RADIOLOGICAL APPARATUS

TECHNICAL FIELD

The present invention relates to a radiological apparatus .

Specifically, the present invention relates to a radiological apparatus of the mobile type which finds advantageous, but not exclusive application, in surgery, to which the following description will refer without however loosing in generality.

BACKGROUND ART

Mobile radiological apparatuses are known which are adapted to be used in operating rooms in order to obtain a stream of images in real time of internal parts or organs of a patient during a surgical procedure, e.g. during endovascular surgery. In this scope, the realtime image stream is a stream of images with a frame rate of 25 F/s (frames per second) according to the European standard, or of 30 F/s according to the United States standard.

Such a radiological apparatus comprises, in general, an X-ray source, a high-voltage generator adapted to electrically supply the X-ray source to generate a beam of X-rays, an X-ray detector arranged in front of the X-ray source to receive the X-ray beam and form a corresponding image, and a mobile base, which is typically provided with wheels for moving on the floor of the operating room and with a container housing the

high-voltage generator and on which a supporting arm for the X-ray source and the X-ray detector is adjustably mounted. The apparatus is further provided with an image acquisition and display unit connected by cable to the X-ray detector.

The X-ray source consists of, for example, an X- ray tube comprising an anode and a cathode, between which the high-voltage generator applies a potential difference of several tens of kV. The X-ray detector consists of, for example, an image intensifier coupled to a digital camera with CCD sensor which allow to obtain a maximum resolution comprised between 2.0 LP/mm (line pairs per millimetre) and 2.5 LP/mm, or of a solid state flat panel, i.e. a flat panel formed by amorphous silicon or selenium technology, which guarantees a higher maximum resolution, specifically comprised between 3.0 LP/mm and 3.1 LP/mm, but which may be more expensive than the image intensifier, and thus less widespread, and is mainly used on fixed radiological apparatuses. The flat panel formed by amorphous selenium is more delicate than the one formed by amorphous silicon, because the individual sensors which form the panel are damaged beyond repair at temperatures higher than 35° C. Furthermore, the quality of the images provided by the amorphous selenium flat panel depend not only on the received X-ray dose but also, significantly, on the thermal adjustment of the individual sensor elements. Both these types of X-ray detectors, i.e. the

type using image intensifier and the type using amorphous silicon or selenium, are used for exposure times comprised between 7 ms and 10 ms .

The detector and the X-ray source are fixed to the opposite ends of the supporting arm so as to be arranged along an optical axis at a reciprocal distance, which is typically of approximately 1 meter. The supporting arm is shaped so as to be positioned straddling a radio-transparent operating table on which the patient lays so that he or she is arranged between the X-ray source and detector. The mobile base allows to rapidly change the position of the X-ray source-detector pair with respect to the patient's body during surgery.

It is known that only a very small percentage of the electricity supplied to the X-ray tube, equal to approximately 1%, is transformed into X-rays, while the remaining part is dispersed in form of heat mainly in the anode of the X-ray tube. For this reason, the X-ray source is normally provided with a cooling system. For the mobile radiological apparatus, to be sufficiently manoeuvrable, it must have small size and light weight, which means that with respect to fixed radiological apparatuses, it must have a high-voltage generator with a lower maximum power and a compacter X-ray source. Consequently, the X-ray source cooling system of mobile radiological apparatuses is necessarily less efficient than the one normally installed on fixed radiological apparatuses. Furthermore, the mobile radiological

apparatus uses an anodic current normally lower than 100 mA, while anodic currents even higher than 800 mA can be used in fixed apparatuses.

In order to obtain real-time images, the radiological apparatus uses a pulsed technique, according to which the high-voltage generator is controlled by means of a periodic control signal to repeatedly turn the X-ray source on and off so as to generate a sequence of X-ray pulses in accordance with the control signal. The duration of each X-ray pulse must be slightly higher than the X-ray detector imaging time. Specifically, the high-voltage generator comprises typically comprises a voltage transformer, which comprises a primary circuit supplied by a periodic high- frequency voltage modulated by the control signal and a secondary circuit connected to a rectifier in order to provide voltage pulses of several tens of kilovolts with a residue ripple equal to approximately 1% and a duration of a few milliseconds. The rectifying circuit comprises a capacitor having a capacitance chosen according to the best accommodation between voltage pulse duration and residual ripple.

