Abstract:
The present invention provides for a method of three dimensional temperature monitoring based in hyperthermic procedures, the heat dissipation, and therefore, temperature variations, does not change dramatically over a spatial region and, second, temperature change, which manifests itself as the phase change, has different characteristics compared to the anatomic magnitude image. Therefore the present invention exploits these facts to provide a method for acquiring the temperature map of three orthogonal planes nearly simultaneously.

Description:
BACKGROUND OF THE INVENTION  
       [0001]     The present invention relates generally to a magnetic resonance imaging (MRI) methods and devices, and more specifically, to a method for using MRI to measure temperature change in either liquid or tissue.  
         [0002]     Thermal energy deposition is often used in medicine as a means of necrosing diseased tissues. Lasers, radio frequency antennas and ultrasonic transducers are examples of devices used for the deposition of thermal energy for therapy. However, regardless of the therapeutic regimen used, it is desirable to have a means of guiding and monitoring this energy deposition to assure the energy is applied in the proper location and to verify that appropriate energy levels are used to prevent undertreatment or overtreatment for two main reasons. The first is to ensure that the diseased tissue has been exposed to an adequate temperature-time treatment to induce necrosis over the entire diseased volume. The second is to ensure that the surrounding healthy tissue is spared excess thermal treatment. Magnetic resonance imaging has been demonstrated as a method for identifying regions of tissue to be treated, guiding therapeutic devices and monitoring the deposition of thermal energy from lasers, ultrasound devices or cryogenic probes.  
         [0003]     When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B 0 ), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B 1 ) which is in the X-Y plane and which is near the Larmor frequency, the net aligned moment, M z , may be rotated, or “tipped”, into the X-Y plane to produce a net transverse magnetic moment M t . A nuclear magnetic resonance (NMR) signal is emitted by the excited spins after the excitation signal B 1  is terminated and this signal may be received and processed to form an image.  
         [0004]     When utilizing NMR signals to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The region of interest may be a small portion of the patient&#39;s anatomy, such as the head or heart, or a much larger portion such as the entire thorax or spine. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well-known reconstruction techniques. To perform such a scan, it is necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G x , G y  and G z ) which have the same direction as the polarizing field B 0 , but which have a gradient along the respective X, Y and Z axes. The magnetic field gradients are produced by a trio of coil assemblies placed around the object being imaged. By controlling the strength of these gradients during each NMR measurement cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.  
         [0005]     MRI has many advantages compared to other imaging modalities such as computed tomography (CT), ultrasound and x-ray. One of these advantages is that MRI can be used to measure temperature change in either liquid or tissue. This allows continuous monitoring of heat dissipation during a hyperthermic procedure such as ablations using various devices. To the inventor&#39;s knowledge, no other major imaging modality has that capability.  
         [0006]     In recent years, there has been rapid development both technically and clinically in the field of thermal imaging with MRI. In general, these techniques are employed to guide non-invasive and minimally invasive interventional procedures. It is expected that this growth trend will only continue with the improvement of the efficacy of these techniques and resulting wider acceptance of these new techniques in the clinical arena. A variety of methods are used to heat bodily tissue during these procedures. However, the most commonly used heating procedures the inventor is aware of include: laser radiation; microwave radiation; radio frequency radiation; and focused ultrasound.  
         [0007]     Several techniques have been used to measure the temperature change including measuring temperature induced change in the longitudinal relaxation time, in the diffusion coefficient, or in the water proton resonance frequency (PRF) shift. The most robust and commonly used technique thus far is based on the PRF shift because of the linearity of the phase change with respect to the temperature change and the near independence of the tissue type.  
         [0008]     MR temperature mapping technique based on PRF was first proposed by Ishihara et al. The Larmor precession frequency depends on the local magnetic field 
 
ω=γB loc  
 
 in which B loc  is determined by 
 
B loc   =B   0 [1+σ 0 +σ( T )]
 
 and γ is the proton gyromagnetic ratio (42.58×10 6  Hz/T). The chemical shift terms σ 0  and σ(T) represent the temperature independent and the temperature dependent contributions respectively and are measure in ppm. The temperature independent term include the effect from B 0  field inhomogeneities. The temperature dependent term is linearly proportional to the temperature within the temperature range that is of interest to most hyperthermic treatment (40-70° C.): 
 
