Technical Note

Magnetic Resonance Angiography WithSliding Interleaved Projection Reconstruction(SLIPR) Acquisition

Dennis L. Parker, PhD,1* John A. Roberts, BS,1 Andrew L. Alexander, PhD,1

K. Craig Goodrich, BS,2 and Jay Tsuruda, MD1

In this paper, we report on the development of a novelmultiple thin-slab projection-reconstruction acquisitiontechnique. To eliminate the slab boundary artifact, theslabs are highly overlapped and only a small fraction of theprojection view angles are sampled at each slab position.After Fourier transformation in the slice direction, thereare sufficient numbers of projection measurements at eachslice position to obtain very high resolution MR angio-grams. The technique presented has all of the advantagesof multiple overlapping thin slab acquisition (MOTSA) withno evidence of slab boundary artifact. J. Magn. Reson.Imaging 1999;10:569–575. r 1999 Wiley-Liss, Inc.

Index terms: magnetic resonance imaging; magnetic reso-nance angiography; projection reconstruction; angiography

THE MULTIPLE OVERLAPPING thin slab acquisitiontechnique (MOTSA) has been shown to combine many ofthe advantages of multiple slice two-dimensional (2D)and single slab three-dimensional (3D) MR angiography(1–6). However, because of the imperfect slab excitationprofile and because of the variation in flow saturationacross the slab, MOTSA produces a distinct artifactknown as the slab boundary artifact (SBA) or, morecommonly, the venetian blind artifact. This artifact canarise from in-slab saturation of the flow, signal varia-tions at the edge slices from imperfect slab selectionprofiles, and subject motion during the acquisition ofadjacent slabs.

The slab boundary artifact can be reduced with theuse of optimized imaging parameters (7). Attempts toeliminate the slab boundary artifact have included theuse of an excitation radio frequency (RF) pulse with aramped or tilted profile (8–11) and the development ofalternative types of spatial encoding, such as the use ofquadratic RF excitation (12,13).

Recently, Liu and Rutt developed a thin slab acquisi-tion technique that greatly reduces the slab boundary

artifact. The technique was originally referred to assliding-slab Interleaved MOTSA (SI-MOTSA) (14) andhas more recently been named SLINKY (Sliding-slabInterleaved ky) (15). In SLINKY, the position of the slabin the slice encoding direction changes as a function ofthe ky phase encoding index. After Fourier transforma-tion in the slice (kz) direction and proper data sorting togroup all measurements from each slice position, acomplete set of kx, ky measurements are obtained forimage reconstruction at all central slice positions. Forslices at the edge of the imaged volume, there aretypically insufficient numbers of measurements in ky toperform the reconstruction.

SLINKY works reasonably well in the elimination ofthe slab boundary artifact. However, the techniqueintroduces phase discontinuities and amplitude varia-tions into the k-space data that can result in ghostingartifacts. To prevent the most serious of these artifacts,the technique requires that the phase across the slabprofile be accurately corrected. One approach for phasecorrection is the collection of a navigator echo andsubtraction of the measured phase for each slice posi-tion within each slab. Even with this first-order phasecorrection, the uncorrected amplitude variations andany residual phase variations in a SLINKY study mayresult in ghosting artifacts. Although the ghosts are oflow amplitude and generally do not interfere with MIPimages, they might cause misinterpretation when sourceimages are viewed.

In this paper, we consider a new sliding interleavedmultislab acquisition technique in which the in-planescanning is performed with projection-reconstructionacquisition. This sliding interleaved projection recon-struction (SLIPR) acquisition builds on some excellentproperties of projection-reconstruction (PR) techniques(16). First, spatial resolution depends primarily on theresolution obtained during signal readout. Second. theartifact caused by a limited number of views is typicallya set of streaks that occur at a distance from sharp edgetransitions. Third, all projections pass through thecenter of kx/ky space, allowing a direct phase measure-ment and correction of each projection. Because ofthese properties, SLIPR can result in excellent imagequality with no visible slab boundary artifacts. Further,

1Department of Radiology, University of Utah, Salt Lake City, Utah,84132.2Department of Radiology, LDS Hospital, Salt Lake City, Utah 84143.*Address reprint requests to: D.L.P., Ph.D., Medical Imaging ResearchLaboratory, Department of Radiology, University of Utah, 729 ArapeenDrive, Salt Lake City, Utah, 84108.Received November 30, 1998; Accepted May 19, 1999.

