WO2009072056A2

WO2009072056A2 - Monolithically integrated crystalline direct-conversion semiconductor detector for detecting incident x-radiation at ultra-fine pitch and method for manufacturing such an x-ray semiconductor detector

Info

Publication number: WO2009072056A2
Application number: PCT/IB2008/055032
Authority: WO
Inventors: Nicolaas J. A. Van Veen; Johannes W. M. Jacobs
Original assignee: Koninklijke Philips Electronics N. V.
Priority date: 2007-12-04
Filing date: 2008-12-01
Publication date: 2009-06-11
Also published as: WO2009072056A3

Abstract

The present invention refers to a monolithically integrated crystalline direct- conversion semiconductor detector for detecting X-radiation incident to a detector surface exposed to an irradiation with X-rays at ultra- fine pitch, a fluoroscopic or radiographic X-ray imaging system, 3D rotational angiography device, X-ray C-arc system, fan or cone beam computed tomography imaging device which comprises a monolithically integrated crystalline direct-conversion semiconductor detector as well as to a method for manufacturing such a direct-conversion semiconductor detector. According to the present invention, said direct-conversion semiconductor detector comprises an unstructured semiconductor layer (303) which is made of a crystalline direct-conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer (304) forming a cathode, wherein said semiconductor layer (303) is glued with its cathode side onto a support substrate (307) with an X-ray transparent intermediate electrically conductive layer (305) made of an anisotropic conductive adhesive film or paste lying in-between. According to the present invention, said semiconductor layer (303) is thinned to a desired thickness of e.g. 300 μm, thereby having a polished surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer (305). Said surface is contacted at ultra-fine pitch with metal bumps (302a, b) of a bumped CMOS wafer (301) given by a crystalline semiconductor layer, wherein said metal bumps (302a, b) are bonded to the anode-sided surface of the semiconductor layer (303).

Description

Monolithically Integrated Crystalline Direct-Conversion Semiconductor Detector for Detecting Incident X-radiation at Ultra-Fine Pitch and Method for Manufacturing such an X- Ray Semiconductor Detector

FIELD OF THE INVENTION

The present invention refers to a monolithically integrated crystalline direct- conversion semiconductor detector for detecting X-radiation incident to a detector surface exposed to an irradiation with X-rays at ultra-fine pitch, to a fluoroscopic or radiographic X- ray imaging system, 3D rotational angiography device, X-ray C-arc system, fan or cone beam computed tomography imaging device which comprises a monolithically integrated crystalline direct-conversion semiconductor detector as well as to a method for manufacturing such a direct-conversion semiconductor detector.

BACKGROUND OF THE INVENTION

The increasingly significant role of radiation imaging in science and technology, with its numerous applications in medical imaging and high-energy particle physics, has led to extensive research for developing new pixel detector materials as well as adequate technologies for image capture and acquisition. The need to replace film radiography by filmless approaches is widely acknowledged in the medical community. Replacing an all-purpose medium used for image acquisition, storage and display with a technology which optimizes each task will result in productivity improvements in radiology. Real-time imaging, elimination of consumables (film, chemicals) and tasks (film handling), facilitation of image display, archiving and transfer are some of the advantages. For the past decades, charge-coupled devices (CCDs) have been unequal leader in the field of electronic image sensors for all kinds of applications. This has been driven by the market demand for ever larger pixel numbers and better image quality. However, interest in image sensors based on CMOS technology has dramatically increased in the past ten years. CMOS-based imaging devices offer significant advantages over CCDs such as system-on-chip capability, low power consumption and possibly lower cost. Intensive research in new generations of pixel detectors, as well as adequate readout methods to achieve optimum imaging performances, is rapidly developing. The combination of different radiation detection methods and image capturing techniques demand adaptation of various technologies for detector processing and read-out electronics. The choice between integrating type and single photon counting readout modes is an important decision. Creating design techniques for pixel circuitries and devising methods to achieve good image properties, such as e.g. high resolution and low noise, are typical challenges, to only name a few.

Today, flat image detectors increasingly are replacing conventional analog and digital film- foil systems used in X-ray radiography, mammography and angiography. By applying digital detector systems, a large number of aims shall be achieved. A fundamental desire consists in enabling clinically important radiographic technologies at reduced radiation doses. For this purpose, absorption of X-rays incident to a detector surface which is exposed to X-radiation should be as high as possible in order to accomplish a transformation of all incident X-ray photons into a usable signal. In addition to that, an undistorted high-quality image shall be available on a viewing display immediately after or during the irradiation so as to be able to use the detector system for real-time imaging, such as e.g. in X-ray fluoroscopy, and to enable a thorough radiological analysis based on medical image editing programs. For achieving a high image quality, the detector system must be able to suppress distorting noise signal portions and exhibit high spatial and energy resolution as well as high contrast.

For evaluating the efficiency when transforming X-radiation into an image signal, the detective quantum efficiency (DQE) can be used, which is defined as the quotient of squared signal-to-noise ratio at the detector output to squared signal-to-noise ratio at the detector input as a function of spatial frequency. The DQE describes the efficiency of a detector system during detection of X-ray photons incident to the detector surface and depends on the detector itself, the quality of X-radiation, the employed radiation dose and the spatial frequency to be detected, thereby taking on a value which is always lower than one. Further performance criteria for a detector system are given by its dynamic range, usable dose range, its achievable image frame rate, pixel size, the quotient of active and passive detector surface area and, finally, geometrical shape and size of the detector.

To obtain an image signal which can be visualized on a display, X-ray photons incident to a detector surface are converted into electric signals. Depending on the applied type of conversion (direct or indirect), detector systems can be subdivided into two different technologies.

