WO1998012964A1

WO1998012964A1 - Apparatus and method for silent magnetic resonance imaging with rotating gradient coil

Info

Publication number: WO1998012964A1
Application number: PCT/US1997/017206
Authority: WO
Inventors: Zang-Hee Cho; William E. Bunney, Jr.
Original assignee: Cho Zang Hee; Bunney William E Jr
Priority date: 1996-09-25
Filing date: 1997-09-25
Publication date: 1998-04-02

Abstract

The present invention describes a new imaging technique using a mechanically rotating DC gradient coil (112), effectively replacing both the phase encoding and readout gradient pulsings, and reducing the induced noise from these coils. The invention comprises a static magnetic field generator (128) for exciting an inspection object (106) disposed in an inspection space (104) with a magnetic field about the (z-) axis, a magnetic resonance signal measurement device (110) for measuring signals from the inspection object (106), a first gradient magnetic field generator (112) for creating a first (x-) gradient magnetic field in the inspection space (104), wherein the first gradient magnetic field is rotatable about a first (z-) axis, and a second gradient magnetic field generator (130) for creating a second (z-) gradient magnetic field in the inspection space (104), wherein the first (x-) gradient magnetic field and the second (z-) magnetic field are substantially orthogonal.

Description

APPARATUS AND METHOD FOR SILENT MAGNETIC RESONANCE IMAGING WITH ROTATING GRADIENT COIL

CROSS REFERENCE TO RELATED APPLICATIONS This application claims the benefit under 35 U.S.C. § 119(e) of U.S. Provisional application serial number 60/027,375, filed September 25, 1996, entitled "APPARATUS AND METHOD FOR SILENT MAGNETIC RESONANCE IMAGING" by Zang-Hee Cho and William E. Bunney, Jr., which application is hereby incorporated by reference.

REFERENCES The following references are incorporated by reference herein, and are referenced throughout by their reference number indicated in [brackets]:

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Reconstruction Methods with Multiple Photon γ-ray Transmission Scanning," Phys. Med. Biol., 19, No. 4, 51 1-522, 1974.

2. Z.H. Cho, "Computerized Tomography," Encyclopedia of Phys. Sci. Technol., 14, 1987.

3. Z.H. Cho, H.S. Kim, H.B. Song, and J. dimming, "Fourier Transform Nuclear

Magnetic Resonance Tomographic Imaging," in Proc. Of IEEE, 78(100), 1152-1173, 1982.

4. Z.H. Cho, D.J. Kim, and Y.K. Kim, "Total Inhomogeneity Correction Including

Chemical Shifts and Susceptibility by View Angle Tilting," in Med. Phys. 15, 7-11, 1988.

5. Z.H. Cho, and E. Wong, "Fringe Field MRI," US Patent No. 5,023,554, Issued in June 11, 1991. 6. Z.H. Cho, and J.H. Yi, "A Novel Type of Surface Gradient Coil (SGC), "Journal of Magn." Res. 94, 471-485, 1991.

7. Z.H. Cho, Y.M. Ro, and T.H. Lim, "NMR Venography Using the Susceptibility Effect by Deoxyhemoglobin," Mag. Reson. Med. 28, 25-38, 1992.

8. K. Kwong, J.W. Belliveau, D.A. Chesler, I.E. Goldberg, R.M. Weisskoff, B.P. Poncelet, D.N. Kennedy, B.E. Hoppel, M.S. Cohen, R. Turner, H.M. Cheng, T.J. Brady, and B.R. Rosen, "Dynamic Magnetic Resonance Imaging of Human Brain Activity During Primary Sensory Stimulation," Proc. Natl. Acad. of Sci. U.S.A. 89, 5675-5679, 1992.

9. S. Ogawa, R.S. Menon, D.W.Tank, S.G. Kim, H. Merkle, J.M. Ellermann, and K. Ugurbil, "Functional Brain Mapping by Blood Oxygenation Level

Dependent Contrast Magnetic Resonance Imaging," Biophy. J. 64, 803-812, 1993.