σ( T ) =αT  
 
 The term α represents the temperature dependence of the water proton resonance frequency and is approximately 0.01 ppm/° C. The change in temperature will cause change in the resonance frequency which in turn leads to change in the phase (Δφ) of the signal: 
 
Δφ=φ−φ ref =(ω−ω ref ) T   E =α( T−T   ref ) T   E   γB   0  
 
 in which T ref  is the reference temperature, normally taken at the start of the procedure, and T E  is the echo time. The temperature can be obtained using:  
       T   =       T   ref     +     Δφ     α   ⁢           ⁢     T   E     ⁢   γ   ⁢           ⁢     B   0               
 
         [0013]     In the past, most temperature mapping has been done in a two dimensional plane with a single slice scan in order to have sufficient spatial and temporal resolution. Even though most of the heat delivery technique causes focused heating at a single spot, the subsequent heat dissipation occurs in the tissues surrounding the heated area and is intrinsically a three dimensional process. The heating pattern in the tissues may or may not be isotropic depending upon tissue composition, blood flow, and diffusion in the tissue. Therefore it is desirable to monitor the heat distribution, i.e. the temperature map in a three dimensional fashion to avoid injury to healthy tissues. However, a regular three dimensional acquisition is time consuming, so the spatial resolution or the temporal resolution will be compromised. In addition, it is not easy to present three dimensional phase information effectively due to effects of phase wraparound and noise.  
       SUMMARY OF THE INVENTION  
       [0014]     Accordingly, what is needed is a method that utilizes an MR imaging system for effectively monitoring the temperature in three dimensions. The present invention provides for such a method of three dimensional temperature monitoring. It has the advantages of being highly efficient in acquisition and being straightforward in presentation. The present invention is based on two general observations. First, in heterogeneous tissues where the hyperthermic procedure is being conducted, the heat dissipation, and therefore, temperature variations, do not change dramatically over a spatial region. Second, temperature change, which manifests itself as the phase change, has different characteristics compared to the anatomic magnitude image.  
         [0015]     In general, methods of heating tissue from a point source, which include laser ablation, focused ultrasound ablation or RF ablation, cause a temperature profile in the tissue wherein the temperature drops gradually from the center while heat is being applied. Therefore the present invention provides a method for acquiring the temperature map of three orthogonal planes nearly simultaneously. Obviously, this will aid the clinician in regulating the hyperthermic process to target the diseased tissue and avoid damaging healthy tissue. The foregoing and other features and advantages of the present invention will be apparent from the description that follows. 
     
    
     BRIEF DESCRIPTION OF THE DRAWINGS  
       [0016]      FIG. 1  is a block diagram of an MRI system used in the present invention.  
         [0017]      FIG. 2  is an electrical block diagram of the transceiver that forms part of the MRI system of  FIG. 1 .  
         [0018]      FIG. 3  is a graphical representation of the preferred pulse sequence used to acquire the phase image data according to the present invention.  
         [0019]      FIG. 4  is a schematic view of the phantom placement of the three orthogonal planes used in experimentation and validation of the present invention.  
         [0020]      FIG. 5  is a pictorial view of the magnitude and phase difference images of the phantoms shown in  FIG. 4 .  
         [0021]      FIG. 6  is a graph showing the temperature change recorded on three different planes.  
     
