WO2010080046A2

WO2010080046A2 - High resolution gamma camera

Info

Publication number: WO2010080046A2
Application number: PCT/PT2010/000001
Authority: WO
Inventors: António Jorge VAZ DUARTE SOARES; Ian Derek Cullum
Original assignee: Isa - Intelligent Sensing Anywhere, S.A.
Priority date: 2009-01-08
Filing date: 2010-01-06
Publication date: 2010-07-15
Also published as: PT104328A; WO2010080046A3

Abstract

The invention described refers to a new Gamma Camera for gamma ray imaging applications, for example in Nuclear Medicine. The new Gamma Camera has an inorganic scintillation crystal (2), whose larger area surfaces are covered by wavelength shifting optical fibres (1). A gamma ray interaction within the crystal leads to the generation of a scintillation light pulse spatially confined. The fibres absorb the scintillation light and subsequently re-emit photons isotropically, with an emission spectrum different from its absorption spectrum. A fraction of the re-emitted photons is trapped within the fibres and channeled along its axis to be detected by silicon photomultipliers (Si-PMT) which are optically coupled to the ends of the fibres (3). The signals of the Si-PMTs, after electronic amplification and digitization, are used to calculate the center of gravity of the light signal detected, and thus estimate the position of interaction of the gamma rays within the crystal. The invention has applications in Nuclear Medicine and industrial applications for spatial mapping of gamma ray radiation.

Description

DESCRIPTION

"HIGH RESOLUTION GAMMA CAMERA"

Technical domain

The invention consists of an instrumentation device for mapping gamma ray radiation, namely for the development of high resolution gamma cameras for Nuclear Medicine or industrial applications.

State of the art and description of problem to be solved

Nuclear Medicine Imaging is divided in two distinct areas - Single Photon Imaging (SPE) and Positron Emission Tomography (PET) . SPE imaging uses lower energy gamma rays and is carried out using a camera gamma. PET uses radioactive markers of higher energy and dedicated PET cameras.

The Nuclear Medicine gamma camera was invented in 1958 by H. 0. Anger (U.S. Patent 3011057) . It is the most widely used imaging device in Nuclear Medicine to generate bi-dimensional images of the distribution of gamma ray emitters previously injected in patients.

In general terms, Anger's initial design remains in use to the current day for the commercially available gamma cameras. Its performance has been improved allowing an intrinsic spatial resolution of the order of 3mm in the current higher resolution cameras. However, these gradual and minor optimizations have had relatively little impact in the overall resolution of the camera.

As a simple description, the gamma camera consists of an inorganic scintillation crystal that absorbs gamma ray radiation converting it in scintillation light pulses. The pulses which are relevant for generating images are those that correspond to localised interactions of gamma rays within the crystal, namely via the photoelectric effect where a photoelectron is generated. The decay of the photoelectron to its stable energy state leads to a series of interactions within the scintillator with generation of scintillation light, whose characteristics are determined by the structure and nature of the scintillating crystal and by the dopants utilised to control its scintillating characteristics. These scintillation pulses are detected by an array of light sensors - photomultipliers (PMTs) . The PMTs convert the scintillation light into a pulse of primary photoelectrons, which is subsequently amplified to create a higher intensity signal, which is then amplified and digitised electronically.

The light spread over various photomultipliers allows calculating the centroid of the scintillation intensity, and thus estimate the position of interaction of the gamma ray within the scintillation crystal. The intrinsic spatial resolution of such a system is defined by the ability of the system to separate two point-like radiation sources impinging on the crystal in 2 distinct positions. The spatial resolution is determined primarily by two factors: it is inversely proportional to the spread of scintillation light generated, and directly proportional to the square root of the number of primary electrons generated by the photomultiplier .

Sufficient light spread is required in order to encompass 3 to 5 photomultipliers and thus allow the calculation of the centroid of scintillation intensity. Given the size of the photomultipliers used in gamma cameras, with diameters between 3 and 5 cm, this spread must be of at least 9cm (defined as the full width at half maximum of the intensity of light along a given direction) . At a first glance, the light spread is determined by the thickness of the scintillation crystal. This thickness is directly linked to the efficiency of absorption of gamma rays. It is important to have a high absorption efficiency, required to reduce as much as possible the radiation dose given to the patient and as well as to allow creating images with statistically relevant quality. For the most widely used scintillation crystal (NaI (Tl) - Sodium Iodide, Tellurium doped) the thickness required to absorb over 90% of the Nuclear Medicine gamma rays is of the order of lcm. This small thickness does not allow the required spread of light and it is therefore common practice to introduce light guides between the crystal and the array of photomultipliers. The use of smaller photomultipliers, which would allow reducing the thickness of the light guides and consequently of the light spread, is not an economically viable solution, since it increases the number of photomultipliers required, and has the the added disadvantage of a considerable increase in the cost per photomultiplier .