10. R. Turner, P. Jezzard, H. Wen, K.K. Kwang, D. LeBihan, T. Zeffiro, and R.S. Balaban, "Functional Mapping of the Human Visual Cortex at 4 and 1.5

Tesla Using Deoxydenation Contrast EPI," Mag. Reson. Med. 20, 277-279, 1993.

11. Z.H. Cho, Y.M. Ro, S.H. Park, and S.C. Chung, "NMR Functional Imaging Using a Tailored RF Gradient Echo Sequence: A True Susceptibility

Measurement Technique," Mag. Reson. Med. 35, 1-5, 1996.

12. P. Mansfield, B.L.W. Chapman, R. Bowtell, P. Glover, R. Coxan and P. Harvey, "Active Acoustic Screening Reduction of Noise in Gradient Coils by Lorentz Force Balancing," Mag. Reson. Med. 33, 271-281, 1995. 13. R. Botwell and P. Manfield, "Quiet Transverse Gradient Coils: Lorentz force Balance Designs Using Geometrical Similitude," Mag. Reson. Med. 24, 494- 497, 1995.

14. L.A. Shepp and B.F. Logan, "The Fourier Imaging of a Head Section," IEEE Tran. Nucl. Sci. NS-21, No. 3, 21-43, 1974.

15. T.F. Budinger and G.T. Gullberg, "Three Dimensional Reconstruction in Nuclear Medicine Emission Imaging," IEEE Trans. Nucl. Sci. NS-21,

No. 3, 2-20, 1974.

16. C.B. Ahn, J.H. Kim, and Z.H. Cho, "High Speed Spiral Echo Planar NMR, Imaging~I," IEEE Med. Imag. Vol. MI-5, No. 1, March 1986.

17. Quirk, M.E., Letandre A.J., Ciottone R.A., Lingley J.F.; "Anxiety in Patients Undergoing MR Imaging, Radiology;" 170:464-466, 1989.

18. Brummett R.E., Talbot J.M., Charuhas P.; "Potential Hearing Loss Resulting from MR Imaging;" Radiology; 169:539-540, 1988.

19. Hurwitz R., Lane S.R., Bell R.A., Brant-Zawadzki M.N.; "Acoustic Analysis of Gradient-Coilnoise in MR Imaging;" Radiology; 173:545-548, 1989.

20. Goldman A., Grossman W.E., Friedlander P.C.; "Reduction of Sound Levels with Anti-Noise in MR imaging," Radiology; 173:549-550, 1989

21. McJury M., Stewart R.W., Crawford D., Toma E., "Acoustic Noise Control in High-field MRI," SMRM 1223, 1995. BACKGROUND OF THE INVENTION

1. Field of Invention

The present invention relates generally to medical imaging devices for clinical diagnosis, and in particular to a method and apparatus for silent magnetic resonance imaging.

2. Description of Related Art

Magnetic resonance imaging (MRI) was developed by researchers in the United States and Britain in the early 1970s. MRI is based on the precessional properties of tiny magnetic moments of atomic nuclei. Hydrogen nuclei in tissue absorb stimulating electromagnetic radiation, and subsequent emissions from the nuclei are detected and used to form detailed and high contrast images far superior to those available by X-ray or other methods. Nuclei emissions are created with a strong magnetic field, a radiofrequency, and a "gradient" to form the image of an object placed inside the magnetic field. When MRI is used for diagnostic purposes, atoms in the human body are excited by an external radiofrequency in conjunction with "gradient pulsings." Images are created from the emissions of these excited atoms using a resonance phenomenon and a computer to transform these emissions into a usable image. Of diagnostic techniques available today, the images obtainable with MRI are among the most detailed available, and the application of MRI in the medical community is growing rapidly. MRI allows visualization of tumors and other abnormalities without the hazardous radiation from X-rays. MRI has also been extensively applied to allow scientists to determine how the brain works by studying the physiological functional behavior of the human brain. Today, there are at least 10,000 MRI scanners in operation throughout the world, each costing nearly $2.0 million dollars or more.

