toward zero above the electron energy. For relativistic electrons ina fully ionized plasma (no screening), the bremsstrahlung loss is
(
)
(
)
dEe
Ee
2 2
−
= 4na Z r0 αc ln+ 0.36 Ee .
dt B
me c2
In the case of total screening, above formula is modiﬁed to
(
)
(
)
dEe
183
1
2 2
−
= 4na Z r0 αc ln −1/3 −
Ee ,
dt B
18
Z

(2.17)(2.18)

2

where na is the target atom density, r0 = mee c2 is the classical radius of the electron, and α isthe ﬁne structure constant.
When an energetic photon collides with a low energy electron, the photon is scattered andsome of the photon energy transfers to the electron. This process is called Compton scattering.
The inverse process also surely exists, and generates a gamma-ray photon: when the low energy
photon collides with an energetic electron, this photon has some energies from the electron via
the collision. This ‘Inverse Compton Scattering’ is important in the regions of the high photon
density.
The energy loss for an electron in the photon ﬁeld energy density wph is written as
(
)
dEe
4
−
= σT cwph γ 2 β 2 .
dt IC3(2.19)A nucleus has several speciﬁc, quantized states of the energy to bind the nucleons. These
nuclear states have typical energy bands of the sub-MeV or MeV scale, and hence the transition
between those states of a nuclear absorbs or emits the MeV gamma-ray photons. Thus, line
gamma rays are emitted from ‘de-excitation’ of a nucleus:X ∗ −→ X + γ,

(2.20)and ‘radioactive decay’:
X −→ Y ∗ + e+ −→ Y + γ.

(2.21)
13

Chapter 2. Science of MeV Gamma Ray

Table 2.1: The gamma rays by nuclear transitions [41, 43, 44, 45]
process
Reaction
Energy [MeV]
half life⋆
12
11∗
De-ExcitationC(p, 2p) B
4.444
14 N(p, p∗ )14 B∗
2.313, 5.105
16 O(p, p∗ )16 O∗
2.741, 6.129, 6.917, 7.117
26 Mg(p, p∗ )26 Mg∗
1.809
56 Fe(p, p∗ )56 Fe∗
0.847, 1.238, 1.811
radioactive decay

56 Ni

60 Fe

0.158, 0.270, 0.480, 0.759, 0.812
0.847, 1.238, 2.598
0.122, 0.136
1.157
1.809
1.173, 1.333

n +1 H →2 D + γ

2.223

56 Co
57 Co
44 Ti
26 Al

capture⋆ d: day, y: year

6.10 d77.2 d271.7 d63 y
7.4 × 105 y
1.5 × 106 y

De-excitation occurs by the energetic collisions of cosmic-rays with the interstellar gas nuclei, inspace. On the other hand, a radioactive decay occurs by nuclearsynthesis in a supernova and acore of heavy stars. The examples of line gamma-ray emissions by nuclear transitions expectedin the universe are listed in Table 2.1. The observation of line gamma rays from this processlead to identify the origin of the nucleus.
Pion is a boson involved in the strong nuclear interaction, and is created during the stronginteraction events such as collisions of cosmic-ray protons with ambient-gas nuclei. The dominant
π-producing channels in hadronic interactions are as follows,p + p −→ p + p + aπ 0 + b(π + + π − ),(2.22)

p + p −→ p + n + π + + aπ 0 + b(π + + π − ),(2.23)

p + α −→ p + p + aπ 0 + b(π + + π − ),(2.24)

p + α −→ p + n + π + + aπ 0 + b(π + + π − ),(2.25)

where a and b are positive integers. π 0 decays rapidly into two gamma rays with a proper timeof 9 × 10−17 sec, and the distribution of the number of photon has a peak at approximately 70
MeV at the center of gravity system, which is the half of the rest mass of pion. Since pions
created by energetic protons has a momentum distribution, the observed gamma-ray spectrumof the pion decay is broadened by Doppler shift.
An electron-positron annihilation is an important source of gamma rays. In the annihilation,
two or more photons are created, and the total energy of electron and positron is distributed tothese photons. A positron and an electron may form a bound system, which is called positronium,
and two diﬀerent states exist. One is the ground state, and a positronium decays to two gamma
rays. In this case, each photon has an energy of 511 keV, which is equal to the rest energy of an
14

Sub-MeV and MeV Gamma-ray astronomy

Table 2.2: Interest of the Sub-MeV and MeV gamma-ray astronomy
physical features
science target
line gamma
Nucleosynthesis (supernovae)
π 0 γ reaction
Origin of the cosmic rays
transparency
Compact object (e.g. AGN, blackhole)
Deep Field (e.g. gamma ray bursts)

electron. Here the two-gamma photons are emitted in the direction opposite (180◦ ) at rest. The
other state is the parallel-spin state, and it decays to three gamma rays having the continuum
spectrum.
Actually, the 511 keV line was observed near the Galactic Center by The Solar Maximum
Mission (SMM), The Oriented Scintillation Spectrometer Experiment (OSSE), and INTEGRAL(Fig 2.6), and implied the annihilation rate of ∼ 2 × 1043 sec−1 . Annihilation photons suggests
the existence of the electron-positron plasma. Low energy positrons are produced by β + -decay,higher energy gamma photon (more than 1.022 MeV). The examples of the radioactive isotopes
causing β + -decay are 26 Al, 44 Ti, 56 Co, and the distinct nova also products 13 N and 18 F.

2.4

Sub-MeV and MeV Gamma-ray astronomy

Sub-MeV and MeV gamma-ray astronomy began in 1950’s when Hayakawa et al. predicted that
gamma photon might be emitted by π 0 decay process which is generated by cosmic rays [46, 47].
After that, the cosmic gamma rays was reported by several detectors on the satellites such as
OSO-3 (1972) [48, 49], SAS-2 (1973) [50], and COS-B (1975) [51].
The physical features (process) of Sub-MeV and MeV gamma-ray region is listed in Table
2.2. The line gamma rays are generated in the supernovae where nucleosynthesis occurs. For
example, radioactive nuclides such as 56 Ni and 44 Ti (Table 2.1), which produced by α process,emit gamma photons with an energy of a few hundred keV.
In 1982, Mahoney et al. detected line gamma-ray emission from 26 Al (1.809 MeV),60 Fe(1.173,1.332 MeV) using germanium-photon counting detectors onboard HEAO 3 [52]. Figure
2.5 shows the all sky map of 1.8 MeV from the decay of 26 Al observed by COMPTEL [41], and
26 Al are mainly distributed in the galactic plane.

COMPTEL’s results reveal that the main

contributors to the synthesis of 26 Al are most likely the massive stars, which contribute throughtheir winds (Wolf-Rayet stars) and through their supernova explosion [53].
511-keV gamma rays are also found to concentrate in the Galactic center by INTEGRAL
shown in Fig. 2.6 [54]. Several groups proposed the models or source objects for 511-keV emission
such as positron production from Sgr A∗ [55], MeV-mass dark matter[54], microquasars [56],etc.
In addition, line gamma rays are generated in solar ﬂare by a neutron capture (2.22 MeV),
and de-excitation process such as 12 C(p, 2p)11 B∗ (4.44 MeV) [57]. These line gamma rays are
15

Chapter 2. Science of MeV Gamma Ray

Figure 2.5: All sky map in 1.8 MeV by COMPTEL [41]Figure 2.6: 511-keV line gamma rays map in the Galactic center with an INTEGRAL [54].

16Sub-MeV and MeV Gamma-ray astronomy

Figure 2.7: Fluxes of diﬀuse radiation produced by both electronic and nucleonic components of cosmic rays in the inner Galaxy [59]. Contributions from π 0 -decay (thin solid line),
bremsstrahlung (dashed), inverse Compton (dot-dashed), and positron annihilation in ﬂight(dotted line) are shown.

produced by proton acceleration in the sun. By the observation of various line gamma rays, thehistory of the nucleosynthesis are would be cleared.
The study of gamma rays from π 0 -decay directly shows the existence of high energy relativistic hadrons [58]. Figure 2.7 shows the ﬂuxes of gamma ray emission from the galactic disk,and π 0 observation is important for the understanding of the sites and acceleration mechanismof galactic cosmic rays [59].
Since the universe are more transparent to MeV gamma rays than other energy band, gamma
rays are a good probe for compact objects and the deep universe. The deep universe survey
provides us the understanding in cosmological parameters and the history of the star formation.
GRB, for example, has a uniform spatial distribution up to high redshift range. Some groups
reports that GRB is a candidate for a standard candle, which is used to measure the far distance[60]. Indeed, Swift satellite discovered a GRB (GRB 090423), which is the most distant cosmic
explosion (at present) ever seen with a redshift of 8.2 [61, 62].
Extragalactic diﬀuse gamma rays come from the outside of the Galaxy, of which the distribution is uniform in the all sky. Figure 2.8 is the spectrum of the extragalactic diﬀuse gamma
rays. The spectrum shows that the majority is considered to be gamma rays by blazars andnon-blazar AGNs in the energy region of MeV gamma rays.
Although the gamma ray imaging in astronomy is important to resolve the above issues, theimaging observation in the sub-MeV and MeV gamma region is not suﬃcient. For example,
Compton Gamma-Ray Observatory (CGRO, 1991-2000) satellite had 4 detectors: COMPTEL,
17

Chapter 2. Science of MeV Gamma Ray

Figure 2.8: Spectrum of extragalactic diﬀuse gamma-ray [63], which is from non-blazar AGNs
and blazars for a blazar gamma-ray luminosity function model. The model curves for the blazar
component (absorbed+cascade), non-blazar AGN component, and the sum of the two are shown.
(detail are described in [63].)18

Sub-MeV and MeV Gamma-ray astronomy
EGRET, OSSE and Burst And Transient Source Experiment (BATSE). COMPTEL found about
30 steady gamma-ray sources in the MeV region [64], while EGRET found about 270 sources inthe sub GeV region [65]. Figures 2.9 and 2.10 show the all sky maps of the gamma-ray photons
obtained by EGRET and COMPTEL, respectively. In addition, the Large Area Telescope (LAT)on the Fermi Gamma-ray Space Telescope (Fermi) released the First Fermi-LAT catalog (1FGL)
containing 1451 sources detected and characterized in the 100 MeV to 100 GeV range during
an approximately 2-year observation (Fig. 2.11). In the present circumstances, fewer sources
are observed in the MeV and sub-MEV region than GeV and sub-GeV region by approximately
two orders of magnitude.
To improve the imaging quality, we have developed an ETCC for the all-sky survey. The
next chapter, I describe the detail of our ETCC and the imaging.

19

Chapter 2. Science of MeV Gamma Ray

Figure 2.9: Distribution of the gamma-ray objects by EGRET [65].

Figure 2.10: All sky map by COMPTEL (1 - 30 MeV) [41].

Figure 2.11: The distribution of the gamma-ray objects by LAT [66].
20Medical Imaging and Proton Therapy

Table 2.3: Isotopes used in nuclear medicine
Isotope
Energy (keV)
Decay
Half-time
99m Tc140
IT⋆
6 hours131 I−
365
β decay
8 day
18 F
511
β + decay 109.8 mim.
15 O
511
β + decay
2 mim.
⋆ :isomeric transition.

2.5

Medical Imaging and Proton Therapy

2.5.1

Medical Imaging

Medical imaging is to image and examine a tumor (cancer) in the human or animal body for a diagnosis and treatment. Although there are some medical imaging techniques such as Ultrasoundimaging, thermal imaging, I brieﬂy introduce the X and gamma-ray imaging techniques. The
both X and gamma-ray are ionizing radiation, however gamma rays have a strong penetration
ability through the human body than the X-rays as described in section 2.1. Then X-rays areused for examination of the density of the tissue in the body. Since tumors tend to be denser
than normal ones, X-ray imaging is able to check the diﬀerence of the density. However a small
or young tumor may be missed by the X-ray photograph.
On the other hand, a nuclear medicine imaging technique displays the distribution of not thedensity of the tissue, but the some biological activity using an agent with some RI. This agent
tends to accumulate in the tumor based on the cellular function and physiology, and gamma
rays from the RI are detected using two or three-dimensional detectors to localize the tumor.
Since the detectors for the medical gamma-ray imaging such as the PET and SPECT have anarrow energy dynamic range, there are energy limits in use of isotopes in the agent. Table 2.3
lists common isotopes used in nuclear medicine.

2.5.2

Proton Therapy

Radiation therapy is a treatment method of a tumor using photons or charged particles, andthe tumor is killed by the energy deposit from the photons or charged particles. Comparedwith other tumor treatment such as surgery and chemotherapy, the patient has only a small
impact and long-term hospitalization is not necessary. In addition, chemotherapy usually has
an adverse drug reaction. However radiation therapy also has a risk to kill the normal cells, if
the Bragg peak does not match with the target.
Figure 2.12 shows relative dose distribution of conventional photon (X-ray) radiation ascompared to proton radiation as the beam penetrates body tissue [67]. X-ray photons depositmuch more energy in front of the tumor region than the tumor region. In addition, there is someenergy deposit beyond the tumor. On the other hand, protons have minimal side scatter in thetissue than X-rays, and protons lose most of their energy at the end of their range corresponding
21

Chapter 2. Science of MeV Gamma Ray

Figure 2.12: Relative dose distribution of conventional photon (X-ray) radiation as compared
to proton radiation (250 MeV) as the beam enters and traverses body tissue. [67]to the tumor. Therefore proton therapy is safer to normal cells than X-rays therapy.
The width of Bragg peak is generally narrow for actual size of the tumor. Then in orderto adjust the Bragg peak position to the tumor position, either a variable ﬁlter in front of
the patient, or a variable energy beam, called Spread out Bragg peak (SOBP) is used. Proton
therapy is applied for the treatment of various disease sites [68] including paranasal sinus tumors[69], chordoma [70, 71], chondrosarcoma [72], lung tumors [73] etc.
In order to ensures that the treatment position matches the tumor, before the treatment, theproton range is commonly calculated based on the information of electron density distributions
measured from the X-ray CT images. However this calculation has uncertainty due to diﬀerent
electron densities tissues such as bone[75]. The proton range uncertain is even 1-2 cm [77].
To verify the Bragg peak position (dose) in the body, gamma rays emitted by proton beam
are studied, because the gamma-ray emission is related to proton energy deposit as shown in
Fig. 2.14, which is a simulated result by GEANT 4.
Here, in proton therapy, the mechanism of gamma-ray emission is as follows.
1. Annihilation
2. Prompt gamma rays (except Annihilation)
The ingredients of the human body are Hydrogen, Oxygen, Nitrogen, Carbon etc., andapproximately 70 % of the component is water. There are only six major channels producing
the positron emitters 11 C, 13 N and 15 O by proton beam in human tissue as listed in Table 2.4[78]. Those positron emitters have diﬀerent half-times, and reactions in Table 2.4, which havediﬀerent cross sections depending on the proton energy (Fig. 2.15). Thus the distribution of the22Medical Imaging and Proton Therapy

Figure 2.13: Combining proton beams with diﬀering energies results in the summation of narrow,monoenergetic Bragg peaks into a broader, more clinically useful ”spreadout” Bragg peak [76].

Radiative dose (normalized)

1
0.8
0.6
0.4
0.2
0
0

2

4

6

8
10
12
depth (cm)

14

16

18

20

Figure 2.14: The simulated distributions of 140-MeV proton dose deposition (dotted line) andof gamma rays emitted in water in the all energy band (solid line) as a function of water depth.
The detail of the simulation, which corresponds to our experiment setup is described in Chapter
5.

23Chapter 2. Science of MeV Gamma Ray

Table 2.4: Positron emitter production in the proton therapy
reactionthreshold [MeV] Half-life [min] Positron Max. energy [MeV]
16 O(p,pn)15 O
16.79
2.037
1.72
16 O(p,2p2n)13 Na)
c)
5.66
9.965
1.19
14 N(p,pn)13 N
11.44
9.965
1.19
12 C(p,pn)11 C
20.61
20.39
0.96
14 N(p,2p2n)11 Ca)
3.22c)
20.39
0.96
16 O(p,3p3n)11 Cb)
c)
27.50
20.39
0.96
a): (p,2p2n) is inclusive of (p,α)b): (p, 3p3n) is inclusive of (p, αpn)c): The listed thresholds refer to (p, α) and (p, αpn)

Table 2.5: Line gamma rays by proton interaction with organic nuclei [45].
ReactionTransition⋆
Energy [MeV] Mean time [sec]12 C(p,x)10 B∗
10 B∗0.718 → g.s.
0.718
1.0 × 10−9
11 B∗1.740 →11 B∗0.718
16 O(p,x)10 B∗
1.022
7.5 × 10−15
12 C(p,2p)11 B∗
11 B∗ → g.s.
4.44
5.6 × 10−19
14 N(p,p∗ )14 N∗
14 N∗5.106 → g.s.
5.105
6.3 × 10−12
16 O(p,x)14 N∗
14 N∗5.106 → g.s.
5.105
6.3 × 10−12
16 O(p,p∗ )16 O∗
16 O∗6.130 → g.s.
6.129
2.7 × 10−11⋆ g.s. : Ground state.
511-keV gamma rays depends on both the incident proton energy and the elapsed time between
the proton irradiation and the gamma-ray imaging.
Figure 2.16 shows the simulated distribution of linear production densities of 11 C, 13 N and
15 O [80], where the 511-keV gamma-ray distribution does not trace the Bragg peak.

Howeversince the PET is a very common imaging technique, the veriﬁcation of the dose by 511-keV
gamma rays has been studied [20].
Also other line, prompt gamma rays are generated during the proton radiation in the body.
Table 2.5 lists example of the prompt lines, and Fig. 2.17 shows the simulated distributionof the depth dose and gamma rays including prompt gamma rays [81] when a water tank isirradiated with the proton beam. Although the peak of the prompt gamma-ray distribution
also does not match the Bragg position, the fall-oﬀ position is matches to the Bragg position.
Moreover, the gamma-ray intensity is correlated with the Bragg-peak intensity, therefore theprompt gamma-ray detector might measure also the proton beam current unlike the PET.
Figure 2.18 shows the ﬂuence of the primary protons, secondary proton, neutron, and thetotal absorbed dose as a reference [74]. Around the Bragg peak, the main contribution to theabsorbed dose is primary protons (from 80 to 90 %), whereas secondary protons created byprimary proton-induced reactions contribute to the dose from 20 to 5 %. The secondary protons
and neutrons such as recoil protons are generated by the nuclear reactions. This ﬁgure indicatesthat neutrons exist beyond the Bragg peak, and the gamma rays from neutrons such as 2.22-MeV
24

Medical Imaging and Proton Therapy

0.12
0.1

]s
nr 0.08ab[noit 0.06ces
ss
roC 0.04
0.02
0

0

50

100

150

200

Incident energy [MeV]

Figure 2.15: Nuclear reaction cross sections as a function of the proton energy [79]. Each markdenotes as followings: cross:12 C(p,pn)11 C, triangle:16 O(p,2p2n)13 N, square:16 O(p,pn)15 O, andcircle:14 N(p,pn)13 N.

Absorbed Energy [MeV/cm]

O Production [/mm]15

C Production [/mm]11

N Production [/mm]13

Figure 2.16: The simulation results of linear production densities of 11 C, 13 N and 15 O as afunction of depth. The absorbed energy by the tissue is superimposed using a right-side verticalscale for comparison [80].

25Chapter 2. Science of MeV Gamma Ray

Figure 2.17: Comparison of the depth dose distribution (left) with the prompt-gamma distribution (right) for a beam energy of 70 MeV [81].

gamma rays (n-capture) is not used for the tracer of Bragg position.

26

Medical Imaging and Proton Therapy

Figure 2.18: Fluences of protons as a function of depth in a water phantom per incident protonwith energy of 150 MeV. (solid line) total ﬂuence of protons; ♢ ﬂuence of primary protons, • and◦ ﬂuences of ”p-secondary protons” and ”n-secondary protons”, which are generated by protonand decrease reactions. Total absorbed proton or electron dose is added for reference [74].

27Chapter 3

Compton and Other Gamma-Ray
Imaging
3.1

MeV Gamma-Ray Imaging Detector in Astronomy

In this section, I brieﬂy summarize several imaging detectors onboard satellites in the subMeV and MeV region. Followings imaging techniques have been used for astronomy or being
developed astronomy.
1. Collimator method
2. Coded aperture imaging
3. Gamma-ray lens method
4. Compton method
5. Pair-creation method
Collimator method is a simple method to select the gamma-ray direction by material shieldas shown in Fig. 3.1. The narrower aperture of a collimator is necessary if a better position
resolution is expected even though its ﬁeld of view is small. Usually, the minimum angularresolution is several degrees under 1 MeV.
The Oriented Scintillation-Spectrometer Experiment (OSSE), which was onboad NASA’s
Compton Gamma-Ray Observatory (CGRO) (1991 - 2000), used the active and passive collimators. Figure 3.1 (b) shows the schematic view of the one detector of OSSE, in which there were
4 identical detectors on CGRO. The main detector is a phosphor-sandwich (phoswich) detector
consisting of the NaI:Tl crystal (diameter: 33 cm, thickness: 10.2 cm) and the CsI:Na crystal(thickness of 7.6 cm). In front of the Nal:Tl crystal, a passive tungsten collimator was mounted.
Around the main detector and the tungsten collimator, there was an annular shield of 8.5 cm
thick NaI:Tl crystal for the anti-coincidence which is called active veto. This active veto countereﬃciently rejects the gamma-ray background from the inner detector by the coincidence between
inner and veto counter hitting. For the rejection of charged particles, a thin plastic scintillator
28

MeV Gamma-Ray Imaging Detector in Astronomy(0.6 cm thick) was also placed on the collimator. OSSE had an energy range from 0.1 MeV to
10 MeV and had a FOV of 3.8◦ × 11.4◦ (FWHM).

(a)

(b)

sourcenoise

collimator
scintillator
PMT

Figure 3.1: Schematic view of the gamma-ray imaging detector with the collimator (a), andOSSE (b) [41].