JOURNAL OF MAGNETIC RESONANCE IMAGING 10:569–575 (1999)

r 1999 Wiley-Liss, Inc. 569

any residual artifacts are generally small amplitudestreaks that occur far from vessels of interest.

We first describe the concept of sliding interleavedprojection imaging. We then describe the reconstructiontechnique and provide examples of the application.

THEORY

In our implementation of multiple slab projection recon-struction imaging, a slice-selective RF pulse is used toexcite the magnetization in a thin slab. A small numberof equally spaced projection angles are selected for datameasurements at each slab position. The transversegradients (x,y) are constant during signal readout toobtain projection measurements in each of the selectedprojection directions. The readout gradient is rotatedbetween the gradient axes transverse to the slice selec-tion direction to obtain measurements at each anglethrough the kx/ky plane as shown. To minimize the echotime (TE), the readout gradient is applied in such amanner that the echo is asymmetric. Fourier phaseencoding, performed in the time between excitation andreadout, is used for spatial encoding in the slice selec-tion direction, and a full set of slice encoded measure-ments are obtained for each projection angle.

As is the case with SLINKY, data measurements inSLIPR are obtained from each possible slice position.Assuming that there are m slices in each slab and n totalviews in a single slice, the first slab position is imagedusing full slice selective phase encoding for all of n/mequally spaced projection measurements. The slab isthen shifted by one slice position in the slice encodingdirection, and the process is repeated for a set of equallyspaced angles that are interleaved with the first set ofangles. The slab shift and interleaved projection acquisi-tion are repeated until the entire imaged volume iscovered. The order of projection angle interleaving isarbitrary, but must be chosen so that the full set ofprojection angles are sampled over the thickness of asingle slab. If n is not evenly divisible by m, the numberof projections acquired for any slab position will dependdirectly on the relative starting projection angle.

Two schematics of data acquisition interleaving areshown in Fig. 1. In the figures, each line segmentrepresents the set of slice encoded projection measure-ments obtained in a single projection direction from asingle slab position. After transformation of the measure-ments in the slice encode direction, the line segmentsrepresent projection measurements for each slice posi-tion along the line segment. Fig. 1a and b illustrate twoways in which the measurements from multiple slabpositions can be interleaved to yield complete projectioncoverage for all the central slices. In both cases, a set ofuniformly spaced projections are obtained for each slabposition. In Fig. 1a the view angle is incremented foreach slab position increment. In this manner a completeset of interleaved projections are obtained for each slicein the center of the imaged volume. In Fig. 1b, the viewangles are chosen to give reasonably uniform angularcoverage for the slices at the edges of the imagedvolume. These edge slices then give reasonable images

of the high contrast blood vessels, with very little artifact(16,17). The artifact diminishes as the number of inter-leaved projections increases.

After transformation in the slice direction, and selec-tion of all measurements for each given slice position,image reconstruction is accomplished by first locatingthe center of each echo, subtracting the phase at theecho center from the entire projection, and then regrid-ding the measurements into the appropriate locationsin a 2D array. A 2D Fourier transform is used for finalimage reconstruction. The sharpness of the recon-structed image depends upon the homogeneity of themain magnetic field and how accurately the field isknown. Error in the assumed field yields the sameblurring that occurs in computed tomography whenthere is an error in the location of the center of rotation.In MR projection imaging, the center of rotation de-pends directly on the homogeneity of the magnetic field,the center of the imaging gradients and the receiverfrequency used. It is possible to adjust the center ofrotation as a function of position in the image in order toobtain sharp image detail throughout the image volume(18).

SLIPR has the same relative noise efficiency as otherPR techniques. The relative noise variance can be deter-mined by comparing the Fourier transform reconstruc-tion with that of projection reconstruction. The follow-ing is a simplified description of a result publishedpreviously (19).