However, one of the primary obstacles to more extensive use of MRI techniques is the audible acoustical or sound noise generated by the MRI scanners. Most of this noise is caused by mechanical forces generated by the interaction of the magnetic field with the gradient or gradient pulsing technique employed by MRI scanners. As current pulses are applied to the gradient coils located within the strong magnetic field, Lorentz forces inducing acoustical vibration in the coils. The resulting rapid mechanical motion generates high frequency sound noises in conduction with other materials surrounding or supporting the coils. The sound noise also becomes proportionally louder as the magnetic field strength of the MRI scanner increases.

Gradient pulsing noise has prevented the use of MRI in the diagnosis and treatment for many patients, including the elderly, severely sick, small children, persons with psychiatric disorders, and those suffering from tinnitus, especially where temporally long scans are required. Besides physical discomfort, this problem also has other negative effects. In some areas of research, such as functional MRI [7, 8, 9, 10, 11], sound noise is particularly troublesome because it can stimulate auditory and sensory portions of the body, particularly the temporal lobes of the brain. In some cases, small functional signals generated by the tiny susceptibility changes (such as magnetically sensitive variations due to the iron in the blood or more specifically due to the generation of deoxyhemoglobin) are disturbed by the large sound noise. Gradient pulsing, however, is a basic ingredient of MRI scanning techniques, because, without it, imaging cannot be realized. In the past, the noise problems associated with MRI scanning have been simply overwhelmed by the powerful imaging capabilities of MRI.

There have been a number of attempts to reduce or eliminate gradient pulsing induced sound noise with MRI scanners. Early efforts included the technique of reducing vibrations tightly binding the coils onto a massive or heavy support. However, this technique has resulted in largely disappointing results. More recently, Mansfield et al. [12] and Botwell and Mansfield [13] have discussed a compensated or balanced type of gradient coil to reduce the Lorentz forces generated by the coils. These techniques, though elegant and successful in reducing induced sound noise intensity up to 18 dB, have not been entirely successful, largely because the actual implementation of these techniques is difficult and complex, and demands excessive power in the gradient amplifier. Moreover, the objectionable noise was not completely eliminated. Therefore, there is a need for a simple and gradient power efficient method for reducing the acoustical noise inherent with pulsed gradient techniques used in magnetic resonance imaging. The present invention satisfies that need.

SUMMARY QF THE DISCLOSURE In practice, among the three gradient coils used in MRI scanners, namely the

X, Y, and Z gradient coils, the most dominant and the largest sound-producing gradient coils are the X and Y coils, which are the readout and phase encoding gradient coils.

The present application describes a new imaging technique using a mechanically rotating DC gradient coil, effectively replacing both the phase encoding and readout gradient pulsings, and reducing the induced noise from these coils. One of the disclosed embodiments comprises a static magnetic field generator for exciting an inspection object disposed in an inspection space with a magnetic field about the (z-) axis, a magnetic resonance signal measurement device for measuring signals from the inspection object, a first gradient magnetic field generator for creating a first (x-) gradient magnetic field in the inspection space, wherein the first gradient magnetic field is rotatable about a first (z-) axis, and a second gradient magnetic field generator for creating a second (z-) gradient magnetic field in the inspection space, wherein the first (x-) gradient magnetic field and the second (z-) gradient magnetic field are substantially orthogonal. The present application also describes a method for inspecting a specimen in an inspection space using nuclear magnetic resonance information, which comprises the steps of applying a static magnetic field about a first (z-) axis to the inspection space, applying a substantially constant read (x-) gradient magnetic field to the inspection space, while rotating the read (x-) gradient magnetic field about the first (z-) axis, applying a first radio frequency (RF) pulse to the inspection space while applying a (z-) gradient magnetic field pulse in the selection (z) direction, and measuring the induced magnetic resonance signal from the specimen.