    
     DETAILED DESCRIPTION  
       [0022]     Referring now to the drawings in detail wherein like numbered elements refer to like elements throughout,  FIG. 1  is a block diagram showing the components of an MRI system. The operation of the system is controlled by a console  100  which includes a keyboard, a control panel  102  and a display. The console  100  communicates through a link  116  with a separate computer system  107  that enables an operator to control the production and display of images on the screen  104 . The computer system  107  includes a number of modules which communicate with each other through a backplane  118 . These include an image processor module  106 , a CPU module  108  and a memory module  113 , known in the art as a frame buffer for storing image data arrays. The computer system  107  is linked to a disk storage  111  and a tape drive  112  through a high speed serial link  115 .  
         [0023]     The system control  122  includes a set of modules connected together by a backplane. These include a CPU module  119  and a pulse generator module  121  which connects to operator console  100  through a serial link  125 . It is through this link  125  that the system control  122  receives commands from the operator which indicated the scan sequence to be performed. The pulse generator module  121  connects to a set of gradient amplifiers  127 , to indicate the timing and shape of the gradient pulses to be produced during the scan. The pulse generator module  121  also receives patient data from a physiological acquisition controller  129  that receives signals from several different sensors connected to the patient, such as ECG signals from electrodes or respiratory signals from bellows. And finally, the pulse generator module  121  connects to a scan room interface circuit  133  that a patient positioning system  134  receives commands to move the patient to the desired position for the scan.  
         [0024]     The gradient waveforms produced by the pulse generator module  121  are applied to a gradient amplifier system  127  comprised of G x , G y  and G z  amplifiers. Each gradient amplifier excites a corresponding gradient coil in an assembly generally designed  139  to produce a magnetic field gradients used for position encoding acquired signals. The gradient coil assembly  139  forms part of a magnet assembly  141  which includes a polarizing magnet  140  and a whole body RF coil  152 . A transceiver module  150  in the system control  122  produces pulses which are amplified by an RF amplifier  151  and coupled to the RF coil  152  by a transmit/receive switch  154 . The resulting signals radiated by the excited nuclei in the patient may be sensed by the same RF coil  152  and coupled through the transmit/receive switch  154  to a preamplifier  153 . The amplified NMR signals are demodulated, filtered, and digitized in the receiver section of the transceiver  150 . The transmit/receive switch  154  is controlled by a signal from the pulse generator module  121  to electrically connect the RF amplifier  151  to the coil  152  during the transmit mode and to connect the preamplifier  153  during the receive mode. The transmit/receive switch also enables a separate RF coil (for example, a head coil or surface coil) to be used in either the transmit or receive mode.  
         [0025]     The NMR signals picked up by the RF coil  152  are digitized by the transceiver module  150  and transferred to a memory module  150  in the system control  122 . When the scan is completed and an entire array of data has been acquired in the memory module  160  in the system control  122 , an array processor  161  operates to Fourier transform the data into an array of image data. This image data is conveyed through the serial link  115  to the computer system  107  where it is stored in the disk memory  111 . In response to commands received from the operator console  100 , this image data may be archived on the tape drive  112 , or it may be further processed by the image process or  106  and conveyed to the operator console  100  and presented on the display  104 .  
         [0026]     Referring particularly to  FIGS. 1 and 2 , the transceiver  150  produces the RF excitation field B 1  through power amplifier  151  at a coil  152 A and receives the resulting signal induced in coil  152 B. As indicated in  FIGS. 1 and 2 , coils  152 A and  152 B may be separate as shown in  FIG. 2  and/or together as shown in  FIG. 1 . The base, or carrier, frequency of the RF excitation field is produced under control of a frequency synthesizer  200  which receives a set of digital signals (CF) from the CPU module  119  and pulse generator module  121 . These digital signals indicate the frequency and phase of the RF carrier signal produced at an output  201 . The commanded RF carrier is applied to a modulator and up converter  202  where its amplitude is modulated in response to a signal R(t) also received from the pulse generator module  121 . The signal R(t) defines the envelope of the RF excitation pulse to be produced and is produced in the module  121  by sequentially reading out a series of stored digital values. These stored digital values may, in turn, be changed from the operator console  100  to enable any desired RF pulse envelope to be produced.  
         [0027]     The magnitude of the RF excitation pulse produced at output  205  is attenuated by an exciter attenuator circuit  206  which receives a digital command, TA, from the backplane  118 . The attenuated RF excitation pulses are applied to the power amplifier  151  that drives the RF coil  152 A.  
         [0028]     Referring still to  FIGS. 1 and 2  the signal produced by the subject is picked up by the receiver coil  152 B and applied through the preamplifier  153  to the input of a receiver attenuator  207 . The receiver attenuator  207  further amplifies the signal by an amount determined by a digital attenuation signal (RA) received from the backplane  118 .  
         [0029]     The received signal is at or around the Larmor frequency, and this high frequency signal is down converted in a two step process by a down converter  208  which first mixes the NMR signal with the carrier signal on line  201  and then mixes the resulting difference signal with the 2.5 MHz reference signal on line  204 . The down converted NMR signal is applied to the input of an analog-to-digital (A/D) converter  209  which samples and digitizes the analog signal and applies it to a digital detector and signal processor  210  which produces 16-bit in-phase (I) values and 16-bit quadrature (Q) values corresponding to the received signal. The resulting stream of digitized I and Q values of the received signal are output through backplane  118  to the memory module  160  where they are employed to reconstruct an image. The reference frequency generator  203  provides a reference phase for received NMR signals.  
         [0030]     The method of the present invention employs the foregoing imaging reconstruction method and apparatus to accomplish MR temperature mapping on the basis of Water Proton Resonance Frequency (PRF) measurement which is well summarized in Quesson, B., Zwart J. A., Moonen, C. T. W.,  Magnetic Resonance temperature imaging for guidance of thermotherapy . J Megn Reson Imaging 2000; 12:525-533. In PRF, the local magnetic field B loc ({right arrow over (r)}) as observed by the spins is a function of the main magnetic field B 0  and the chemical shift o(T({right arrow over (r)})): 
 