Coded aperture imaging camera consists of a position-sensitive photon camera and a codedmask in front of the camera (Fig. 3.2 (a)). The Coded mask has holes, which are arranged
according to some rules, and the diﬀerent gamma-ray sources make the diﬀerent shadows owing
to the holes pattern. In other words, the shadow pattern solves the gamma-ray direction by aninverse fast Fourier transform (FFT) method.
The INTErnational Gamma-Ray Astrophysics Laboratory (INTEGRAL), which is a missionof the ESA launched in 2002, has two gamma-ray telescopes with coded masks: SPI and IBIS.
The SPectrometer on Integral (SPI) is based on coded aperture imaging. The detector of
SPI consists of the array of 19 Ge crystals cooled to 85 K, and thus has a good energy resolutionof 2 keV for 1 MeV. The mask of SPI is made by a 3 cm thick tungsten plate, and the distance
between the mask and the detector is 1.71 m. The whole detector is surrounded by the 511kg active BGO shield viewed by 181 PhotoMultiplier Tubes (PMT). Due to the high energy
resolution, the SPI’s sensitivity for narrow-line gamma-rays is improved at least by a factor of
10 better than other gamma-ray telescopes.
Also, the Imager on Board of the Integral Satellite (IBIS) uses a coded aperture imaging.
The coded mask of IBIS is a 16 mm thick tungsten mask. The detector consists of 2 layers. The
lower layer is the array of 4096 CsI scintillators viewed by the silicon PIN photodiodes (each CsI
size: 9 cm×9cm ×30 cm, total area: 3318 cm2 ), named PICsIT. The upper layer is the array of
16384 CdTe pixels (pixel size: 4cm×4cm×2 mm, total area: 2621 cm2 ), named ISGRI. ISGRI
detects from 15 keV to 400 keV, and PICsIT covers from 200 keV to 10 MeV. The detector
29Chapter 3. Compton and Other Gamma-Ray Imaging

(b)

(a)

Flux2

Flux1

Mask

Counts
Detector

Figure 3.2: Schematic view of the Coded aperture imaging (a) [82], and photograph of Coded
Mask in Swift/BAT (b) [83].

layers are surrounded by the BGO active shield, and a passive tungsten collimator is placed
between the mask and BGO shield. The distance between the tungsten mask and the upper
layer detector is 3.2 m, so that IBIS has a good angular resolution of 12 arc sec.
In addition, Burst Alert Telescope (BAT) loaded on Swift satellite consists of 2.7-m2 areamask made from lead (Fig. 3.2 (b)) and a CdZnTe detector with a size of 4 mm ×4 mm ×2mm.
BAT has an energy dynamic range of 15-150 keV, and a half-coded ﬁeld of view of 1.4 str.
Here, both collimator and coded aperture except BAT/Swift imaging are required with vetocounters surrounding the detector for background rejection, and therefore, the whole detector
system becomes so heavy. In addition, when a mask is placed at a longer distance from thephoton detector, the angular resolution is better up to a few arcmin, while more photon events
are required to keep an image quality. Therefore, only bright objects are imaged using a codedmask detector with a good-angular resolution.
Gamma-ray lens method is a new technique to detect the focused gamma rays using Bragg
reﬂection as shown in Fig. 3.3. Owing to use of Bragg reﬂection, the energy band is narrow thanother detectors, and the focus length is so long. For example, CLAIRE experiment, which is theballoon-born experiment in 2001, had an energy band of 169 - 171 keV, an energy resolution of 2
keV at 170 keV (FWHM), and a focus length of approximately 2.8 m. However, the ﬁeld of view
was quite narrow 45” [84]. Here, when a detector has an energy range of a few hundred keV,
the focus length should be ∼ 100m, Although the lens method is expected to be an excellent
the position and energy resolution, which is suitable for a pointing observation, we must resolve
diﬃcult techniques such as formation ﬂight between a lens and a photon detector in the space.
30

MeV Gamma-Ray Imaging Detector in Astronomy

Figure 3.3: Schematic view of gamma-ray lens [85].

Compton imaging uses a Compton scattering which is a main interaction with the matterin the sub-MeV and MeV regions. The COMPTEL based on classical Compton imaging, which
was onboad CGRO, was the ﬁrst Compton telescope onboard a satellite. The schematic viewof COMPTEL is shown in Fig. 3.4. The low-Z detector of it was a liquid organic scintillator
NE213A (geometrical area: 4188 cm2 ), and the high-Z detector was a NaI(Tl) crystal (geometrical area: 8744 cm2 ). Although each detector was surrounded by an anti-coincidence counter
dome for the rejection of charged particles, the veto counter for gamma rays is not used. For the
background rejection, COMPTEL used the time of ﬂight (TOF) technique. COMPTEL had an
FWHM angular resolution of 1 - 3◦ typically for 0.75 - 30 MeV gamma-ray [42]. In addition,
the energy resolution and the ﬁeld of view were 5-10% and ∼ 1 str (FWHM), respectively, at
0.75-30 MeV.
The Energetic Gamma-Ray-Experiment Telescope (EGRET) aboard CGRO was based ona pair-tracking imaging detector. The schematic view of EGRET is shown in Figure 3.5. The
EGRET consisted of a spark chamber for direction measurement, an anti-coincidence counteras a veto counter, and a NaI(Tl) calorimeter. The spark chamber had interleaved tantalum
foils and tracking layers. The trigger counter had the plastic scintillator sheets inserted into thelower part of the spark chamber, and the thickness of the NaI(Tl) crystal was 20 cm. For therejection of charged particles, the anti-coincidence hood with the 2 cm thick plastic scintillator
surrounded the spark chamber.
Here, EGRET detected approximately 270 steady gamma-ray sources in the GeV region,
while COMPTEL detected only approximately 30 steady sources in the MeV region. The reason
is probably that COMPTEL could not reject backgrounds completely by the TOF of both
detectors even though the sub-MeV and MeV bands have much larger backgrounds than the31Chapter 3. Compton and Other Gamma-Ray Imaging

Figure 3.4: Schematic view of COMPTEL [42].

GeV band. Actually, COMPTEL had various backgrounds as shown in Figure 3.6, of which
details were described in [86]. Each process is:
A Internal Single Photon: Gamma rays, produced at the inside of the detector by neutron
capture or by radioactive isotopes (like 40 K), may make Compton scattering in the forward
detector, and its scattered gamma-rays are absorbed in the backward detector.
B External Single Photon: Gamma rays are produced at other detectors or the satellite
platform by neutron capture, decay of radioactive isotopes and scattering, and then they
may make Compton scattering in the forward detector and its scattered gamma-rays areabsorbed in the backward detector.

Figure 3.5: Schematic view of EGRET [41].

32Backward Peak

pa

C

sti
c

A

rti

cle

MeV Gamma-Ray Imaging Detector in Astronomy

re

lat

ivi

D1
B

CountsD
F

E

Forward Peak
D2Signal

F
D

EGRET

E

COMPTELOSSE

A&C

Continuum
E+F

CGRO Platform−20

Figure 3.6: Various background of COMPTEL [86].

−15

−10

−5

0
ToF [nsec]

5

10

15

20

Figure 3.7: TOF distribution [86].

C Internal Multi-photon: Multi-photons are produced by the interactions, such as 27 Al(n,α)24 Na
or 27 Al(n;n′ , · · · )27 Al. Then some photons may hit the forward detector and also the backward detector simultaneously.
D External Multi Photon: In other detectors or the satellite platform, some gamma rays may
be created by the same processes of C, and some photons may hit the forward detector
and also the backward detector simultaneously.
E Random Coincidence: Sometimes, the diﬀerent photons produced by the independentinteractions hit the forward detector and also the backward detector accidentally.
F Cosmic-Ray Interaction: It interacts with material at the diﬀerent positions, generates
some photons simultaneously, when cosmic rays come into the satellite. Then those photons hit both the forward and backward detectors.
G Other Process: As the other backgrounds, neutron, electron and atmospheric gamma-rays
were measured.
For the background rejection, COMPTEL used the time of ﬂight (TOF) between the low-Z
detector and the high-Z detector. Figure 3.7 shows the distribution of TOF by simulation.
Although the most events are backward events in which gamma-ray interacts in the backward
detector and then in the forward detector, these events can be easily rejected using the timing
above 1 MeV. However D, E and F events make continuum components, and the forward peakincludes the one third of them. Since the interaction of A and B events are the real Compton
interaction, they are the intrinsic background and hardly be rejected. Moreover, if there is adecay in the forward detector, C events have no diﬀerence with celestial gamma rays in TOF
distribution. Therefore, even if we pick up the forward peak, the most of the selected events arestill due to background events as shown in Figure 3.7.
From the reason described above, the actual sensitivity of COMPTEL was lower than thedesign sensitivity. However, in the MeV region where the main interaction of photon with matteris Compton scattering, the most suitable method with better sensitivity is Compton imaging.
In fact, in the MeV region, COMPTEL with Compton imaging had a better sensitivity than33

Chapter 3. Compton and Other Gamma-Ray Imaging
SPI and IBIS with coded aperture imaging as shown in Fig. 3.8, which shows the sensitivities
of several detectors launched so far for the continuum spectrum. However, the sensitivity of
COMPTEL is worse by two orders of magnitudes than that in the X-ray and GeV gamma-ray
bands. In order to obtain the better sensitivity for the all sky-survey in the sub-MeV and MeV
region, we should use the Compton imaging method, and the incident gamma-ray direction isdetermined by not accumulating some photons, but a single photon.

Fermi

Figure 3.8: Continuum sensitivities of X/Gamma-ray observatories [87].

3.2

MeV Gamma-Ray Imaging Detector in Medical Imaging

For medical imaging, two detectors are used generally. The ﬁrst is Single photon emission
computed tomography (SPECT) using a collimator method. The part of the photon detector
commonly consists of scintillator. Figure 3.9 (a) shows the schematic view of the SPECT, anddouble and triple-SPECT are also used in the medical imaging in order to obtain the higher
sensitivity as shown in Fig. 3.9 (b,c). SPECT has a typical spatial resolution of ∼ 4 − 6 mm
(FWHM) depending on the size of the collimator aperture, and an available energy dynamic
range is approximately 80 − 300 keV by changing the collimators (ﬁne collimator of ∼ 4 mm
resolution in ≤ 200 keV) [88]. The energy range is limited by background from collimator with
Compton scattering and the penetration of higher energy gamma rays.
The other is a Positron emission tomography (PET). Figure 3.10 shows the schematic view of
PET which consists of pairs of segmented detector on the opposite side. Using the pair detector,
back-to-back 511-keV gamma rays from a positron-emitting RI are detected. The position of
the positron-emitting RIs is obtained by the crossing point of several back-to-back gamma-raylines. PET has been able to image three-dimensionally by moving the detector along the axis
34

MeV Gamma-Ray Imaging Detector in Medical Imaging

(b)

(a)
sourcegamma rays
collimator
Photondetector(c)Figure 3.9: Schematic view of (a) one- (b) double- (c) triple- SPECTs.

Table 3.1: Calculated positron range [89].
18 F
11 C
13 N
positron emitter
FWHM [mm]0.102 0.188 0.282
FWTM⋆ [mm]
1.03
1.86
2.53
⋆ FWHM: full width at tenth maximum.

15 O

0.501
4.14

of the body. Commonly, as the segmented detector, scintillator is used due to a high stopping
power.
Here, the FWHM spatial resolution of the PET ∆x/x has a principle limit as describedfollowing equation:√( )d 2∆x/x =+ r2 + b2 + (0.0022D),
2

(3.1)

where d, r, b, and D denote the dimension of the square face of a crystal, positron range, a segmented scintillator size (block eﬀect), and diameter of the PET (non-collinearity), respectively.
Positron range depends on the radio isotopes which have diﬀerent energies, and the range belongs to Bethe-Bloch formula. Figure 3.11 shows a calculated distribution of positron-annihilation
positions in water projected onto a plane for 18 F and the distribution proﬁle 3.11 [89]. Table
3.1 lists the positron range for several positron emitters.
The block eﬀect b represents the uncertainty associated with identifying individual crystals
with secondary detection devices such as photomultiplier tubes. This factor typically contributes
only a fraction of a millimeter to the spatial resolution. Non-collinearity is to have an error in
the angle between the back-to-back gamma photons, because of the probability of positrons
annihilating while in motion. For the large-size PET for human body, the non-collinearity eﬀect
35

Chapter 3. Compton and Other Gamma-Ray Imaging

Detector(array)(511 keV)t = t1γelectron(511 keV)t = t2γe- + positrone(511 keV)t = t2γ(511 keV)t = t1γ

Figure 3.10: Schematic view of PET. Back-to-back gamma-ray emission at the same time (t = t1
or t2 ) are detected.

Figure 3.11: Calculated distribution of positron annihilation from 18 F source in water phantom(left), histogram of x coordinates from positron annihilation point distribution (right) [89].

is not negligible. Thus the typical spatial resolution for the human PET is ∼ 5 mm (FWHM),
and small PET for the animal has a better spatial resolution of ∼ 2 mm (FWHM).
For medical requirement, the spatial resolutions of the SPECT and PET are sometime not
suﬃcient. To recover the resolution of the PET and SPECT, usually in the medical diagnosis,
X-ray CT or magnetic resonance imaging (MRI) also is used for the compensation.
The SPECT and PET have a narrow energy dynamic range, and available RIs is limited.
If an imaging detector with a wide-energy dynamic range is available, it is possible to choicevarious RIs. For example, zinc is necessary to divide cells, and a 65 Z (1116 keV) ion might trace
the position of growing up in the living body such as root in the plant or tumor [90, 91]. In
addition, the wide energy range enables us to obtain a multi-RIs imaging simultaneously. Table
3.2 shows the RIs which are used at present or expected to be put to practical use for the medical
imaging.
36

Electron-Tracking Compton Imaging

IRenergy [keV]

Table 3.2: Example of RIs for the imaging
131 I
198 Au
22 Na
18 F
54 Mn
65 Zn
354
364
412
511, 1275 511 835 1116

133 Ba59 Fe

1099, 1292

⃗s, E0ψ

ϕ⃗e, Ke⃗g , EγαFigure 3.12: Parameters in Electron Tracking
Compton Imaging

Figure 3.13: Deﬁnition of ARM & SPD

To obtain a wide energy dynamic range, a Compton camera has been studied for medical
imaging so far [92, 93, 94]. However, most studies are still at simulation steps, and their Compton
cameras were a classical imaging type. On the other hand, our electron-tracking Compton
camera, as described in next section and Chapter 4, succeeded in obtain the animal imaging
with RIs.

3.3

Electron-Tracking Compton Imaging

The electron tracking Compton imaging measures both the energies and the directions of theincident gamma-ray by obtaining with the energy and direction of the scattered gamma photon
and the recoil electron. In Fig. 3.12, the energy and direction of the scattered gamma-ray are
Eγ and ⃗g , the energy and direction of the recoil electron are Ke and ⃗e, the scattering angle is ϕ,
the recoil angle is ψ and the diﬀerential angle between ⃗g and ⃗e is α, where ⃗g and ⃗e are the unit
vectors. Then, the incident energy E0 and the incident direction ⃗s of the initial gamma-ray aredescribed, respectively:
E 0 = E γ + Ke ,
(
)
sin ϕ
sin ϕ
⃗srcs = cos ϕ −
⃗g +
⃗e,
tan α
sin α
√
Ke (Ke + 2me c2 )
Eγ
=
⃗g +
⃗e.
Eγ + KeE γ + Ke

(3.2)
(3.3)
(3.4)

where the scattering angle is ϕ, which is written as follows,cos ϕ = 1 −me c2 Ke,Eγ + Ke Eγ

(3.5)
37

Chapter 3. Compton and Other Gamma-Ray Imaging
and the recoil angle ψ is
(
)√
me c2
Ke
cos ψ = 1 +.
E γ + Ke
Ke + 2me c2(3.6)

The diﬀerential angle α between ⃗g and ⃗e is described by the deﬁnition:cos αgeo = ⃗g · ⃗e.

(3.7)

On the other hand, α is described with Eγ and Ke by the Compton kinematics:
(
)√
me c2
Ke
.
cos αkin = 1 −EγKe + 2me c2(3.8)

The angle α is a characteristic parameter of the electron tracking Compton imaging. The
α makes it possible to select the Compton scattering events from the backgrounds. The α
is obtained from the equation (3.7) and (3.8), where (3.7) depends on only the geometricalinformation, and (3.8) depends on only the kinematic information. Thus, the αgeo is independent
from the αkin . For this reason, we can select good events, in which Compton the scattering occurs
in the tracker and scattered photon is absorbed perfectly in the absorber, by requiring the follows
αgeo = αkin .

(3.9)

In the Electron Tracking Compton Imaging, the accuracy of the event reconstruction isestimated using two parameters. One is the Angular Resolution Measure (ARM), which is theaccuracy of the scattering angle:(∆ϕARM = arccos (⃗s · ⃗g ) − arccos 1 −me c2 KeEγ + Ke Eγ).

(3.10)The other is Scatter Plane Deviation (SPD), which is the accuracy of the determination of thescattering plane:
( (
))
(
)
⃗s × ⃗g⃗srcs × ⃗g⃗s × ⃗g ⃗srcs × ⃗g
∆νSPD = sign ⃗g ·×arccos·.
|⃗s × ⃗g | |⃗srcs × ⃗g ||⃗s × ⃗g | |⃗srcs × ⃗g |(3.11)

where ⃗s is the real direction, and ⃗srcs is the reconstructed direction of the incident gamma-ray.
Thus, the error region of one event is a sector shape, as shown in Fig. 3.13.
In the previous sections, we assume that gamma-ray photon scatters with a free electron.
That process is called unbound Compton scatter. However, in actual detectors, electrons arebounded in the potential energy of the nuclear or molecular orbits. Thus, electrons have akinematic energy before the Compton scattering.
In Compton imaging, the scattering angle is obtained by the energy of scattered gamma-ray
and the energy of recoil electron. Therefore, the accuracy of the scattering angle (ARM) is
certainly aﬀected by Doppler broadening, and Doppler broadening makes the intrinsic limit for
the angular resolution of the ARM.
Since Doppler broadening is caused by the motion of the electron before the scattering, theenergy distribution of scattered gamma-ray is changed at which orbit the electron recoils. Figure
38

Electron-Tracking Compton Imaging

Figure 3.14: The ARM distribution of each Figure 3.15: The Z dependence of the ARM[95]electron orbit (E0 = 200keV, Si) [95]3.14 shows the uncertainty of the ARM of each orbit. Because the electron at the outer orbit
has the lower kinematic energy, the eﬀect of Doppler broadening is smaller, and the uncertaintyof the ARM is the smaller. The dependence of the uncertainty of the ARM on Z is shown in
Fig. 3.15. Figure 3.16 shows the energy dependence of Doppler broadening. The energy of anincident gamma-ray is higher, the eﬀect of Doppler broadening is smaller because the initialenergy of an electron is smaller enough.

Figure 3.16: The energy dependence of Doppler broadening [95]Due to the Doppler broadening, the gaseous detector has an advantage of the use for converterof the Compton camera than solid state detector such as semi-conductor. In addition, the gaseous
detector have a better capability of electron-tracking than a solid state. Although, the densityof solid is higher by an order of 3 than that of gas, a gaseous detector including a micro patterngas detector is easy to expand the volume with a low cost. On the other hand, as an absorber,scintillator has better stopping power than semi-conductor.
Our aim to develop the ETCC is obtaining the higher sensitivity by a factor of 10 than CGRO.
For the ﬁrst step, we demonstrated the balloon-bone experiment in 2006, and we acquired thecosmic and atmospheric gamma-ray photons as expected. This balloon-bone experiment and theprevious studies suggest the feasibility of being our expected sensitivity. Also our ETCC also
39Chapter 3. Compton and Other Gamma-Ray Imaging
has been applied to medical imaging. In fact, we have already demonstrated a RI imaging withsmall animals as the target. Since the ETCC is thus expected to useful for various applications,
we have developed an ETCC with the improvement of performances such as detection eﬃciency
and angular resolution.

40

Chapter 4

Gaseous TPC and Scintillator
4.1µ-PIC (Micro Pixel Chamber)

The ETCC consists of a gaseous time projection chamber (TPC) as a converter, and a scintillator
as an absorber. One of the features in the ETCC is to measure three-dimensional electron trackin the TPC, which needs an enough position resolution of several hundred µm and time resolutionof ∼ 10 nsec Thus, we have developed a µ-TPC consisting of a micro pattern detector (µ-PIC),
and a Gas Electron Multiplier.
Micro Pixel Chamber (µ-PIC) is a one of Micro Pattern Gaseous Detector (MPGD) [97].
Figure 4.1 shows the schematic view of µ-PIC. The µ-PIC has a pixel electrode like a sliced
proportional counter. The substrate is made by polyimide, and the electrodes are Cu coatedwith Ni. µ-PIC is manufactured by the print circuit board technology, and thus the device witha very large area can be easily developed. Actually, a large µ-PIC with an area of 30 cm ×30 cm
is already applied to several ﬁelds. Each pixel is aligned with the pitch of 400 µm. µ-PIC has∼ 65000 pixels in an area of 10 cm ×10 cm and ∼ 6 × 105 pixels in an area of 30 cm ×30 cm,
respectively. The anode and cathode electrodes are connected on the strip, respectively, andthese electrodes run perpendicularly. Therefore, µ-PIC has a 2 dimensional position sensitivity
with a ﬁne position resolution.
The small pixel structure causes a high gas gain (20000 at maximum, and 6000 in stable
operation) and a long stability (with gain of 5000 for several years by exchanging the gas once

Figure 4.1: The structure of µ-PIC (left), and the cross section of the anode and cathode (right).
41

Chapter 4. Gaseous TPC and Scintillator

(b)
(a)
(c)
cathodeanode

Figure 4.2: (a) Photograph of 10 cm ×10 cm µ-PIC. (b) and (c) are micrographs of anode andcathode.

approximately 1-2 month), because the pixel electrode is hardly damaged by discharge, which
was a serious problem for micro strip gaseous chambers. Also µ-PIC is able to work under a()
high intensity of ∼ 107 count/ sec mm2 by the smallness of the pixel.
Since a µ-PIC has a lot of readout strips (a 10 cm ×10 cm µ-PIC has 512 ch, and a 30
cm×30 cm µ-PIC has 1536ch), we have developed a readout system for µ-PIC. The schematic
view of this system is shown in Fig. 4.3.
Each of all strips connects to a preampliﬁer chip one-by-one, and the ampliﬁer feeds bothan analog signal and a discriminated digital signal. The analog signal is summed with every
16 channels on the board, and the wave form is digitized by Flash ADC (FADC) or Charge
ADC. On the other hand, discriminated digital signals are sent to the position encoding module
consisting of Field Programmable Gate Arrays (FPGA) with an internal clock of 100 MHz. The
FPGA calculates the anode and cathode coincident position within one clock interval (10 ns),
and here we say ”position encoding mode-I”. The number of the hit electrode is sent from theposition encoder to a memory module (”VME Memory Board TMB-2” produced by ARKUS) on
a VME system. We applied the low threshold level for 25 mV (corresponding to 3.6 mV at thecomparator input) supplied to the ASDs. Then input charges sometimes cause the propagation
delay at the comparator output. This propagation delay is the main reason for the counting loss
due to the narrow coincidence of 10 nsec. Thus, the small gain variation leads to non-uniformityof the signal detection eﬃciency. Then we have improved a novel position encoding, mode-II,which measures all but points of both of anodes and cathodes without coincidence. Then the
coincident position is calculated at the oﬀ-line [98]. Although we used mode-II, in the oﬀ-line
analysis, we selected only coincident event within one clock interval as the same as mode-I,
42

µ-PIC (Micro Pixel Chamber)

Figure 4.3: Data Acquisition System for µ-PIC [28]because an analysis method for mode-II is not completed.
As a preampliﬁer for µ-PIC, we use ASD (Ampliﬁer-Shaper-Discriminator) chips, which weredeveloped by KEK for Thin Gap Chamber of ATLAS experiment in LHC at CERN [99]. The
ASD chip has 4 input lines per 1 chip. Each line has a preampliﬁer, main-ampliﬁer, discriminator,
analog output and digital outputs. The preampliﬁer has a time constant of 16 nsec, and theanalog signal is an output of this preampliﬁer. The digital outputs are the discriminator outputs
in LVDS.

43

Chapter 4. Gaseous TPC and Scintillator
The gas gain of µ-PIC is shown in Fig. 4.4, where µ-PIC has a maximum gain of ∼ 20000.
Also µ-PIC have realized a good stability with the gain of ∼ 6000 for more 1000 hours. Thus,

effective gas gainµ-PIC is a stable detector with a high gas gain.