Data obtained with this technique provides a set of projection data which can be used for projection reconstruction [1, 2, 14, 15] or conventional fast Fourier transform (FFT) image reconstruction using interpolation techniques [3].

Experimental results obtained with a 2.0 Telsa whole-body MRI system indicates that virtually all the gradient pulsing sound noises are effectively eliminated by the present invention. Specifically, measured sound noise signals with the present invention are approximately 684 times less than that of obtained with conventional imaging techniques.

A simplified formula for estimating the power attenuation factor F based on the power of the sound noise signals is given by;

E = 101og_]0 ^ [1]

where S_∞nv and S_sM are the sound noise signals of the conventional spin echo sequence and the silent MRI technique of the present invention. Based on equation [1], the actual sound noise level obtained with the new technique is approximately -

( ^<? 28.35 dB -101og₁₀ - ^ = - 101og₁₀ 684 lower than the sound noise of the

conventional technique. This measured remaining noise is also only about 5 dB above the resting noise level (without application of the imaging pulse sequence).

BRIEF DESCRIPTION OF THE DRAWINGS Fig. 1 (a) is an illustration of a side view of one embodiment of the present invention;

Fig. 1(b) is an illustration of a rear or front view of one embodiment of the present invention; Fig. 2(a) is a diagram showing pulse sequences for the silent MRI spin-echo imaging of the present invention, illustrating the absence of the phase encoding gradient and the application of a DC gradient in the readout gradient;

Fig. 2(b) is a diagram showing pulse sequences for conventional spin-echo imaging;

Fig. 3 is a flow chart describing the method steps performed in one embodiment of the present invention;

Fig. 4(a) is an image obtained from computer simulation of the silent MRI technique provided by the present invention; Fig. 4(b) is an experimentally obtained image and sinogram obtained using a

2.0 Telsa whole body scanner with a hand-rotated platform;

Fig. 5(a) is a plot of sound noise amplitude obtained with the silent MRI technique of the present invention;

Fig. 5(b) is a plot of sound noise amplitude obtained from a conventional spin-echo technique;

Fig. 6(a) is a plot of the noise spectra obtained by Fourier transform of the sound signal presented in Fig. 5a;

Fig. 6(b) is a plot of the noise spectra obtained by Fourier transform of the sound signal presented in Fig. 5b; and Fig. 7 is a diagram showing a further embodiment of the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT In the following description of the preferred embodiment, reference is made to the accompanying drawings which form a part hereof, and in which is shown by way of illustration, a specific embodiment in which the invention may be practiced. It is to be understood that other embodiments may be used and changes may be made without departing from the scope of the present invention. 1. Elimination of Phase and Frequency Encoding Gradients

As described above, the phase and frequency encoding gradient coils are the largest contributors to induced acoustical noise in the conventional MRI scanner. To eliminate the phase encoding gradient pulsing, the phase encoding (Y) gradient coil is replaced by a rotating gradient coil (e.g. x- or y- gradient coil) similar to an X-ray CT [1, 14]. As a result, only the X-gradient (readout gradient) is now rotated with DC voltage applied rather than the pulsed gradient as in the case with conventional MRI imaging. The simultaneous application of the rotation of X- gradient and a constant DC voltage on it allows frequency encoding to be achieved and projection data to be obtained at any given time or angle without pulsed gradients. The rotation corresponds to the Y-gradient previously used for phase encoding, and the constant DC gradient corresponds to the read gradient or frequency encoding.

