In the related art, an X-ray imaging device is employed as a medical X-ray image photographing device.  In addition, as an X-ray detector employed in the X-ray imaging device, for example, a flat panel type X-ray detector (hereinafter, abbreviated as an “FPD”) is known in the related art. 
As illustrated in FIG.  24(a), an FPD 101 of the related art has a stacked structure obtained by stacking a panel-like base material 103, a scintillator panel 105, and a photodetection panel 107 sequentially in this order.  The scintillator panel 105 has a scintillator element that absorbs an X-ray and converts it into light.  The photodetection panel 107 has a substrate 109 and pixels 111 arranged in a two-dimensional matrix shape.  Each of the pixels 111 has a photoelectric conversion element and an output element (not illustrated). 
Referring to FIG.  24(a), an X-ray incident to the FPD 101 from a direction indicated by a reference symbol R is converted into light in the scintillator element provided in the scintillator panel 105, and the light is emitted as scintillator light.  The scintillator light emitted from the scintillator panel 105 is transmitted to the pixel 111.  In addition, the photoelectric conversion element provided in the pixel 111 photoelectrically converts the scintillator light to output an X-ray detection signal as an electric signal from the output element.  Furthermore, an X-ray image is created on the basis of the output X-ray detection signal. 
In recent years, in the field of medical industry, so-called dual energy imaging is performed to photograph the same part of the inspection target object at different tube voltages.  In the dual energy imaging, X-ray images based on X-rays having different energy distributions are created independently, so that differences between elements of the inspection target object can be visualized.  For example, by taking a difference between an X-ray image based on the low-energy X-rays and an X-ray image based on the high-energy X-rays, an X-ray image of a hard-tissue such as bones and an X-ray image of a soft tissue such as muscles can be distinguished. 
When the photographing in the dual energy imaging is performed by changing the tube voltage, the X-ray irradiation is performed twice.  For this reason, diagnostic performance of the X-ray image is degraded due to a body movement of the inspection target object.  In this regard, an X-ray detector capable of obtaining a pair of X-ray images including the X-ray image based on low-energy X-rays and the X-ray image based on high-energy X-rays out of the irradiated X-rays by irradiating X-rays onto the inspection target object once has been proposed (for example, see Patent Literature 1).  Such an X-ray detector capable of obtaining the X-ray image based on low-energy X-rays and the X-ray image based on high-energy X-rays through a single X-ray irradiation will be referred to hereinafter as a “dual energy type” X-ray detector.  In addition, X-ray detectors other than the dual energy type X-ray detector will be distinguished as a “typical” X-ray detector. 
Such a dual energy type X-ray detector 201 has a structure in which a pair of FPDs 101 illustrated in FIG.  24(a) are overlapped.  That is, as illustrated in FIG.  24(b), a first FPD 203 for detecting relatively low-energy X-rays and a second FPD 205 for detecting relatively high-energy X-rays are stacked in an X-ray irradiation direction indicated by the reference symbol “R.” A low-energy X-rays P1 are converted by a scintillator element 207 provided in the FPD 203 into scintillator light Q1 and are converted into electric signals in pixels 209. 
Meanwhile, a high-energy X-rays P2 pass through the scintillator element 207 and are converted into scintillator light Q2 in a scintillator element 211 provided in the FPD 205.  The scintillator light Q2 is converted into the electric signal in a pixel 213.  The FPDs 203 and 205 are stacked while base materials A provided in each FPD face each other.  Note that it is necessary to make the thickness of the scintillator element relatively thick in order to absorb and detect high-energy X-rays.  That is, compared to the scintillator element 207, the scintillator element 211 typically has a larger thickness. 
In this regard, a structure in which the scintillator elements are partitioned by partitioning walls in a typical FPD structure different from the dual energy type X-ray detector has been proposed (for example, see Patent Literature 2).  A typical FPD 301 in which the scintillator elements are partitioned by the partitioning walls will now be described with reference to FIGS.  25(a) and 25(b).  Similar to the FPD 101 of FIG.  24(a), the FPD 301 has a stacked structure obtained by stacking a base material 303, a scintillator panel 305, and a photodetection panel 307, and the photodetection panel has a substrate 309 and pixels 311 (FIG.  25(a)). 
As illustrated in FIG.  25(b), the scintillator panel 305 has grid-like light blocking walls 313 and scintillator elements 315.  Each of the scintillator elements 315 is filled in a cell space partitioned by the light blocking wall 313.  In general, the pitch of the light blocking wall 313 is set to be substantially equal to (or an integer multiple of) the pitch of the pixel 311. 
In this manner, the scintillator panel 305 is shaped such that the scintillator elements 315 arranged in a two-dimensional matrix shape are partitioned by the light blocking walls 313.  Since the scattering scintillator light is blocked by the light blocking walls 313, it is possible to prevent the scattering light generated in the scintillator element 315 from reaching a neighboring scintillator element 315.  Therefore, by partitioning the scintillator elements 315 using the light blocking walls 313, it is possible to avoid degradation of the resolution of the X-ray image even when the scintillator element 315 is thickened.  Such a configuration is useful particularly in the X-ray detector for detecting high-energy X-rays. 
In addition, in a structure relating to Patent Literature 2, the pitch of the partitioning wall can be reduced to a short distance such as 60 to 150 μm.  For this reason, using the X-ray detector of Patent Literature 2, it is possible to avoid degradation of the resolution of the X-ray image even when an X-ray image having a smaller pixel pitch is required, such as X-ray CT imaging.