On the one hand, X-ray detection (such as e.g. needed for high-resolution X- ray mammography, coronary angiography, etc.) may be achieved by indirect conversion using a scintillator detector, which is a two-stage process. Conventional scintillator detectors as known from the prior art, such as e.g. thallium-activated cesium iodide scintillators, as used in conventional X-ray detectors for radiographic diagnosis of a patient to be non- invasively examined, work by converting X-radiation which has been applied to the patient's body to photons of a light which may be composed of wavelengths from the ultraviolet or visible spectrum and then detecting the emitted light with a photomultiplier tube (PMT) or photodiode array consisting of a number of photodiodes arranged in a pixel matrix which convert these scintillation light pulses into a set of electrical charges forming an electric signal to be amplified and measured. An X-ray photon arriving on such a scintillator detector deposits all or at least a part of its energy in the scintillator material in the form of the kinetic energy of numerous electrons, depending on the type and number of interactions. These electrons are able to excite to the conduction band other electrons which can be captured by a trace impurity (an activator atom) and cause transitions leading to the emission of visible light. The role of the activator is to generate meta- states between the pure crystal valence and conduction bands, so that an electron excited to the conduction band can drop in one of this meta-states and de-excite from it to the valence band. This has the advantages of being a more efficient mechanism with respect to the normal de-excitation from crystal conduction band and leads to the emission of visible light photons because of the lower energy of meta- states with respect to the conduction band. Scintillating materials can also be hybridized with a semiconductor photosensor for radiation imaging when being used as a coating layer on a pixel matrix to convert incident radiation into light which may be composed of wavelengths from the ultraviolet or visible spectrum. This light is then detected in the respective photosensor pixel. In this context, it should be mentioned that the coating layer itself can also be pixilated so as to achieve better image resolution.

X-ray detectors using indirect conversion are mainly based on scintillators made of a material which may e.g. be given by an organic crystal dissolved in an X-ray transparent polymer or by an anorganic materials as given by alkaline-earth chalcogenide crystals, such as e.g. terbium-doped gadolinium oxysulfide (Gd₂O₂S(Tb)), or crystalline alkali-halide composite materials, such as e.g. thallium-doped sodium iodide (NaI(Tl)), thallium-doped cesium iodide (CsI(Tl)), cerium-doped lanthanum bromide (LaBr₃(Ce)) or barium fluoride (BaF₂). Scintillator layers made of these materials may thereby be applied to a photodiode read-out sensor substrate made of amorphous silicon (α-Si).

For being applied in X-ray radiography, caesium- iodide scintillator layers are usually manufactured with a thickness of approximately 500 μm. For the radiation spectrum used in X-ray mammography, a thickness between 100 μm and 200 μm is normally sufficient. In the relevant literature it is described that DQE values of more than 60 % can be achieved for a CsI/α-Si-based flat image detector system.

As known from literature, however, scintillator detectors typically exhibit the problem of "sideways leaking" incident X-radiation, which consequently leads to a blurring of the obtained X-ray image. Although this blurring can be overcome by using crystalline direct conversion materials, these materials may exhibit low image frame rates, which should be avoided in modern X-ray imaging.

When applying a direct-conversion detector, on the other hand, incident X-ray photons are directly transformed into electrical charges by means of a semiconductor material, which leads to an image which is obtained as a distribution of charges over a number of pixels arranged in a pixel matrix. For this purpose, a read-out sensor substrate is used which comprises a number of switching transistors, photodiodes and storage capacitors. The maximum spatial resolution and image size of the detector thereby depends on the size and number of the pixels. Using flat image detectors that are based on the direct conversion principle poses high demands on the converter material. The converter material must exhibit good X- ray absorption, such as e.g. lead (II) iodide (PbI₂), cadmium telluride (Cd^Te_x), cadmium zinc telluride (Cdi ^Zn_xTe) or mercury (II) iodide (HgI₂), and generate a high number of charge carriers during the conversion of incident X-ray photons. These charge carriers should be detected as completely as possible by means of a read-out electronics. Therefore, the applied converter material must have a high charge collection efficiency and a small dark current. One of the most promising materials for direct-conversion semiconductor detectors is a ternary II- VI compound semiconductor made of crystalline cadmium zinc telluride (Cdi_ ^Zn_xTe, in the following also referred to as "CZT"). Thereby, x represents the zinc fraction of the semiconductor material, which preferably lies within a range between 0.01 and 0.10.

Detector systems using direct conversion technology, contrary to those which are based on indirect conversion techniques, have the advantage that their energy resolution is not affected by diffusion of secondary light. In addition to that, direct conversion technology leads to a filling factor of almost 100 % and makes realization of small pixel sizes possible.

In medical imaging, direct-conversion detectors made of amorphous selenium (α-Se) are well known and used in the scope of flat image detectors. The amorphous selenium can extensively be deposited and thus directly applied to a read-out sensor substrate which is made of amorphous silicon on glass or a crystalline silicon CMOS wafer. For being applied in X-ray radiography, layers with a thickness of up to one millimeter are needed owing to the low atomic number of selenium (Z = 34). An X-ray quantum which is absorbed by an α-Se layer generates some hundred to more than thousand charge carriers. For a high charge collection efficiency, however, electric field strengths of at least 10 V-μm^"1 are needed. This means that for a 1-mm thick α-Se layer a high voltage of 10 kV is required. Owing to their increased capability and their wide-spread availability, the interest in using direct-conversion semiconductor detectors with large band gap energies has strongly increased in the recent years. This is because they provide high resolution, small size, low current consumption, application within a large range of operation temperature, insensitivity to magnetic fields and, dependent on the respectively applied semiconductor material, good radiation robustness due to their large band gap energy.

For detecting X-radiation, however, the employed semiconductor materials have to meet specific requirements. For example, they shall provide a high detection efficiency, exhibit a low leakage or dark current, have a good charge collection efficiency and high charge carrier mobilities and show a stable behavior. These demands emerge from the principle of radiation detection and shall at least briefly be mentioned in the following sections, beginning with the underlying construction and functional principle of planar semiconductor detectors.

A planar hybrid semiconductor pixel detector generally consists of a thin single crystal which, on one side, is covered with a flat metallization layer which forms the electric rear-side contact. The opposite side (front side) of the semiconductor is provided with a structured metallization layer whose individual subsections are denoted as "pixel contacts" which, in their entirety, constitute the pixel matrix. The size and distances of the pixel contacts along with further material- specific parameters of the semiconductor define the maximum spatial resolution capability of the detector, which typically lies in a range between some ten to a few hundred micrometers. Contrary to flat α-Si based image detectors, the use of directly converting semiconductor pixel detectors, due to the use of separate read-out chips, leads to having the entire pixel area available for the pixel contact, except for a narrow intermediate space to its neighboring pixels. Thus, a comparatively high geometrical fill factor is achieved, and almost the entire pixel area can contribute to charge collection.