As to the number of primary photoelectrons produced by the photomultiplier, it depends on three main factors:

1. the gamma ray energy - it depends on the radiopharmaceutical used as marker for the specific diagnostic required. The great majority of the radiopharmaceuticals use the Tc^99m radioisotope which emits 140KeV gamma rays.

2. the intrinsic conversion efficiency of the scintillation crystal used, which determines the number of photons generated by the interaction of a gamma ray of a particular energy. This is an intrinsic characteristic of the scintillation crystal.

3. the "quantum efficiency" of the photomultiplier, defined as the probability for creating a primary photoelectron by a photomultiplier for each incident light photon.

The two latter factors have been optimised over the past 50 years with only marginal gains on the spatial resolution obtained. On the other hand, this geometry creates a "dead space" at the periphery of the camera field of view, since the correct calculation of the centroid requires the inclusion of signals from multiple photomultipliers around the point of interaction on the scintillation crystal.

New solutions are therefore needed to improve the spatial resolution of gamma cameras, to address the clinical needs and with a reduced production cost.

One alternative solution for the design of a Gamma Camera is through the detection of light signals from scintillation crystals using wavelength shifting optical fibres. This design reduces significantly the light spread since the fibres are directly in contact with the surfaces of the scintillation crystals. Given the small dimension of the fibres (lmm diameter) it is not required to create additional light spread other than that one generated by the thickness of the scintillation crystal itself.

A similar concept was described by W. Worstell in 1994 (U.S. Patent 5600144). A scintillation crystal with a configuration similar to that of a gamma camera is covered over its larger surfaces by two layers of wavelength shifting optical fibres. In one possible configuration, the upper surface of the crystal is covered by fibres oriented along one direction, whereas the lower surface is covered by fibres oriented along a direction perpendicular to the former (Figure 1) .

The interaction of a gamma ray photon within the crystal generates an isotropic pulse of scintillation. Typically, this scintillation occurs within a narrow range of wavelengths according to the crystal's characteristics. A fraction of these scintillation photons hits the optical fibres coupled to the upper and lower surfaces. Depending on the refractive indices of the crystal, of the fibres and of the optical coupling between these, a fraction of these photons enters the optical fibre. The optical fibres should have an absorption spectrum matching the emission spectrum of the scintillation crystal, as well as a high absorption coefficient for that range of wavelengths. The photons that enter the optical fibre are converted into scintillation photons of higher wavelength which are isotropically re- emitted inside the fibres. Photons re-emitted along directions near the fibre axis will be optically trapped and channeled along the fibre to one of its extremities, according to the refractive characteristics of the optical fibre. The number of photons captured and channeled to the extremities of the fibres depends on several coefficients:

1. number of primary photons generated within the scintillation crystal following the conversion of gamma rays into scintillation;

2. number of primary photons hitting the optical fibre;

3. number of primary photons entering the optical fibre; 4. absorption coefficient of the optical fibre for the primary photons (as a function of wavelength) ;

5. re-emission coefficient of the fibre;

6. numeric aperture of the optical fibre (depends on the refractive indices of the fibre core and its cladding layers) ;

7. re-absorption coefficient of the re-emitted photons (as a function of the wavelength and of the length of the optical fibre) .

Given the various losses introduced by the above mentioned coefficients, the maximum number of photons detected at the extremities of one fibre for each gamma ray interaction is a small percentage of the number of primary photons generated, of the order of 0.2% to 0.5%, considering the scintillators and fibres currently available for Nuclear Medicine applications. Studies have shown (A.J. Soares et al, IEEE Trans. Nucl. Sci . , 46(3), pp. 572-582, 1999) that using the most appropriate scintillators for the detection of 140KeV (Tc^99m) gamma rays - NaI(Tl) and CsI(Na) - coupled to blue-to-green wavelength shifting optical fibres (e.g. the Bicron BCF-91A, manufactured by Saint-Gobain Crystals, and the YIl manufactured by Kuraray Corp., Japan) a maximum number of 10 photons would be collected at each fibre extremity per gamma ray interaction.