In vascular surgery, endovascular, cardiological and haemodynamic applications the maximum image rate, i.e. 25 F/s or 30 F/s, is often required to reduce the "dragging/lagging" effect which appears on the image sequence and which is due to the movement of the patient's internal organs. Furthermore, in order to

obtain high-quality images or sequences, as required above all in cardiological surgery and haemodynamic applications, the high-voltage generator is controlled to generate the maximum permissible anodic current. The two-fold need for real-time, high-quality images causes the rapid overheating of the X-ray tube, with the risk of taking it to thermal saturation before the surgical procedure is over. With this regard, it is worth mentioning that in some countries it is expressly forbidden to carry out some surgical applications with mobile radiological apparatuses because, today, they do not guarantee that the surgery can be safely concluded without the X-ray tube reaching thermal saturation, which, if it occurs, makes sure that the surgeon can no longer "see" what he or she is doing, and thus it is a potential condition of severe danger for the patient.

Finally, it must be remembered that the European Community directives in the matter of protection from damage caused by exposure to ionizing radiations for medical purposes establish that the exposure doses, except for radio-therapy treatments, must be the lowest possible compatible with obtaining the required radiological information. Therefore, in order to protect patients and medical staff, it is desirable for any type of radiological apparatus, either mobile or fixed, to use pulsed technology and to generate the shortest possible X-ray pulses during surgical procedures. DISCLOSURE OF INVENTION

It is the purpose of the present invention to make a mobile radiological apparatus, which is adapted to reliably acquire radiological images in real time for the entire duration of a surgical procedure, produces X- ray dose rates lower than those of the known radiological apparatuses and, at the same time, is easy and cost-effective to make.

According to the present invention a radiological apparatus as claimed in the attached claims is provided. BRIEF DESCRIPTION OF THE DRAWINGS

For a better understanding of the present invention, there will now be described preferred embodiments, only by way of non-limitative example and with reference to the accompanying drawings, in which: - figure 1 shows, according to a simplified side elevation view, the radiological apparatus made in accordance with an embodiment of the present invention;

- figure 2 shows a block diagram of a part of the radiological apparatus of figure 1 for generating X- rays ;

- figure 3 shows the pattern in time of some signals exchanged between some units of the radiological apparatus in figures 1 and 2; and

- figure 4 shows a block diagram of the X-ray generating part according to a further embodiment of the present invention.

BEST MODE FOR CARRYING OUT THE INVENTION

In figure 1, numeral 1 generically indicates, as

a whole, a mobile radiological apparatus comprising an X-ray source 2, a high-voltage generator 3 adapted to electrically supply the X-ray source 2 so that this generates an X-ray beam 4, and an X-ray detector 5 arranged in front of the X-ray source 2 to receive the X-ray beam 4 and form at least one image of a patient 6 laying on a radio-transparent bed 7 arranged between the X-ray source 2 and the X-ray detector 5. The radiological apparatus 1 comprises a mobile base 8 provided with wheels 9 for moving on the floor (not shown) of an operating room and a container 10 adapted to house the high-voltage generator 3. A C-shaped arm llsupports the X-ray source 2 and the X-ray detector 5 and is adjustably mounted on the mobile base 8 to allow the positioning thereof with respect to the bed 7. Specifically, corresponding containers 12 and 13, one of which houses the X-ray source 2 and the other which houses the x-ray detector 5, are fixed to the opposite ends 11a and lib of the arm 11. The radiological apparatus 1 shown in figure 1 is thus made, from the point of view of the arrangement of the high-voltage generator 3, according to a so-called "housing" solution, in which the high-voltage generator 3 is housed in a different container from the one which houses the X-ray source 2.