Data acquisition sampling for conventional Fourierimaging can be written:

s( j8Dkx, k8Dky) 5 e m(x, y)e2pi(xj8Dkx1yk8Dky)dx dy (1)

The Fourier reconstruction algorithm is then:

mf ( j, k) 5 oj8k8

sj8k8e22pijj8/Nxe22pikk8/NyDkxDky (2)

Projection-reconstruction sampling can be written:

s( j8Dkr, k8Du) 5 em(x=) e2pi(x=

·j8Dkq

r(k8Du)) dx dy (3)

And the corresponding projection reconstruction algo-rithm is:

mp(r, f) 5 oj8k8

sj8k8 e22pirj8Dkq

rcos(k8Du2w) 0 j 8 0Dkr2Dku (4)

The variances can be predicted from the above equa-tions, recognizing that Dkx 5 1/Nx, Dky 5 1/Ny, Dku 5

p/Nu,and Dkr 5 1/Nr. For equal imaging time andresolution, we let Nx 5 Nr and Ny 5 Nu. The noisevariances obtained using the Taylor series approxima-tion become:

s2 (mf ) 5 s2/NxNy (5)

s2 (mp) 5p2

12

s2

NrNu

(6)

These relations indicate that PR techniques shouldhave a slightly lower noise variance than Fourier tech-

570 Parker et al.

niques. This reduced noise variance is probably due tothe fact that the PR technique does not include thecorners of k-space. This results in a slight increase inthe effective voxel volume with the conversion of theconventional Sinc point spread function (PSF) of rectilin-ear Fourier imaging into the circularly symmetric JincPSF (20):

Jinc(r) 5 aJ1(2par)

r(7)

where a is the radius of measurements in k-space.This signal-to-noise ratio (SNR) analysis has ignored

the fact that there may be signal variation as a functionof the position of measurements within a slab. For eachreconstructed SLIPR slice, measurements are obtainedfrom all possible positions over the slab profile, while forMOTSA the signal intensity depends upon the sliceposition within the slab. Thus, for MOTSA, the expectedSNR will depend upon the position of the slice within theslab. For SLIPR the SNR will depend less on slabposition, but will be reduced to the extent that measure-ments from low intensity edge positions are included inthe reconstruction.

METHODS

Human Volunteers

Informed consent was obtained from two normal hu-man volunteers used in this study.

Data Acquisition, MOTSA

For MOTSA, four slabs of 16 slices/slab were obtainedwith slab spacing equivalent to 10 slice thicknesses,such that six edge slices of each 16 slice slab overlappedwith the six edge slices of the neighboring slab. In eachslab, two slices were blanked (discarded) off either end,and the remaining two overlapping slices were com-bined, taking the maximum pixel value for each corre-sponding slice. In this manner, a total of 42 slices werereconstructed. Imaging parameters were field of view(FOV) 200 3 200 mm, 1-mm slice thickness, 152readout samples (0.6 asymmetry 3 256), 256 y phaseencodings, and TR/TE/u 5 33/2.7/25°. For the firstcarotid study (Fig. 2), images were obtained using a

Figure 1. Diagram of sliding interleaved projection-reconstruc-tion (SLIPR) acquisition. (a) Linear data acquisition order. Theprojection angles are incremented by one unit for each sliceincrement in the slab position. (b) Distributed data acquisitionorder. The projection angles are ordered to give a more uniformdistribution of view angles during acquisition for each sliceincrement in the slab position.

Figure 2. Example MIP images from a carotid study. Left is a4-slab MOTSA study, with 16 slices/slab, a spacing of 10 slicesbetween slabs, two slices discarded off each slab end and theremaining two overlapping slices combined, taking the maxi-mum pixel value for each corresponding slice. In this manner, atotal of 42 slices were reconstructed. On the right is an imagefrom a SLIPR study obtained with repetition time, flip angle,and imaging time identical to the MOTSA study. The SLIPRslab spacing of 12 slices between slab positions is slightlylarger than that of the MOTSA study and no overlap processingis used. Images for all slice positions, including end slices, arereconstructed, resulting in a total of 64 reconstructed images.Imaging parameters are given in the text.

MRA with Sliding Interleaved Projection Reconstruction Acquisition 571

truncated Sinc RF excitation pulse identical to thatused for the SLIPR acquisitions. For the second carotidstudy (Fig. 3), the MOTSA images were obtained using aramped RF excitation to reduce the apparent slabboundary artifact.