Nowadays, silicon (Si), germanium (Ge), gallium arsenide (GaAs), cadmium telluride (CdTe), cadmium zinc telluride ((Cd, Zn)Te), mercury (II) iodide (HgI₂), lead (II) iodide (PbI₂), lead (II) oxide (PbO) and thallium bromide (TlBr) are commonly applied as direct-converter semiconductor materials for X-ray detector applications. However, the demand for an operation at room temperature thereby restricts the number of useable semiconductor materials since e.g. germanium can only be used with an adequate cooling means. When additionally demanding a high likelihood of mutual interaction, only those ones under the aforementioned semiconductor materials are left which are characterized by a high atomic number, such as e.g. HgI₂, PbI₂, PbO, CdTe and (Cd, Zn)Te. Only with these materials it is possible to achieve a comparatively good absorption of incident X-ray photons, even for the small detector thicknesses of only a few millimeters which are required for achieving a good energy resolution. Under the high-Z semiconductor materials, mercury (II) iodide, lead (II) oxide (PbO), cadmium telluride and cadmium zinc telluride are most frequently applied for applications at room temperature.

Due to its high atomic numbers of Z = 48 for cadmium (Cd) and Z = 52 for tellurium (Te) and its high density of p = 5.85 g-cm^"3, cadmium telluride (CdTe) provides for an excellent absorption rate, and its large band gap energy of 1.52 eV allows operation at room temperature since thermal generation of charge carriers is kept to a minimum. Due to its low ionizing energy of 4.43 eV for generating an electron-hole pair, cadmium telluride can optimally be applied for a spectroscopic operation with very high energy resolution. It is commercially available as a photodetector material for detection of X-radiation, γ- and/?- radiation as well as for detecting thermal neutrons and is frequently used in photon-counting and spectroscopic operation. Depending on the quality of the applied semiconductor material, geometrical shape and size of the detector and its electrical contact structures, radiation with energy in the range of a few kiloelectronvolts up to a maximum of 1 MeV can be detected. Due to these properties and its stable behavior, cadmium telluride can suitably be applied for pixel imaging.

Furthermore, the applied metal/semiconductor contacts play an important role in determining the performance of a CZT -based X-ray detector device. A good metal/semiconductor contact preferably has all of the following properties, especially for a segmented detector: good adhesion, capability of preventing charge injection, capability of preventing the inclusion of "oxides" beneath the metal (which is because an intermediate oxide layer sandwiched in between the metal and the CZT will lower the barrier height and potentially cause polarization, thus having a negative effect on detector performance) and reliability for assembly processes, including reflow soldering at low temperatures. The choice of a suitable material for fabricating the electrical contacts is therefore an essential prerequisite for obtaining a good signal-to-noise ratio, which is a necessity in modern X-ray detectors as required by current standards of performance. In addition to that, the type of fabricated metal- semiconductor contacts affects the lines of feree of the electrical field in the interior of a direct-conversion semiconductor detector. On cadmium telluride, two types of metal-semiconductor contacts which lead to a detector operation with ohmic contacts (photo resistors) or Schottky contacts (Schottky diodes) can principally be realized: Ohmic metal-semiconductor-contacts on cadmium telluride are characterized by a non-directional passing of charge carriers and an undisturbed electric field profile. For cadmium telluride, contacts exhibiting an ohmic behavior, which allows the use of the entire crystal volume for detection of incident X-radiation, can only be realized with noble metals such as gold (Au) and platinum (Pt). A necessary prerequisite for the detector operation is thereby a sufficiently high specific resistance of the employed cadmium telluride.

A Schottky contact is a metal-semiconductor contact with a potential barrier which exhibits a rectifying effect. Schottky-contacted CdTe detectors are exclusively operated in reverse direction. Owing to the high specific resistances of Schottky-contacted CdTe detectors lying in a range of up to approximately 10¹⁰ Ωcm, very small leakage currents can be achieved even with not very high-ohmic cadmium telluride crystals.

Due to the use of zinc (Zn), the band gap energy of cadmium zinc telluride lies at a higher level than that of cadmium telluride. For a zinc portion of 10 % there exists a band gap energy of e.g. 1,57 eV. Thereby, smaller leakage currents and an improved mechanical stability can be achieved, which is due to the higher specific resistances of cadmium zinc telluride in the detector operation. Owing to the lower atomic number of zinc (Z = 30), the effective cross-section for an absorption is reduced as opposed to CdTe crystals.

Since cadmium zinc telluride typically exhibits a low n-doped zone, Schottky contacts are obtained when using platinum electrodes on a CZT semiconductor. With this metal, contacts with ohmic behavior can be realized due to the lower work function of gold. The use of hybrid pixel X-ray detectors requires a mechanical and electrical connection of the semiconductor detector with the read-out chip. For imaging applications, a matrix of contacts is needed for electrically connecting image elements of a detector with the input pins of a read-out chip. Unfortunately, these contacts can not be fabricated with conventional wire bonding technology. The principle of arranging contacts by using the entire contact surface is realized in flip-chip mounting technology. This means mounting unhoused semiconductors with their structured side down onto the contacts of a substrate. One of the joining partners which are to be fixedly joined together must be equipped with bumps which are mirror-symmetrically arranged with respect to the contacts of the other joining partner. Due to this arrangement of the bumps on the structured bottom surface of the mounted chip, a minimum occupation of surface area is achieved and contacting of contact matrices is made possible.

The different variations of the flip-chip mounting technology perform differently when being used for hybridization of cadmium telluride and cadmium zinc telluride semiconductor based pixel X-ray detectors with the read-out chips. They can be classified according to the employed materials, methods, processes and connection types. Basically, it can be distinguished between adhesive bonding techniques, thermo compression procedures and reflow soldering processes.