Ar 80% + C2H6 20%
Ar 90% + C2H6 10%
104

103

400

420

440

460

480

500

520

540

560

580 600 620
anode HV [V]Figure 4.4: Eﬀective gas gain as a function of the anode voltage (SN040223-1) [96].

Since a 10 cm ×10 cm µ-PIC has 65536 pixels (= 256 ch × 256 ch) and a 30 × 30 cm2
one has 589824 pixels (= 768 ch × 768 ch). The non-uniformity of the pixel structure causes a
non-uniformity of the gain on µ-PIC.
Figure 4.5 shows the gain map of the 10 cm ×10 cm µ-PIC. The Root Mean Square (RMS)of the gain variation on the whole area is ∼ 5 %, which surely indicates that µ-PIC has a good
uniformity. On the other hand, Figure 4.6 shows the gain map for the 30 cm ×30 cm µ-PIC.
The ratio of the gain between the minimum gain area and the maximum gain area is 2.2, andeffective gas gain

the gas gain uniformity is 16.7 % at the RMS.

3500
3000
2500
2000
1500

4000
3500
3000
2500
2000
1500

1000

1000

500

500

0
10
An 8
od
e[ 6
cm
]

0
700
an 600od 500
e 400
300
200
100

4
2
0 0

2

4

10
8
m]
c
[
6
e
od
Cath

0

700 600
de
catho

300
500 400

200 100

0

Figure 4.5: Eﬀective gain map of 10 cm ×10 Figure 4.6: Eﬀective gain map of 30 cm ×30
cm (SN040426-1) [96].
cm (SN041129-1) [96].

µ-PIC is used as an X-ray imaging detector. Figure 4.7 shows a spectrum of 55 Fe measured
44Gas Electron Multiplierby 10 cm ×10 cm µ-PIC. The energy resolution for X rays of 5.89 keV is 30 % at FWHM forthe whole area. Figure 4.8 is an X-ray image of a test chart irradiated by the X-ray generator
where the shadow of 2 slits per 1mm can be distinguished clearly. Using the image of the edgeof this test chart, we obtained the position resolution is ∼ 120 µm at σ. This value is quite√
equal to the predicted one from statistics, which is 400µm≃ 115 µm.
12MnKαpedestal

Ar-escFigure 4.7: Spectrum of 55 Fe (whole area of 10
Figure 4.8: X-ray imaging of a test chart withcm ×10 cm: Ar 90 % + C2 H6 10 %, 1 atm).
10 cm ×10 cm µ-PIC [26].

4.2

Gas Electron Multiplier

In order to obtain the three-dimensional track of the electron in Compton scattering, the tracking
capability for a Minimum Ionizing Particle (MIP) is required at least. The energy depositof a MIP in the argon gas is 2.54 keV/cm at the normal pressure, in which a MIP makes
approximately 3.9 electrons per 400 µm. For the detection of MIP tracks, we must detect suchfew electrons in one pixel. Although the detection of MIP is needed with a high gas multiplicationof 2 × 104 at least with the µ-PIC system, the stable gas gain of µ-PIC is only approximately
6 × 103 . Then we use a Gas Electron Multiplier (GEM) [31, 32] to compensate this gain gap.
A GEM has been developed by Sauli and other groups, which is a polyimide foil with Cuplated electrodes at both sides as shown in Fig. 4.9. A GEM foil has a lot of small holes asshown in right of the ﬁgure. If the diﬀerent voltage are supplied to Cu electrodes on both sides,
a strong electric ﬁeld is generated inside of this hole. When the seed electron drifts into thishole, the gas multiplication is caused in such a strong electric ﬁeld. The foil is very thin witha thickness of 50 µm and consists of the low-Z material. But the distance of both electrodes is
too short for a GEM foil to get a high gain alone, and hence the typical gain of single GEM isapproximately 100.
We used this GEM foil as a pre-multiplier for µ-TPC. A seed electron is multiplied by a45

Chapter 4. Gaseous TPC and Scintillator

(a)

(b)

140µm

m
c
0
1

70µm
Figure 4.9: The photograph and microphotograph GEM. The hole diameter is 70 µm, and thepitch of holes is 140 µm.

strong-electric ﬁeld between the hole of the GEM (∼ 5 kV/cm) with a gain of approximately 10(Fig. 4.10). After that, the multiplied electrons travel into µ-PIC, and second multiplication iscaused by µ-PIC with a gain of thousands. Therefore, we can take an enough signal with a gas
gain of above 2 × 104 . Figure 4.11 shows the gas gain of the GEM + µ-PIC system (Fig. 4.10).

4.3

Time Projection Chamber with µ-PIC (µ-TPC)

When a charged particle passes through a gas, the gas around the particle track is ionized,
in which a group of ionized electrons are called ”electron cloud”. If an adequate electric ﬁeld
in the gas are applied from the drift space to the µ-PIC, the electron cloud drifts along theelectric ﬁeld. At this time, the drift time from the ionization point to the termination on theanode can be transferred to the distance between the ionization point and the termination
point. Therefore, if we know the start timing of the drift, the distance between the ionization
point and the anode (z-position) is measured by the drift time. Such a system is called a Time
Projection Chamber (TPC). A two-dimensional detector, for example a Multi Wire Proportional
Chamber (MWPC), is usually used to obtain x, and y positions, and the TPC can measure
three-dimensional positions of a track of a charged particle.
µ-PIC is a two-dimensional gaseous detector, and is surely used as a readout detector of the
TPC. A TPC with µ-PIC is called µ-TPC, and the schematic view is shown in Figure 4.12. In
our advanced Compton camera, this µ-TPC is used as a tracker, which detects a ﬁne recoil trackof a recoil electron.
The electron drift velocity v is described such as:
v− =

eE
τ,
me

(4.1)
46

Time Projection Chamber with µ-PIC (µ-TPC)

(b)

(a)
drift plane
-HV

(5µm)

electron cloud
GEM

Cu
polyimide(50µm)µ-PIC-HV

Figure 4.10: Schematic view of the GEM + µ-PIC system (a), and the line of electric force in
the GEM hole (b).

Gas gain

105

104

103
430

440

450

460 470 480 490µ-PIC anode voltage [V]

500

510

520

Figure 4.11: The comparison of the gain curve of a µ-PIC with a GEM (ﬁlled circles) and thatof µ-PIC alone (open circles). (Ar 90 % + C2 H6 10 % at 1 atm, ∆VGEM = 300 V, 330 V/cm at
Drift region).

47

Chapter 4. Gaseous TPC and Scintillator

Figure 4.12: Schematic view of µ-TPC.

where, E, P , τ are the electric ﬁeld intensity, the gas pressure, and the mean free time, respectively. Figure 4.13 shows the electron drift velocity as a function of the electric ﬁeld. Since
τ also depends on E/P , v depends on E/P . This ﬁgure says that the behavior of v strongly
depends on the type of the gas even at the same E/P .
10
9

1 atm CF4

Drift velocity [cm / µsec]

8

2 atm CF47
6

2 atm Ar (90%) + C2H 6 (10%)

5
4
3

1 atm Ar (90%) + C2H6 (10%)

2
1
0
0

100

200

300

400
500
600
E field [V / cm]

700

800

900

Figure 4.13: The electron drift velocity as a function of the electric ﬁeld calculated by magboltz7.1.

Drifting electrons, which at ﬁrst localize in the ionization point at time t = 0, graduallydiﬀuse by the collisions with a gas molecule. At the time t and position x, the transverse
diﬀusion of electrons in the width of dx is described by()
N
x2
dN = √
exp −
dx,
4Dt
4πDt

(4.2)

where N is the number of electrons, D is a diﬀusion constant. The diﬀusion constant depends onthe gas condition, and also it depends on the electric ﬁeld in the case of an electron. Figure 4.14(a) and (b) show the standard deviation of the longitudinal diﬀusion and transverse diﬀusion,respectively. Since the diﬀusion of electrons limits the position resolution for the TPC, a gas
having a small diﬀusion, like CF4 , is better as a TPC gas.
48

Time Projection Chamber with µ-PIC (µ-TPC)(b)
Diffusion (transverse) [µ m / cm]

Diffusion (longitudinal) [µ m / cm]

(a)
10

3

10

1 atm Ar (90%) + C2 H6 (10%)
2 atm Ar (90%) + C2H6 (10%)

102

1 atm CF4
100

200

300

400
500
600
E field [V / cm]

700

800

1 atm Ar (90%) + C2H6 (10%)

2 atm Ar (90%) + C2 H6 (10%)
1 atm CF4102

2 atm CF4

0

3

900

2 atm CF4
0

100

200

300

400
500
600
E field [V / cm]

700

800

900

Figure 4.14: The standard deviations of longitudinal diﬀusion (a) and transverse diﬀusion (b)calculated by magboltz-7.1.

Drift cage
GEM
µ-PIC

Figure 4.15: Photograph of the TPC without gas vessel.

Figure 4.15 shows the internal parts in the TPC consisting of a GEM and a µ-PIC. The
TPC system has an enough gas volume to detect comparatively high energy X-rays, and X-ray
spectra were measured with a wider energy range than only a µ-PIC detector.
The spectrum of TPC irradiated with X-rays from a 109 Cd source is shown in Fig. 4.16(a). In this ﬁgure, (A), (B) denote ClKα and CuKα ﬂuorescence lines, respectively. Chlorine isincluded in insulation tape which is used for prevention of discharge in the TPC, while copper
is used in the GEM and µ-PIC. CdKα line is shown in (C).
Figure 4.16 (b) shows the FWHM energy resolution as a function of the energy, and thebest-ﬁt line for the data points are represented the following equation:∆E= (29.5 ± 0.8) ×
E

(

E
22keV

)−0.508±0.008
.

49

(4.3)

Chapter 4. Gaseous TPC and Scintillator

(a)

(C)

(A)

(b)

(B)

Figure 4.16: Energy spectrum of the TPC irradiated with X-rays from a 109 Cd source (a), andthe FWHM energy resolution as a function of the energy (b). (A), (B), and (C) in (a) indicate
2.6, 8.0, and 22 keV peak.

4.4

Scintillation Camera

Scintillation camera for the absorber of the ETCC is required to measure the energy and hit
position of the scattered gamma rays. Thus we selected the pixel array camera in which every
pixel is read with a position-sensitive PMT. The reason of the use and the introduction of othertypes of position-sensitive camera are described in Chapter 6.
Table 4.1 lists the properties of various scintillators. As the scintillator for an absorber, therequired properties are a good energy resolution and a high stopping power. For the use of
the pixel structure, the scintillator can be easily handled. Also, for the balloon loading, and asatellite in the future experiment, the scintillator is required to have radiation hardness. For
these purposes, we selected GSO:Ce scintillator. Table 4.1 indicates that GSO:Ce scintillator
has several good features such as a high eﬀective atomic number, a high density, a high stopping power, reasonable energy resolution, fast decay time, strong radiation hardness and nohydroscopic. Therefore GSO:Ce is suitable for the array camera as an absorber of our ETCC.
For matching the PMT pixel size with the scintillator pixel size, we selected a 5.9 ×5.9 mm
GSO:Ce pixel. Usually, we use a 13 mm long pixel due to the radiation length of GSO:Ce,while a 26 mm long one were adopted in the proton beam test in order to measure higher energy
gamma rays (1 MeV). A GSO:Ce array consists of 8×8 pixels, and a reﬂector is inserted between
each pixel (Fig. 4.17).
The GSO:Ce array was then coupled to a 64-channel multi-anode PMT H8500 with opticalgrease (OKEN6262A) (Fig. 4.17). The H8500 had an eﬀective area of 49 mm ×49 mm, an external size of 52 mm×52 mm, with a 1.5-mm thick borosilicate glass window. The photocathodein the PMT was bialkali, and the quantum eﬃciency is over 20 % at 300 - 450 nm (24 % at 420
50

Scintillation Camera

Figure 4.17: Hamamatsu multi-anode PMT H8500 (left) and GSO:Ce pixel array (right)

(a)

(b)

Y

n
ia
g

Y

X

X

Figure 4.18: (a) Top view of H8500, (b) map of the anode gain for each pixel.

51

Chapter 4. Gaseous TPC and Scintillatornm). The typical gas gain is ∼ 1.5 × 106 with a supply voltage of −1, 000 V at 25 ◦ C, and thegain uniformity of H8500 among 64 anodes is approximately 3 (Fig. 4.18).

52

NaI(Tl)
CsI(Tl)CsI(Na)BGO
LSO:CeYSO:Ce
YAP:CeYAG:Yb
GSO:CeLGSO
LaBr3 :Ce
LaCl3 :Ce
CWO
PWO
BaF2

scintillator

Density[g/cm3 ]
3.67
4.53
4.53
7.13
7.4
4.45
5.35
4.57
6.71
7.2
5.3
3.79
7.9
8.28
4.89

Table 4.1: The properties of various scintillators [100]Abs. coeﬃcient 511keV Decay time wave length light yield hydroscopic[1/cm][ns][nm](relative)
0.34
230
415
1
strong
0.44
1050
550
0.85
weak
0.44
630
420
0.90
strong0.92
300
4800.07-0.12
no
0.87
40
420
0.4-0.75
no
0.38
40
420
0.3-0.45
no
0.46
28
370
0.4
no
3
330/500
0.02
no
0.70
30-60
440
0.18
no
43
420
0.4
no
0.45
28
385
1.30
strong0.34
26385
0.7-0.9strong
0.92
5000
475
0.3-0.4no1.15< 3 < 20
430
0.04
no
0.48
0.6/620220/310
5/16
no

53

10
105−6
104−5

> 106

102−3
105
104
102−3

radiation hardness[Gray]
10
10

1.9
1.9
2.3
2.2
2.2

1.85
1.8
1.8
2.15
1.82
1.79
1.94
1.82
1.85

refraction index

Scintillation Camera

Chapter 4. Gaseous TPC and Scintillator
The 64-anode signals of the PMT were reduced to four channels with a resistor-chain board
connecting the 64 anodes with 100 Ω chips as shown in Fig. 4.19 [101]. The output signal fromeach of the four corners was preampliﬁed with an integrating time of 5.4 µs, and the signals
were shaped in a shaping time of 0.5 µs (Hoshin Electronics Co., LTD. N016-V). Then they
were digitized with a peak-hold analog-to-digital converter (ADC) (Hoshin V014). Four signals(a, b, c, and d) were used to calculate the two-dimensional positions (x, y) and the light yieldcorresponding to the energy of an incident gamma ray (E) using the following equations:(c + d) − (a + b)
a+b+c+d(a + c) − (b + d)y =a+b+c+d
E = a+b+c+d
x =

(4.4)
(4.5)
(4.6)

Here, the relation between (a, b, c, d), and (x, y) is deﬁned in Fig. 4.19.
For the balloon experiment a compact high voltage power system for PMTs is inevitable.
We selected EMCO Q12N-5 as a direct current - high voltage (DC-HV) converter (Figure 4.20),
and its speciﬁcation is listed up in Table 4.2. We control this DC-HV converter by supplying
DC level from DAC, which is operated by a VME CPU. We connect this converter to PMT
one-by-one, and control the HV of each PMT individually.
In order to extend a detection area, we construct a scintillation wall consisting of GSO:Cearrays. For example, 3 × 3 arrays were installed in the ETCC for proton beam test (see Chapter
5). The details of the performance of the GSO:Ce array camera such as reconstructed image
and the energy resolution are described in chapter 6.

54

Scintillation Camera

(c)

(a)

100Ω

(b)

y

x

Figure 4.19: Photograph and schematic view of the resistor-chain board as shown in (a) and(b), respectively. (c) shows the chain board coupled to H8500.

Figure 4.20: DC to HV convertor (EMCO
Q12N-5)55

Table 4.2: The properties of EMCO Q12N-5
parameter
value
Input Voltage
0.7 - 5 V
Input Current< 50 mA (No load)< 175 mA (Full load)
Output Voltage
0 - −1.2 kV
Output Current< 0.4 mA
Conversion Factor ∼ 60 %Ripple Noise< 0.25 %Weight4g
Size
1.27 × 1.27 × 1.27 cm3
Operating Temp.
−25 - +70 ◦ CChapter 5

Performance of a Real-Time Monitor
of Continuum Spectral Gamma Raysin Proton Therapy
5.1

Verification of Dose in Proton therapy

Radiation therapy using protons can damage a tumor eﬃciently due to its Bragg peak, while
normal cells are safe. In order to ensure this advantage and to prevent unwanted exposure, theveriﬁcation of the energy deposit only on the tumor is strongly important. At the moment,
there is no established method to observe the proton dose in the body directly.
In the current treatment planning, the calculation of the proton dose is based on the X-ray
CT data or PET imaging. However these calculations still include the uncertainty of 1 - 2 cm,
while a 1-cm-size tumor is treated.
For the safer treatment, several groups studied the application of PET imaging to the veriﬁcation whether the tumor position is irradiated with the proton beam correctly. However the
distribution of positron emitters did not match the Bragg peak position as shown in Fig. 5.1,
because the positron emitters (e.g. 11 C, 15 O, 45 Ca) have diﬀerent cross sections as described inchapter 2. In addition, each tissue consists of diﬀerent ratio of the elements. For example, bonehas more calcium(Ca) than other tissues, and much more positron emitters of 45 Ca tend to becreated. In spite of the mismatch between the annihilation distribution and the Bragg curve,
several groups have shown the rough feasibility of the PET for monitoring of proton therapy
using a phantom [21, 20, 103, 104, 105]. Figure 5.2 shows the PET imaging for a polymethyl
methacrylate (PMMA) target irradiated with 172.63-MeV proton beam, which was calibratedusing a time-activity curve for relative isotope abundances from the measurement after the endof the irradiation. Although the results showed a good agreement between measurement andprediction, the acquisition with the PET was in general prolonged over 10 to 20 min after stopping the irradiation in order to collect suﬃcient statistics up to 106 to 107 true coincidences[104].
Moreover, generated annihilation gamma rays is fewer than other gamma rays which generated by nuclear reaction during proton irradiation. This was why it could not be realistically56

Verification of Dose in Proton therapy

Figure 5.1: Measured β + -activity depth proﬁles (solid line) for proton irradiation of PMMA
targets at 140 MeV initial energy. The dotted line shows the calculated relative dose distributions[104].
(a)(b)

Figure 5.2: (a)Measured (solid) and predicted (dashed) β + -activity depth proﬁles irradiated with(maximum) energy of 171.62-MeV proton. The dotted line depicts the expected dose, calculated
by means of FLUKA. The thin solid line shows the expected range from [106], corrected for thematerial in the beam path as the simulated data. (b) PET image reconstructed from the data
exclusively measured during the about 300 sec delivery of the proton [104].

57

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy
applied in clinical use [107].
Prompt-gamma ray imaging reported as a good candidate to be a dose monitor for proton
therapy [22, 109], because the distribution of the prompt gamma rays correlates closely with theproton dose distribution, in the end of the range (fall-oﬀ) especially. Moreover use of the prompt
gamma rays is expected to provide us the accurate information of the proton dose, because the
gamma-ray emission yields are the same order of magnitude as the number of fragmentation
events [110].
Several groups measured a prompt gamma-ray distribution for the proton therapy using
not a imaging detector but only a photon counting detector with a collimator [22, 109]. For
example, Fig. 5.3 shows a prompt gamma scanner (PGS) which had three layers of shielding
against neutrons and gamma rays generated from a water phantom regarded as human body.
The paraﬃn layer moderated the high energy neutrons, and the B4 C powder captured theneutrons by the B(n,γ) reaction. Finally the lead layer prevented from the unwanted gammas.
The gamma detector was a CsI(Tl) scintillator. The PGS measured gamma ray distributions in
the water phantom irradiated with three diﬀerent energy protons (100, 150, and 200 MeV), anda parallel-plate ionization chamber (IC) also measured the beam intensity in the water in order
to ﬁnd out the correlation between the actual dose distribution and the response of the PGS asshown in Fig. 5.4 (a). Here the PGS data was selected with a gamma-ray energy of over 4 MeV,
and it was noticed that the gamma ray distribution with an energy of over 4 MeV was more
sharp as shown Fig. 5.4 (b). Fig. 5.4 (a) shows that the peaks of prompt gamma distributions
were well correlated with the Bragg peaks of the dose distributions within 1 - 2 mm at 100 MeV[22].
However, the scan by the PGS took much time to measure the dose distribution, and it was
not a practical detector for the medical imaging. Thus, a detector with a wider energy dynamic
range up to a few MeV is required for the measurement of continuum gamma rays above 1 MeV.
Therefore we need the imaging detector which can measure the continuum gamma rays.
Here, the PET does not have the ability to obtain the continuum gamma rays. Although the
SPECT is studied as prompt gamma-ray monitor using simulations [111], no one demonstrate
the monitor. In fact, the SPECT is also diﬃcult to image the continuum gamma rays, andusually their energy range is only below ∼ 300 keV. In order to detect the prompt gamma rays,
we must measure the gamma rays during the beam therapy, and there are a lot of backgrounds
out of the target owing to the nuclear interaction or scattering. Thus, classical Compton cameras
are also hard to image due to the low gamma-ray background rejection. Although Richard et al.
proposed a double scattering Compton camera as a simulation study [107], any detectors have
not yet developed.
Our ETCC is expected to satisfy the above requirement, because the ETCC has imaging
ability with a wide energy dynamic range in the sub-MeV and MeV region and the rejection ability of background gamma rays. In addition, the ETCC is expected to identify the particle fromthe measurement of the energy deposit (dE/dx) of the charged particle and three-dimensional
58

Verification of Dose in Proton therapy

Protonbeam

Collimation
hole

Scintillation
Scan PGS detector
Along zWater
phantom

B4C

Figure 5.3: Schematic view of PGS [22].

(a)100 MeV

150 MeV(b)

200 MeV

)d
ez
lia
rm
on
(
es
od
ev
it
lae
R

)d
ez
lia
m
ro
n(
es
od
ev
it
lae
R

Depth (cm)Depth (cm)Figure 5.4: The results of the PGS. (a) Comparisons of the depth-dose distributions measured
by the ionization chamber with the PGS measurements at three diﬀerent proton energies of 100,
150, and 200 MeV. (b) Distributions of gamma-ray counts as a function of depth with diﬀerent
minimum gamma energies as indicated [22].

59Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

Drift plane
Aluminum vessel
1.5-mm thickness
Aluminum vessel
17-mm thickness

Ar(90%) + C2H6(10%) 1atm
sealing

GEM ∆300 V/(50 µm)

100 mmµ-PIC

150 mm 330 V/cm

GEM

10 mm 1.0 kV / cm

50 mm
26 mm
Multi-anode PMT
Hamamatsu H8500

GSO:Ce array
166 mmFigure 5.5: Schematic view of the ETCC for the proton therapy monitor.

tracking in the gaseous TPC, which enables to reject the background under the proton irradiation. Therefore, the ETCC is expected to be used for the real-time monitoring system in proton
therapy [108], and then we tested the performance of the ETCC for the proton therapy using
a water phantom irradiated with proton beam, and we compare the gamma rays distribution
between the measurement and simulation.

5.2

Setup of the Beam Experiment Using the Compton Camera

In order to demonstrate feasibility for proton dose monitor, we measured the continuous gamma
rays from a water phantom irradiated with proton beam. Several groups reported the necessityof a detector of higher energy gamma rays over ∼ 1 MeV described in chapter 2 and theprevious section. However, our ETTC had not been studied for a higher energy gamma-ray
imaging above 1.2 MeV from an 59 Fe source. Therefore, we ﬁrst tested the high energy gamma
ray imaging using a gamma-ray point sources up to 1.8 MeV (88 Y) in the experiment room at
Kyoto University.