These photons must be detected by appropriate photo-sensors, typically photomultipliers or silicon photodiodes (e.g. APDs) . The detection efficiency of these photo-sensors for the scintillation photons, the so-called quantum efficiency, introduces yet another limiting factor to the overall detection capability. The quantum efficiency of PMTs for the emission spectrum of these fibres is typically lower than 15%, which means that no more than 1.5 primary photoelectrons are generated at the fibre closest to the position of gamma ray interaction, i.e., the fibre with the highest light signal. The APDs (Avalanche Photo- Diode) are not particularly suited for the detection of low light levels. Despite the higher quantum efficiency of silicon, which could reach as high as 70% for green light, APDs generate background noise at room temperature, whose level is too high for an efficient detection of only a few photons (e.g. up to ten) .

Therefore this technique has never been used for Single Photon Emission Nuclear Medicine applications, being used only for PET (US Patent 7115875), where the gamma ray energy is 511KeV. It is also used in High Energy Physics experiments .

Thus, it was not possible up to now to apply this technology for the development of Gamma Cameras. For the gamma ray energy mainly used in Nuclear Medicine - 140KeV - the number of photons generated is not sufficient for efficient detection with traditional photomultipliers, nor with traditional silicon photodiodes. Description of the invention

The recent invention and development of silicon photomultipliers (Si-PMT, P. Buzhan et al., Nucl. Instrum. Methods A502, 48 (2003) ) has radically changed this situation. These devices combine the high quantum efficiency of silicon for converting photons into primary photo-electrons, with high electronic signal amplification which is a characteristic of photomultipliers, enabling this way a dramatic improvement in the overall detection efficiency of photons.

The present invention proposes therefore the development of a high resolution Gamma Camera based on the imaging technique described above, i.e. using converters of gamma ray radiation into scintillation light (scintillators) directly coupled to wavelength shifting optical fibres, and using silicon photomultipliers for the readout of the light signals trapped inside the fibres. The energy signal readout will be carried out by photomultipliers which collect light not absorbed by the optical fibres.

A Gamma Camera based on this concept will have a significantly better spatial resolution comparing to current Gamma Cameras, mostly due to the reduction of the light spread - the light sensors in this case are the optical fibres, lmm diameter, comparing to the array of photomultipliers currently used in Gamma Cameras (diameters of 3 to 5 cm) . It enables also the construction of a camera with a small "dead area" in the periphery around its field of view, which is particularly relevant for the development of compact gamma cameras. The reduction of the "dead area" is particularly important in Nuclear Medicine as it allows generating images in the most favourable projections via a more flexible positioning of the camera around the object being imaged.

Silicon photomultipliers are made of arrays of small diodes - each has between 20x20 and 100x100 μm - where each individual diode (pixel) operates in high gain mode (Geiger mode) . Creating a large array of these pixels, it is possible to manufacture diodes of 1x1, 2x2 and 3x3 mm, with high detection efficiency for the green light (wavelength approximately 500 nm) , and high gain, allowing therefore excellent single photoelectron resolution.

Description of drawings

- Figure 1 shows a gamma camera with two layers of wavelength shifting optical fibres (1) covering the two largest surfaces of a scintillation crystal (2), positioned orthogonally, and coupled to silicon photomultipliers (3). The energy signal is measured by an array of photomultipliers (4) that collect the light not trapped by the optical fibres.

- Figure 2 shows the coupling between one wavelength shifting optical fibre (10) extremity and one silicon photomultiplier (11) . - Figure 3 shows the two extremities of an optical fibre (20) coupled to two silicon photomultipliers (21) in order to detect all the photons trapped inside each optical fibre .

- Figure 4 shows one possible representation of this invention which is described in detail below.

Detailed description of ways to implement the invention

One representative configuration of the present invention is shown in Figure 4.

The scintillation crystal (30) may be NaI(Tl), CsI(Na) or LaBr3(Ce), whose scintillation light emission spectrum is around the blue region (wavelength of around 420nm) , have high light yield (measured in number of photons generated per KeV of deposited energy) , and have a high density allowing high probability of total absorption of all Tc^99m (140 KeV) gamma rays for small scintillator thickness (e.g. 5 to 8 mm).

In each of its two larger surfaces, which define the field of view and imaging plane, the crystal is completely covered by two layers of blue-to-green wavelength shifting optical fibres with lmm diameter (31) . These layers cover the two opposing surfaces of the crystal, with the fibres positioned orthogonally in the two layers. The light coupling between the crystal and the fibres should take into consideration the different refractive indices in order to optimise the number of blue photons entering the fibres. Silicon gel is a good example of a good optical coupler typical for these materials, but others may be used as well.

This setup allows generating images of the gamma ray interactions in this crystal. The extremities of each fibre are coupled to silicon photomultipliers (32) (e.g. Hamamatsu MPPC S10362) whose high quantum efficiency for green light and high gain allows an accurate estimation of the number of photons trapped and channeled within the optical fibres.