Adhesive bonding techniques: Mounting unhoused ICs based on adhesive bonding techniques can be differentiated according to the properties of the employed adhesives. Today, isotropic and anisotropic electrically conductive adhesives as well as nonconducting adhesives are used. Isotropic electrically conductive adhesives can be used where the use of solder is not possible due to low maximum process temperatures. The electrically conductive adhesive is structured, applied to the contact surfaces by means of silk-screen printing and constitutes the actual electrical connection of the bumps. An anisotropic electrically conductive adhesive is amorphously applied to one of the joining partners which are to be fixedly joined together. The anisotropic conductive film (ACF) consists of a number of polymer bullets having a diameter between here and five micrometers. These bullets are galvanized with a nickel-gold layer covered by a further film consisting of an isolating polymer. The polymer bullets which are placed between the contacts are wedged in between the contacts due to the approximation of the joining partners which are to be joined together. Owing to the pressure which is exerted on them, an electric contact between the contact surfaces of the chip and the substrate is obtained. Non-conducting adhesives are only used in combination with stud bumping. In this flip-chip variant, said bumps are applied by means of a modified wire bonder (such as e.g. given by a gold wire) onto the aluminum contacts of the joining partners. Subsequently, the non-conducting adhesive is applied at a temperature which is required for the liquefaction of the adhesive. The adhesive contained between the contacts is thereby displaced by pressure on the joining partners, and an electrical connection is obtained. Thermo compression procedure: A further flip-chip mounting technique is the thermocompression method, where both joining partners are provided with gold (Au) contacts. Thereby, gold bumps are generated from a gold wire by means of a modified ball- wedge bonder by applying ultrasound and temperature or deposited by electrogalvanics onto one of the joining partners. The actual fabrication of the electrical connections between the joining partners is thereby realized by applying pressure and temperature during the placement procedure.

Reflow soldering: Soldering processes represent the oldest flip-chip technology. The bumps thereby consist of metals or metal alloys having a melting point which lies below a temperature that is crucial for the joining partners. For a protection of the joining partners against impurities and for guaranteeing a good adhesion as well as a good electric connection, a complex under-bump metallization layer is needed between the actual contact surfaces and the solder volume of the bumps. The actual connection is fabricated during the reflow soldering process. Thereby, adhesion forces emerging due to a wetting of the contacts with the solder lead to a mechanical connection of the joining partners. If no complete covering is obtained during the alignment of the joining partners' contact surfaces, asymmetric solder bodies are generated during the reflow soldering process. The forces which result from the surface tension lead to a self-alignment of the joining partners until a complete covering of the contacts is reached. Differences in the soldering processes thereby result from the method to produce the solder bumps.

For the flip-chip mounting of cadmium telluride and cadmium zinc telluride semiconductors, reflow soldering techniques are best suited. At moderate process temperatures, low contact resistances can principally be achieved by the wetting of the entire contact surface with a suitable solder material. The different methods of solder bump generation, however, are not all equally suited for the hybridization of semiconductor pixel X-ray detectors.

SUMMARY OF THE INVENTION

Unfortunately, the performance of direct-conversion semiconductor detectors which are used in imaging applications and fabricated from CZT crystals is often limited. This is due to the fact that conventional fabrication processes do not achieve all four of the above-mentioned contact properties. Conventional direct-conversion semiconductor detectors are typically characterized by pixilated electrode arrays that are fabricated from various deposition and lithography processes with a gap between adjacent pixels, which is known as the "interpixel gap" or region. Interpixel leakage currents act as a source of noise that reduces the ability to spectrally resolve the X-ray photons emitted by the X-ray tube and passing through the patient's body tissue - the energy resolution (ER). The so-called interpixel resistance is a key limitation to performance and is typically much lower than the overall device resistivity. Thus, in order to improve the spectral resolution capability of direct- conversion semiconductor detector devices using CZT crystals, it is desirable to decrease interpixel leakage currents and attendant detrimental noise effects. Furthermore, CZT semiconductor layers are extremely fragile at a desired thickness of 300 μm, which burdens the fabrication process of CZT -based direct-conversion semiconductor detectors. It is thus an object of the present invention to provide an X-ray detector which enables a radiologist to generate high-quality X-ray images with a reduced blurring effect (particularly at high frame rates) and which is fabricated in a way which reliably avoids the manufacturing problems mentioned above.

To address this object, a first exemplary embodiment of the present invention is dedicated to a monolithically integrated solid-state direct-conversion semiconductor detector for detecting X-radiation which is incident to a detector surface that is exposed to an irradiation with X-rays. Said direct-conversion semiconductor detector thereby comprises an unstructured semiconductor layer which is made of a crystalline direct-conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer forming a cathode, wherein said semiconductor layer is glued with its cathode side onto a support substrate with an X-ray transparent intermediate electrically conductive layer, which may e.g. be made of an anisotropic conductive adhesive film or paste lying in-between. In this context, it should be mentioned that adhesives with polymer balls coated with gold should be preferred over solid gold filler. Aside therefrom, conductive adhesives using carbon particles are good candidates. According to the present invention, said semiconductor layer has an uncovered surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer. Said surface is contacted at ultra-fine pitch with metal bumps on a read-out sensor substrate, wherein said metal bumps are bonded to the anode-sided surface of the semiconductor layer. According to the present invention, the support substrate may thereby be made of a dielectric low-Zbulk material having an atomic number (Z) which is much lower than the atomic number of the semiconductor layer, thus leading to minor X-ray absorption, as well as having a low coefficient of thermal expansion (CTE) matched to the thermal expansion coefficient of the semiconductor layer. The read-out sensor substrate may thereby be given by a pixel matrix array on a CMOS silicon wafer, and the semiconductor layer may be made of a crystalline direct- conversion semiconductor material thinned to a desired thickness (e.g. 300 μm), thereby having the surface of the anode-faced side being polished. The crystalline direct-conversion semiconductor material of said semiconductor layer may particularly be given by a CZT semiconductor, and said metal bumps may be made of gold (Au) contacts with additional indium (In) caps placed at the semiconductor layer sided top of said gold contacts. The unpatterned, non-pixilated electrically conductive layer forming said cathode may be made of indium. To stabilize the interconnect structure of the solid-state direct-conversion semiconductor detector, an underfill material may be applied.

A second exemplary embodiment of the present invention refers to an X-ray imaging system, 3D rotational angiography or computed tomography imaging device comprising a mono lit hically integrated solid-state direct-conversion semiconductor detector for detecting X-radiation which is incident to a detector surface that is exposed to an irradiation with X-rays. Again, said direct-conversion semiconductor detector thereby comprises an unstructured semiconductor layer which is made of a crystalline direct- conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer forming a cathode, wherein said semiconductor layer is glued with its cathode side onto a support substrate with an X-ray transparent intermediate electrically conductive layer, which may e.g. be made of an anisotropic conductive adhesive film or paste lying in-between. As already described above with reference to said first exemplary embodiment, said semiconductor layer has an uncovered surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer. Said surface is contacted at ultra-fine pitch with metal bumps on a read-out sensor substrate, wherein said metal bumps are bonded to the anode-sided surface of the semiconductor layer.