 B   loc ( T )=[1 +o ( T )] B   0  
 
 The chemical shift field (in ppm) is the sum of temperature-independent contributions, for example, those originating from B 0  field inhomogeneities, represented by σ 0 , and a temperature-dependent contribution σ T (T): 
 
σ( T )=σ 0 +σ T ( T ) 
 
 The chemical shift field can be calculated from the phase information in RF-spoiled gradient-echo images: 
 
Φ( T )=λσ( T ) T   E   B   0  
 
 where Φ is the image phase, λ is the gyromagnetic ratio of the observed nucleus (42.58×10 6  Hz/T for protons), and T E  is echo time. To measure temperature-dependent changes in chemical shift, the term σ 0 ({right arrow over (r)}) must be eliminated, which is typically accomplished by subtraction of the field distribution measured at a given reference temperature from the field distribution measured at temperature T, leading to:  
         Δ   ⁢           ⁢   T     =       T   -     T   ref       =         Φ   ⁡     (   T   )       -     Φ   ⁡     (     T   ref     )           αγ   ⁢           ⁢     T   E     ⁢     B   0               
 
 where α is the temperature-dependent water chemical shift in ppm/° C. In principle, any gradient-echo method can be used for PRF-based MR thermometry, so long as contributions from simulated echoes can be neglected. RF spoiling of fast gradient echoes is thus necessary when flip angles close to the Ernst angle are used for optimal signal to noise ration (SNR) for short T(r). 
 