Figs. 1(a) and 1(b) are illustrations of one embodiment of the present invention. This embodiment of the silent MRI apparatus 100 comprises a static magnetic field generator 102, which creates a static magnetic field in inspection space 104 in which a specimen or inspection object 106 is inserted. The static magnetic field generator 102 is similar to those used in other MRI devices. The silent MRI apparatus 100 also comprises a radio frequency (RF) magnetic field generator 108 for exciting the inspection object 106 disposed in the inspection space 104 using the methods described herein. A magnetic resonance signal measurement device 110 measures magnetic resonance signals induced into the specimen by the static magnetic field generator 102, the RF magnetic field generator 108, and the first and second gradient magnetic field generators 112, 128 as described herein and processes these measurements to form images of the specimen 106.

The silent MRI apparatus 100 also comprises a first gradient magnetic field generator 112, which is rotatable about the z axis 114, thereby creating a rotating first (x-) gradient magnetic field in inspection space 104. The first gradient magnetic field generator 112 comprises a number of coils 1 13 disposed to generate the first gradient magnetic field when coupled to a source of electrical energy via line 122. In one embodiment, a substantially constant or DC voltage signal is applied to these coils, thereby creating a substantially constant gradient magnetic field, which is rotated along the Z axis by a flywheel 124 coupled to the first gradient magnetic field generator 112, and driven by drive wheel 126 and drive shaft 127. The outer case 116 of the silent MRI apparatus 110 comprises support members 118 which accept support inserts 120 coupled to the first gradient magnetic field generator 112, to allow it to rotate freely. Rotation of the magnetic field can either be accomplished mechanically, or electrically, using a controllable coil matrix or other means. The illustrated embodiment also comprises a second gradient magnetic field generator 128, which comprises one or more coils 130, which when coupled to a power source by electrical line 130, creates a second (z-) gradient magnetic field in the inspection space 104 that is substantially orthogonal to the gradient magnetic field created by the first gradient magnetic field generator 112. In one embodiment, the second gradient magnetic field generator 128 is substantially stationary relative to the first gradient magnetic field generator 112.

After the X-gradient is rotated 180° with a suitable angular increment, a set of projections are obtained [4, 5, 6]. This set of data is then used to reconstruct the slice image using a projection reconstruction algorithm [14], or other suitable technique.

This method effectively eliminates the two largest gradient pulsing sound noise sources, namely the phase encoding and frequency encoding gradient pulsings. The selection (or Z-) gradient is still used in the same manner as with a conventional spin echo sequence, and will generate acoustical or sound noise similar to that from the reading and phase encoding gradient coils. However, as described above, the selecting gradient coil generates very small acoustic or sound noise.

The pulse sequences employed in one embodiment of the present invention for single-slice spin-echo imaging are further illustrated in Fig. 2(a). As shown, the phase encoding (Y-) gradient and the reading (X-) gradient are replaced by rotating and constant DC gradient coil, respectively. However, the selection (Z-) gradient coil remains and is used in the same manner as in a conventional spin echo sequence. For comparison purposes, a conventional spin echo sequence for all gradients is presented in Fig. 2(b).

Fig. 3 presents a flow chart of the method steps of one embodiment of the present invention. In the illustrated embodiment, a static magnetic field is applied 302 to an inspection space about a first axis. Next, a substantially constant read gradient magnetic field is applied 304 to the inspection space while rotating the read gradient magnetic field about the first axis. Then, a first RF pulse is applied 306 to the inspection space while applying a (z-) gradient magnetic field pulse in the selection (z) direction. This induces a magnetic resonance signal from the specimen 106, which is measured 308, and used to process the measurements to present the desired information to the user. In one embodiment, the first RF pulse is a 90 degree RF pulse, and a 180 degree pulse is also applied to reduce undesirable measurement errors.

2. Mathematical Construct for Silent MRI Apparatus and Method

The mathematical construct for substituting the rotating and DC gradient for the X and Y gradients is now described.

Suppose a slice is selected at z=Zo using conventional spin echo imaging. In this case, the echo signal obtained can be noted as

S(t) = J )M(x,y;z₀) e^ e^' dxdy , n = l, 2, ..., N, [2]

m = 1, 2, ..., M

where, in accordance with the conventional notations used in MR imaging analyses; ΔG_y represents an increment in Y-gradient; T_P represents Y-gradient pulsewidth; G_x represents X-gradient amplitude; m and n represent integer numbers; M_G represents a number of gradient steps;

Δt is the time increment; and γ is the gyromagnetic ratio characteristic of the nuclear species.