In the camera described in the scope of this invention, the signals channeled through the optical fibres are read out at one of the fibre's extremities using an Si- PMT, whereas the other extremity is polished and coated with a reflective layer to re-direct the light travelling in that direction to the opposite extremity and thus be detected by the Si-PMT.

The above mentioned readout of the optical fibre's light signals may as well be realised at both ends of the fibres using two Si-PMTs, one in each fibre extremity, and adding the Si-PMT signals in order to obtain the total signal of light trapped inside the optical fibres .

In a Gamma Camera the energy signal (i.e., the total signal generated) is important for the rejection of partial energy deposition events, as well as of Compton scattering of gamma rays inside the patient's body, which add background noise to the final image and therefore degrade image quality. In this configuration the energy signal is generated in traditional photomultipliers (33) coupled to one of the optical fibre layers. The fraction of scintillation light that is not trapped inside the fibres is detected by these photomultipliers, producing the energy signal, as well as a trigger to discriminate valid events.

Moreover, the high resolution gamma camera of this invention is equipped with an Si-PMT cooling system to reduce noise.

The electronic signals of the silicon photomultiplier and of the energy photomultipliers are similar and should therefore be treated similarly with respect to the readout electronics. The electronic signals readout of each of these sensors should be carried out in a first amplification stage (34), followed by signal digitisation (35) . The digitised signals are readout by software programs which will calculate the centre of interaction and allow as well the real-time display of the image being generated (36) .

As with a traditional gamma camera, a gamma ray collimator is used (37) . A high resolution collimator should be used to take advantage of the high spatial resolution expected of the camera. Clinical and industrial application of the invention

A gamma camera based on the concept here described will have a high spatial resolution and will have many potential clinical applications, such as in Nuclear Medicine for the early detection of tumours using radiopharmaceuticals specific for the detection of cancerous tissues. Furthermore, this concept enables the construction of compact and highly portable gamma cameras, widening the scope of application within the Nuclear Medicine centres and even outside these. Industrial applications for gamma ray radiation mapping may as well benefit from the portability and high resolution of the camera.

Lastly, the proposed technology differs from the commonly used methods for the development of compact high resolution gamma cameras, where small pixel devices are used in bi-dimensional arrays, requiring readout of n x n pixels, and leading to high cost. The use of optical fibres as the detection method introduces multiplexing on the light signals readout requiring the readout of n + n pixels of Si-PMTs, with a consequent cost reduction.

Claims

1. High resolution gamma camera using converters of gamma ray radiation into scintillation light directly coupled to wavelength shifting optical fibres (1), characterised by the readout of the light signals trapped within the fibres using silicon photomultipliers Si-PMT (32) and by the readout of the energy signals using photomultipliers (33) which collect all the light not absorbed by the fibres.

2. High resolution gamma camera according to claim 1, characterised by the said Si-PMT photomultipliers (32) readout to be carried out at one of the extremities of the optical fibres, while the other extremity of the fibre is polished and coated with a reflecting layer to allow redirecting the light travelling towards that extremity to be reflected and directed to the opposite extremity and consequently be detected by the Si-PMT.

3. High resolution gamma camera according to claim 1, characterised by the readout of the light signals at both extremities of the optical fibre (1) using two Si- PMT (32), one at each extremity of the optical fibre, and adding the signals generated in both Si-PMT in order to obtain the overall light signal trapped within the optical fibre .

4. High resolution gamma camera according to claim 1, characterised by a scintillation crystal whose larger area surfaces are fully covered by two orthogonal layers of optical fibres, and taking into consideration the various refracting indices in order to optimise the optical coupling and therefore maximise the number of photons entering the fibres.

5. High resolution gamma camera according to claim 2, characterised by silicon gel optical coupling.

6. High resolution gamma camera according to claim 1, characterised by the energy signal being readout by the photomultipliers (33) and used for the rejection of partial energy deposition events, as well as for the rejection of Compton scattering events taking place within the patient's body, since both types of events degrade the final image quality, having the said photomultipliers coupled to one of the layers of optical fibres.

7. High resolution gamma camera according to claim 1, characterised by an identical treatment of the electronic signals of the photomultipliers (32 and 33), with the readout carried out with a first amplification stage (34), followed by signal digitisation (35) and by the digitised signals being used by software programs to calculate the centre of interaction and to display the image in real-time as it is formed (36) .

8. High resolution gamma camera according to claim 1, characterised by the usage of a gamma ray radiation collimator (37) .

9. High resolution gamma camera according to claim 1, characterised by a cooling system for the Si-PMT for the reduction of noise.