The radiological apparatus 1 is provided with an image acquisition and display unit 14 constituted, for example, by a personal computer communicating with the

X-ray detector 5 and appropriately configured for receiving and displaying the images formed by the X-ray detector 5 itself.

According to the present invention, the X-ray detector 5 comprises a CMOS flat panel with direct deposition of a scintillator.

Specifically, the CMOS flat panel, which will be hereinafter indicated again by numeral 5, is a panel sensor of a known type formed on a monolithic silicon support, which comprises a photodiode matrix formed by CMOS technology and on which a layer of scintillator material, e.g. caesium iodide (CsI), was directly deposited. This technology allows to obtain a so-called fill factor higher than that of the amorphous silicon or selenium flat panel, which is obtained by composing and connecting a plurality of individual X-ray sensors . Indeed, the fill factor expresses the percentage of active area by unit of surface of the panel, where the active area is the surface which is physically occupied by the sensors and not by other circuit parts needed for the reciprocal connection of the sensors themselves . Specifically, the CMOS flat panel 5 presents a fill factor comprised between 71% and 87%, and the amorphous silicon or selenium flat panel present a fill factor comprised between 37% and 55%. A higher maximum spatial resolution of the flat panel derives from a higher fill factor.

Furthermore, the CMOS flat panel 5 presents an

unknown advantageous feature, and thus not exploited so far in all radiological applications, in both surgical and diagnostic contexts, and namely an imaging time equal to approximately 20 μs, i.e. essentially much shorter than that of the amorphous silicon panels which are currently used with exposure times equal to or higher than 7 ms, as previously mentioned.

In order to fully exploit the features of the CMOS flat panel 5, and specifically the imaging speed, the rest of the radiological apparatus 1 is made as described below. Specifically, the radiological apparatus 1 according to the present invention aims at overcoming the incapacity of high-voltage generators of the prior art to produce high-voltage pulses lasting less than a few milliseconds. Furthermore, the construction of the radiological apparatus 1 according to the present invention aims at preventing the CMOS flat panel 5 from detecting, by virtue of its speed, the residual ripple in the high-voltage pulses, transforming the ripple into noise on the radiographic image. Such a residual ripple would, indeed, be normally filtered by the X-ray detectors used until now because they are slower in imaging.

With reference to figure 2, the X-ray source 2 comprises an X-ray tube with control grid. More precisely, the X-ray source 2 comprises an X-ray tube, which comprises a cathode 15, an anode 16 and a grid 17 arranged between the cathode 15 and the anode 16.

Advantageously, the X-ray tube is of the rotating anode type having an anode angle equal to or higher than 10°. Advantageously, the anode 16 of the rotating anode X-ray tube is adapted to rotate at a speed equal to or higher than 3000 rpm. The X-ray tube presents a focal spot comprised between 0.1 mm and 0.3 mm to exploit the high spatial resolution of the CMOS flat panel 5.

The radiological apparatus 1 comprises a control unit 18, which is adapted to internally generate a signal PS consisting of a sequence of pulses and to output a low-voltage control signal CS obtained according to the signal PS. The radiological apparatus 1 further comprises a driving unit 19, which is connected to the control unit 18 to receive the control signal CS and is adapted to transform such control signal CS into a corresponding intermediate-voltage driving signal DS adapted to drive the grid 17 so that the X-ray beam 4 is generated intermittently, i.e. so as to generate a sequence of X-ray pulses according to the signal PS. The CMOS flat panel 5 is adapted to provide a respective image for each X-ray pulse. A low-voltage signal means a voltage signal the maximum amplitude of which is either lower than or equal to 300 V, and an intermediate- voltage signal means a voltage signal with negative values presenting a maximum amplitude, referred to the electric potential of the cathode 15, comprised between -1000 V and -3000 V.