Noting that the Y-gradient (mΔG_yT_P) is used for the phase encoding by varying m in discrete steps while X-gradient (G_xnΔt) is used as the frequency encoding by increasing nΔt in a continuous manner, respectively, an equivalent form of the projection data can be expressed as:

or

<«"Λ_fφ oo

S(k_x,) = ∑>Δφ lM(x,y-z₀)e^,2nx'"'^' dx'

1M-0 _o

where x , y' , k_x. , Mφ and Δφ are given by

Y^" cosφ sinφ X

= y. -sinφ cosφ y_ •>

nM k, = y JG,. (t)dl = γG,«Δ(φ = k, (φ) , [4] φ = — , M = 1, 2, ..., 90 (Δφ = 2°) or 180 (Δφ = 1°) steps and Δφ is the M angular increment.

G_x. (t) or G_φ (t) is a combined X and Y gradient amplitude at an angle φ in the coordinate. Note, however, that in Eq. [3], G_x,(t) is a variable usually given by

G .(t) * Gφ(/) = G_;[-nιax cosφ(t) + G,_,max sinφ(t) = /(φ) . [5]

To implement the silent MRI capability of the present invention, k_x( ) is replaced by a constant k (or G), i.e.,

πΔ. where k = γ f Gdt and k_r = ^k + k

Equation [6] represents the echo signal obtainable with the silent MRI pulse sequence of one embodiment of the present invention with a number of angular rotations mΔφ covering 180° and also carried out by a mechanical rotation instead of gradient pulsings given by equation [5]. This technique, although conceptually with respect to projection reconstruction, can be used for image construction as well.

This can be accomplished by appropriately sampling the projection data in k-space, and converting this data into an appropriate coordinate system, such as Cartesian coordinates. 3. Experimental Results

Both computer simulations and physical experiments were performed to verify the operation and performance of the present invention. Imaging experiments were performed using rotating gradient coil applied with a DC current (readout gradient coil) and a phantom. The rotating gradient coil was rotated by had from the side. The phantom is located inside the gradient coil and is isolated from the gradient coil and the rest of its movement. The hardware structure used to perform implement the rotating gradient system and perform the physical is illustrated in Figs. 1(a) and 1(b). The rotating gradient system constructed to implement the silent MRI technique includes a 35 cm diameter gradient coil set (x-, y-, and z- coils) within a hand-rotated 50 cm diameter circular tube.

After a 180° rotation with an angular increment of 1° or 2°, a projection data set of 180° (1° interval) or 90° (2° interval) views is obtained, and using a projection reconstruction algorithm such as that described in [14], this data is used to reconstruct the image.

For image reconstruction, the k-space projection data obtained was first Fourier transformed and then converted to a projection data. Reconstruction was then performed using a conventional projection reconstruction algorithm. Image reconstruction with the silent MRI of the present invention mirrors that of conventional MRI, using suitable interpolation and sampling in k-space, and two dimensional Fourier transform techniques. Other image reconstruction techniques may also be employed, including Fourier (gradient), Hadamard (RF) encoding techniques for multislice imaging, or DANTE ultrafast imaging techniques and its derivatives. Figs. 4(a) and 4(b) show the reconstructed images and their corresponding sinogram obtained by a computer simulation using Bloch equations [3] and experimental data using a 2.0 Telsa whole body MRI system, respectively. The experimentally obtained image shows good agreement with that predicted by simulation, except at the 180° boundaries, where the image is distorted somewhat, a result which is probably due to a small mismatch of the mechanical rotation center. Figs. 5(a) and 5(b) shows the sound noise amplitudes recorded while obtaining images with MRI techniques. Fig. 5(a) shows the performance obtained with the present invention, implementing the silent MRI pulse sequence technique shown in Fig. 2(a). Fig. 5(b) shows the performance obtained with conventional spin-echo pulse techniques, such as those depicted in Fig. 2(b). These measurements were performed with a 16 bit analog to digital converter card coupled to a personal computer and a non-magnetic acoustic transducer, or microphone. These measurements were performed on the KAIS 2.0T MRI system as well as on the GE 1.5T MRI system. Noting that the vertical scale of the plot presented in Fig. 5(a) is increased by a factor of 15, it is apparent that the sound noise amplitude obtained with the silent MRI techniques of the present invention are negligibly small compared to that of the conventional pulse sequence (shown in Fig. 2(b)). Tests were also performed using other conventional imaging techniques including gradient echo, EPI, fast spin echo, inversion recovery, and other sequences used in conventional MRI systems. This allowed noise level comparisons for each method.