         [0035]     The method of the present invention is based on the observation that, even in heterogeneous tissues, heat dissipation does not changed dramatically over a spatial region. Therefore, temperature variation does not change dramatically over a spatial region. Unfortunately, the temperature variation does vary as to direction. The differences in temperature variation in regards to direction may depend on the orientation of the heating device as well as the composition of the tissue being heated.  
         [0036]     However, it is not absolutely necessary to get the temperature distribution of the entire imaging volume. It would instead be more useful to provide a temperature map of three orthogonal planes, as shown in  FIG. 4 . There are three basic ways of accomplishing this. First, one can use regular multi-phase scan in which one plane is acquired after another and the images of three planes are acquired in a cyclic fashion. The drawback of this technique is that the temperature maps represented by three planes are not truly simultaneous because they are not acquired simultaneously. Second, one can also partition the k-space in each imaging plane into small regions and update one region on one plane at a time. In general, k-space is a device for mathematically defining the imaging volume. Each individual point in the image is reconstructed from every point in the k-space representation of the image. For example, if 256×256 in matrix size, then the k-space is also 256×256. The images are reconstructed by combining the segmented k-space data for each plane. This method provides data that is close to simultaneous or at least better than that of the first method. One drawback to this method is that dummy pulses (disdaqs) are needed before the acquisition of each region in order to avoid image artifacts. This increases acquisition time. Lastly, and as provided for in the method of the present invention, every k-space line in one plane is acquired after another k-space line in another plane in the same cyclic fashion described in the first two methods. This interleaved acquisition can provide temperature map of all three orthogonal planes simultaneously. The advantages of this method include reduced imaging time and truly simultaneous acquisition of a temperature map in all three planes.  
         [0037]      FIG. 3  shows an example of the pulse sequence diagram for the method of interleaved 3-plane acquisition. The pulse sequence diagram shows the alternating slice select/phase encode/readout direction between the three axes.  
         [0038]     Software has also been developed to implement the method of the present invention. The image acquisition sequence follows as such: 
        A(1),B(1),C(1),A(2),B(2),C(2) . . . ,A(25), B(25),C(25) . . . ,A(256), B(256),C(256) 
 
 in which P(i) denotes the i&#39;th k-space line in plane P. The reconstruction software then sorts through these k-space lines to form complete data sets for each individual plane before Fast Fourier Transform (FFT). In the event of radial sampling, also called the projection reconstruction method of the k-space, the reconstruction sorts through these k-space lines before regrinding to form complete data sets for each plane. Lastly, the software reconstructs the phase difference images. 
       
 
         [0041]     The inventor carried out a series of phantom experiments to demonstrate the feasibility of this technique. The experiments were performed using a Signa 1.5 scanner with a BRM gradient system. A head coil was used for the image acquisition. The following parameters were used during image acquisition:  
                                               TR    11.9 msec   Flip Angle   30       TE    3.7 msec   Total Imaging Time   9.87 sec       Slice Thickness      3 mm    Field of View   20.0 cm       Receiver Band Width   31.25 kHz       Acquisition Matrix   256 × 256                  
 
         [0042]     A number of small Agarose phantoms were made with T1 shortening Gd-DTPA contrast agent mixed. The contrast agent improves the signal to noise ratio and thus the sensitivity of the MR imaging machine to temperature variation. The phantoms were placed in a holder and located on three orthogonal axes  301 ,  302 ,  303  as shown in  FIG. 4 . First, a set of mask images was taken as the reference for later phase subtraction. The phantom at the center  311  was removed and heated before it was put back in the holder. The temperature of the phantom was raised to about 70° C. During the next 25 minutes images were taken every minute while the center phantom cooled towards room temperature. The other phantoms  310  were used as references for monitoring the system phase drift.  
         [0043]      FIG. 5  shows the phase difference images and the magntitude image of three orthogonal planes  301 ,  302 ,  303  of the heated phantom both at the beginning of the cooling and at the end of the cooling period. The top row shows the magnitude image on three planes  501 ,  502 ,  503 , A1 through A3. The middle row shows the phase difference images of the heated phantom as seen from three corresponding planes. The bottom row  511 ,  512 ,  513  shows the same images  521 ,  522 ,  523  as the middle row after 25 minutes of cooling.  
         [0044]     Beneficially, the cross pattern of phantom images in the magnitude images is almost absent in the phase-different image. These patterns are caused by the interleaved acquisition method, which saturates the spins in the three orthogonal planes throughout the scan. It is an important advantage of the present invention to not have these effects in the phase different image as it would not provide any temperature information for the along these dark bands.  
         [0045]     Analysis of the regions towards the center of the heated phantom was performed in all three planes. Mean values were then calculated. Similar analysis was then performed on each of the reference phantoms to determine the phase drift over the length of the experiment. It was found that a substantial amount of phase drift was experienced over the length of the experiment. The phase drift was then subtracted from the mean values of the temperature change. The results are shown graphically in  FIG. 6 . As is shown, change in the phase angle corresponds to a change in the temperature of the sample. The present invention provides a new and unique method observing the temperature change using an MR imaging system. It further provides for a method for using simultaneous three plane, two dimensional acquisition to represent three-dimensional temperature changes. It further provides for a method that is very efficient and effective in terms of acquisition and presentation. The method of the present invention accomplishes this by producing an image indicative of temperature change in a sample positioned in an MR imaging system wherein the MR imaging system acquires data from a plurality of k-space points. The method includes the steps of first performing an NMR pulse sequence to acquire phase reference images from the sample; wherein the NMR pulse sequence could be an RF-spoiled gradient echo pulse sequences. Second, a reference phase image is constructed from the sample. Third, a second NMR pulse sequence is used to acquire measurement NMR, wherein the NMR pulse could be an RF-spoiled gradient echo pulse sequence. The second NMR pulse sequence could further include the steps of: acquiring a first k-space line from a series of k-space points in a first k-space plane; acquiring a second k-space line from a series of k-space points in a second k-space plane; and acquiring a third k-space line from a series of k-space points in a third k-space plane. The method of the present invention could further include an image acquisition sequence that follows as such: 
    A(1), B(1),C(1),A(2),B(2),C(2) . . . ,A(25),B(25),C(25) . . . ,A(256),B(256),C(256) 
 