As already described above, the support substrate of the direct-conversion semiconductor detector may thereby be made of a dielectric low-Z bulk material having an atomic number (Z) which is much lower than the atomic number of the semiconductor layer, thus leading to minor X-ray absorption, as well as having a low coefficient of thermal expansion (CTE) matched to the thermal expansion coefficient of the semiconductor layer. Again, said read-out sensor substrate may thereby be given by a pixel matrix array on a CMOS silicon wafer, and the semiconductor layer may be made of a crystalline direct-conversion semiconductor material thinned to a desired thickness (e.g. 300 μm), thereby having the surface of the anode-faced side being polished.

The crystalline direct-conversion semiconductor material of said semiconductor layer may particularly be given by a CZT semiconductor, and said metal bumps may be made of gold (Au) contacts with additional indium (In) caps placed at the semiconductor layer sided top of said gold contacts. The unpatterned, non-pixilated electrically conductive layer said cathode may be made of indium. To stabilize the interconnect structure of the solid-state direct-conversion semiconductor detector, an underfill material may be applied. Finally, a third exemplary embodiment of the present invention is directed to a method for manufacturing a monolithically integrated solid-state direct-conversion semiconductor detector for detecting X-radiation incident to a detector surface which is exposed to an irradiation with X-rays. As already described above with reference to said first and second exemplary embodiments, said direct-conversion semiconductor detector may comprise a semiconductor layer made of a crystalline direct-conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer forming a cathode. According to the present invention, said method thereby comprises the steps of gluing said semiconductor layer with its cathode side onto a support substrate with at least one X-ray transparent intermediate layer made of an anisotropic, electrically conductive adhesive film or paste lying in-between, thinning and polishing the semiconductor layer at its anode-side surface on the opposite side of said electrically conductive intermediate layer while being applied to and supported by the substrate and contacting said surface at ultra-fine pitch with metal bumps of a bumped CMOS wafer given by a crystalline semiconductor layer. Furthermore, said manufacturing method may comprise the step of stabilizing the interconnect structure of the obtained solid-state direct-conversion semiconductor detector by an underfill material.

In this connection, it should also be noted that said contacting step may further consist in that said bumped CMOS wafer is bonded to the anode-sided surface of the semiconductor layer in a sequential plating process using the same photolithography mask. Thereby, a low temperature thermocompression bonding technique with or without additional ultrasonic agitation may be applied.

BRIEF DESCRIPTION OF THE DRAWINGS

These and other advantageous features and aspects of the invention will be elucidated by way of example with respect to the embodiments described hereinafter and with respect to the accompanying drawings. Therein,

Fig. 1 shows a direct-converter semiconductor detector with an unpixilated cathode and a pixilated anode for detecting X-radiation, wherein each pixel contact of a pixel matrix is connected to a charge-sensitive preamplifier of an electronic read-out circuitry,

Fig. 2a-i show various fabrication stages of for manufacturing tri-layer metal contacts on a semiconductor substrate at given positions (pixels) for defining radiation detector cells with an interpixel gap and high resistivity between the detector cells as known from the prior art, illustrated by nine schematic cross-sectional side views of a detector substrate with gold contacts on a Cdl-xZnxTe semiconductor layer, Fig. 3 shows a cross-sectional side view of a direct-conversion semiconductor detector according to the present invention, and

Fig. 4 shows the method for manufacturing a mono lit hically integrated solid-state direct-conversion semiconductor detector for detecting X- radiation incident to a detector surface which is exposed to an irradiation with X-rays as claimed in the present invention.

DETAILED DESCRIPTION OF THE PRESENT INVENTION

In the following sections, an exemplary embodiment of the claimed direct- conversion semiconductor detector according to the present invention as well as the corresponding method of manufacturing such a direct-conversion semiconductor detector will be explained in more detail, thereby referring to the accompanying drawings.

A sectional view of a direct-converter semiconductor X-ray detector with a pixilated anode for detecting X-radiation is shown in Fig. 1. Thereby, each pixel contact of a pixel matrix is connected to a charge-sensitive preamplifier of an electronic read-out circuitry. X-radiation which is incident onto the pixel detector (see left pixel in Fig. 1) thereby leads to generation of electron-hole pairs, which is due to mutual interactions of absorbed X-ray photons with the solid-state semiconductor (see middle pixel in Fig. 1). A voltage which is applied to the contacts of the detector generates an electrical field which serves for transporting the generated charge carriers to the anode or cathode contact, respectively (see right pixel in Fig. 1). The movement of the electrons and holes in the semiconductor induces an electric charge pulse in the detector electrodes which is proportional to the absorbed energy and can be processed by a subsequent read-out electronics. Within the semiconductor layer, incident X-ray photons having an energy of typically less than 200 keV electromagnetically interact with the electrons and holes in the atomic orbitals of the atoms of which the semiconductor photodetector material is composed. Predominating effects are interactions with the valence electrons in the outermost atomic orbitals of the absorber atoms, and basically the following four types of mutual interaction can be observed: photoelectric absorption (photoelectric effect), Rayleigh scattering (coherent scattering), Compton scattering (incoherent scattering) and the pair generation effect.

In a photoelectric absorption process (also referred to as "photoelectric effect"), the whole energy of an incident X-ray photon is transferred to a valence electron (photo electron) of the photodetector material, and the X-ray photon is completely absorbed when the energy E₁ of the X-ray photon exceeds the binding energy E_b of the potential photoelectron. After the X-ray photon has been absorbed, the photoelectron which results from this process has the energy of the absorbed X-ray quantum decreased by the binding energy of the valence electron. This kinetic energy is then released again due to multiple mutual interactions with the surrounding semiconductor lattice under generation of a plurality of electron-hole pairs. Since the entire energy of the incident X-ray photon is deposited within the semiconductor detector due to the photoelectric effect, the initial energy of the X-ray photon can be determined. Therefore, photoelectric absorption is the ideal process for measuring the energy of an X-ray photon. The predominating mutual interaction takes place at energies below some hundred kiloelectronvolts. Rayleigh scattering (also referred to as coherent scattering) occurs when particles have a small size as compared to the wavelength of the radiation. Scattering of X- ray photons thereby takes place at all electrons of an absorber atom's atomic orbitals due to stimulation of dipole oscillations without energy transfer to the atom (elastic collision). The weakening of the X-radiation occurs due to the direction change of a part of the radiation. The Compton effect (inelastic X-ray scattering) describes the scattering of particularly higher energetic X-ray photons at free or weakly bound (quasi- free) electrons of an absorber atom's atomic orbital. The X-ray quantum thereby collides with a valence electron of the absorber atom and transfers a part of its energy to a Compton electron which is then emitted. The X-ray quantum is scattered and moves further in a changed direction as given by the scattering angle θ. As the Compton electron leaves the atomic orbital of the absorber atom, the Compton effect is a ionizing process. For the energies which are applied in medical X-ray imaging, the Compton effect plays a dominating role, especially for the scattering of the X-radiation at the human body. Since only a part of the energy of Compton- scattered X-ray photons is deposited by the Compton effect, these photons can cause multiple further interactions by further Compton scatterings or photoelectric absorption. In particular at energies of about 100 keV, the likelihood of multiple mutual interactions caused by the Compton effect is increased.