Figs. 6(a) and 6(b) depict spectra obtained by Fourier transform of the sound noise signals depicted in Figs. 5(a) and 5(b), respectively. Note that the spectral distribution for the two are substantially identical, except for the strengths. Some of the spectra obtained from the commercial MRI scanners show much wider spectral distributions and suggests that acoustic noise suppression techniques such as the acoustic noise control (ANC) may not be effective, since these techniques apply only for the specific frequency only [21].

Because of a small DC gradient applied continuously during the application of the selection gradient (typically the x-gradient), each selected slice is tilted an amount θ= tan^"'(G_x/G_z) which can be a few degrees from the rotating (Z) axis. When this tilted slice is rotated 180 degrees, the result is a nutation of the selected slices or sum of the nutated slices which will result in a slice that is not fully overlapped, thus resulting in reduced resolution at the periphery. In practice, the image resolution degradation is not significant as long as the tilting angle is not large. However, this problem can be ameliorated by a synchronized motion of the specimen or inspection object 106 so that the tilted angle θ is compensated at each projection data collection. Where the specimen 106 is a patient, this can be accomplished by a movable patient bed. The synchronized rotation of the patient bed is small enough as to be unobjectionable to the patient, and is mechanically simple to implement with the addition of a ball bearing device a few millimeters in diameter. One possible implementation of this embodiment is illustrated in Fig. 7, which shows a patient bed 704 placed within the first gradient magnetic field generator 112. Rotation of the first gradient magnetic field generator 112 is accomplished via a plurality of gradient rotation bearings 702. Patient bed 704 is fixed at point 706 by the imaging plane where the imaging object or specimen 106 is placed. Rotation of the patient bed 704 is performed according to the trajectory indicated in Fig. 7. Only a small synchronous rotation with the first gradient magnetic field generator 112 is required. Accordingly, the specimen 106, or patient is synchronously rotated according a computed timing and magnitude along the same axis as the first gradient magnetic field to cancel or reduce residual measurement errors and distortion.

CONCLUSION In conclusion, the present application discloses a method and apparatus which effectively eliminates sound noise emanating from MRI systems arising from gradient pulsings. By replacing the phase encoding gradient with a mechanical rotation of readout gradient with an application of DC voltage, the two most dominant sources of sound noise, namely the phase encoding and readout gradient pulsing are virtually eliminated, with the remaining noise level almost unnoticeable. Although the sound noise due to Z- or selection gradient pulsing still remains, it Lorentz force is not significant enough to generate a tangible sound noise large enough to disturb children or sick patients.

The method and apparatus of the present invention fulfills an important need, particularly with respect to pediatric, psychiatric, geriatric, and functional MRI. With respect to functional MRI, the subtle BOLD effect signals, long been helplessly disturbed by the sound noises, can now be measured because the interfering sound noises are effectively eliminated.