Data Acquisition, SLIPR

Our version of the SLIPR technique was implementedusing simple modifications to a conventional multislab3D time-of-flight MRA pulse sequence. For the cervicalcarotid images shown, each fully sampled slice has 256view angles, projections were performed over a full 360°and 12 of the 16 slices excited are used for imagereconstruction. Thus, at each slab position, measure-ments are obtained from projection angles that arespecified by 12 3 360°/256 5 16.875° increments. Theview angles in subsequent slab positions are interleavedto avoid multiple sampling within any given slice. De-pending on the starting view angle, some slab positionssample 21 projections and some sample 22. All 16 sliceencodings are sampled for each projection angle. Theorder of projection angle interleaving is arbitrary, butmust be chosen so that the full set of projection anglesare sampled over the thickness of a single slab. Auniformly distributed interleaving, such as that shownin Fig. 1b was used in this study. After transformationin the slice encoding direction, the measurements aresorted by slice position and standard projection recon-struction techniques can be used.

Other imaging parameters for the SLIPR carotid acqui-sitions are: FOV 200 3 200 mm, 1-mm slice thickness,256 projection angles for each slice position, 16 slices/slab. For the study of Fig. 2, 192 readout samples (0.75asymmetry 3256), and TR/TE/u 5 33/6.2/25° wasused. For the study of Fig. 3, 132 readout samples (0.53asymmetry 3 256), TR/TE/u 5 33/2.2/25° was used.For the images in both studies a truncated Sinc RFexcitation pulse was used, the SLIPR slabs were spacedby 12 slice thicknesses, and the outer pair of overlappedmeasurements were discarded for each slab. Imagesfrom all slice positions, including the edge slices, werereconstructed, yielding a total of 64 images.

We have also performed one intracranial SLIPR studywith slightly different acquisition parameters: FOV 220 3220 mm, 1-mm slice thickness, 192 projection anglesfor each slice position, 16 slices/slab, 304 readoutsamples (0.6 asymmetry 3 512), and TR/TE/u 5 39/5.8/25°.

Image Reconstruction, SLIPR

After initial Fourier transformation in the slice-encodedirection, the results obtained are sets of projectionmeasurements as a function of slice position. By combin-ing measurements for the same slice position frommultiple slab positions, a full set of projections can beobtained. For the slices at either end of the imagedvolume, the full set consists of only the projectionsobtained for a single slab position. For internal slicepositions, there are many measurements from multipleangles. To remove phase errors due to the slab selectionprofile and the multiple positions of the slab select pulse

that are used for each slice position, a phase correctionis performed for each set of projection measurementsbased upon the phase measured at the center of thekx/ky-plane. This correction requires that the center ofk-space be accurately determined. Once the phase iscorrected the measurements are regridded into a 2Dk-space array, corrected for the number of measure-ments at each position, and reconstruction is per-formed with a standard 2D Fourier transform. Forsimplicity we have used a very simple nearest neighborregridding algorithm. This is equivalent to convolutionwith a Rect function in k-space or the multiplication bya Sinc function in the reconstructed image space.Because all of the vessels imaged in this study are farfrom the edge of the FOV, there is minimal intensityvariation apparent in the final image. It will be a simplematter to incorporate other regridding algorithms suchas that of Jackson et al. (21).

RESULTS

A comparison of MOTSA and SLIPR images are shown inFig. 2. Both studies were acquired with the samerepetition time (TR), flip angle and the same number ofkz phase encodings. Because the number of projectionangles acquired in the SLIPR study is the same as thetotal number of ky phase encodings in the MOTSAstudy, both studies took exactly the same imaging time.The slab boundary artifact is clearly visible in theMOTSA study, although it is still possible to see thevessel anatomy. There is no slab boundary artifactvisible in these SLIPR images.

In our experience, the results of MOTSA in the carotidcan be improved by using a ramped RF excitation pulse.The images in Fig. 3a were obtained with identicalimaging parameters to those of Fig. 2 except that aramped RF excitation pulse was used for the MOTSAstudy. This MOTSA study demonstrates less slab bound-ary artifact than that of Fig. 2. In Fig. 3b we display six ofthe SLIPR slices from the upper edge of the imagedvolume. All are displayed with the same intensity scal-ing and show that there is little decrease in intensity asthe number of projection angles included in the recon-struction decreases. In fact, the decrease in intensity islikely due primarily to the RF excitation profile, whichfalls off near the edge of each slab. Although some viewaliasing artifact is visible, this artifact is of significantlylower intensity than the arteries themselves.