In the pair generation process, the energy of an X-ray quantum is completely transformed into an electron-positron pair within the Coulomb field of an atomic nucleus. That is, the energy of the X-ray quantum is converted into mass and it dominates the absorption of high-energetic X-radiation. After having released the kinetic energy of the positron which has been generated during the pair generation process, this positron recombines with an electron while emitting an annihilation radiation. Although the pair generation process is not relevant for the energy range which is radiated by X-ray tubes within an X-ray imaging procedure, it is important for the detectors which are used in the scope of hybrid positron emission and computed tomography (PET-CT) devices.

One of the most important assumptions for operating a semiconductor detector is an efficient mutual interaction of atoms with the incident X-radiation. Therefore, high-Z semiconductor materials, which means materials having a comparatively high atomic number (Z), are needed, which is due to the fact that the effective cross section of a semiconductor detector is a function of its material's atomic number:

With increasing atomic number Z, the likelihood of mutual interaction between the incident photons and the photodetector material and thus the probability of absorption increases. As shown in the above table, the atomic number of the applied semiconductor material influences the effective cross-section σ of photoelectric absorption as Z raised to the power of/? (with n being a real value lying between 4 and 5). With increasing atomic number Z, its cross-section therefore increases much faster than the effective cross- section in case of an inelastic X-ray scattering or pair generation. For this reason, high-Z semiconductor materials can preferably be applied as absorption materials. In comparison with low-absorptive semiconductor materials, semiconductor materials with high atomic numbers can be used to realize smaller detector thicknesses at the same high mutual interaction likelihood.

Fig. 2a-i illustrate an example of a detailed fabrication method of forming tri- layer metal contacts on a semiconductor substrate at positions (pixels) for defining radiation detector cells with an interpixel gap with high resistivity between the detector cells as disclosed and claimed in US 2007 / 0194243 Al, which is herewith incorporated by reference. Thereby, it is assumed that the semiconductor substrate is made of cadmium zinc telluride (Cdi ^Zn_xTe) or cadmium telluride (CdTe), although it will be appreciated that other semiconductor materials, for example lead iodide, thallium bromide, gallium arsenide or silicon, can be used. Also, it will be assumed that the metal used for the metallization layer and the contacts is gold, although it will be appreciated that other metals, alloys or other conductive materials, for example platinum or indium, could be used.

Thus, Fig. 2a-i are schematic cross-sectional views from the side of a detector substrate at various stages in the formation of gold contacts on a Cd^Zn_xTe substrate. The detailed features and structure at each step of the process are shown, resulting in an array of contact pixels on the rear surface of the CZT (drawn as facing up in this illustration), protective side coatings, and a single electrode on the front surface of the CZT tile (drawn as facing down in this illustration). In this example, two additional contact layers are added on to the pixilated primary contact layer on the rear side, for improved device assembly. The process can be applied to any array size and pixel configuration for CZT devices. A typical device size for application in PET imaging is a 20x20x5 mm detector, having 8x8 pixels or 11x11 pixels. As a precursor to contact fabrication, the CZT wafer is polished and etched such that high quality clean crystal surfaces are prepared for the deposition process.

The direct lithography fabrication process is described with reference to Figs. 2a- i, and for the case of the primary contact being gold, with two additional contact layers and for simultaneous forming of the cathode contact on the opposing side of the CZT tile. In a first step, a primary layer of gold 200 is deposited on the CZT tile 203. The devices described used electroless deposition, but alternatively the gold may be deposited by known techniques, such as sputtering. The CZT tiles are first cleaned in acetone, as is well known. The clean CZT tiles 203 are dipped in an electroless gold solution for several minutes depositing a gold layer 200, then the tile is removed and rinsed with methanol. Typical thickness of deposition is equal to or greater than 100 nm. The deposited gold may be annealed at 90 ⁰C for 15 minutes to increase adhesion to the substrate. An adhesion test can be done after a few hours using Scotch tape to confirm quality of the adhesion.

In a second step (an optional step), two additional contact layers are deposited onto the rear (to be pixilated) side of the tile, over the primary contact on the rear side. In this example, a nickel (Ni) layer 205 is deposited using sputtering or a thermal evaporation process to a thickness of less than 100 nm and nominally 50 nm. Then, a gold (Au) layer 201 is deposited using sputtering, thermal evaporation and/or an electroless process to a thickness of less than 50 nm and nominally 20 nm. Alternative conductive contact material can be substituted for either or both of the additional contact layers. In a third step, a photoresist 201 is applied over the contact layer(s). Tiles 203 are dipped in resist, such as e.g. a Shipley 1805 resist. Excessive resist is removed if necessary from the edge using a Q-tip, making sure the resist does not form any edge bead (especially on the pixilated face) as this would be detrimental for the pixel quality. Generally, the least possible amount of resist should remain on the pixilated face. The resist should be dried out for 10 minutes with the pixilated face kept up and horizontal.

The resist coating is hardened in a fourth step by baking for ten minutes at 90 ⁰C. This step is done to drive excess solvent out of the resist. The tile is now prepared for lithography exposure.

In a fifth step, a pixel pattern is formed on the rear side of the tile 203 by photolithography. An UV mask 201 is aligned over the CZT tile surface, and the negative resist is exposed to UV. The direct lithography mask shades regions of the resist in a selected pixel pattern and exposes interpixel gaps to UV radiation. A contact mask is used but other methods will work as well, such as proximity and projection masks. A glass plate is placed on top making sure that the glass plate is horizontal. This ensures uniform contact between the tile and the mask. For the example resist, exposure by a UV lamp for several minutes is suitable. If desired, a positive resist may be used instead of the negative resist (in which case, the exposure mask's transparent and opaque regions are reversed).