5.2.1

Performance of the ETCC

We have prepared an ETCC consisting of 3 × 3 GSO:Ce arrays, and a gaseous TPC as shownin Fig. 5.5. The TPC had a detection area of 10 cm×10 cm and a height of 15 cm, and the gas
volume was enclosed in a sealed aluminum vessel and ﬁlled with an argon-ethane gas mixture(9:1) at 1 atm. A drift voltage of −5500 V was supplied to the drift plane. The 50-µm-thick
GEM had an eﬀective area of 10 cm ×10 cm. The voltages of the top and bottom planes of the
GEM were set at −830 V and −530 V, respectively, for a gas gain of approximately 10. The
60

Setup of the Beam Experiment Using the Compton Camera

(a)

(b)

59.5 keV

Figure 5.6: (a) TPC energy spectrum irradiated with X-rays from an 241 Am source. (b) Energy
resolution as a function of the energy.

separation between the GEM and µ-PIC was 0.5 cm.
The µ-PIC, with an eﬀective area of 10 cm ×10 cm, was operated at 470 V with a gain of
approximately 3500. The total gas gain of the TPC was ∼ 3.5 × 104 . Figure 5.6 (a) shows theenergy spectrum irradiated with X-rays from an 241 Am source using a charge ADC, where a59.5 keV line is measured. We also irradiated the TPC with 109 Cd (22 keV), 133 Ba (31 keV),and then the energy resolution was shown in Fig. 5.6 (b). The best-ﬁt line for the data points
were represented as the following equation:()−0.32±0.01∆EE= (30.0 ± 0.9) ×
.
E
22keV

(5.1)

To obtain the high energy gamma rays eﬃciently, we used 26-mm-thick GSO:Ce arrays,
while other structure of the GSO:Ce array camera was the same as a 13-mm thick prototype
array of which detail are described in Chapter 6. 3 × 3 GSO:Ce arrays were set at 5 cm front of
the µ-PIC and the center of pixel. The average and standard deviation of energy resolution was
12.6 ± 1.4% at 662 keV (FWHM) for the all pixels, and the dependence of the energy resolution
as a function of the energy is shown in Fig. 5.7. The energy resolution was worse than the
13-mm-thick ones with a typical energy resolution of ∼ 11% at 662 keV (FWHM), because thegain of the preampliﬁer was set lower than the typical use by a factor of approximately 0.5 inorder to obtain a wider energy range up to 2 MeV at least. In addition, aspect ratio of the pixel
scintillator, which deﬁned as the ratio of the pixel-height h to the reading area A (h/A) here,
was longer than the 13-mm-thick one, which also degraded the energy resolution. The best-ﬁtline in Fig. 5.7 was represented the following equation:()−0.28±0.2∆EE= (12.7 ± 3.4) ×.
E
662keV

Using this ETCC, gamma-ray images were obtained by following analysis.
61

(5.2)

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

Energy Resolution (FWHM) [%]

20

10

7
500

600 700 800
1000
Energy [keV]

2000

Figure 5.7: Average-energy resolution (FWHM) of the 26-mm thick GSO:Ce arrays as a functionof the energy. The error bars denote the standard division of the energy resolutions among
GSO:Ce 64 pixels.

Trigger number To obtain a real Compton scattering event, one electron track in the TPC
and one hit in the GSO:Ce unit individually are required. When the two or more GSO:Cearray hits in the one Compton event, this event would be the background such as a charged
particle event, a multiple scattering event or a random coincidence. Thus, we used thescintillator information from only one GSO:Ce array units in each Compton event.
Hardware threshold In this beam test, our target energy band was more than ∼ 400 keV as
described previously,since 511-keV gamma rays are necessary to compare with the PET
imaging. Thus, we set a hardware threshold for the scintillator cameras corresponding to∼ 350 keV, while a threshold of TPC analog signal was ∼ 1 keV.
N-hit cut To obtain the electron tracking, We required three or more than three hit points ofthe electron track in the TPC (Nhit ≥ 3).
Fiducial cut The edge of the gas detection region near the drift cage or GEM tend to generatesome backgrounds. For example, the Compton scattering in the drift cage is hard toreconstruct. Then we selected an event in which all hit points in the TPC are within aﬁducial volume with an area of 7 cm ×7 cm and from 2 to 14 cm above the µ-PIC. Thiscut is also eﬀective to select the escape event of the recoil electron from the TPC.
dE/dx cut For the electron-tracking measurement in the TPC, dE/dx is calculated based on
the Bethe-Bloch formula. Figure 5.8 shows the recoil-electron track length as a functionof the energy when the ETCC is irradiated with 662-keV gamma rays. (a) and (b) are
simulation (GEANT4) results of all hit events, and only Compton events in which therecoil-electron energy is measured with the TPC, respectively. These ﬁgures suggest that62

Setup of the Beam Experiment Using the Compton Camera

(a)

(b)

]
m
c[
h
tg
n
e
l
kc
raT

]
m
c[
th
g
n
le
kc
ar
T

Electron energy [keV]Electron energy [keV]
(c)

]
m
c[
th
g
n
le
kc
ar
T

Electron energy [keV]

Figure 5.8: Recoil-electron track length as a function of the energy when the ETCC is irradiatedwith 662-keV gamma rays (simulation data) or gamma rays from a 137 Cs source (real data). (a)Simulated hit events in the TPC, (b) Compton events selected from (a), (c) raw data of measured
events with the ETCC.

the restriction on the length l and energy E of the electron (dE/dx cut) is eﬃcient for
the rejection of gamma-ray backgrounds such as random coincidence and escape events ofa recoil-electron escaping from the TPC. (c) shows the dE/dx for the experimental data,
where the selection region is drawn by red liens as represented by following inequalities:
1.671 × E 8.056×10−3+ 0.5 ≥ l

(5.3)

1.779 × E 2.100×10−3− 1.0 ≤ l,

(5.4)

where these parameters are determined by the simulation study.
Using above selected events, we obtained resultant gamma-ray images. Figures 5.9 (a) and(b) show the reconstructed images for an 88 Y point source (1836 keV) located 34 cm from thecenter of the µ-PIC surface and also at 5-cm oﬀ-center, where we reconstructed gamma rays withan energy band between 1712 - 1918 keV. These images show the ETCC had imaging ability forhigher energy gamma rays than 1 MeV.
The ETCC had an energy resolution of 15.3 ± 0.5% at 662 keV (FWHM). ARM and SPD
were (5.1 ± 0.2)◦ and (119 ± 2)◦ at 662 keV (FWHM), respectively. Figure 5.10 shows the energy
resolution as a function of the gamma-ray energy, and the angular resolutions (ARM and SPD)are shown in Fig. 5.11. This ARM ∆ϕ corresponded to an FWHM spatial resolution ∆x of
2.2 cm at 662 keV when the distance between the center of 15-cm-high TPC and the target
was 25 cm 1 . In addition, the position accuracy in the center position was determined to be1∆x = 25 × tan(∆ϕ/2) × 2

63

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

(a)

(b)

]
cm
[
Y

]
m
c[
Y

X [cm]

X [cm]

Figure 5.9: Gamma-ray images obtained with the ETCC for an 88 Y point source at the center(a) and 5-cm oﬀ center. Cross marks denote the target point.

approximately 3 mm (FWHM) at 835 keV by accumulating a few hundreds events from a 54 Mn

200

20

Angular resolution (FWHM) [degree]

Energy Resolution (FWHM) [%]point source.

100

10

20

10

4

9
500

600 700

1000
Energy [keV]

500

2000

600

700

800 900 1000
Energy [keV]

2000

Figure 5.11: FWHM angular resolution of the
Figure 5.10: FWHM energy resolution of the
ETCC as a function of the energy. (ﬁlled cirETCC as a function of the energy.
cle):ARM, (open circle):SPD.

5.2.2

Setup of Beam Test

Figures 5.12 and 5.13 show the setup for the proton beam test using ”ES Course” at ResearchCenter for Nuclear Physics (RCNP) of Osaka University. The beam energy was 140 MeV. The
proton beam target was a water phantom with a volume of ϕ20 cm ×30 which made frompolymethyl methacrylate (PMMA) with a thickness of 0.5 cm except a ϕ20-cm surface in the
opposite side of beam entrance (1-cm thickness). Roughly the water phantom was regarded as
human body which consists of approximately 70 % water. From GEANT4 calculation, the water
depth in the Bragg peak position was 12 cm, taking into account the obstacles such as the wallof water tank and the ﬂange of the vacuum tube made of PMMA. Then the TPC center was set
in front of the Bragg peak.

64

Setup of the Beam Experiment Using the Compton Camera

140 MeV Proton
Vacuum tube
PMMA seal
Water tank
(PMMA)

1 cm
0.5 cmWater
12 cm phantom
30 cm
φ20 cm1 cm

3×3
GSO arrays

TPC
15 cm

ϕ

z

eGEM µ-PIC GSO array

ARM drift plane

x

17 cm

10 cm

34 cm

5 cm

Figure 5.12: Schematic view of the setup for proton beam test at RCNP.

(a)

(b)
y

protonWaterphantom

Protonz

Waterphantom

TPC
x

TPC

Figure 5.13: (a) Photograph of the setup, (b) deﬁnition of the axis.

65

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

8 cm

10 cm

1 cm x 2

Figure 5.14: Photograph of plastic scintillators for a beam current monitor.

5.2.3

Beam Current Monitor

Since a beam current monitor as a laboratory equipment in the beam line does not work below
the sub nA range, we set two monitor sets named #1 and #2. Each set consists of two plastic
scintillators (ELJEN EJ-200) with a size of 8 cm ×10 cm ×1 cm read by a single anode PMT(Hamamatsu R1166A, ϕ 19-mm tube) which was operated with a gain of ∼ 106 (Figs. 5.14 and5.15). In each monitor set, the coincidence event of two plastic scintillators was counted, andthey measured the relative proton beam current indirectly.
These monitors were installed at two points as shown in Fig. 5.16. Here #1 usually measured
the relative beam current during a Compton camera imaging, while #2 was used for a high countrate proton beam for which a current monitor equipped in RCNP was available. In order tocalibrate the relation between the coincidence count rates of each plastic scintillator monitor andthe proton beam intensity, at ﬁrst #2 was calibrated by the monitor equipped in RCNP. Then
using the correlation between #1 and #2, we roughly estimated the absolute current during theCompton imaging roughly.
Before the beam experiment in RCNP, we tested the beam monitor in Kyoto University.
Each monitor set was irradiated with gamma rays from a 60 Co source (1173, 1333 keV) withseveral intensities by use of sources with diﬀerent activities or change of the distance. Figure5.17 shows the count rate as a function of estimated gamma-ray intensity. Best-ﬁt lines for thedata points in the Figs. 5.17 (a) and (b) were represented the following equations:count rate = (2.591 × 104 ) · I 1.228(for #1)(5.5)

count rate = (3.267 × 104 ) · I 1.306(for #2),(5.6)

where I is gamma-ray intensity. Although the gamma-ray intensity and count rate were out of66

Setup of the Beam Experiment Using the Compton Camera

Plastic scintillator

discriminatorcoincidence

scaler#1#2

Figure 5.15: Schematic view of set up.

Water tank

12cm
Bragg peak
Plastic
Scintillator set #1

23.5 cm

6 cm

stage

Protpn beam
Vacuum tube
flange cap(PMMA)

40 cm
25cm
56.5cmPlastic
Scintillator set #2

Figure 5.16: Schematic view of set up in the RCNP.

67

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

(b)

104

104

3

10

10

Count rate [Hz]Count rate [Hz]

(a)

102

3

102

10

10
10-2
10-1
Gamma-ray intensity [MBq]

10-2
10-1
Gamma-ray intensity [MBq]Figure 5.17: Count rate as a function of the gamma-ray intensity for the beam experimentmonitor set # 1(a) and 2 (b).

proportion with a linear function, this results result showed a strong correlation between them.
Thus these scintillators had the feasibility for the beam monitor.

68

Results of the Beam Test

Table 5.1: Beam time schedule
day
Beam start Beam Stop time
April 28
10:02
12:1
2:10
12:22
13:09
0:46
13:19
16:32
3:13
17:42
18:52
1:10
19:10
19:24
0:13
20:04
20:16
0:12
22:30
1:34
3:04
April 29
1:42
7:05
5:22
7:22
8:48
1:26
8:54
13:51
4:56
14:00
14:44
0:44
14:56
20:01
5:04
20:09
20:26
0:17
0:37
1:23
0:45
April 30
1:30
8:59
7:29

5.3

Results of the Beam Test

We irradiated the water tank with a 140-MeV proton beam on April 28 - 30 in 2010 as shownin Table 5.1, where the beam did not continue the whole time due to the maintenance of thebeam accelerator or entrance into the beam room by other groups. Figure 5.18 shows the countrate of the plastic scintillators (the proton-beam current monitors) #1 and #2 during the beamon and oﬀ as a function of elapsed time from the beam start at 10:02 on April 2010, and thesampling time was 10 sec per a one measurement. Sometimes data acquisitions were in failure,and spike peaks were seen in the Fig. 5.18. These monitor #1 and #2 had averages and standard
deviations of count rates of 107 ± 13 and 2.1 ± 0.5 Hz, respectively.
The current monitor # 2 were calibrated by a laboratory equipment of the current monitor
in the RCNP using ∼ 1-nA proton beam (Fig. 5.19), and the count-rate correlation between
monitor # 1 and 2 are shown in Fig. 5.20. Best-ﬁt lines for the data points in the Figs. 5.19
and 5.20 were represented the following equation:
p2 = (4.535 × 103 ) × R1.366
p2 = (2.732 × 10−2 ) × p10.8326(for Fig.5.19)(for Fig.5.20),

(5.7)
(5.8)

where p1, and p2 Hz denote count rate of # 1, 2, respectively, and R nA is proton beam current
measured by the RCNP’s current monitor. This results showed that the beam current during
the measurement with the ETCC was ∼ 2 pA.
For the ETCC, the beam duration time was 32 hours (live time), and dead time due tothe data acquisition was 6 %. Figure 5.21 shows the energy spectrum of the ETCC during thebeam time. This spectrum had a 511-keV peak. However other peaks such as 2.2-MeV neutron
capture peak were not detected, because of large background gamma rays.
69

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

140Count rate [Hz]

120
100
80
60
40
20
0
0

10

20

30
Time [hours]

40

50

Figure 5.18: Count rates of plastic scintillator monitor #1. start time (time 0) is 10:02 on April
28.

103

#2 Count rate [Hz]#2 Count rate [Hz]

105

104

103

102

10

1
1

2
3
4
Beam current by RCNP current monitor [nA]

5

103

104
#1 Count rate [Hz]105

Figure 5.19: Count rates of plastic scintillator Figure 5.20: Count rates of plastic scintillator#2 as a function of the proton beam current #2 as a function of the count rates of plastic
measured by RCNP’s current monitor.
scintillator #1.

70Results of the Beam Test

511 keVFigure 5.21: Energy spectrum of the ETCC.

Figure 5.22 shows the water phantom imaging with the ETCC. In the region from 463 to 559
keV, which corresponds to the twice width of a FWHM energy resolution of the ETCC at 511
keV, we obtained 890 ﬁnal events, while in the 0.8 - 2 MeV range, 1288 events were detected.
We succeeded in obtaining the individual images for two diﬀerent energy band (463 - 559 keV
and 0.8 - 2 MeV).

71

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

(a)

(b)

20

Water tank
]
m
c[ 0
Y

-20
-20

0
X [cm]

)t
i
n
u
yr
ar
ti
b
ra
(
t
n
u
o
C

-20

20

(c)

0
X [cm]

20

0
X [cm]

20

(d)

20

Water tank
]
m
c[ 0
Y

-20
-20

0
X [cm]

20

)t
i
n
u
yr
ar
ti
b
ra
(
t
n
u
o
C

-20

Figure 5.22: Imaging results of the proton beam test. Left side images are reconstructed images
with an energy band of 463 - 559 keV (a) and 0.8 - 2 MeV (b), respectively. (b) and (d) are image
proﬁles on the beam axis, indicated by the dashed square in (a) and (c), respectively. Dottedlines and arrows denote the Bragg peak and the direction of the proton beam, respectively. X
and Y axes are deﬁned in Fig. 5.13 (b).

72Analysis of Imaging

5.4

Analysis of Imaging

5.4.1

Gamma-ray Distribution in the Water Phantom

For the discussion of the imaging results, we used the GEANT 4 simulation (version 9.3.patch01). Here our simulation code was based on the hadron therapy as an example consisting of thehadron physics package, ”G4HadronElastic” and ”Binary Cascade model”. In addition, we used
Livemore Library as a low energy model, because low energy Compton scattering (∼ 100 - 300keV) also occurs in the ETCC. The set up for the simulation was drawn in Fig. 5.12 which isschematic view of a real set up.
Since a full-simulation was estimated to take so long time of more than a year using our fourPCs (CPU ∼ 3 GHz × 4), not super computers, we only simulated the interaction between theincident proton beam and some atoms in the air, the water phantom, and the ETCC. Here, thefull-simulation includes also the reconstruction of the image with the ETCC using its imaging
method and performances such as energy resolutions and some noises in the electronics.
When the water phantom was irradiated with 140-MeV proton beam, the gamma rays weregenerated as shown in Fig. 5.23, which was distributions of generated gamma rays in the allenergy, 6 - 10, 0.8 - 2, and 0.50 - 0.52 MeV. Slice images of the Figs. 5.24 are shown in Fig.
5.23. Although higher energy gamma rays formed a sharper peak near the Bragg peak, thedistribution of 0.8 - 2-MeV gamma rays, especially the fall-oﬀ, also correlated to the Bragg
peak. However, the distribution of around 511-keV gamma rays was found to be a more diﬀuse
image.

5.4.2

Method of Gamma-ray Imaging with the ETCC

In order to obtain the realistic results for the ETCC measurement, we processed the position
data of incident simulated gamma rays in the TPC using some statistical weights or detector
performance as followings:
1. energy resolution of the ETCC
2. detector responses function
3. point-spread function
4. detection eﬃciency as a function of energy
The response function represented the sensitivity as a function of the detection position in
the ETCC. The center of a ﬁeld-of-view was higher sensitivity than the edge, and the countrate depends on the source position. To make a response function, I used the real point source
images measured by the ETCC. Figures 5.25 and 5.26 showed the ETCC response functions
along x and y axes, respectively. Those functions represented the dependence of count rate on
the position, and they were made by a 54 Mn point source imaging in the several points on the73Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

(b)

(a)

10

8

16000

8

7000

6

14000

6

6000

4

12000

4

2

10000

0

X [cm]

X [cm]

10

8000

-2
6000

-4

4000

-2

3000

-4

4000

-8

2000

-8

-10 0

0

-10 0

10
15
20
Y (Depth) [cm]

25

30

2000

-6
1000

5

10
15
20
Y (Depth) [cm]

25

30

0

(d)

(c)

10

10

1600

8

8
1400

6
4

1200

4

2

1000

2

0

800

-2

80

60

0
-2

600

40

-4

-4
-6
-8
-10 0

100

6

X [cm]

X [cm]

0

-6

5

5000

2

5

10
15
20
Y (Depth) [cm]

25

30

400

-6

200

-8

0

-10 0

20

5

10
15
20
Y (Depth) [cm]

25

30

0

Figure 5.23: Distributions of generated gamma rays with an energy of all energy (a), 6-10 (b),0.8 -2 (c), and 0.50 - 0.52 MeV (d). Dotted lines indicate the Bragg peak position.

105
Counts (arbitary unit)

(A)

(B)

104

(C)

(D)
103
0

5

10
15
20
Depth in the water [cm]25

30

Figure 5.24: Slice image in the Fig. 5.23 from x = −9 to 9. (A) to (D) denote the data of (a)to (c) in the Fig. 5.23, respectively. Dotted line indicates the Bragg peak position.

74

Analysis of Imaging

x-axisy-axisTable 5.2: Fitting parameters in equation 5.9
a
b
c
d
e
f
0.002096 -0.4345 5.952 0.00136 -0.05418 19.92
0.001791
1.198
6.86 0.00087
1.496
15.67

g
3.96 × 10−5
3.989 × 10−50.0035Count rate [Hz]0.003
0.00250.002

0.00150.001
0.0005

-100

-50

0
X [cm]

50

100

Figure 5.25: Response function for x-axis.

0.003

Count rate [Hz]0.00250.002

0.00150.001

0.0005

-100

-50

0
Y [cm]

50

100

Figure 5.26: Response function for y-axis.

x and y axes. Phenomenological best-ﬁt lines for the data points in Figs. 5.25 and 5.25 wererepresented by the following equation which is double-Gaussian function:
(
(
(
) )
(
) )
x−b 2
x−e 2
count rate = a exp −0.5
+ d exp −0.5
+ g,
c
f

(5.9)

where a - g are ﬁtting parameters, and each value was listed in Table 5.2.
Using the two best-ﬁt function fx and fy , we assumed that the two-dimensional response
function f2 is represented as following equation and shown in Fig. 5.27 Then simulated point
was multiplied by the two-dimensional response function.
()(fx (r) + fy (r))|x| − |y|f2 =+ 0.5 (fx (r) − fy (r)) ·,2|x| + |y|√where r = (x2 + y 2 ).

(5.10)

Point-spread function (PSF) represents the performance of the imaging. Using this function
as a probability density function, a simulated point position was spread into the real image.
75

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

]
cm
[
Y

X [cm]

Figure 5.27: Map of the two-dimensional response function.

PSF was also made by point source imaging. Figure 5.28 shows the reconstructed imaging ofa 54 Mn point source at the center of the TPC (x = y = 0), and histogram of its slice image.
Here, the slice image was ﬁtted with a best-ﬁt function fcenter phenomenologically as describedfollowing equation consisting of Gaussian G and Lorentzian L
(
(
) )
)
(
p3 /2
p1
x − p5 2
fcenter =+ p4 exp −0.5
π (x − p2 )2 + (p3 /2)2
p6:= L(x, p1,2,3 ) + G(x, p4,5,6 ),(5.11)(5.12)

where p1 to p6 are ﬁtting parameters.
However, Fig. 5.29 shows 54 Mn point source images from the various positions along x-axis,
and their functions had an asymmetry form except the center position (x = y = 0) of thepoint source in Fig. 5.28. Thus phenomenological ﬁtting was used following equation instead ofequation 5.12.
()
p7
fpointspread = p8 t1 L(x, p1,2,3 ) + t2 L(x, p1,2,7 ) + G(x, p4,5,6 ) ,p3(5.13)

where (t1 , t2 ) are (1, 0) for the 0 < x < p2 or 0 > p2 > x, and (0, 1) for the otherwise.
Here, when we ﬁtted the slice images in Fig. 5.29, we assumed that only p7 and p8 are freeparameters and the other are ﬁxed parameters which is obtained by ﬁtting in the center position
source imaging using the equation 5.12. In addition, the components of the Gaussian ﬁtting was
ﬁxed.
76

Analysis of Imaging

(a)

(b)
FWHM：9.2㎝

)t
in
u
rya
rt
ib
ra
(
tsn
uo
C

]
m
c[
Y

X [cm]

X [cm]

Figure 5.28: (a) Reconstructed image of a 54 Mn point source and (b) a projection image fromsquare region in the (a).
(a-1)

(b-2)

(b-1)

(a-2)

30

30

0.05

0.05
20

20
0.04

-10

0.03
0.02

-30
-30

-20

-10

0
X [cm]

10

20

-20

-10

0
10
X [cm]

20

30

-20

-10

0
10
X [cm]

20

0
-30

30

(d-1)

-10

0
X [cm]

10

20

30

-10

0
10
X [cm]

20

30

-10

0
10
X [cm]

20

30

(d-2)

0.05

0.05
20

20

-10

0.03
0.02

-30
-30

-20

-10

0
X [cm]

10

20

-20

-10

0
10
X [cm]

20

0.02
0.01

-30
-30

30

(e-2)

(e-1)

0.03

-20

0
-30

30

0

-10

0.01

-20

Counts (a. u.)