Axial, sagittal, and coronal MIP images from an intra-cranial study are shown in Fig. 4. To reduce clutter inthe sagittal and coronal reprojections, the images weresubregioned to show only the middle and anteriorcerebral circulations. There is no visible slab boundaryartifact.

DISCUSSION

In this paper, we have presented the novel concept ofsliding interleaved projection reconstruction (SLIPR)acquisition. We have shown that the SLIPR techniquemaintains the advantages of multiple overlapping thinslab acquisition, but yields very little slab boundary

572 Parker et al.

Figure 3. (a) Example MIP images from a second carotid study on the same human volunteer as Fig. 2. A ramped RF excitationpulse was used for the MOTSA study to reduce the slab boundary artifact in the arteries. Imaging parameters given in the text. (b)Cross-sectional images from the top of the reconstructed SLIPR volume of part a. The slice number (out of 63 reconstructed slices)and the number of projection angles (out of 256) is indicated on each image. (c) Measurements of SNR along the length of theinternal carotid artery for MOTSA (thin line) and SLIPR (thick line).

MRA with Sliding Interleaved Projection Reconstruction Acquisition 573

artifact. Because projection reconstruction techniquescan reconstruct images from a very small numbers ofviews, the end slices in the SLIPR technique can be usedto generate useful images. In general, this is not thecase with SLINKY. In SLIPR the artifacts in the endslices are minimal, and in our implementation, theintensity of these slices decreases faster than the arti-fact amplitude. Because less overlap is required andbecause all slices are reconstructed in SLIPR the FOV inthe slice encoding direction is significantly longer thanthe FOV obtained from the corresponding MOTSA study.

In SLIPR, every projection measurement passesthrough the center of the kx/ky plane. Therefore, thecenter value can be used as a navigator to remove anyphase difference resulting from the phase variationacross the RF pulse profile, the phase difference causedby the transmit frequency difference between slab posi-tions, or phase differences caused by any other factor.

This study represents only a very preliminary attemptto investigate the properties of sliding interleaved projec-tion-reconstruction acquisition. In general, the moder-ate resolution SLIPR carotid artery images obtained todate have been comparable to the corresponding MOTSAimages with a significant decrease in the slab boundaryartifact. The higher resolution images obtained havebeen slightly noisier and less sharp than the correspond-

ing Fourier imaging counterpart images. Neither theSLIPR technique, as implemented, nor the reconstruc-tion algorithm used are optimal. Care was taken inthese studies to determine the local resonant frequency.Errors in the resonant frequency determination corre-spond to shifts in the center of rotation. Such errorstranslate into blurring, which, if large enough, caneliminate small vessel detail. When the main field is nothomogeneous, the blurring will become spatially depen-dent. The blurring can be greatly reduced by increasedsampling bandwidth, at the expense of increased imagenoise.

CONCLUSIONS

In summary, we have presented the novel concept ofSLIPR acquisition. The results of our imaging experi-ments demonstrate the significant potential of SLIPR forvascular imaging. The technique retains the advantagesof MOTSA while at the same time allowing an inter-leaved acquisition that greatly reduces any slab bound-ary artifact. A number of experiments remain to beperformed to develop an understanding of the tech-nique’s full potential and to address limitations associ-ated with MRI projection-reconstruction techniques,

Figure 4. Example images from an intracranial study. See methods for parameters used.

574 Parker et al.

such as local image blurring in the presence of fieldinhomogeneity.

ACKNOWLEDGMENTS

We acknowledge the assistance of Henry Buswell, BrianChapman, and others in the Medical Imaging ResearchLaboratory and Dr. Duane D. Blatter of LDS Hospital.This work is supported in part by NIH grants R01HL48223 and R01 HL57990.

REEFERENCES1. Marchal G, Bosmans H, Van Fraeyenhoven L, Wilms G, Van Hecke

P, Plets C, Baert AL. Intracranial vascular lesions: optimization andclinical evaluation of three-dimensional time-of-flight MR angiogra-phy. Radiology (QSH), 1990;175(2):443–448.

2. Parker DL, Yuan C, Blatter DD. MR angiography by multiple thinslab 3D acquisition. Magn Reson Med 1991;17:434–451.

3. Blatter DD, Parker DL and Robison RO. Intracranial MR Angiogra-phy by Multiple Thin Slab 3D Acquisition: 1. Quantitative analysisof vessel visibility. Radiology 1991;179:805–811.