In a sixth step the exposed photoresist is developed. The resist developer (for example Microposit developer, MF-319) should cover the tile(s). The tiles are placed into the developer with the pixilated side facing up, developed for 2 minutes and the tile(s) are removed from the developer and rinsed in de-ionized water. The UV exposed resist is removed, in preparation for creating the interpixel gap.

In a seventh step the remaining resist (pixel pattern) is baked for 20 minutes at 90 ⁰C. This step is done to harden the resist further. In an eighth step, the exposed contact regions 207 (not covered by the pixel resist pattern 206) are etched. For the example contact materials, the following etching solution is suitable for etching through either just the primary contact layer or the optional three-layer contact. A 2 % bromine ethanol glycol (BrEG) solution is prepared by pouring a 25 ml of ethylene glycol into a plastic beaker, then 0.5 ml of bromine is added using a disposable pipette. Using the same pipette, the solution is mixed thoroughly until it becomes uniform. However, a different pipette or mixing device may also be used. Etching is conducted for approximately three minutes. This etching is done to remove unmasked interpixel contact material. To open the interpixel gap to achieve clean interpixel gaps, active spray agitation is performed. Disposable pipettes can be used to create Br-EG constant flow to agitate for better etching. However, a different pipette or agitation or mixing device may also be used. The spray etching technique should rapidly remove contact material flakes from the interpixel gaps, resulting in high interpixel resistance. The tiles are removed from the etchant and rinsed in deionized water. In a ninth step, the remaining resist is stripped using an acetone bath, resulting in tile 208 with a pixel array of contacts. No photoresist therefore remains on the CdTe/CdZnTe detector since it is usually a hydroscopic material that in time would absorb humidity and deteriorate the detector performance.

The overall combination of depositing the metal layer over the entire substrate surface at once, direct photolithography and the etching process results in the improved device interpixel resistance and performance.

In a tenth step, the primary contact material (in this example gold) on the sides of the fabricated CZT device 209 is removed by side polishing. For example, the sides of the tile(s) are first polished with 1200 grit then with 0.3 micron as fine polish. An alternate embodiment could, in said first step, mask the sides of the CZT tile instead of depositing gold on all sides. For this reason, the side contact removal step (step 10) may be optional. The resulting fabricated CZT device has a cathode contact 200 remaining on the front side, a pixilated anode contact array formed of a primary contact 200, and secondary contact layers 205 and 201, separated by interpixel gap 207. Figs. 2a-i illustrate the multi-layer pixels as being identical width in cross-section for illustrative purpose. The preferred embodiment is that the secondary contact layers are smaller in area than the primary contact pixel. This can be realized by applying the secondary contacts via sputtering.

Typically, as is common in device fabrication, a protective coating is applied to the polished side edges. The CZT tile is dipped in a protective coating (such as e.g. Humiseal) to cover the exposed sides and dried for at least five hours. Fig. 3 shows a cross-sectional side view of a direct-conversion semiconductor detector according to the present invention. As depicted, said direct-conversion semiconductor detector comprises an unstructured semiconductor layer 303 which is made of a crystalline direct-conversion semiconductor material supplied with an unpatterned, non- pixilated electrically conductive layer 304 forming a cathode, wherein said semiconductor layer 303 is glued with its cathode side onto a support substrate 307 with an X-ray transparent intermediate electrically conductive layer 305 made of an anisotropic conductive adhesive film or paste lying in-between. According to the present invention, said semiconductor layer 303 is thinned to a desired thickness of e.g. 300 μm, thereby having a polished surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer 305. Said surface is contacted at ultra-fine pitch with metal bumps 302a, b of a bumped CMOS wafer 301 given by a crystalline semiconductor layer, wherein said metal bumps 302a, b are bonded to the anode-sided surface of the semiconductor layer 303. According to the depicted embodiment, said support substrate 307 is made of a dielectric low-Z bulk material having an atomic number (Z) which is lower than the atomic number of the semiconductor layer 303, thus leading to minor X-ray absorption, and a low coefficient of thermal expansion (CTE) matched to the thermal expansion coefficient of the semiconductor layer 303. As described above, the crystalline direct-conversion semiconductor material of said semiconductor layer 303 is preferably given by a ternary II -VI compound semiconductor such as cadmium zinc telluride (Cdi ^Zn_xTe) with a zinc fraction (x) lying in a range between 0.05 and 0.95.

Said metal bumps 302a, b are made of gold contacts 302a with additional indium caps 302b placed at the semiconductor layer sided top of said gold contacts 302a, and the unpatterned, non-pixilated electrically conductive layer 304 forming said cathode is made of indium.

The interconnect structure of the solid-state direct-conversion semiconductor detector is thereby stabilized by an underfill material. Fig. 4 shows the claimed method for manufacturing a mono lit hically integrated solid-state direct-conversion semiconductor detector for detecting X-radiation incident to a detector surface which is exposed to an irradiation with X-rays according to the present invention. After gluing (Sl) said semiconductor layer 303 with its cathode side onto a support substrate 307 with at least one X-ray transparent intermediate layer 305 made of an anisotropic, electrically conductive adhesive film or paste lying in-between, the semiconductor layer 303 is thinned (S2) and polished (S3) at its anode-side surface on the opposite side of said electrically conductive intermediate layer 305 while being applied to and supported by the substrate 307. The surface is then contacted (S4) at ultra- fine pitch with metal bumps 302a, b of a bumped CMOS wafer 301 given by a crystalline semiconductor layer. Finally, the interconnect structure of the obtained solid-state direct-conversion semiconductor detector may be stabilized (S5) by an underfill material. In this connection, it should also be noted that said contacting step may further consist in that said bumped CMOS wafer 301 is bonded (S6) to the anode-sided surface of the semiconductor layer 303 in a sequential plating process using the same photolithography mask. Thereby, a low temperature thermocompression bonding technique with or without additional ultrasonic agitation may be applied.