0

0.04
10
Y [cm]Counts (a. u.)

0.04
10

-20

-10

0
10
X [cm]

20

0
-30

30

0.05

0.05
20

20
0.04

-10

0.03
0.02

-20

-10

0
10
X [cm]

20

30

0
-30

0

-10

0.01

-20

Counts (a. u.)

010
Y [cm]Counts (a. u.)

0.04
10

-30
-30

-20

(f-2)

(f-1)
30

30

Y [cm]

-20

30

30

Y [cm]

0.02
0.01

-30
-30

(c-2)

(c-1)

0.03

-20

0
-30

30

0

-10

0.01

-20

Counts (a. u.)

Y [cm]

0

10
Y [cm]Counts (a. u.)

0.04
10

-10

0
10
X [cm]

20

-30
-30

30

0.02
0.01

-20
-20

0.03

-20

-10

0
10
X [cm]

20

30

0
-30

-20

Figure 5.29: 54 Mn point source images from various position and their projection images. The
setting points of the 54 Mn point source were (a):(x [cm],y [cm]) = (5, 0), (b):(10, 0), (c):(15, 0),(d)(-6,0), (e)(-10, 0), (f)(-15,0).

77

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

(a)

(b)
22

18
20
16

18
16
p7

p7

14
12

14
12

10
10
8

8

6

6
-15

-10

-5

0
X [cm]

5

10

-15

15

-10

-5

0
Y [cm]

5

10

15

Figure 5.30: Parameter p7 s as a function of the source position for x (a) and y-axis (b).

From the ﬁtting results, the parameter p7 s as a function of the source position along both x
and y-axes (Fig. 5.30) were ﬁtted using quadratic functions which are represented the followings
p7 (x) = 0.0446x2 + 0.114x + 6.36

(for x − axis)(5.14)

p7 (y) = 0.0586y 2 + 0.182y + 5.51

(for y − axis).

(5.15)

Using the equations 5.14, 5.15, distribution function was described all position along x andy-axes, in fact, Fig. 5.29 shows the ﬁtting results with only a free parameter p8 . In addition,using an assumption of equation 5.10 for the equation 5.14, 5.14, simulated gamma-ray position
was spread to the realistic image.
Figure 5.31 shows the energy dependence of the point image for 22 Na (511 keV), 137 Cs (662keV), 54 Cs (835 keV), (d) 22 Na (1275 keV), and (e) 88 Y (1836 keV). Like a 54 Mn, these images
were ﬁtted with fcenter as shown the equation 5.12, where the Gaussian G part was ﬁtted byﬁxing the same parameters as the ﬁtting for a 54 Mn source image. Figure 5.32 shows the ﬁtting
results for slice image of various point sources along x and y-axes. Here, in some energy except
the point source energy (511, 662, 835, 1275, and 1835 keV), these parameters were obtainedby the arithmetic mean of the parameters between the two point source images. Using theparameters (p1 to p6 ) and equations 5.14, 5.14, a simulated point was also spread to realistic
image.

78

Analysis of Imaging

30
20

30

(a)

20

Y [cm]

0

0

-10

-10

-20

-20

-30
-30

Y [cm]

10

-20

-10

0
X [cm]

10

20

-30
-30

30

30

30

(c)
20

20

10

10
Y [cm]

Y [cm]

10

0

-10

-20

-20

-20

-10

0
X [cm]

10

20

30

-10

0
X [cm]

10

20

30

-20

-10

0
X [cm]

10

20

30

-10

0
X [cm]

10

20

30

(d)

0

-10

-30
-30

(b)

-30
-30

-20

30
20

(e)

Y [cm]

10
0
-10
-20
-30
-30

-20

Figure 5.31: Several point source images at x = y = 0. (a) 22 Na (511 keV), (b) 137 Cs (662 keV),(c) 54 Cs (835 keV), (d) 22 Na (1275 keV), and (e) 88 Y (1836 keV).

79

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy

0.05
0.045

(a-X)

0.045

(a-Y)

0.04

(b-X)

0.05

(b-Y)

0.04
0.04

0.035

0.04

0.02

0.03

0.02

0.01

0.02

0.01

0.01

0.03

0.02

0.01

0.005

0.005

0.05

0.025

0.015

0.015

0
-30

0.03

Counts (a. u.)

0.03
0.025

Counts (a. u.)Counts (a. u.)Counts (a. u.)

0.035

-20

-10

0
X [cm]

10

20

0
-30

30

0.07

(c-X)

-20

-10

0
Y [cm]

10

20

0
-30

30

(c-Y)

0.03

-20

-10

0
X [cm]

10

20

0
-30

30

(d-X)

0.03

-20

-10

0
Y [cm]

10

20

30

-10

0
Y [cm]

10

20

30

(d-Y)

0.06

0.025

0.025

0.02

0.04
0.03
0.02

0.01
0.01
0
-30

0.024
0.022

-20

-10

0
X [cm]

10

20

0
-30

30

-20

-10

0
Y [cm]

10-100
Y [cm]

10

20

0.02
0.015

0.02
0.015

0.01

0.01

0.005

0.005

0
-30

30

Counts (a. u.)

0.03Counts (a. u.)

0.05
Counts (a. u.)Counts (a. u.)

0.04

-20

-10

0
X [cm]

10

20

30

0
-30

-20

0.025

(d-Y)

(d-X)

0.02

0.02

0.016

Counts (a. u.)Counts (a. u.)

0.018

0.014
0.012
0.01

0.015

0.01

0.008
0.006
0.005

0.004
0.002
0
-30

-20

-10

0
X [cm]

10

20

30

0
-30

-20

20

30

Figure 5.32: Slice image of various point sources along x (X) and y-axes (Y) and their ﬁtting
results. (a) 22 Na (511 keV), (b) 137 Cs (662 keV), (c) 54 Cs (835 keV), (d) 22 Na (1275 keV), and(e) 88 Y (1836 keV).

80

Analysis of Imaging

(a)

(b)

10

10

]
m
c[0
Y

]
m
c[0
Y

-10
-20

-10

0

X [cm]

10

-10
-20

20

-10

0

10

20

X [cm]

Figure 5.33: simulated images for water phantom irradiated with 140-MeV proton beam. Energy
bands are 0.8 - 2 MeV (a) and 463 - 559 keV (b). Broken lines denote the Bragg peak position.

5.4.3

Gamma-ray Imaging

We obtained the simulated images, and Figs. 5.33 (a) and (b) show the images in the gamma-ray
energy band of 0.8 - 2 MeV, and 463 - 559 keV, respectively. Their slice images is shown in
Fig. 5.34. These results indicated that a peak of gamma-ray distribution for 0.8 - 2 MeV wassharper than for 463 - 559 keV, and this result was consistent with the real data. In addition,
both the real data and simulation were found that the position of the gamma-ray peaks for 0.8- 2 MeV were in upstream of the proton beam at ∼ 1 cm from the Bragg peak, while the ﬂatpeak (∼ 4-cm width) for 463 - 559 keV was around the TPC center. However, these gamma-ray
distributions between the real data and simulation were not the same perfectly. The simulated
distributions were sharper than real ones, because we did not consider the random coincidenceby neutrons or gamma rays in the simulation. Now we optimize the simulation including therandom coincidence.
In Fig. 5.33, we considered reaction functions of the ETCC, F (X, Y, E) including PSFs, andthe reaction functions depended on the emission position and the gamma-ray energy (Fig. 5.35).
Conversely, the real map of the gamma-ray emission distribution can be reconstructed fromthe measured images with the ETCC using the inverse reaction function F −1 (X, Y, E). In deed,some reconstruction methods corresponding to F −1 (X, Y, E) such as the maximum likelihood
expectation maximization (ML-EM) algorithm [162] is available in medical imaging such as PET
and SPECT. Figure 5.22 was not applied to F −1 (X, Y, E), and the deconvolution to the ETCC
image is also required like PET and SPECT images.
In medical imaging, the line gamma-ray source is used, while we need detect continuum
gamma rays. Thus the inverse function F −1 (X, Y ) for medical imaging dose not depend on
only the position. However, these reconstruction methods have not been applied to continuum81Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy(a)Bragg peak

Bragg peak

(b)
12

)t
in
u
yr 8
ar
ti
br
a(
st4
nu
oC

t)i 3000nuyr
ar 2000ti
bra(
st 1000
nu
oC
0

0

-20

-10

0

X [cm]

10

20

-20

-10

0

X [cm]

10

20

Figure 5.34: (a)simulated slice images for Fig. 5.33. (b) measured slice images for Fig. 5.22.
(dotted line:0.8 - 2 MeV, solid line:463 - 559 keV).

gamma-ray image, and it is diﬃcult to optimize the inverse function F −1 (X, Y, E). Therefore
I study the inverse function F −1 (X, Y, E) for proton therapy using point gamma-ray source
images with various energies that I set and measured in various positions.
Due to a wide energy range, we succeeded in imaging for higher energy band and around
511 keV simultaneously, and we obtained the diﬀerent images between the high energy prompt
gamma rays and low energy including annihilation gamma rays. However the spatial resolutionof the ETCC was 2.2 cm (FWHM) at 835 keV in this experiment, and it was not suﬃcient toestimate quantitative position and to use in medical. At least the spatial resolution is required tobe 1.0 cm in medical, Therefore we should improve the angular (spatial) resolution of the ETCC,
and we have developed an ETCC with a better angular resolution as described in Chapter 6.

5.5

Future Works

For proton therapy and astronomy, we need an ETCC with a higher detection eﬃciency. In
addition, the ETCC should measure the higher energy gamma rays (≥ a few MeV) with asuﬃcient detection eﬃciency in order to verify the proton dose more clearly. However, we did
not measure higher energy gamma rays more than 2 MeV, because of the low eﬃciency of the
ETCC. Then we improve several factors for the detection eﬃciency as written followings.
Multi-head imaging Measurement of a target using several ETCCs simultaneously is not only
eﬀective to have higher detection eﬃciency, but to obtain a three-dimensional imaging.
Larger size ETCC We have developed a larger size TPC up to 40 cm-size by combining four
20-cm-size µ-PICs.
82

Future Works

F-1(X, Y, E)
]
m
c[
Y

]
m
c[
Y

Bragg peak

protonBragg peak

protonX [cm]

X [cm]

F(X, Y, E)
Figure 5.35: Ideal gamma-ray and simulated image with the reaction function of the ETCC.

Optimization of the gas Optimizations of the gas such as CF4 -based gas, and its pressure
are also eﬀective to obtain the higher detector eﬃciency. Actually, a detection eﬃciencies
of 356-keV gamma rays were 2.44 × 10−5 , 3.51 × 10−5 , and 1.81 × 10−5 for CF4 -based
mixture gas (Ar (54% ) + CF4 (40% ) + isoC4 H10 (6% )) at 1, 1.4 atm, and Ar (90% ) +C2 H6 (10% ) gas at 1 atm, respectively [113].
Optimization of data acquisition logic The coincidence between the anode and cathode hit
information in the TPC is sometimes not eﬀectively detected, and we improve a novel logic
as shown in section 4.1.
In this beam test, we used low current proton beam than usual current in the therapy (∼ 1nA) by an order of 2 to 3, because of keeping imaging quality due to the random coincidence.
In this test, we operated with an ETCC count rate of approximately 20 − 30 Hz, and the deadtime of the ETCC is over 50% when the count rate become approximately 50 Hz. In order to
use the ETCC as clinical applications in proton therapy, we must overcome to endure the highcount rate, and reduce the random coincidence. Then the solutions and improvement points asa future works are shown as followings.
(a) Optimization of the gas and data acquisition logic The drift time of the recoil electron is up to 4 - 5 µsec in the TPC (10 cm ×10 cm ×15 cm), Thus in the FieldProgrammable Gate Array (FPGA) code of the encoder, we now set a time window of
8 µsec after the trigger to determine (hereafter referred to as ”drift waiting”). The total
dead time, deﬁned as the veto time duration, was 10.5 µsec per a trigger. CF4 -based gas(e.g. CF4 + Ar + C2 H6 ) has a faster drift velocity by roughly 2 times as compared to thecurrent setup (Ar + C2 H6 gas). Thus, we can reduce the ”drift waiting” in the TPC by afactor of ∼ 2.
83

Chapter 5. Performance of a Real-Time Monitor of Continuum Spectral Gamma Rays in Proton Therapy(b) Pixel reading for scintillation array In order to prevent random coincidence, it would
be eﬀective to read each pixel scintillator without a resister chain board and charge-division
method described in section 4.4 .
(c) Optimization of the signal digitalization of scintillation A charge analog-digital converter (CADC) is expected to read signals from scintillator faster than a peak-hold ADC(PHADC), because the gate duration of CADC is shorter. (PHADC: a few µsec, CADC:∼ 100 nsec). Although the PHADC is used generally in our ETCC, we will test the CADC.
(d) Hardware threshold / Lead shield for the monitor of proton therapy, the target gamma
rays has higher energies (≥ 1 MeV). Thus rejection of the lower energy gamma rays with ahardware threshold or lead shield in front of the TPC is eﬀective to reduce the dead time.
(e) Detection of the energy distribution along the electron track The energy of the recoilelectron is measured only total energy for a one Compton event now. When the energies
are measured at each hit point in the µ-PIC, we obtain the Bragg curve of the recoilelectron Compton event-by-event. Thus, the µ-PIC and TPC would discriminate between
the recoil-electron event and the random coincidence, and then the tracking ability is
expected to improve.
(f ) Trigger Since scintillation hit is used for an ETCC trigger, it is likely to be contaminationof the background gamma rays coming out of the TPC. Here, Gas scintillation light in the
TPC might be used as the trigger.
We already began to test the solutions (a) and (b). As a ﬁrst step of (a), we measured apoint source with an ETCC containing the above CF4 -based at 1 and 1.4 atm with the samedata acquisition logic which is ”drift waiting” of 8 µsec. The FWHM ARMs were 11.7 ± 0.4,
11.2 ± 0.3, and 10.4 ± 0.7 for the CF4 -based gas at 1, 1.4 atm, and Ar (90% ) + C2 H6 (10% )gas at 1 atm, respectively [113]. The ETCC with CF4 -based gas did not have a worse angularresolution remarkably. Therefore, these gas is expected to reduce the dead time.
For the solution (b), we have developed a 64-ch read out system named Head Amp Unit
CP80190 (Clear Pulse) as shown in Fig 5.36, and we measure each pixel signal from an 8 × 8
LaBr3 :Ce array coupled to an 8 × 8 multi-anode PMT. The FWHM energy resolution was 5.4%
at 662 keV, and it is similar resolution as compared to a conventional reading as described in
Chapter 4 and 6. As a next step, we estimate an ETCC using the this system. The solution (c)is not diﬃcult technique. However, (c) is not needed when we succeeded in the improvementwith (b), because the 64-ch read out system include the ADC.
(d) is also not diﬃcult technique, and we tested it in the next experiment. The solution (e)have been developed, and we have already applied to the neutron imaging [114]. Then we apply
the solution (e) to the ETCC. The solution of (f) is just diﬃcult technique, because the gas
scintillation light is so week, and the spark light near the GEM or µ-PIC might contaminate the84Future Works

140 mmFigure 5.36: Photograph of the 64-ch read out system named Head Amp Unit CP80190 (Clear
Pulse).
gas scintillation light. Due to the discrimination of the gas scintillation light eﬃciently, many
photon detectors might be required in the TPC, because the distribution of the gas scintillation
emission is diﬀerent from the that of other noise. Thus we research and develop the a novel
trigger using some photon detectors such as a PMT, an avalanche photodiode, and Multi-Pixel
Photon Counter.
From the above improvement, the ETCC is expected to endure the beam current with over
150 times higher than that of current measurement. This will allow the ETCC to monitor thelocation after a few seconds at the beam current of ∼ 0.5 − 1 nA. As a result, the ETCC would
be used in proton therapy after the further developments.

85

Chapter 6

Improvement of Scintillator
6.1

Position-Sensitive Scintillation Camera

In this section, I have developed a new scintillation camera with position sensitivity, in orderto obtain a better angular resolution of the ETCC. At ﬁrst, I brieﬂy mention which type of theposition-sensitive scintillation camera is suitable for the ETCC.
Position-sensitive scintillation cameras are used in various ﬁelds such as medical imaging [115]and gamma-ray astronomy [116, 117, 118]. Figure 6.1 shows the various types of a positionsensitive scintillation camera as listed below.
(a) Anger-logic camera [119, 120](b) Photomultiplier tube (PMT)-Quadrant-Sharing technology [121, 122](c) Planar type [123](d) Array camera consisting of a scintillator array in which every pixel is read with a positionsensitive PMT [124, 125]
The Anger logic camera consisting of a large one crystal plate and PMTs as shown in Fig. 6.1(a), measures the position by calculating a weighted centroid of the charge deposited in several
photo-sensors such as a PMT. This algorithm looks simple, but causes several problems of thespatial non-linearity, especially near the image edges [125]. Figures 6.1 (b) and (c) are one of
the improved Anger-logic cameras, which have pixellation structures. These structures limit the
scintillation light path, and hence improves the spatial non-linearity at some level. This crystal
with the structures is called ”planar” (Fig. 6.1 (c)). In particular, when the positioning system,
which uses four single-anode PMTs on the four corners of the crystal, is called PMT-QuadrantSharing technology (Fig. 6.1 (b)). Pixel array type measures the position of a hit pixel directlyfrom the hit position of a position-sensitive detector as shown in Fig. 6.1 (d).
In this paper, I describe a crystal without pixellation structures as a ”continuous crystal”.
The continuous crystal is read with Anger-logic, and it has the advantages of structural simplicity, and lower cost. On the other hand, a pixel array camera has the advantages of an intrinsic
spatial resolution dictated by crystal pixel width. Emitted light was focused within the pixel86

Position-Sensitive Scintillation Camera

PMT

Figure 6.1: Schematic view of position-sensitive scintillation camera. (a)and (a’) are Anger logic
system using single anode PMTs and a multi-anode PMT, respectively. (b), (c), and (d) are
PMT-Quadrant-Sharing, planar, and pixel type, respectively.

area via a reﬂector [126]. In addition, a thick continuous crystal has a three-dimensional position
sensitivity because PMT signals in an Anger camera have a marked dependence on the depth of
gamma-ray interactions [127]. However, a thick continuous crystal generally has a worse spatial
resolution near the edges of the detector than near the center [125, 128, 130]. Approximately,
30 % area in the continuous crystal is not an eﬀective area keeping a good and uniform spatial
resolution [125, 130]. On the other hand, an array camera has a uniform spatial resolution over
the whole area [128].
In our Compton camera, we adopted the array type scintillation camera because of theuniform spatial resolution over the entire area. The array camera is able to avoid the necessityof the position calibration in space. The array also has the ability to discriminate multi-hit
events when multiple scattering occurs between diﬀerent pixels by Compton scattering in the
pixels or the detector is under the high count rate condition especially in medical imaging. Alsofrom the limitation of the weight in space, an array camera is superior to an Anger camera dueto the use of the whole area of the scintillator.
87

Chapter 6. Improvement of Scintillator

6.2

Optimizing the Scintillator and the Pixel Size

6.2.1

Use of Hygroscopic Scintillator

Owing to the feasibility of the array type camera for the ETCC, we ﬁrst developed an arraycamera consisting of an 8 × 8 array of GSO:Ce scintillator pixels using a 3M Vikuiti ESRT M
reﬂector. The pixel size was 5.9 mm ×5.9 mm ×13.0 mm, and a pixel pitch was 6.1 mm.
GSO:Ce has some merits for the ETCC as followings. GSO:Ce has a shorter radiation lengthof 13.8 mm and a shorter decay time of 30 - 60 ns (depending on the amount of doped Ce) than
NaI:Tl and CsI:Tl [132, 133, 134, 135]. In addition, GSO:Ce is nonhygroscopic, and no hermetic
package is needed. Therefore, an assembly of an array is simpler. On the other hand, NaI:Tl and
CsI:Tl are hygroscopic and needed a hermetic package. Moreover, GSO:Ce is less susceptible
to radiation damage than either NaI:Tl or CsI:Tl [136]. These are inevitable requirements for
detectors used in space.
Here, we have to optimize the angular resolution measure (ARM) of a Compton camera,
which depends on the the energy resolution of the scintillator. In principle, a better energy
resolution scintillator certainly improves the angular resolution of the ETCC.
Owing to pixellation structures, the array camera tends to have a lower energy resolution
than a continuous crystal. Especially, the energy resolutions of hygroscopic scintillators are more
likely to be worse. However, some array cameras already provided a better energy resolution.
For example, Lee et al. have reported that a NaI:Tl array with a pixel size of 2 mm ×2 mm
×10 mm achieved a FWHM energy resolution of 7.37% at 662 keV using a crossed-wire anode
PMT (Hamamatsu R3292)[137]. In contrast, our GSO:Ce array coupled to a multi-anode PMT(H8500) had an FWHM energy resolution of ∼ 10.5% at 662 keV [131]. Although the NaI:Tl
has hygroscopic nature, and the pixel size is smaller than ours, the NaI:Tl array keep energy
resolution that monolithic one has. In addition, it has a better energy resolution than GSO:Cearray.
Since NaI:Tl scintillator has a week radiation hardness, the use of the NaI:Tl array is diﬃcult
in the space. In addition, because a decay constant time of NaI:Tl is long (approximately 250
nsec) than other scintillators, the signal from the NaI:Tl is likely to pile-up under the high-countrate condition, especially for medical imaging. Therefore we tried to develop a new scintillator
array with a better energy resolution except NaI:Tl, even if the crystal has hygroscopic nature.