More in detail, the control unit 18 comprises a

periodic signal generator 20 for generating a periodic voltage VS, e.g. a square wave voltage, a pulse generator 21 for generating the signal PS, a modulator 22 for obtaining the control signal CS by means of an on/off modulation of the periodic voltage VS with the signal PS so that the periodic voltage VS is off, and thus the control signal CS is null, during each pulse of the signal PS. In other words, the control signal CS consists of a sequence of period voltage segments alternated with no-signal periods, each of which has duration equal to that of a corresponding pulse of the signal PS. Advantageously, the period voltage VS oscillates at a frequency comprised between 400 kHz and 600 kHz. The driving unit 19 comprises a voltage transformer 23 for taking the amplitude of the control signal CS to the values of said intermediate voltage and a rectifier block 24 in cascade to the voltage transformer 23 for obtaining, as driving signal DS, a sequence of intermediate-voltage periods, each of which corresponds to a corresponding periodic voltage segment of the control signal CS, alternated with a sequence of no-voltage periods .

Figure 3 shows an example (not in scale) of the time pattern of the signal PS, of the control signal CS generated according to the signal PS and of the driving signal DS obtained from the control signal CS. Letter P indicates the pulses of the signal PS. Advantageously, the pulses P have same duration equal to a determined

pulse duration D. The pulses P follow each other according to a pulse period T, i.e. according to a pulse rate PR equal to 1/T. Each pulse P of the signal PS has an essentially rectangular shape, i.e. presents up and down times considerably lower than the pulse duration D. Specifically, the up and down times are approximately one hundred times shorter than the pulse duration D. The control signal CS has an amplitude (in modulus) of 300 V. Letters PG indicate intermediate voltage portions of the driving signal DS. The portions PG have an essentially constant amplitude equal to -2000 V. As apparent from figure 3 , the driving signal DS has an essentially rectangular waveform which is complementary, from the time point of view, to the signal PS because the absence periods of the portions PG coincide with the presence of the corresponding pulses P, and vice versa. Each portion PG of the driving signal DS is adapted to set the grid 17 at negative potential with respect to that of the cathode 15, so as to inhibit the generation of the X-ray beam 4 during the pulse PG. During the absence of portions PG, i.e. when the driving signal DS is null, the grid 17 is at the same potential as the cathode 15, thus allowing the generation of the X-ray beam at the pulses P. In this manner, a sequence of sequence of X-ray pulses is generated, each of which has essentially the same duration as the pulse duration D. The pattern of the sequence of X-ray pulses is shown and indicated by RX in figure 3.

With reference again to figures 1 and 2, the X- ray source 2 and the driving unit 19 are mounted on an electronic module 25 which is housed in the container 12 fixed to the arm 11. The driving unit 19 is connected to the grid 17 by means of an intermediate-voltage, low- parasitic-capacity connection 26 made on the electronic module 25 itself. Such a connection 26 consists, for example, of a printed circuit and/or a short intermediate voltage cable segment, i.e. a cable segment for intermediate voltage of length shorter than 50 cm. The control unit 18 is connected to the driving unit 19 by means of a common high-frequency, low-voltage coaxial cable 27, so as to allow the control unit 18 itself to be housed in the container 10 of the mobile base 8. The high-voltage generator 3 is connected to the cathode 15 and to the anode 16 by means of high-voltage cables 28 and is adapted to supply a continuous potential difference obtained in a known manner between the cathode 15 and the anode 16 starting from a supply voltage having periodic waveform. Advantageously, the generated potential difference has a selectable amplitude within a range comprised between 40 kV and 120 kV, according to the part of the body of the patient 6 to be penetrated. The high-voltage generator 3 presents an output stage comprising a capacitor (not shown) adapted to level out the generated potential difference. According to the present invention, because the X-ray beam 4 is turned on and off exclusively by driving the