4. Blatter DD, Parker DL, Hou P, Bahr AL, Robison RO, Johles SA,Boyer RS, Schwartz R. Cerebral MR Angiography with multiple thinslab acquisition: Part II. Earlier clinical experience. Radiology1992;183:379–389.

5. Davis WL, Warnock SH, Harnsberger HR, Parker DL, Chen CX.Intracranial MRA: single volume vs. multiple thin slab 3D time-of-flight acquisition. J Comp Assist Tomog 1993;17(1):15–21.

6. Blatter DD, Bahr AL, Parker DL, Robison RO, Kimball JA, Perry DM,Horn S. Cervical carotid MR angiography with multiple overlappingthin-slab acquisition: Comparison with conventional angiography.Am J Radiol 1993;161:1269–1277.

7. Robison RO, Blatter DD, Parker DL, Barney WW, Perry DM. Cervicalcarotid MR angiography with multiple overlapping thin slab acquisi-tion (MOTSA). Part 1: optimization of imaging parameters. J MagnReson Imag 1994;4:529–535.

8. Tkach JA, Masaryk TJ, Ruggieri PM, Ross JS, Dillinger J, Purdy D,Laub G. Use of tilted optimized nonsaturating excitation (TONE) RFpulses and MTC to improve the quality of MR angiograms of the

carotid bifurcation. Proceedings of the Society of Magnetic Reso-nance in Medicine, 1992; p 3905.

9. Nagele T, Klose U, Grodd W, Petersen D, Tintera J. The effects oflinearly increasing flip angles on 3D inflow MR angiography. MagnReson Med 1994;31:561–588.

10. Dagirmanjian A, Ross JS, Obuchowski N, Lewin JS, Tkach JA,Ruggieri PM, Masaryk TJ. High resolution, magnetization transfersaturation, variable flip angle, time-of-flight MRA in the detection ofintracranial vascular stenoses. J Comput Assist Tomogr 1995;19(5):700–706

11. Goodrich KC, Blatter DD, Parker DL, Du YP, Meyer KJ and Bern-stein MA. A quantitative study of ramped radio frequency, magneti-zation transfer, and slab thickness in three-dimensional time-of-flight magnetic resonance angiography in a patient population.Invest Radio 1996;31(6):323–332.

12. Pipe, JG. Spatial encoding and reconstruction in MRI with qua-dratic phase profiles. Magn Reson Med 1995;33(1):24–33.

13. Pipe, JG. Analysis of localized quadratic encoding and reconstruc-tion. Magn Reson Med 1996;36(1):137–146.

14. Liu K, Rutt BK. Sliding interleaved ky (SLINKY) acquisition: A novel3-D MRA technique with suppressed slab boundary artifact. JMagn Reson Imag 1998;8:905–911.

15. Liu K, Lee DH, Rutt BK. Systematic assessment and evaluation ofsliding-interleaved ky (SLINKY) acquisition for 3D MRA. J MagnReson Imag 1998;8:912–923.

16. Peters D, Mistretta C, Korosec F, Holden F, Kelcz F, Wedding K,Grist T. Using projection reconstruction with a limited number ofprojections to increase image resolution or acquisition speed.Proceedings ISMRM Sixth Scientific. Session, Sydney, 1998; p 182.

17. Peters DC, Grist TM, Korosec FR, Holden JE, Block WF, WeddingKL, Carroll TJ, Mistretta CA. Undersampled projection reconstruc-tion applied to MR angiography. Submitted to Magn Reson Med.

18. Noll DC, Pauly JM, Meyer CH, Nishimura DG, Macovski A.Deblurring for non-2D Fourier transform magnetic resonance imag-ing. Magn Reson Med 1992;25:319–333.

19. Callaghan PT. Principles of nuclear magnetic resonance micros-copy, Oxford Science Publications. Oxford: Clarendon Press; 1991.p 184.

20. Bracewell RN. The Fourier transform and its applications. NewYork: McGraw-Hill; 1986. p 249.

21. Jackson JI, Meyer CH, Nishimura DG, Macovski A. Selection of aconvolution function for Fourier inversion using gridding. IEEETrans. Med. Imaging 1991;10:473–478.

MRA with Sliding Interleaved Projection Reconstruction Acquisition 575