6.3

ARM Calculation for the Selection of a Scintillator and the
Pixel Size

For the designing the new array, I studied the detail of the ARM of an ETCC by simulation.
The ARM depends on not only the energy resolution, but also the pixel size of the scintillator.
Therefore, I estimated which factor mentioned below is more eﬃcient for the improvement of
the angular resolution.
88

ARM Calculation for the Selection of a Scintillator and the Pixel Size

Figure 6.2: Dependence of the angular resolution measure (ARM) on the incident gamma-ray
energy (solid line), using a scintillator with an FWHM energy resolution of 10.5 % at 662 keV.
Filled circles denote the lower limit by Doppler broadening for argon gas [95]

The ﬁrst factor is the pixel pitch. Figure 6.2 shows the simulated ARM as a function of theincident gamma-ray energy for 6 mm ×1 mm ×13 mm pixels. Here, this calumniation has been
done based on the following assumptions:
1) Compton scatterings occur in the TPC with a volume of 10 cm ×10 cm ×10 cm, ﬁlled withargon (90%) and ethane (10%) gas at 1 atm.
2) We arrange 3 × 3 arrays, which consist of 8 × 8 pixels, below 5 cm from the bottom of the
TPC.
3) The energy resolution of the scintillator is 10.5% at 662 keV (FWHM), which is typical for
GSO:Ce.
4) The energy resolution of the recoil electron in the TPC is 30% at 22 keV (FWHM).
5) There is no error in the measurement of the Compton scattering position with the TPC.
6) Doppler broadening is taken into account.
Table 6.1 summarizes the dependence on the spatial resolution of the GSO:Ce scintillator
array. The dependence of the ARM angular resolution on the pixel pitch shows that even if the
pixel size is 1 mm ×1 mm ×13 mm, the ARM is improved only from 5.2◦ to 4.9◦ at 662 keV(FWHM).
Second, the ARM dependence on the thickness of (height) the scintillator is listed in Table
6.2, for 6-mm pixel pitch. It is found that thinner pixel has a better angular resolution of89

Chapter 6. Improvement of Scintillator

Table 6.1: ARM dependence on the spatial resolution of the GSO:Ce arraypixel size
FWHM ARM
length × width × thickness
at 662 keV[ mm × mm × mm ][degrees]
6 × 6 × 13
5.2
3 × 3 × 13
5.0
2 × 2 × 13
4.9
1 × 1 × 13
4.9

Table 6.2: ARM dependence on the pixel height with an FWHM energy resolution of 10.5 % at
662 keV.
pixel height [mm] FWHM ARM at 662 keV [degrees]
15.0
5.4
13.0
5.2
6.5
4.9
3.0
4.7

the Compton camera, though a thinner crystal have a lower detection eﬃciency. Here, theattenuation length of the GSO:Ce is 14.2 mm at 511 keV [151], and a height shorter than ∼10
mm is not enough to detect gamma rays with an energy of over 500 keV.
The dependence of the ARM angular resolutions on the energy resolution of the scintillator
is listed in Table 6.3, assuming a 6 mm ×6 mm ×13 mm scintillator pixel and the above
assumptions 1), 2) and 4) - 6). For the scintillator with an energy resolution of 3.0 %, the ARM
is improved to 3.8◦ at 662 keV (FWHM).
In the above calculations, we ignored the contribution of the spatial resolution of the TPC
on the ARM. The contribution of the TPC spatial resolution of ∼ 0.4 mm (1 σ) [38] to the
ARM was estimated to be ∼ 0.3◦ at 662 keV, which is small compared to the contribution to
scintillator here.
Tables 6.1, 6.2 and 6.3 show that the improvement of the energy resolution of the scintillator
is more eﬀective in obtaining a better angular resolution than improving the position resolutionof the scintillator. Table 6.2 suggests that the improvement due to the determination of thedepth of interaction (DOI) [138, 139, 140, 141], looks small for GSO:Ce. Table 6.3 also shows

Table 6.3: ARM dependence on the energy resolution of the scintillator
FWHM energy resolution of FWHM ARM at 662 keVscintillator at 662 keV [% ][degrees]
10.5
5.2
7.0
4.5
5.0
4.0
3.0
3.8

90

LaBr3 :Ce Camerathat the eﬀect of the improvement is small when a scintillator has an energy resolution of better
than ∼ 5 % (FWHM) at 662 keV.
Therefore, those studies indicate the necessity of a new scintillator with a better energy
resolution and the goal of the energy resolution is ∼ 5 % (FWHM) at 662 keV in order toachieve an ARM of approximately 4◦ (FWHM) at 662 keV. Additionally the pixel pitch isenough to be 6 mm which is the same size as a GSO:Ce array.

6.4

LaBr3 :Ce Camera

In the chapter 4, Fig. 4.1 shows the features of various scintillators. A LaBr3 :Ce looks a better
candidate as a replacement for a GSO:Ce, because of its excellent energy resolution. The typicalenergy resolution is 3% (FWHM) at 662 keV [142, 143, 144, 145, 146], which is independent of
the crystal size from 4 mm ×4 mm ×5 mm up to ϕ 51 mm ×76 mm [144, 147].
LaBr3 crystals doped with 0.5 mol% and 5 mol% Ce have primary (∼ 90 %) decay times of
26 and 15 nsec, respectively. In addition, they have high light yields of ∼ 6 × 104 photons/MeV(∼ 160% of NaI:Tl; ∼ 820 % of GSO:Ce) [142, 143, 144, 145, 146, 148]. The radiation hardnessof LaBr3 :Ce is also very good [149, 150].
On the other hand, LaBr3 :Ce has several disadvantages. One is hygroscopic, and another isa long attenuation length of LaBr3 :Ce of 2.1 cm at 511 keV, which is 1.4 times longer than thatof GSO:Ce [151, 152]. Therefore the detection eﬃciencies for gamma rays in LaBr3 :Ce throughphoto-absorption and Compton scattering are 0.21 and 0.51 at 500 keV for a thickness of 5 and15 mm, respectively. However, the main aim is to improve the angular resolution of ETCC.
Although the detection eﬃciency is worse, we have developed a LaBr3 :Ce array camera without
DOI as a ﬁrst step.
Position-sensitive detectors made of LaBr3 :Ce have been reported by several groups [153,129, 150, 155, 130, 156, 157, 158]. Kuhn et al. measured a LaBr3 :Ce 10 × 10 array with a pixel
size of 4 mm ×4 mm ×30 mm in an Anger-logic detector [154]. Pani et al. [129] and Bloser et al.
[153] measured a 10-mm thick continuous LaBr3 :Ce crystal with 3-mm and 1-mm thick optical
windows, respectively, using a multi-anode PMT. Case et al. [157] and Budden et al. [158] usedlayers of wavelength-shifting ﬁbers arranged parallel to the surface of the LaBr3 :Ce crystal tomeasure a two-dimensional position. Two orthogonal layers were used and read out at the endof the ﬁbers by a multi-anode PMT. On the other hand, to measure the energy, the LaBr3 :Cecrystal was read out using a single-anode PMT viewing through the ﬁber layers. However, a
LaBr3 :Ce array has not yet been measured with a multi-anode PMT which has the same areaof the anode as the pixel.
To improve both the energy and angular resolutions of the ETCC, we have assembled a
LaBr3 :Ce array. The array has the same pitch of crystal pixel as that of the anode array ina multi-anode PMT, and a thickness of pixel was chosen to be 15 mm for detecting sub-MeV
gamma rays. Here, I describe the evaluation of the energy resolutions of the LaBr3 :Ce array
91

Chapter 6. Improvement of Scintillator
coupled to a multi-anode PMT (H8500) for gamma-ray energies between 122 and 835 keV. Also
obtained results are compared to those obtained with the GSO:Ce array described above using
the same PMT and readout system.

92

Materials for LaBr3 :Ce Array Camera

38 mm

Figure 6.3: Photograph of LaBr3 :Ce crystals #1 (left) and #2 (right) with sizes of ϕ38 mm ×38
mm.

6.5

Materials for LaBr3 :Ce Array Camera

6.5.1

Energy resolutions of the monolithic LaBr3 :Ce crystals by whole irradiation

Our array was constructed from two blocks of LaBr3 :Ce crystal, hereafter #1 and #2, of a 5mol% Ce-doped monolithic LaBr3 BrilLanCe 380T M [159] crystal with a size of ϕ38 mm ×38
mm (Fig. 6.3).
First we checked the energy resolutions of those monolithic crystals. Two monolithic crystals
were measured using a 2-inch single-anode PMT (Hamamatsu R6231 [160]) with an optical grease(OKEN 6262A [69]). for irradiation of the entire crystal by gamma rays from radioisotopes.
Table 6.4 lists the FWHM energy resolutions of #1 and #2. Figure 6.4 shows the FWHM
energy resolutions of the crystals from 81 to 1275 keV. Best-ﬁt lines for the data points of #1
and #2 are represented as the following equation 6.1 with α = 3.10 ± 0.03, β = 0.545 ± 0.002,
and α = 3.06 ± 0.03, β = 0.543 ± 0.002, respectively,(
resolution(%) = α ×E662keV)−β(6.1)

There was no signiﬁcant diﬀerence in the energy resolutions between #1 and #2. In addition,
this results is the same as the typical FWHM energy resolution of the LaBr3 , which shows the
goodness of those two monolithic crystals.
93

Chapter 6. Improvement of Scintillator

(a)

(b)

) 10
%
(
)
M7
H
W
F(
n5
o
it
u
l 4
o
s
e
R
yg 3
r
e
n
E

) 10
(%
)
M7
H
W
F(
n5
o
it
lu 4
o
s
e
R
yg 3
r
e
n
E
2
70 1002
70 100200
Energy (keV)

1000

200
Energy (keV)

1000

Figure 6.4: Energy resolutions of LaBr3 :Ce crystals #1 (a) and #2 (b), which were cut into
pixels in an array.

Table 6.4: FWHM energy resolutions of monolithic LaBr3 :Cecrystal number Energy (keV) Energy resolution (% )#1356
4.19 ± 0.02
662
3.06 ± 0.01
#2356
4.12 ± 0.02
662
3.01 ± 0.01

6.5.2

Uniformity of Light Output and Energy Resolution

Since those crystals have a large size of ϕ38 mm ×38 mm, there is a possibility that they
have a position dependences of the light output and energy resolution. Thus we measured theuniformity of the light output and energy resolution by irradiating the crystals with gamma raysfrom a collimated 133 Ba source. The beam was formed using a 70-mm long Pb collimator witha 2-mm diameter opening. The points from d1 to d6 in Fig. 6.5 were irradiated with gamma
rays from the vertical direction to the side of the crystals.
The ratios of the light output at each point to the value at point C of Fig. 6.5 are shownin Fig. 6.6. The diﬀerence was plotted within 2%. The average and standard deviation (SD) of#1 and #2 are listed in Table 6.5. The FWHM energy resolutions at 356 keV are shown in Fig.
6.7. The average and SD values of #1 and #2 are also listed in Table 6.5. The SD of the light
output are approximately 0.5 % of the mean at 356 keV, while Iltis et al. reported that the SD
94

Materials for LaBr3 :Ce Array Camera

(a)

O1

O2
O3

(b)

内

I33
18mm

O8 8

I1
9mm

I7

O7

O4

O6

O7

O3
O5

0mm

d1

10mm

d5

d2

25mm

d6

d3

35mm

d4

C
I5

O1

O5

PMT side

Top view

side view

Figure 6.5: Filled circles denote points where collimated gamma rays from a 133 Ba source were
irradiated. (a) top view, (b) side view.

Table 6.5: Uniformity of Monolithic LaBr3 :Cemean ± SD
Light output at 356 keV #1 1.003 ± 0.005#2 1.006 ± 0.005
FWHM Energy resolution at 356 keV #1
4.10 ± 0.07
#2
4.08 ± 0.07

is 0.4 % at 662 keV.
Although we and Iltis et al. used diﬀerent gamma-ray sources and the materials including
PMT, the both variations for the light output are similarly small [147]. Also the energy resolutions of #1 and #2 have small variations. Thus our crystals are considered good to make apixel material for an array. In addition, the array is expected to have a uniformity of energy
resolution within a few percent, if we make a pixel ideally.

95

Chapter 6. Improvement of Scintillator)it 1.02nuyra
rti 1.01
b
ra
(t
u
p
tu1.00othgiL0.99C

I1 3 5 7

O1 2 3 4 5 6 7 8
Positiond1 2 34 5 6

Figure 6.6: Uniformities of light outputs of the LaBr3 :Ce crystals #1 (squares) and #2 (ﬁlled
circles) at positions shown in Fig. 6.5.

6.6

Assembly of LaBr3 :Ce and GSO:Ce arrays

These ϕ 38-mm LaBr3 :Ce crystals were cut into 64 pixels with a size of 5.8 mm ×5.8 mm ×15mm. At the same time, we studied what kind of crystal surface is able to keep the energy
resolution. In this process, we had to develop the cutting and polishing method by ourselves,
though we cannot describe the details for claiming the patents. In this study, we adopted arough surface of the pixel. Figure 6.8 shows the photograph of LaBr3 :Ce pixels cutted by ourown technique. Each pixel is wrapped with TETRATEXT M as a reﬂector. Figure 6.9 shows
an FWHM energy resolution of LaBr3 :Ce pixel without hermetic package as a function of theenergy measured by the 2-inch single-anode PMT (Hamamatsu R6231). Owing to hygroscopic
nature, the pixel was measured under dry condition. Best-ﬁt line for the data points in Fig. 6.9
is represented as α = 3.5 ± 0.5,and β = 0.47 ± 0.02, where α and β are described in equation 6.1.
We succeeded in keeping the energy resolution of LaBr3 :Ce after the cutting. For comparison
with a LaBr3 :Ce, we also measured a GSO:Ce pixel with a size of 5.9 mm ×5.9 mm ×13 using the2-inch single-anode PMT (Hamamatsu R6231). Best-ﬁt line for the FWHM energy resolution
as a function of the energy was represented as α = 8.2 ± 0.1, β = 0.497 ± 0.003.
Next the pixels were assembled into an 8 × 8 array with a pixel pitch of 6.1 mm using a
TETRATEXT M reﬂector as shown in Figs. 6.10, and 6.11. The total volume of the crystal was
49 mm ×49 mm ×15 mm. Due to hygroscopicity, the array was sealed in a hermetic package
96Assembly of LaBr3 :Ce and GSO:Ce arrays

(a))

4.5
(%
) 4.4
M
H
W
F( 4.3
no4.2
tiu
lo4.1
se
R4.0
yg
re4.1
nE
4.2

(b)

) 4.5
%
()
M4.4
H
W
4.3
(F
no4.2
it
luo4.1
se
R4.0
yg
re4.1
nE
4.2

C

I1 3 5 7

O1 2 3 4 5 6 7 8
Positiond1 2 34 5 6

Figure 6.7: Uniformities of energy resolutions of the LaBr3 :Ce crystals #1 (a) and #2 (b) at
positions shown in Fig. 6.5.

with a volume of 62 mm ×62 mm ×20 mm, which had a quartz window with a volume of 58
mm ×58 mm ×3 mm on the PMT side, and aluminum plates with a thickness of 0.5 mm on
the remaining sides. Because sealing the LaBr3 :Ce crystal requires an adhesive-bonding surface
between the window and the aluminum package, the area of the window was larger than that of
the crystal array.
For comparison with the LaBr3 :Ce array, we also assembled an 8 × 8 array of 0.5 mol%Ce-doped GSO pixels with a size of 5.9 mm ×5.9 mm ×13 mm with a pitch of 6.1 mm using
a 3M Vikuiti ESRT M reﬂector (Fig. 6.12) [131]. The GSO:Ce array had neither a hermetic
package nor a quartz window without hygroscopicity.
Since the anode gains of a multi-anode PMT H8500 was varied by a factor of approximately
3 among 64 anodes, we measured the light output and energy resolution of each pixel in the
LaBr3 :Ce and GSO:Ce arrays using a 2-inch square single-anode PMT (Hamamatsu R6236)
with optical grease (OKEN 6262A). Here each pixel was irradiated with gamma rays from thecollimated 133 Ba source as described before. The R6236 was operated with high voltages of 800
and 1000 V for the LaBr3 :Ce and GSO:Ce arrays, respectively. Figure 6.13 shows the energy
spectra of the center pixel in the LaBr3 :Ce and GSO:Ce arrays. Moreover, Figs. 6.14 (a) and97

Energy Resolution (FWHM) [%]Chapter 6. Improvement of Scintillator

15mm

10
9
8
7
6
5
4
3

80 100Figure 6.8: Photograph of LaBr3 :Ce pixel

200
300 400
Energy [keV]

1000

Figure 6.9: FWHM Energy resolution of
LaBr3 :Ce pixel as a function of the energy.

(b) show light outputs normalized to each maximum pixel in the LaBr3 :Ce and GSO:Ce arrays,
respectively. The average and SD values of the FWHM energy resolutions of all 8 × 8 LaBr3 :Ceand GSO:Ce pixels were 5.8 ± 0.9% and 11.3 ± 0.9% at 356 keV, respectively (Fig. 6.14 c, d).
This results showed that the FWHM energy resolution of LaBr3 :Ce was worse by approximately 1 % point, after the assembly. On the other hand, that of GSO:Ce is not changed. It
suggests that the LaBr3 :Ce crystals might be damaged due to moisture during the assembly or
glass window worsens the resolution.

98

Assembly of LaBr3 :Ce and GSO:Ce arrays

Glasswindow

5.8mm (pixel size)
62mm

Hermetic
packageof aluminum

20mm
Figure 6.10: Photograph of an 8 × 8 array of LaBr3 :Ce pixels in a hermetic package.

bonding
agent

62 mm
58 mm
49 mm
PMT side

Glass window(3-mm thickness)
Aluminumpackage

20 mm

LaBr3 array(13-mm thickness)

Gamma rays

Figure 6.11: Schematic view of an 8 × 8 array of LaBr3 :Ce pixels.

99

Chapter 6. Improvement of Scintillator

5.9 mm (pixel size)

49
mm

13 mm
Figure 6.12: Photograph of an 8 × 8 array of GSO:Ce pixels in a hermetic package.

6.7

Performance of Array Camera

To estimate the energy resolution of LaBr3 :Ce array, the LaBr3 :Ce array was then mounted
to a 64-channel multi-anode PMT H8500 with optical grease (OKEN6262A) as shown in Fig.
6.15. The H8500 had an eﬀective area of 49 mm ×49 mm, an external size of 52 mm × 52
mm, and a 1.5-mm thick borosilicate glass window. The outer part of the quartz window of
LaBr3 :Ce array than the PMT surface was covered with a TeﬂonT M reﬂector. The PMT had
a bialkali photocathode and 8 stages of metal channel dynodes for electron multiplication. The
relative uniformity of the PMT anode-gain were 2.4 among 64 anodes. To reduce the numberof signals, the 64-anode signals of the PMT were reduced to four channels using a resistor-chainboard connecting the anodes with 100Ω chips [101]. Each output signal of the four corners
was preampliﬁed with an integrating time of 5.4 µsec, shaped in a shaping time of 0.5 µsec
with a module (Hoshin Electronics Co., LTD. N016-V). This shaped signal was digitized with apeak-hold analog-to-digital converter (ADC) (Hoshin V014). The values of the four signals (a,b, c, and d) were used to calculate the both two-dimensional positions (x, y) and the light yieldcorresponding to the energy of an incident gamma ray (E). The detail of the calculation, called
the charge-division method, is as shown in Chapter 4.
Although we have calculated the position of the hit anode without any correction of theanode gain uniformity, it was shown that all pixels in 9 GSO:Ce arrays were found to be clearly
resolved in the reconstructed images [131]. We used the same PMT H8500 in the measurements
100Performance of Array Camera

Figure 6.13: Energy spectra of the center pixel (x = 4, y = 4 in Fig. 10) in the LaBr3 :Ce (solidline) and GSO:Ce (dotted line) arrays coupled to a single-anode PMT (Hamamatsu R6236) andirradiated with collimated gamma rays from a 133 Ba source.

of the LaBr3 :Ce and GSO:Ce arrays. The H8500 was operated with a gain of ∼ 5 × 105 at ahigh voltage of ∼ −800 V for the LaBr3 :Ce array, and with a gain of ∼ 1 × 106 at a high voltageof ∼ −900 V for the GSO:Ce array. These gains were adjusted within the dynamic range of thepreampliﬁer considering a larger light output of LaBr3 :Ce than that of GSO:Ce.
The arrays were irradiated with gamma rays from 57 Co (122 keV), 356 Ba (81, 356 keV), 22 Na(511 keV), 137 Cs (662 keV), and 54 Mn (835 keV) sources. Figs. 6.16 and 6.17 show ﬂood-ﬁeld
irradiation images taken with the LaBr3 :Ce and GSO:Ce arrays, histograms of events with pixels
in the central row of the array (image proﬁle), and spectra from the center pixel in the arrayfor a 137 Cs source. Here, we used the imaging method as mentioned in Chapter 4. The image
proﬁle was obtained using approximately 50-150 keV energy deposit events depending on theanode gain. In addition, an image proﬁle in ﬂood-ﬁeld irradiation from a 57 Co source is shownfor the events above approximately 25-75 keV. The peak-to-valley ratios of the image proﬁlesof the LaBr3 :Ce array were 21 and 36 on the average for 137 Cs and 57 Co sources, respectively,
whereas those of the GSO:Ce array were 20 and 17 using the same sources.

6.7.1

Energy Resolution

The FWHM energy resolutions at 662 keV of the LaBr3 :Ce and GSO:Ce arrays are also shownin Figs. 6.18 (a) and (b). There found a signiﬁcant diﬀerence in the energy resolutions between
101Chapter 6. Improvement of Scintillator

(a)

(b)

8
) 7
w
o
r( 6
n5
o
tii 4
s
o3
p
-Y
2
1

Li
gh
t
O
u
tp
u
t
(a
.u
.)

8
) 7
w
o
r( 6
n 5
o
it 4
is
o 3
p
-Y
2
1

1 2 3 4 5 6 7 8
X-position (column)

Li
gh
t
O
u
tp
u
t
(a
.u
.)