grid 17 by means of the driving signal DS, the frequency of the supply voltage and the capacity of the capacitor are dimensioned to reduce the residual ripple on the generated potential difference as much as possible, compatibly with the small size that a high-voltage generator 3 must have because the residual ripple is detected by the CMOS flat panel 5 as noise. Advantageously, the supply voltage oscillates at a frequency comprised between 60 kHz and 300 kHz. Advantageously, the capacity of the capacitor is comprised between 4 nF and 1 nF in order to obtain a strongly stabilized potential difference, i.e. a potential difference having a residue ripple lower than 0.5% of the amplitude. The pulse duration D of each pulse P is essentially determined only by the imaging time of the CMOS flat panel 5. Specifically, the pulse duration D is higher than the imaging time of the flat CMOS sensor and advantageously the pulse duration D is selected in a range comprised between 20 μs and 1 ms . Choosing such a pulse duration D is possible by adopting an X-ray tube with control grid, which grid 17 can work as an on/off switch of the X-ray beam 4 at practically unlimited speed, and because the highest parasitic capacities present in parallel along the cascade comprised between the control unit 18 and the grid 17 are those due to the coaxial cable 27, and are in fact negligible. Indeed, the connections 26 and 27 do not contemplate the use of

long high-voltage or intermediate-voltage cable segments, which cables would typically present parasitic capacities in the order of 150 pF/m and therefore would not allow to obtain such short pulse durations D. The pulse rate PR can be selected in a range comprised between 0 Hz and 30 Hz. The pulse rate PR equal to 0 Hz corresponds to a degenerate sequence of pulses P, so to speak, i.e. to a signal PS comprising a single pulse P. This configuration is adapted to obtain static images, and thus for diagnostic applications. When real-time images must be obtained, e.g. during a surgical procedure, the pulse rate PR is higher than 0 Hz. For example, by selecting a pulse rate PR equal to 12.5 Hz or to 25 Hz, or to 30 Hz, a frame rate respectively equal to 12.5 F/s, 25 F/s or 30 F/s is obtained.

According to a further embodiment of the present invention shown in figure 4, in which the corresponding elements are indicated with the same numbers and letters as figure 2, the electronic module 25, which comprises the X-ray source 2 and the driving unit 19, and the high-voltage generator 3 are housed in the container 12 fixed to the arm 11 and the high-voltage generator 3 is connected to the cathode 15 and to the anode 16 by means of a high-voltage, low-parasitic-capacity connection 29 consisting, for example, of high-voltage cables shorter than 50 cm. The radiological apparatus shown in figure 4 is thus made, from the point of view of the arrangement of the high-voltage generator 3 , according to a solution

known as "monotank" , in which the high-voltage generator 3 and the X-ray source 2 are housed in the same container to prevent the use of the cables 28 in figure 2. The main advantage of the above-described radiological apparatus 1 is to drastically reduce the doses of X-rays to the patient and to the medical staff in any type of radiological application, and specifically in applications in which a real-time image stream is required, by virtue of the CMOS flat panel 5 used as X-ray detector and by driving the grid of the X- ray source 2, which allows to obtain very short X-ray pulses. Some tests, during which the data shown below were collected, confirm the dose rate reduction obtained by the radiological apparatus 1 of the present invention.

Test 1) Continuous radioscopy (prior art)

• anode current = 3 mA cathode 15-anode 16 voltage = 100 kv • dose rate = 20,886 mS/min

Test 2) Pulsed radioscopy (prior art)

• anode current = 60 mA cathode 15-anode 16 voltage = 100 kV

• pulse duration = 2.5 ms • frame rate = 12.5 F/s

• dose rate = 26,669 mS/min

Test 3) Pulsed radioscopy (invention)

• anode current = 60 mA

cathode 15-anode 16 voltage ^' = 100 kV

• pulse duration = 100 μs

• frame rate = 12.5 F/s

• dose rate = 1,066 mS/min Another advantage which is obtained using the radiological apparatus of the present invention is averting the risk of overheating of the X-ray source 2 , always by virtue of very short X-ray pulses which allow a shorter utilization of the X-ray source 2 in the unit of time.

Furthermore, from the above it is inferred that the use of the CMOS flat panel 5, combined with driving the grid of the X-ray source, is advantageous also for a radiological apparatus of the fixed type having the same X-ray generation layout as the radiological apparatus 1

(figure 2 or 4), but which comprises a generally more powerful high-voltage generator 3, i.e. capable of generating a potential difference between cathode 15 and anode 16 of up to 150 kV and an anode current of up to 800 mA.