APPLICATIONS OF THE INVENTION An immediate application of the claimed direct-conversion semiconductor detector consists in an improvement in the manufacturing process and image quality of direct-conversion X-ray semiconductor detectors as particularly used in the field of medical imaging based on X-ray radiography (such as e.g. needed for generating high-resolution X- ray mammographic or angiographic images with the aid of a computed tomography system or 3D rotational angiography device). When being applied in the scope of an X-ray imaging system, 3D rotational angiography or computed tomography imaging device, the proposed direct-conversion semiconductor detector as described above would provide for generating high-quality X-ray images that are free of blurring (particularly at high frame rates). Applying the claimed method of manufacturing reliably thereby avoids the problem of getting fragile CZT layers when being thinned to a thickness of about 300 μm, which has already been mentioned above.

While the present invention has been illustrated and described in detail in the drawings and in the foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive, which means that the invention is not limited to the disclosed embodiments. Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure and the appended claims. In the claims, the word "comprising" does not exclude other elements or steps, and the indefinite article "a" or "an" does not exclude a plurality. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures can not be used to advantage. Furthermore, any reference signs used in the claims should not be construed as limiting the scope of the invention.

Claims

CLAIMS:

1. A mono lit hically integrated solid-state direct-conversion semiconductor detector for detecting X-radiation incident to a detector surface which is exposed to an irradiation with X-rays, said direct-conversion semiconductor detector comprising an unstructured semiconductor layer (303) made of a crystalline direct-conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer (304) forming a cathode, said semiconductor layer (303) being glued with its cathode side onto a support substrate (307) with an X-ray transparent intermediate electrically conductive layer (305), wherein said semiconductor layer (303) has an uncovered surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer (305), wherein said surface is contacted at ultra- fine pitch with metal bumps (302a, b) on a read-out sensor substrate (301) and wherein said metal bumps (302a, b) are bonded to the anode-sided surface of said semiconductor layer (303).

2. The mono lit hically integrated solid-state direct-conversion semiconductor detector according to claim 1, wherein the support substrate (307) is made of a dielectric low-Zbulk material having an atomic number (Z) which is lower than the atomic number of the semiconductor layer (303), thus leading to minor X-ray absorption, and a low coefficient of thermal expansion (CTE) matched to the thermal expansion coefficient of the semiconductor layer (303).

3. The mono lit hically integrated solid-state direct-conversion semiconductor detector according to claim 2, wherein the read-out sensor substrate (301) is given by a pixel matrix array on a CMOS silicon wafer.

4. The mono lit hically integrated solid-state direct-conversion semiconductor detector according to claim 3, wherein said semiconductor layer (303) is made of a crystalline direct-conversion semiconductor material thinned to a desired thickness, thereby having the surface of the anode-faced side being polished.

5. The mono lit hically integrated solid-state direct-conversion semiconductor detector according to claim 4, wherein the crystalline direct-conversion semiconductor material of said semiconductor layer (303) is given by a ternary II -VI compound semiconductor such as cadmium zinc telluride (Cdi ^Zn_xTe) with a zinc fraction (x) lying in a range between 0.01 and 0.10.

6. An X-ray imaging system, 3D rotational angiography or computed tomography imaging device comprising a mono lit hically integrated solid-state direct- conversion semiconductor detector for detecting X-radiation incident to a detector surface which is exposed to an irradiation with X-rays, said direct-conversion semiconductor detector comprising an unstructured semiconductor layer (303) made of a crystalline direct- conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer (304) forming a cathode, said semiconductor layer (303) being glued with its cathode side onto a support substrate (307) with an X-ray transparent intermediate electrically conductive layer (305), wherein said semiconductor layer (303) has an uncovered surface on an anode-faced side opposite to the side which is attached to said electrically conductive intermediate layer (305), wherein said surface is contacted at ultra- fine pitch with metal bumps (302a, b) on a read-out sensor substrate (301) and wherein said metal bumps (302a, b) are bonded to the anode-sided surface of the semiconductor layer (303).

7. The X-ray imaging system, 3D rotational angiography or computed tomography imaging device according to claim 9, wherein the support substrate (307) of the direct-conversion semiconductor detector is made of a dielectric low-Zbulk material having an atomic number (Z) which is lower than the atomic number of the semiconductor layer (303), thus leading to minor X-ray absorption, and a low coefficient of thermal expansion (CTE) matched to the thermal expansion coefficient of the semiconductor layer (303).

8. The X-ray imaging system, 3D rotational angiography or computed tomography imaging device according to claim 10, wherein the read-out sensor substrate (301) is given by a pixel matrix array on a CMOS silicon wafer.

9. X-ray imaging system, 3D rotational angiography or computed tomography imaging device according to claim 11 , wherein said semiconductor layer (303) is made of a crystalline direct-conversion semiconductor material thinned to a desired thickness, thereby having the surface of the anode-faced side being polished.

10. The X-ray imaging system, 3D rotational angiography or computed tomography imaging device according to claim 12, wherein the crystalline direct-conversion semiconductor material of said semiconductor layer (303) is given by a ternary II-VI compound semiconductor such as cadmium zinc telluride (Cdi ^Zn_xTe) with a zinc fraction (x) lying in a range between 0.05 and 0.95.

11. A method for manufacturing a mono lit hically integrated solid-state direct- conversion semiconductor detector for detecting X-radiation incident to a detector surface which is exposed to an irradiation with X-rays, said direct-conversion semiconductor detector comprising an unstructured semiconductor layer (303) made of a crystalline direct- conversion semiconductor material supplied with an unpatterned, non-pixilated electrically conductive layer (304) forming a cathode, wherein said method comprises the steps of: - gluing (Sl) said semiconductor layer (303) with its cathode side onto a support substrate (307) with at least one X-ray transparent intermediate layer (305) made of an anisotropic, electrically conductive adhesive film or paste lying in-between, thinning (S2) and polishing (S3) the semiconductor layer (303) at its anode- side surface on the opposite side of said electrically conductive intermediate layer (305) while being applied to and supported by the substrate (307) and contacting (S4) said surface at ultra- fine pitch with metal bumps (302a, b) of a bumped CMOS wafer (301) given by a crystalline semiconductor layer.

12. The method of manufacturing according to claim 17, comprising the step of stabilizing (S5) the interconnect structure of the obtained solid-state direct-conversion semiconductor detector by an underfill material.

13. The method of manufacturing according to claim 18, wherein said contacting step (S4) consists in that said bumped CMOS wafer (301) is bonded (S6) to the anode-sided surface of the semiconductor layer (303) in a sequential plating process using the same photolithography mask.

14. The method of manufacturing according to claim 19, wherein a low temperature thermocompression bonding technique with or without additional ultrasonic agitation is applied.