1 2 3 4 5 6 7 8
X-position (column)

1 2 3 4 5 6 7 8
X-position (column)
(c)

8
)
w7
o
r( 6
n5
o
tii 4
s
o3
p
-Y
2
1

En
e
rg
y
re
so
lu
ti
o
n
(F
W
H
M
)
[%
]

(d)

8
) 7
w
ro
( 6
n5
o
tii 4
s
o3
p
-Y
2
1

1 2 3 4 5 6 7 8
X-position (column)

En
e
rg
y
re
so
lu
ti
o
n
(F
W
H
M
)
[%
]

Figure 6.14: Light output of each pixel of the LaBr3 :Ce (a) and GSO:Ce (b) arrays (each maximum value is normalized to 100), and FWHM energy resolution at 356 keV of each pixel of the
LaBr3 :Ce (c) and GSO:Ce (d) arrays, which were coupled to a single-anode PMT (Hamamatsu
R6236).

the inner and outer pixels for the LaBr3 :Ce array. Fig. 6.18 (c) shows the average and SD
values of the FWHM energy resolutions at several energies. The results of a ﬁt by equation 6.1
are represented by the lines in Fig. 6.18 (c) and the resulting parameters are listed in Table
6.6. The LaBr3 :Ce array had a higher energy resolution than the GSO:Ce array by a factor of
approximately 1.5.
There are two possible reasons why the outer 28 LaBr3 :Ce pixels had poorer energy resolutions than the inner ones. First, the intensity of light emitted in the outer pixels is lower thanthat of the inner pixels. Or second, although the outer light intensity is the same as the innerintensity, a smaller fraction of light from the outer pixels enters the PMT window.
The second reason is supported by the measurement of the array using a single-anode PMT
R6236 with a larger window than the quartz window of the scintillator array, as described in
section 6.5.2. This measurement shows that the energy resolutions were uniform among all 8 × 8
pixels. To improve the energy resolution of the outer pixels of the LaBr3 :Ce array, we will needto develop a new package for the LaBr3 :Ce array with a quartz window the same size as the
PMT window (see section 6.8).
We also measured the energy resolution as an array, where sum of all 8 × 8 pixels were used.
In addition, the sum of the inner 6 × 6 LaBr3 :Ce pixels is also measured, because the energy
102Performance of Array Camera

Resistorchain
PMT
anode
PMT
Glass
LaBr3 pixels

49mm
52mm
62mm

Hermetic
package

Figure 6.15: Schematic view of the LaBr3 :Ce scintillation camera.

resolution of the inner ones is better clearly. Fig. 6.19 (a) shows the sum of energy spectra
taken with all LaBr3 :Ce pixels, inner 6 × 6 LaBr3 :Ce pixels, and all GSO:Ce pixels in ﬂood-ﬁeld
irradiation from a 137 Cs source. The FWHM energy resolutions of these spectra at 662 keV are
6.4 ± 0.2% with all LaBr3 :Ce pixels, 5.4 ± 0.3% with inner 6 × 6 LaBr3 :Ce pixels, and 10.4 ± 0.2%
with all GSO:Ce pixels. Figure 6.19 (b) shows the dependence of the resolutions on the energy.
The parameters for the best-ﬁt lines in Fig. 6.19 (b) are listed in Table 6.6.
Although the outer pixels have poorer energy resoluteness than the inner ones, the energy
results of LaBr3 :Ce array is still better than GSO:Ce array. For the use of LaBr3 :Ce array in
the ETCC, the FWHM ARM at 662 keV is expected to be 4.6◦ and 4.3◦ for the LaBr3 :Ce array

Table 6.6: Parameters of best-ﬁt lines
crystal
Average of 64 pixels
LaBr3 :Ce All pixels
inner 6 × 6 5.8 ± 0.4
GSO:CeAll pixels
Sum of 64 pixels
LaBr3 :Ce All pixels
inner 6 × 6 5.4 ± 0.3
GSO:CeAll pixels⋆ :α and β are deﬁned by equation 6.1103α⋆
7.0 ± 0.5
0.47 ± 0.03
10.8 ± 0.3
6.5 ± 0.5
0.60 ± 0.01
10.4 ± 0.9

β⋆0.44 ± 0.01
0.48 ± 0.01
0.57 ± 0.01
0.51 ± 0.01

Chapter 6. Improvement of Scintillator

Figure 6.16: Reconstructed images and energy spectra with the LaBr3 :Ce array coupled to a
PMT H8500. (a) 2-D image in ﬂood-ﬁeld irradiation of gamma rays from a 137 Cs source. Image
proﬁle with pixels in the center row of the array, indicated by the dashed square, when the array
was irradiated with gamma rays from 137 Cs (b) and 57 Co (c) sources. (d) Energy spectrum of
137 Cs from the central pixel in the array indicated by the circle in (a).

with all pixels and the inner 6 × 6 pixels, respectively. On the other hand, ETCC using the
GSO:Ce array produce the resolution of 5.3◦ (FWHM) at 662 keV.

6.8

Improvement of a LaBr3 :Ce ArrayAs described in the previous section, we have developed the LaBr3 :Ce array (hereafter ﬁrst
generation). However, the outer pixels in the ﬁrst generation has lower energy resolution thaninner. In order to improve energy resolution of the outer pixels, we have developed a secondgeneration LaBr3 :Ce array with a compact hermetic package. Here, we changed the design ofthe hermetic package as shown in table 6.7. We developed two second generation arrays.
As materials of the two arrays, we also used four monolithic LaBr3 BrilLanCe 380T M [159]
crystals (hereafter I describe #3 - #6 ) with a size of ϕ38 mm ×38 mm as the same as the ﬁrst
generation. We measured the energy resolutions and the light-output uniformities of #3 to #6
104

Improvement of a LaBr3 :Ce Array

Figure 6.17: Images and energy spectra with the GSO:Ce array coupled to the PMT H8500.
Captions are the same as in Fig. 6.16.

using the same method when we measured the #1 and #2. The parameters of the best-ﬁt lines
for the FWHM energy resolution of #1 to #6 as a function of the energy are listed in Table 6.8.
Additionally, the uniformities of the light output and the energy resolution are listed in table
6.9. The light-output uniformity of #3 was worse than the others. However, this uniformity isenough to assemble the array, because this uniformity is better than the light-output uniformityof the LaBr3 :Ce array measured by single anode PMT as described in section 6.6. Therefore,both #3 and #6 were used as an appropriate material for the new array.
Thus we have developed two arrays (hereafter array A and B). The array A and B were madefrom #3, #4 and #5, #6, respectively. Figure 6.20 shows the second generation LaBr3 :Ce array

Table 6.7: Parameters of the structures in the LaBr3 :Ce arraygeneration
whole volume
glass volume
1st.
62 mm ×62 mm ×20 mm 58 mm ×58 mm ×3.0mm
2nd.
54 mm ×54 mm ×19 mm 52 mm ×52 mm ×2.3mm

105Chapter 6. Improvement of Scintillator

Table 6.8: Parameters of best ﬁt for the monolithic LaBr3 :Ce crystals
crystalα⋆β⋆#13.10 ± 0.03 0.545 ± 0.002#2
3.06 ± 0.03 0.543 ± 0.002#3
3.25 ± 0.05 0.501 ± 0.002#4
3.08 ± 0.05 0.504 ± 0.002#5
3.11 ± 0.06 0.511 ± 0.003#6
3.17 ± 0.05 0.499 ± 0.002⋆:α and β are deﬁned by equation 6.1

Table 6.9: Uniformity of the monolithic LaBr3 :Cemean ± SD
Light output at 356 keV#1 1.003 ± 0.005#2 1.006 ± 0.005#3
0.99 ± 0.02
#4 0.999 ± 0.002#5 1.001 ± 0.002#6 1.000 ± 0.003
FWHM Energy resolution at 356 keV #1
4.10 ± 0.07
#2
4.08 ± 0.07
#3
4.18 ± 0.04
#44.06 ± 0.06
#5
4.10 ± 0.05
#6
4.14 ± 0.06

106

Improvement of a LaBr3 :Ce Array

Figure 6.18: Energy resolution of each pixel of the LaBr3 :Ce (a) and GSO:Ce (b) arrays at 662
keV using the H8500. (c) Averages and SDs of energy resolutions of the pixels as a function ofenergy. Solid line denotes the best ﬁt to the LaBr3 :Ce array data with all 64 pixels (triangles).
Broken line denotes that of the LaBr3 :Ce array with inner 6 × 6 pixels (circles) represented by
the non-shaded portion in (a). Dotted line denotes that of the GSO:Ce array (open squares).
Error bars represent SDs.

compared with the ﬁrst generation.
We measured the second generation arrays using as doing the ﬁrst generation, and we obtained the images and energy spectra (Fig. 6.21). Figure 6.22 shows FWHM energy resolutionsof each pixels for array A and B. Fig. 6.23 shows the FWHM energy resolutions of the average
and SD of all 64 LaBr3 :Ce pixels (array A) as a function of the energy. For comparison, thoseof a GSO:Ce array and ﬁst generation are also shown in Fig. 6.23. The best-ﬁt lines for thedata points listed in table 6.10 including array B, and the point of inner 6 × 6 pixles for only
LaBr3 :Ce arrays of the both ﬁrst and second generations.
The results show that the diﬀerences of the energy resolution between the outer and inner
pixels was considerably improved in the second-generation array. This diﬀerence is similar levelto that of GSO:Ce array. The diﬀerence in GSO:Ce array and second generation is thought thatlight emitted from the pixels on the corner and edge of the array spreads over a few anodes of
the PMT.
107

Chapter 6. Improvement of Scintillator

Figure 6.19: (a) Energy spectra taken with all GSO:Ce pixels (dotted line), all LaBr3 :Ce pixels(solid line), and inner 6 × 6 LaBr3 :Ce pixels (broken line), which were coupled to the PMT
H8500 and irradiated with gamma rays from a 137 Cs source. (b) Energy resolutions and best-ﬁtlines with all GSO:Ce pixels (open squares and dotted line), all LaBr3 :Ce pixels (triangles andsolid line), and inner6 × 6LaBr3 :Ce pixels (circles and broken line) as a function of energy.

Although the array B of the second generation did not have a good energy resolution than
the array A, it is still better than GSO:Ce array. Thus, we succeeded in the development of thearray to use for the absorber in the ETCC.

108

Improvement of a LaBr3 :Ce Array

62mm

54mm

Figure 6.20: Comparison of LaBr3 :Ce array between (a) ﬁrst and (b) second generation (array
A).

800

0.6

700

0.4

600
0.2

500

0

400

-0.2

300

-0.4

200

-0.6

100

-0.8

-1

-0.5

0

0.5

1

Counts (arb. unit)

(b)

0.8

Counts (arb. unit)

y position (arb. unit)

(a)

250
200
150
100
50
0 200 300 400 500 600 700 800
Energy [keV]0x position (arb. unit)

Figure 6.21: (a) Two-dimensional image of the LaBr3 :Ce array A in ﬂood-ﬁeld irradiation of
gamma rays from a 137 Cs source. (b) Energy spectrum of 137 Cs with the central pixel indicatedby the circle in the left image.

(a)
8
7
6

no
iit 5
so
p 4
Y
3
2
1

1 2

3 4

5

X position

6 7

8

E
ne
12 rgyR
esol
10 utio
n
(F
W
8 H
M
)a
t6
6
6 2k
eV
[%
4 ]

(b)

E
ne
r
R
es
ol
u
10 tio
n
(F
W
H
8 M
)a
t6
6
6 2k
eV
[%
4 ]

12 gy

8
7
6

no
iit 5
so
p 4
Y
3
2
1

1 2

3

4

5

X position

6

7

8

Figure 6.22: Map of the FWHM energy resolution of each LaBr3 :Ce pixel at 662 keV. (a) and(b) are ﬁrst and second generation, respectively.
109

Chapter 6. Improvement of Scintillator

Figure 6.23: FWHM energy resolution of the average and SD of all 64 LaBr3 :Ce (2nd generationarray A, circles) as a function of energy, and that of GSO:Ce (squares) and 1st generation(triangle).

Table 6.10: FWHM energy resolution of LaBr3 :Ce arrays
generation
α⋆
β⋆
1st
all pixels
7.0 ± 0.5 0.47 ± 0.03
inner 6 × 6 5.8 ± 0.4 0.44 ± 0.01
2nd arrayA
all pixels
5.7 ± 0.5 0.47 ± 0.01
inner 6 × 6 5.4 ± 0.3 0.44 ± 0.01
2nd arrayB
all pixels
6.4 ± 0.6 0.61 ± 0.02
inner 6 × 6 5.8 ± 0.7 0.58 ± 0.02
GSO:CeAll pixels 10.8 ± 0.3 0.48 ± 0.01
inner 6 × 6 10.0 ± 0.2 0.46 ± 0.03⋆:α and β are deﬁned by equation 6.1

110

Comparison with Other Position-Sensitive LaBr3 :Ce Camera

6.9

Comparison with Other Position-Sensitive LaBr3 :Ce Camera

We adopted the array type scintillation camera because the spatial resolution was dictated by
the pixel width and the uniformity over the entire area, which do not need position calibration in
space even after the deterioration of the scintillator by cosmic rays irradiation. In addition, somemultiple hit events occur in the scintillator, and, for the Compton camera, we need to measure
the ﬁrst hit point of a gamma ray in the scintillator. Although it is diﬃcult to discriminate
between multiple Compton hits and a single hit in the continuous crystal, it is feasible with anarray camera to judge ﬁrst hit point. Indeed, events with multiple hits are not negligible for
gamma rays above a few hundred keV up to MeV where the main interaction process is Compton
scattering. For example, for 662 keV gamma rays vertically incident on the scintillator, GEANT4
simulation show that probability of the multiple hits in our LaBr3 :Ce array is only 7.5% of all
events between the pixels.
For the continuous type, spatial resolution (FWHM) is deﬁned as a half width of image
proﬁle when the crystal is irradiated with collimated gamma rays. On the other hand, the
FWHM resolution of array camera △x is deﬁned as △x = 2.35 × √d12 , where d is pixel pitch.
Here our array camera has the resolution of 4.1 mm. The spatial resolution of continuous typeis several times better than that of the array, and it is eﬀective if it is important to obtain thespatial resolution. However, table 6.1 shows that it is enough with an FWHM spatial resolutionof array camera (4.1 mm) for Compton camera.
Next, I describe the detail of comparison between the array and continuous type. Kuhn et al.
have developed Anger-logic detectors with LaBr3 :Ce pixellated arrays that achieved an FWHM
energy resolution of ∼ 5.3% at 511 keV [154]. On the other hand, that of our detector was 6.3%
at 511 keV using all 8 × 8 LaBr3 :Ce pixels in array A (second generation). Kuhn et al. obtained
a better energy resolution because they used single-anode PMTs. However, the peak-to-valley
ratio of their detectors was approximately 3 above 450 keV, whereas that of our array was morethan 20 from approximately 100 to 900 keV.
Bloser et al. measured a continuous LaBr3 :Ce with a volume of 5 cm ×5 cm ×1 cm using a
3” square position-sensitive PMT (Hamamatsu R2487), irradiated with collimated gamma rays[153]. This crystal had an FWHM spatial resolution of ∼ 3 mm and ∼ 6 mm at 122 keV, andan FWHM energy resolution of 13.5% and ∼ 18% at 122 keV near the center and at a diagonal
distance of ∼ 28 mm from the center, respectively. On the other hand, Pani et al. also reported
that a LaBr3 :Ce detector, consisting of a continuous LaBr3 :Ce crystal with an area of 5 cm ×5
cm, a thickness of 1 cm, and a PMT (H8500), had a spatial resolution of 1.80 mm at 122 keV
in an eﬀective area of 36 mm ×36 mm [15]. Their spatial resolutions were better than Bloser et
al. Because our array had a spatial resolution of 4.1 mm (FWHM) over the entire area of thearray, the continuous LaBr3 :Ce detector had a better spatial resolution than our array in thecenter region. However, the continuous LaBr3 :Ce detector had a poorer spatial resolution of up
111

Chapter 6. Improvement of Scintillator
to 6 mm outside the eﬀective area, whereas ours remained 4.1 mm even near the edges. The
continuous LaBr3 :Ce detector of Pani et al. also had better energy resolutions than Bloser et al.,∼ 12 and ∼ 15% (FWHM) at 122 keV inside and outside the eﬀective area, respectively, whereas
our array had 12.6 ± 2.0% and 17.7 ± 4.3% (FWHM) at 122 keV with inner 6 × 6 and outer 28
LaBr3 :Ce pixels, respectively. Because the eﬀective area of the continuous LaBr3 :Ce detector
was the same as the area of the inner 6 × 6 pixels in our LaBr3 :Ce array, there is no diﬀerencein the energy resolutions over the inner region of approximately 3.6 cm ×3.6 cm between thecontinuous LaBr3 :Ce detector and our array, whereas the outer pixels in our array have poorer
energy resolutions near the edge than the continuous LaBr3 :Ce.
Case et al. [157] and Budden et al. [158] showed position-sensitive detectors using continuous LaBr3 :Ce crystals and wavelength-shifting ﬁbers. Case et al. reported that a continuous
LaBr3 :Ce crystal with a 2.5-cm ϕ and a thickness of 2.5 cm had an FWHM energy resolutionof 5.6% at 662 keV using a single-anode PMT viewing through a layer of 2-mm ϕ wavelengthshifting ﬁbers (Saint-Gobain BCF-99-33A). This energy resolution is similarly to the inner 6 × 6
pixels in our LaBr3 :Ce array. On the other hand, Budden et al. measured the position resolution
in an array of four continuous LaBr3 :Ce crystals each with a size of 3.8 cm ×3.8 cm ×1 cm glued
together into a 7.6 cm ×7.6 cm ×1 cm mosaic. The position measurement was achieved using alayer of wavelength-shifting ﬁbers arranged parallel to the surface of the crystal and coupled toa multi-anode PMT. Two orthogonal layers were used to measure a two-dimensional position.
The position resolution was 2.7 mm (1σ, FWHM: 6.3 mm) irradiated by gamma rays from a
137 Cs source, and it is the worse resolution than our array.

6.10

Performance of the Compton Camera with a LaBr3 Array

Using ﬁrst and second generation of LaBr3 :Ce arrays, we measured the angular resolution of the
ETCC. Figure 6.24 shows the set up for the performance test of the ETCC consisting of ﬁrst andsecond-generation LaBr3 :Ce arrays. Here we also used a 10 × 10 array with a pixel size of 4 mm
×4 mm ×30 mm made by Saint-Gobain [154]. The FWHM energy resolutions were 5.2 ± 0.7%
at 662 keV using the same multi-anode PMT as described in the previous chapter. However,this array had several bad pixels which is hard to separate between pixels in the ﬂood-ﬁeld
irradiation image (Fig. 6.25), and these pixels were not used for the analysis of the ETCC data.

Although we used a similar small ETCC of proton beam experiment,the drift height in the
TPC was diﬀerent from the ETCC of proton beam as followings. TPC had an detection areaof 10 cm×10 cm ×7.5 cm, and a drift voltage of −3300 V was supplied to the drift plane. Theµ-PIC was driven at 490 V with a gain of approximately 3500, and the total gas gain of the TPC
was ∼ 3.5 × 104 . The FWHM energy resolution of the TPC at 22.2 keV (109 Cd) was 27.7 ± 0.6%.
The ETCC was irradiated with gamma rays from 137 Cs, and 54 Mn point sources at 20 cm fromthe center of the µ-PIC (Fig. 6.24). In addition, an 131 I (364 keV) liquid source, which was a112

Performance of the Compton Camera with a LaBr3 Array

Figure 6.24: Schematic view of the setup of the ETCC and Labr3 :Ce arrays including secondgeneration arrays

(a)

1.0

500

).
u. 0.5
a(
no0.0
tii
so
p--0.5
Y
-1.0

C
400 o
un
t
300 s(
a.
200 u
.)
100
-1.0

-0.5

0.0
0.5
X-position (a. u.)

1.0

-1.0

-0.5

0.0
0.5
X-position (a. u.)

1.0

(b)

0

800

).
u. 600
a(
st
nu400
o
C200
0

Figure 6.25: Bad pixels in Saint-Gobain Labr3 :Ce arrays. (a)ﬂood-ﬁeld irradiation image and(b) histograms of events with pixels in the square as shown (a).

113

ARM (FWHM) [degree]

Chapter 6. Improvement of Scintillator

10
9
8
7
6
5
4
3

300

400 500

1000
Energy [keV]

2000

Figure 6.26: ARM of ETCC with a LaBr3 :Ce array (solid circles) and GSO:Ce arrays used for
RCNP (open circles) as a function of the energy.
drop with a size of less than ϕ3mm, was also used.
The FWHM ARM of the ETCC with LaBr3 :Ce arrays was 4.2 ± 0.3◦ at 662 keV as shownin Fig 6.26.
For the comparison, Fig. 6.26 also shows the FWHM ARM of the ETCC with 26-mm thick
GSO:Ce arrays which was used for the RCNP beam test. The FWHM energy resolution of
the GSO:Ce array was 12.7 ± 3.4% at 662 keV, and that of the TPC for RCNP beam test was
30.0 ± 0.9% at 22.2 keV (chapter 5). The FWHM ARM at 662 keV was 5.1 ± 0.2◦ . We succeeded
in improvement of angular resolution of the ETCC using the LaBr3 :Ce array.
Here, the contributions to the ARM were itemized as followings:• Energy resolution of scintillator (σs,e )• Energy resolution of TPC (σt,e )• Position resolution of scintillator (σs,p )• Position resolution of TPC (σt,p )• Doppler broadening (σd )• Other, unknown factor (σo )Table 6.11 summarized the above factors for this experiment using the calculation as described in section 6.3.
Here, σs,e and σs,e were assumed that the energy resolution of the TPC or scintillator had noerror, respectively, and σe was calculated with the both energy resolutions. σe was not described114

Possibility of Improvement in Angular Resolution

Table 6.11: Contribution to ARM (FWHM) at 662 keV [deg.]
σs,e σt,e σea σs,p σs,p σpb
w/ GSO:Ce array
4.0 1.1 4.5 2.8 0.3 2.8
w/ LaBr3 :Ce array
2.0 0.9 2.7 3.3 0.3 3.3
2 + σ2a: σe = f (σs,e , σt,e ), b: σp2 = σs,p
t,p

σd
0.9
0.9

Table 6.12: Contribution to ARM (FWHM) at 662 keV [deg.]
ARM
σtot σe σp σd(measurement)(calculation)
w/ GSO:Ce array
5.1 ± 0.2
5.3 4.5 2.8 0.9
w/ LaBr3 :Ce array
4.2 ± 0.3
4.3 2.7 3.2 0.9

2 + σ 2 , because the energy of scattered gamma ray and Compton recoil electron areas σe = σs,et,e

not independence and the sum of these energy is the incident gamma-ray energy. Since theCompton recoil-electron energy is commonly smaller than scattered gamma-ray energy by anorder of 1, the contribution of the recoil-electron energy resolution to the ARM is small thanthat of scattered gamma-ray energy resolution.
σp was calculated assuming the position resolutions of the TPC and scintillator. σp was
2 + σ 2 , because σdescribed as σp = σs,ps,p and σt,p are independent factor. Here, the positiont,presolution of the TPC was assumed to be 0.4 mm (1 σ) [161], while the volume of the pixel
scintillator was large than the TPC position resolution by an order of more than 1. Thus the
position resolution of scintillator dominated the contribution of position resolution to the ARM
σp .
The measured resolution σtot is described as2σtot= σe2 + σp2 + σd2 + σo2 .

(6.2)

where σe , σp , σd , and σo are independent factor each other. Table 6.12 lists the measurement
and calculation of ARMs, where σo = 0.
The measurement of ARM was consistent with σtot , and σo s were negligible for the ETCC
with GSO:Ce and LaBr3 :Ce arrays.
Using the ETCC with LaBr3 :Ce arrays, the spatial resolution of the ETCC is expected tobe 1.8 and 1.1 cm when the center of 15-cm-high TPC and the target have a distance of 25 and15 cm, respectively, at 662 keV. This performance satisﬁes the medical requirement.

6.11

Possibility of Improvement in Angular Resolution

In the previous section, Table 6.12 suggests that the item of the position resolution is larger thanthat of the energy resolution if LaBr3 :Ce arrays are used in the ETCC. Thus, improvement of
the position resolution is eﬀective in order to obtain a better ARM. Especially, the improvement
115Chapter 6. Improvement of Scintillator

ガンマ線

Incident

Gamma rayComptonscattering
Scattered

TPC

Gamma ray

Scintillator

ΔX

Figure 6.27: Schematic view of the ETCC. The scattered gamma ray is measured with someposition error depending on the thickness of the pixel scintillator.

1st layer

2nd layer

Multi-anode PMT

Figure 6.28: Schematic view of the 2-layer pixel array camera.
of the position resolution is limited by the thickness of the pixel scintillator as shown in Fig.
6.27. One of the solution is use of a larger size TPC than 10-cm scale TPC, because the size ofthe pixel scintillator become smaller relatively than that of the TPC. However, use of a larger
TPC lead to be bad spacial resolution, and it is not advantage to medical imaging.
For medical imaging, the determination of the depth of interaction (DOI) in the scintillator
is eﬀective to improve the position resolution of the scintillator and the ARM of the ETCC.
Although there are some methods of DOI, 2-layer LaBr3 :Ce pixel array (Fig. 6.28) may be
applied to the ETCC because they are easy to keep the energy resolution. Other methods
would lead to worsen the energy resolution, because other scintillator, which has worse energy
resolution than LaBr3 :Ce, is used. When the depth-spatial resolution is ∼ 6 mm in each pixel
scintillator, it is expected to be an angular resolution of ∼ 3.5◦ (FWHM) at 662 keV calculatedusing the method as described in section 6.3.