Existing MRI systems in the field can be adapted to implement the present invention with minor changes, for example, by simple replacement of the existing gradient coil system, and minor software changes. These changes will be unnoticeable to the imaging patients and the entire system noise will be much like the X-ray CT. Because the gradient coil is ordinarily much lighter and a circular cylindrical shape, its mechanical part will be much easier to install and operate than the heavy X-ray tube in the X-ray CT.

Further, the principles of the present invention are easily extended to other areas of imaging, including 3-D volume imaging and other faster imaging systems applications. Acquisition times can be reduced to acceptable levels for these other imaging methods with improvements in the rotating mechanism. Further, other imaging sequences other than spin echo imaging such as spiral scan [16] and echo planar imaging [10, 12] can be implemented. For the fast spin echo, images with much faster rotational speeds of the coil that the conventional rotational speeds (many rotations per second) could lead to an induction of current in the rotating gradient coils. However, this problem can be ameliorated with the use of the phased array gradient coil, for example. Image reconstruction may also be improved using simpler interpolation techniques to covert radially sampled data into Cartesian coordinates, the image can be reconstructed directly by fast Fourier transform (FFT).

Claims

WHAT IS CLAIMED IS:

1. A magnetic resonance (MR) imaging system, comprising: a static magnetic field generator for creating a magnetic field in an inspection space; a radio frequency (RF) magnetic field generator for exciting an inspection object disposed in the inspection space; a first gradient magnetic field generator for creating a first gradient magnetic field in the inspection space, wherein the first gradient magnetic field is rotatable about a first axis; a second gradient magnetic field generator creating a second gradient magnetic field in the inspection space, wherein the first gradient magnetic field and the second gradient magnetic field are non-parallel; and a magnetic resonance signal measurement device for measuring magnetic resonance signals from the inspection object.

2. The apparatus of claim 1 wherein the first gradient magnetic field and the second gradient magnetic field are substantially orthogonal.

3. The apparatus of claim 1 , wherein the second gradient magnetic field is stationary relative to the first gradient magnetic field.

4. The apparatus of claim 1 , wherein the first gradient magnetic field is substantially constant.

5. The apparatus of claim 1 , wherein the first gradient magnetic field generator comprises a coil and a substantially constant signal is applied to the coil.

6. The apparatus of claim 1 , wherein the static magnetic field generator generates a magnetic field about the first axis.

7. The apparatus of claim 1, wherein the first gradient magnetic field is a readout gradient.

8. The apparatus of claim 1 , wherein the second gradient magnetic field is a slice selection gradient.

9. The apparatus of claim 1 , wherein the gradient magnetic field is mechanically rotatable.

10. The apparatus of claim 1 , wherein the gradient magnetic field is electrically rotatable.

11. The method of claim 1 , wherein the inspection object is rotatable along the first axis.

12. A method for inspecting a specimen in an inspection space using nuclear magnetic resonance information, comprising the steps of:

(a) applying a static magnetic field about a first (z-) axis to the inspection space; (a) applying a substantially constant read (x-) gradient magnetic field to the inspection space, while rotating the read (x-) gradient magnetic field about the first axis;

(a) applying a first radio frequency (RF) pulse to the inspection space while applying a (z-) gradient magnetic field pulse in the selection (z) direction; and (b) measuring an induced magnetic resonance signal from the specimen.

13. The method of claim 12, wherein the first RF pulse is a 90 degree RF pulse.

14. The method of claim 13, further comprising the step of applying a 180 degree RF pulse to the inspection space while applying a second gradient magnetic field pulse in the selection (z) direction.

15. A method for inspecting a specimen in an inspection space using nuclear magnetic resonance information, comprising the steps of:

(a) applying a first radio frequency (RF) pulse to the inspection space while applying a (z-) gradient magnetic field pulse in the selection (z) direction; and

(b) measuring an induced magnetic resonance signal from the specimen, wherein step (a) is performed while applying a substantially constant signal to a rotatable magnetic field generator and rotating the magnetic field generator to rotate a read gradient (x) in the inspection space.