 in which P(i) denotes the i&#39;th k-space line in plane P. The fourth step in the method of the present invention includes measuring the signal phase shift. This fourth step could further include measuring the change in the resonance frequency of the water proton. The step wherein the signal phase shift is measured might further include the step of correlating the change in the resonance frequency of the water proton to a change in temperature. The last step of the method of the present invention is to produce a temperature map based on the phase differences. The method of the present invention also provides for periodically updating the reference phase image using measurement NMR data acquired during the scan. The method of the present invention could also include the step of repeating the steps of the image acquisition portion of the present invention so as to provide a plurality of additional temperature maps. 
   
 
         [0048]     Additional advantages and modifications will readily occur to those skilled in the art. Therefore, the invention in its broader aspects is not limited to the specific details disclosed and described herein. Accordingly, various modifications may be made without departing from the spirit or scope of the general inventive concept as defined by the appended claims and their equivalents.  
         [heading-0049]     Parts List:  
         [none]    
       
           100  Console  
           102  Control Panel and Display  
           104  Screen  
           106  Image Processor Module  
           107  Computer System  
           108  CPU Module  
           111  Disk Storage/Memory  
           112  Tape Drive  
           113  Memory Module  
           115  High-speed Serial Link  
           116  Link  
           118  Back Plane  
           119  CPU Module  
           121  Pulse Generator Module  
           122  System Control  
           125  Serial Link  
           127  Gradient Amplifier  
           129  Physiological Acquisition Controller  
           133  Scan Room Interface Circuit  
           134  Patient Positioning System  
           139  Gradient Coil Assembly  
           140  Polarizing Magnet  
           141  Magnet Assembly  
           150  Transceiver Module  
           151  RF Amplifier  
           152  Whole Body RF Coil  
           152   a  Coil  
           152   b  Coil  
           153  Preamplifier  
           154  Transmit/Receive Switch  
           160  Memory module  
           201  RF Carrier Signal Output  
           202  Up Converter  
           204  Reference Signal  
           207  Receiver Attenuator  
           208  Down Converter  
           209  Analog-to-Digital Converter  
           210  Digital Detector and Signal Processor  
           230  Frequency Synthesizer  
           301  First Orthogonal Axis  
           302  Second Orthogonal Axis  
           303  Third Orthogonal Axis  
           310  Reference Phantoms  
           311  Center Phantom  
           501  First Plane Magnitude Image  
           502  Second Plane Magnitude Image  
           503  Third Plane Magnitude Image  
           511  First Plane Phase Difference Image  
           512  Second Plane Phase Difference Image  
           513  Third Plane Phase Difference Image  
           521  First Plane Phase Difference Image  
           522  Second Plane Phase Difference Image  
           523  Third Plane Phase Difference Image