116

Chapter 7

Conclusion
We have developed an Electron Tracking Compton Camera (ETCC) consisting of a gaseous time
projection chamber and a position-sensitive scintillation camera as an absorber. This camera is
expected to be applied to various ﬁelds such as MeV gamma-ray astronomy and medical imaging.
We have already succeeded in imaging the cosmic and atmospheric gamma rays using our small
size camera µ-PIC onboard a balloon in 2006. Then we have developed a larger camera with asize of 30 cm ×30 cm ×30 cm in order to detect the bright objects such as Cyg X-1. We havealso tested for medical imaging using small animals with radio isotopes.
In this thesis, my works were the following two developments: (i) a continuum gamma-ray
imaging detector for proton therapy, which was also important for astronomy and (ii) a novel
position-sensitive scintillation camera with a better energy resolution than the GSO:Ce arraycamera.
As described in section 5.1, measurement of the prompt gamma rays is important in the
proton therapy, because the distribution of the prompt gamma rays is correlated with the protondose, especially near the end of the proton range. Although a scintillation detector with acollimator scanned the distribution of prompt gamma rays from a water phantom irradiatedwith a proton beam, no imaging detector was available for the real time monitor due to the
gamma-ray backgrounds. Since the ETCC has the background rejection ability and the wide
energy range, it is expected to verify the proton dose even during the proton irradiation using
the continuum gamma rays.
We constructed an ETCC for the monitor in proton therapy. As a ﬁrst step, our aim was
to show the diﬀerence of distribution between the annihilation and prompt gamma rays, and toverify that the prompt gamma ray distribution is useful for the detection of Bragg peak position.
Since higher energy gamma rays were reported to trace the proton dose, an ETCC was tuned
with an energy range of 0.4 to 2 MeV. Obtained energy resolution was 15.3 ± 0.5 % at 662 keV(FWHM), and ARM and SPD were (5.1±0.2)◦ and (119±2)◦ at 662 keV (FWHM), respectively.
Using this ETCC, we tested the gamma-ray imaging from a water phantom irradiated with
140-MeV proton beam in the Research Center for Nuclear Physics (RCNP) of Osaka University
in 2010. The beam current was ∼ 2 pA, and measurement duration was 32 hours (live time), and117

Chapter 7. Conclusion
dead time due to the data acquisition was 6 %. Finally, we succeeded in obtaining the diﬀerence
images between 463 - 559 keV and 0.8 - 2.0 MeV bands which include mainly annihilationand prompt gamma rays, respectively. The annihilation gamma-ray distribution had a ﬂatdistribution (peak width: ∼ 4 cm), while prompt gamma rays (0.8 - 2.0 MeV continuum gammarays) had sharp peak at ∼ 1cm-upstream of proton beam from the Bragg peak. Those results
near the peak of the gamma-ray distribution were consistent with the prediction using GEANT4.
This experiment may be ﬁrst one to carry out the imaging for continuum gamma rays using aCompton camera in the laboratory. In addition, this results obviously indicated the ability of
the ETTC for observation of continuum gamma rays in space.
However, in order to verify the proton dose more clearly, the ETCC should measure the higher
energy gamma rays (≥ a few MeV) with a suﬃcient detection eﬃciency. Thus we develop an
ETCC with a higher detection eﬃciency. For example, we test new gas with an high pressure
for the TPC such as 1.4-atm CF4 based gas.
In addition, the angular resolution of the ETCC at 662 keV was not suﬃcient for medical
imaging and astronomy. Actually, an FWHM spatial resolution was 2.2 cm at 662 keV in the
beam experiment, and medical requirement is ∼ 1 cm. Here, use of the scintillator with ahigher energy resolution is eﬀective to for the improvement of angular (spatial) resolution of the
ETCC. Therefore, we have developed a position-sensitive scintillation camera using a LaBr3 :Cecrystal instead of a GSO:Ce, because the LaBr3 :Ce has an excellent energy resolution. The
monolithic LaBr3 :Ce had an energy resolution of ∼ 3 % at 662 keV (FWHM) with a singleanode Photomultiplier Tube (PMT).
Although this crystal has a hygroscopic nature, we developed an 8 × 8 LaBr3 :Ce pixels
array with a pixel size of 5.8 mm ×5.8 mm ×15.0 mm using our own technique. Finally, we
have developed three arrays. Then FWHM energy resolution of all 64 LaBr3 :Ce pixels reached5.7 ± 0.5 % at 662 keV when coupled to a multi-anode PMT (Hamamatsu H8500), while thatof a GSO:Ce array was 10.8 ± 0.3 % at 662 keV. Therefore, we succeeded in development of thenovel position-sensitive scintillation camera with a better energy resolution than the GSO:Cearray.
As a results, the ETCC with the LaBr3 :Ce arrays had an FWHM angular resolution of
4.2 ± 0.3◦ at 662 keV, and we succeeded in improvement of the angular resolutions. Thus, thespatial resolution of the ETCC is expected to be 1.8 and 1.1 cm at 662 keV (FWHM) when thecenter of 15-cm-high TPC and the target have a distance of 25 and 15 cm, respectively. Since
the medical requirement of the spatial resolution is ∼ 1 cm, our detector have an ability to use
for the monitor of dose in proton therapy. In astronomy, the ETCC is expected to have a better
angular resolution of ∼ 3.5◦ (FWHM) at 1.27 MeV than COMPTEL which had an FWHM
ARM of 4.74◦ at 1.27 MeV [112].

118

reference[1] V. Schönfelder; Astron. Nachr., 323 (2002), 524.
[2] G. K. von Schulthess; ”Molecular anatomic imaging” (2006).
[3] V. Schönfelder; New Astronomy Reviews, 48 (2004), 193.
[4] V. Schönfelder; A & AS 143 (2000), 45.
[5] R. C. Hartman et al.; ApJS, 123 (1999), 79.
[6] W. N. Johnson et al.; ApJS, 86 (2001), 693.
[7] A. M. Levine et al.; ApJ, 469 (1996), 33.
[8] F. Tokanai et al.; Proc. SPIE, 2808 (1996), 563.
[9] R. Jager et al.; A & A Supp., 125 (1997), 557.
[10] S. D. Barthelmy et al.;Space Science Reviews, 120 (2005), 143.
[11] C. Winkler et al.; A & A, 411 (2003), 1.
[12] D. E. Kuhl and R. Q. Edwards; Radiology, 80 (1963), 653.
[13] R. J. Jaszczak; Phys. Med. Biol., 51 (2006), 99.
[14] M. M. Ter-Pogossian et al.; Radiology, 114 (1975), 89.
[15] M. E. Phelps et al.; J. of Nucl. Med., 16 (1975), 210.
[16] G. W. Bennett; Science 200 (1978), 1151.
[17] R. R. Wilson; Radiology, 47 (1946), 487.
[18] G. T. Y. Chen et al.; Int. J. Radiat. Oncol. Biol. Phys., 5 (1979), 1809.
[19] B. Schaﬀner, and E. Pedroni; Phys Med Biol., 43 (1998), 1579.
[20] K. Parodi and T. Bortfeld; Phys. Med. Biol. 51 (2006), 1991.
[21] T. Nishio et al.; Med. Phys., 33 (2006), 4190.
119

REFERENCE[22] C. Min et al.; Applied Phys. Lett., 89 (2006), 183517.
[23] J. Ködlseder; Proc. SPIE, 6266 (2006), 623.
[24] R. Orito et al.; Nucl. Instr. and Meth. A, 513 (2003), 408.
[25] T. Tanimori et al.; New Astron. Rev., 48 (2004), 263.
[26] A. Takeda et al.; IEEE Trans. Nucl. Sci., 51 (2004), 2140.
[27] A. Takada et al.; Nucl. Instr. and Meth. A, 546 (2005), 258.
[28] H. Kubo et al.;Nucl. Instr. and Meth. A, 513 (2003), 94.
[29] K. Miuchi et al.; IEEE Trans. Nucl. Sci., 50 (2003), 825.
[30] K. Ueno et al.;Nuc.r Sci. Symp. Conf. Rec., 2008. NSS ’08. IEEE,(2008), 3470.
[31] F. Sauli; Nucl. Instr. and Meth. A, 386 (1997), 531.
[32] M. Inuzuka et al.; Nucl. Instr. and Meth. A, 525 (2004), 529.
[33] A. Ochi et al.;Nucl. Instr. and Meth. A, 471 (2003), 264.
[34] T. Nagayoshi et al.; Nucl. Instr. and Meth. A, 525 (2003), 20.
[35] H. Kubo et al.; Nucl. Sci. Symp. Conf. Rec., 1 (2006), 406.
[36] A. Takada et al.; J. Phys. Soc. Jpn., 78 (2009), 161.
[37] S. Kabuki et al.; Nucl. Instr. and Meth. A, 580 (2007), 1031.
[38] S. Kabuki et al.;Nucl. Sci. Symp. Conf. Rec., 2008. NSS ’08. IEEE, (2008), 3937.
[39] XCOM; Photon Cross Section Database(http://physics.nist.gov/PhysRefData/Xcom/Text/XCOM.html).
[40] G. F. Knoll; “Radiation Detection and Measurement 3rd edition”, WILEY (2001).
[41] V. Schönfelder; “The Universe in Gamma Rays”, Springer (2001).
[42] V. Schönfelder et al.; ApJS, 86 (1993), 657.
[43] K. S. Cheng, and G. E. Romero; “Cosmic Gamma-Ray Sources”, Kluwer Academic publishers (2004).
[44] P. V. Ballmoos; Experimental Astronomy, 6 (1995), 85.
[45] B. Kozlovsky et al.; ApJS, 141 (2002), 523.
[46] S. Hayakawa, Prog. Theo. Phys. 8 (1952), 571.
120

REFERENCE[47] P. Morrison, Nuovo Cimento 7 (1958), 858.
[48] W. Kraushaar et al.ApJ 177 (1972), 341.
[49] G. Badhwar et al.; Astrophysics and Space Science, 27 (1974), 147.
[50] C. E. Fichtel et al.; ApJ 186 (1973), 99.
[51] G. Bignami et al.; Space Sci. Instrum 1 (1975), 245.
[52] W. A. Mahoney et al.; ApJ, 262 (1982), 742.
[53] A. Palacios et al.; Proceedings of the 5th INTEGRAL Workshop on the INTEGRAL Universe, (2004), 111.
[54] J. Knödlseder et al.; A & A, 441 (2005), 513.
[55] K. S. Cheng et al.; ApJ, 645 (2006), 1138.
[56] N. Guessoum et al.; A & A, 457 (2006), 753.
[57] M. J. Harris et al.; A& A, 461 (2007), 723.
[58] F. W. Stecker;Ap & SS, 6 (1970), 377.
[59] F. A. Aharonian and A. M. Atoyan; A & A 362 (2000), 937.
[60] G. Ghirlanda et al.; ApJ, 613 (2004), 13.
[61] N. Tanvir et al.; Nature, 461 (2009) 1254.
[62] N. Salvaterra et al.; Nature, 461 (2009) 1258.
[63] Y. Inoue et al.; astro-ph/1001.0103[64] V. Schönfelder et al.; Astron. Astrophys. Suppl. Ser., 143 (2000), 145.
[65] R. C. Hartman et al.; ApJS, 123 (1999) 79.
[66] A. A. Abdo et al.; ApJS, 188 (2010) 405.
[67] B. Reiner et al.; Chest, 120 (2001), 1803.
[68] T. F. Delaney et al.; Cancer Principles & Practice of Oncology, 17 (2003), 1.
[69] T. F. Delaney et al.; Int J Radiat Oncol Biol Phys. Supplement, 42 (1995), 222.
[70] V. Benk et al.; Int J Radiat Oncol Biol Phys., 31 (2003), 577.
[71] A. Terahara et al.; Int. J. Radiat. Oncol. Biol. Phys., 45 (1999), 351.
[72] A. E. Rosenberg et al.; American J.l of Surgical Pathology, 23 (2003), 1370.
121

REFERENCE[73] Y. Shioyama et al.; Int. J. Radiat. Oncol. Biol. Phys., 56 (2003), 7.
[74] Y. Matsuzaki et al.; Radiol. Phys. Technol. 3 (2010), 84.
[75] Y. K. Lim et al.; Medical Dosimetrists, in press.
[76] J. J. Coen and A. L. Zietman;Nature Reviews Urology, 6 (2009), 324.
[77] U. Schneider and E. Pedroni; Med. Phys., 22 (1995), 353.
[78] J. Beebe-Wang et al.; Nuclear Science Symposium Conference Record, 2003 IEEE , 4(2003), 2496.
[79] NNDC, http://www.nndc.bnl.gov/.
[80] J. Beebe-Wang et al.; Proceedings of EPAC 2002, Paris, France, (2002), 2721.
[81] K. S. Seo et al.; J. Korean Phys. Soc., 48 (2006), 855.
[82] Jean

in’t

Zand

Coded

Aperture

ImaginginHigh-Energy

Astronomy(http://lheawww.gsfc.nasa.gov/docs/cai/coded.html).
[83] S. D. Barthelmy et al.; SSR, 120 (2005), 143.
[84] Peter Von Ballmoos; New Astronomy Reviews, 48 (2004), 243.
[85] Le fonctionnement d’une lentille gamma（http://www.cesr.fr/~pvb/lentille_gamma/L_fon_e.html）[86] G. Weidenspointner et al.; A& A, 368 (2001), 347.
[87] V. Schönfelder; New Astronomy Reviews, 48 (2004), 193.
[88] G. K. Loudos et al.; Computerized Medical Imaging and Graphics, 27 (2003), 307.
[89] C. S. Levin and E. J. Hoﬀman; Phys. Med. Biol., 44 (1999), 781.
[90] S. Oishi et al.; J. Radioanal. and Nucl. Chem., 239 (1999), 411.
[91] K. Matsumoto et al.; Biol. Pharm. Bull., 28 (2005), 2029.
[92] M. Singh and D. Doria; Med. Phys., 10 (1982), 428.
[93] S. Motomura et al.; Pure Appl. Chem., 80 (2008), 2657.
[94] A. Studen et al.; Nucl. Instr. and Meth. A, 501 (2003), 273[95] A. Zoglauer, and G. Kanbach ; Proceeding of SPIE, 4851 (2003), 1302.
[96] A. Takada; Ph. D. Thesis, Kyoto Univ. (2007).
122

REFERENCE[97] T. Nagayoshi; Ph. D. Thesis, Kyoto Univ. (2004).
[98] K. Hattori; Ph. D. Thesis, Kyoto Univ. (2009).
[99] O. Sasaki; “Ampliﬁer-Shaper-Discriminator ICs and ASD Boards”(http://online.kek.jp/~sosamu/ASD-PRR.pdf)[100] R. Orito; Ph. D. Thesis, Kyoto Univ. (2005).
[101] H. Sekiya et al.; Nucl. Instr. and Meth. A, 563 (2006), 49.
[102] G. R. Lynch, O. I. Dahl; Nucl. Instr. and Meth. B, 58 (1991), 6.
[103] K. Parodi et al.; Phys Med Biol., 47 (2002), 21.
[104] K. Parodi et al.; IEEE Trans. Nucl. Sci., 52 (2005), 778.
[105] K. Parodi et al.; Med. Phys., 34 (2007), 419.
[106]

Stopping Powers and Ranges for Protons and Alpha Particles,

ICRU, Rep. (993) 49..

[107] M.-H. Richard et al.; 2009 IEEE Nuclear Science Symposium Conference Record, (2009),
4172.
[108] J-W. Kim and D. Kim; Radiation Measurements, 45 (2010), 1417.
[109] J. C. Polﬀ et al.; Phys. Med. Biol., 54 (2009), 519.
[110] F. L. Foulher et al.; IEEE Trans. Nucl. Sci., 57 (2010), 2768.
[111] M.-S. Park et al.; App. Phys.letter, 97 (2010), 153705.
[112] V. Schönfelder; Advances in Space Research, 11 (1991), 313.
[113] M. Takahashi, et al.; Nucl. Instr. and Meth. A, in press.
[114] J. D. Parker et al.; Nuclear Science Symposium Conference Record (NSS/MIC), 2009
IEEE , (2009), 1107.
[115] C. Trotta et al.; Nucl. Instr. and Meth. A, 577 (2007), 604.
[116] Y. Chou et al.; ApJ, 618 (2005), 856.
[117] L. Bouchet et al.; ApJ, 548 (2001), 990.
[118] S. M. Schindler et al.; Nucl. Instr. and Meth. A, 384 (1997), 425.
[119] H. O. Anger; Rev. Sci. Instr., 29 (1958), 27.
[120] S. Surti et al.; IEEE Trans. Nucl. Sci., 50 (2003), 1357.
123

REFERENCE[121] W.-H. Wong; IEEE Trans. Nucl. Sci., 40 (1993), 962.
[122] S. Kim et al.; IEEE Trans. Nucl. Sci., 53 (2006), 40.
[123] R. Pani et al.; Nucl. Instr. and Meth. A, 477 (2002), 72.
[124] S. Siegel et al.; IEEE Trans. Nucl. Sci., 43 (1996), 1634.
[125] R. Pani et al.; Nucl. Instr. and Meth. B, 213 (2004), 197.
[126] C. S. Levin; Nucl. Instr. and Meth. A, 497 (2003), 60.
[127] W. R. Cook et al.; IEEE Trans. Nucl. Sci., 32 (1985), 129.
[128] T. Y. Song. et al.; IEEE Trans. Nucl. Sci., 52 (2005), 1396.
[129] R. Pani et al.; Nucl. Instr. and Meth. A, 576 (2007), 15.
[130] R. Pani et al.; Nucl. Instr. and Meth. A, 571 (2007), 187.
[131] H. Nishimura et al.; Nucl. Instr. and Meth. A, 573 (2007), 115.
[132] K. Takagi, and T. Fukazawa ; Appl. Phys. Lett, 42 (1983), 43.
[133] H. Ishibashi et al. ; IEEE Trans. Nucl. Sci., 38 (1989), 170.
[134] R. Hofstadter ; Phys. Rev., 74 (1948), 100.
[135] W. van Sciver, and R. Hofstadter ; Phys. Rev., 84 (1951), 1062.
[136] M. Kobayashi et al. ; Nucl. Instr. and Meth. A, 330 (1993), 115.
[137] W. Lee, D. K. Wehe, and B. Kim ; IEEE Trans. Nucl. Sci., 52 (2005), 1119.
[138] K. Shimizu et al.; IEEE Trans. Nucl. Sci., 35 (1988), 717.
[139] J. Rogers et. al.; IEEE Trans. Nucl. Sci., 36 (1989), 993.
[140] W-H. Wong; IEEE Trans. Nucl. Sci., 33 (1986), 591.
[141] J. S. Karp; Nucl. Instr. and Meth. A, 260 (1987), 509.
[142] E. V. D. van Loef et al.; Appl. Phys. Lett., 79 (2001), 1573.
[143] E. V. D. van Loef et al.; Nucl. Instr. and Meth. A, 486 (2002), 254.
[144] P. R. Menge et al.; Nucl. Instr. and Meth. A, 579 (2007), 6.
[145] K. S. Shah et al.; IEEE Trans. Nucl. Sci., 50 (2003), 2410.
[146] K. S. Shah et al.; IEEE Trans. Nucl. Sci., 51 (2004), 2395.
124

REFERENCE[147] A. Iltis et al.; Nucl. Instr. and Meth. A, 563 (2006), 359.
[148] J. Glodo et al.; IEEE Trans. Nucl. Sci., 52 (2005), 1805.
[149] P. F. Bloser et al.; Nucl. Sci. Symp. Conf. Rec., 3 (2006), 1500.
[150] S. Normand et al.; Nucl. Instr. and Meth. A, 572 (2007), 754.
[151] J. Seguinot et al.; Il Nuovo Cimento C, 29 (2006), 429.
[152] W. M. Higgins et al.; J. of Crystal Growth, 287 (2006), 239.
[153] P. F. Bloser et al.; Nucl. Sci. Symp. Conf. Rec., 2 (2006), 1204.
[154] A. Kuhn et al.; IEEE Trans. Nucl. Sci., 51 (2004), 2550.
[155] R. Pani et al.; Nucl. Instr. and Meth. A, 567 (2006), 294.
[156] M. Kokubun et al.; IEEE NSS-MIC 2006, October 29 Nov. 4, 2006, at San Diego, N30-92(2006).
[157] G. L. Case et al.; Proceeding of SPIE, 5898 (2005), 58980K.
[158] B. Budden et al.; Nucl. Sci. Symp. Conf. Rec, 2 (2006), 943.
[159] C. M. Rozsa et al.; Saint Gobain Crystals, available at www.detectors.saintgobain.com.
[160] Hamamatsu Photonics K.K.; R6231 PMT data sheet; Japan.
[161] K. Hattori; Master Thesis, Kyoto Univ. (2006).
[162] L. A. Shepp, and Y. Vardi; IEEE Trans. Med. Imag., MI-1 (1985), 113.

125

Acknowledgments
First and foremost I would like to thank my advisor, Professor Toru Tanimori. He has taught
me how good experimental physics is done. I appreciate all his contributions of time and ideas
to make my Ph.D. experience productive and stimulating. This thesis would not be completed
without his support. Also, I would like to thank Dr. Hidetoshi Kubo for his advice. He taught
me how good experimental physics is done.
In the medical imaging study, I greatly thank the collaborators in Korea, Dr. J-W. Kim(Korea national cancer center) and Mr. D. Kim (Seoul Univ.) for their supports. In addition,
they was very helpful in my Korean trip. For the preparation of the proton beam test, I would
like to tank Mr. K. Ueno (RIKEN) Dr. S. Kabuki, Mr. S. Iwaki, Mr. M. Takahashi, Mr. K.
Taniue, and Mr. N. HIgashi (Kyoto Univ.).
For the development of the LaBr3 arrays, I would like to thank you Dr. Y. Yanagida and
Mr. T. Kadono. My works would not be completed without their support.
I specially thanks to our MeV-gamma ray collaborators Dr. K. Miuchi, Dr. J. D. Parker,
Dr. A. Takada, Dr. Y. Kishimoto, Mr. T. Sawano, Mr. K. Nakamura Mr. Y. Matsuoka Mr.
M. Aono (Kyoto Univ.), Dr. K. Hattori (Okayama Univ.) and Dr. H. Nishimura. I also thankmy colleagues and friends: Mr. M. Nobukawa, S. Nakano, and Y. Takikawa.
This work was supported by the Korea Japan Collaborative Research Project funded by
National Research Foundation of Korea, Japan Society for the Promotion of Science (JSPS) ,Scientiﬁc Research of the Japan Ministry of Education, Culture, Science, Sports, Technology(MEXT), ”SENTAN” by the Japan Science and Technology Agency (JST), the Global COE
Program ”The Next Generation of Physics, Spun from Universality and Emergence” from the
MEXT, and JSPS Research Fellowships for Young Scientists.
In addition, Dr. M. Ishiura and Dr. T. Kouyama (Nagoya Univ.) taught me the techniqueof the experiment and support my study life. Finally, I am forever indebted to my family andcaretaker of my house in Kyoto, who is 101 years old !, for their understanding, endless patienceand encouragement.
KUROSAWA, Shunsuke
Kyoto University
January 2011126