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MPHY501Part1Lecture9Ch10.pptx

Diagnostic Imaging Physics

MPHY-501

Essential Physics of Medical Imaging

By, Bushberg, et.al.

Chapters 10

Computed Tomography

During the next two years you will learn the basics of Diagnostic Imaging. Ultimately after this period you will enroll in a Medical Physics Residency Program and finally take the Board examination that, once you pass, will certify you as a medical physicist.

1

History

1st scan took 4.5 minutes

Rotation times are 500 faster and multi receive channels combine to render images 27,000 faster than in the 1970’s/

Pixel depth of 3 bits to 16bits

1980’s Helical scanning (speed increase)

1990’s Multi Channel (speed increase)

From 1980 – 2009 5 million/year – ~75 million/year

Slip Ring

component in all modern CT scanners

concentric metal bands connected to series of gliding contacts.

allows the rotating gantry to have electrical connections to the stationary components.

enables helical (spiral) CT

major reductions in CT scan times

Imaging hardware rotates inside gantry cover,

Table and patient translate through bore of gantry

image is circle inside square 512 × 512 image matrix

Circular FOV over length of patient = volume FOV cylindrical shape.

Most CT scanners: maximum circular FOVof 50 to 70 cm.

Scan length (in z) restrictions

Terms

“Projection” = data collected at a specific angle

Synonyms: “Projection”, “Profile”, “View”, collection of “rays”

“Ray” = individual attenuation measurement, correspond to a line through the object

Line is defined at one end by x-ray source and at the other by a detector.

Parallel Beam Geometry

CT no longer use this acquisition geometry.

Parallel beam geometry remains useful in the science of reconstruction.

A projection is a collection of rays.

Fan Beam Projection

Fan beam projection: a fan of data that converges on a vertex

Using rotate-rotate geometry (shown here)

apex of the fan is the x-ray tube

Individual rays correspond to each detector measurement

fan beam projection

collection of rays in this geometry

Fan Beam & Cone Beam

As modern scanners use more detector arrays,

their width in the z-axis dimension gives rise to a narrow cone beam geometry, as illustrated here.

Cone beam reconstruction algorithms are used on all modern MDCT scanners.

Fan Beam

Most clinical CT scanners have detector arrays that are arranged in an arc relative to the x-ray tube

This arc is efficient !!

By mounting detector modules on a support structure aligned along radius of curvature emanating from the x-ray source, there is very little difference in fluence to the detectors due to the inverse square law.

Very importantly, the primary x-rays strike the detector elements in a nearly normal manner, and so there are effectively no lateral positioning errors due to x-ray beam parallax.

Cone Beam

Some cone beam CT scanners use flat panel detectors

Considered as true cone beam scanners

Cone half angles approaching 10 degrees

Used for:

Image guidance (radiation therapy machines)

Cone beam dental CT

Niche CT applications:

Breast,

Extremity

SPECT/CT

Full Cone Beam Geometry

Cone angle is almost as great as fan angle.

Planar flat panel detector system (2D bank of detectors) has no curvature,

Correction methods needed for different SID (source to dexel)

Inverse square law

Heel effect differences in fluence to each detector elements in 2D array

Parallax occurs in this geometry, cone angle can be appreciable

Example:

30 × 40 cm flat panel detector

SID 90 cm

fan angle = 25 degrees

cone angle = 19 degrees

If,

Then,

Basic Concepts and Definitions

Isocenter is:

Center of rotation of the CT gantry,

in most cases is also the center of the reconstructed CT image

i.e. pixel (256, 256) on the 512 × 512 reconstructed CT image.

The maximum FOV is defined by the physical extent of the curved detector arrays (the fan angle).

By convention;

most CT-dimensions are related to plane of isocenter.

Source-to-isocenter distance is illustrated as A

Source-to-detector distance is labeled B.

Magnification factor is M = B/A.

If (minimum) CT slice thickness is 0.50 mm, actual detector width is larger by M.

Example:

B ≈ 95 cm and A ≈ 50 cm, M = 1.9

Width of detector arrays is M × 0.50 mm = 0.95 mm.

Modern MDCT:

rotate-rotate geometry = 3rd generation

x-ray tube & detector arrays

mount rigidly on same rotating platform

rotate in unison

Fixed Geometry

Allows use of an antiscatter grid in detector array

Grid septa align with dead space between individual detectors thus preserve geometrical detection efficiency

Increasing x-ray beam width, multidetector array CT (40mm, 80 mm, & larger), there is greater need for more aggressive scatter suppression

At least one manufacturer uses a 2D grid on large MDCT system

Torque & Heal Issues Influence Design

Gantry rotation speeds ~ 5 rps (0.20s/rotation)

Angular velocities = enormous g-forces

Plane of anode disk is parallel to plane of gantry rotation to reduce gyroscopic effects & torque to rotating anode if mounted otherwise.

Torque & Heal Issues Influence Design Continued

Anode-cathode axis & heel effect run parallel to the z-axis of the scanner

Eliminates heel effect-induced spectral changes along the fan angle

The angular x-ray output from an x-ray tube can be very wide in the dimension parallel to the anode disk but is quite limited in the anode-cathode dimension

This x-ray tube orientation is necessary given the approximately 60-degree fan beam of current scanners.

Continuous Output X-Ray Source

x-ray beam is not pulsed during the scan (some exceptions are described later).

Detector array sampling time becomes acquisition interval for each CT projection.

Sampling dwell times are between 0.2 and 0.5 ms, meaning that between 1,000 and 3,000 (2,500) projections are acquired per 360-degree rotation for a 0.5 second gantry rotation period [ # of samples = 1/(2*Dwell Time) ].

For a typical CT system with about a 50-cm source-to-isocenter distance, the circle that defines the x-ray tube trajectory is about 3,140 mm (2πr) in circumference.

For a 2-kHz detector sampling rate, x-ray focal spot moves 3,140mm/2,000 ≈ 1.5 mm along the circumference of the circle per sample.

X-ray detectors move in unison with x-ray source, patient is stationary, thus 1.5-mm circumferential displacement of the x-ray source during time it takes to acquire one projection can lead to motion blurring and a loss of spatial resolution.

To compensate, some CT use magnetic steering of electrons between cathode and strike anode.

Beam Steering

(A) x-ray tube & detector rotate around the table

They can travel on the order of 1 to 2 mm per acquisition cycle, leading to potential motion artifacts.

(B) Some manufacturers compensate by steering electron beam between cathode and anode thus rotating focal spot in opposite direction, partially correcting for gantry motion.

Another approach to focal spot steering

Conventional x-ray tube geometry

Electrostatic collimators are used to steer the electron beam to alter its trajectory from cathode to anode.

Alignment of Anode & Cathode

Using N actual detectors arrays (in z), with two focal spot positions (about a half detector width apart), a total of 2N detector channels are formed… remember that the z-direction is the cone beam direction.

As table advances through the gantry, x-ray beam is rapidly shifted along z

Magnetic steering coils shift electron beam trajectory to different locations of the focal spot

Thus oversampling in z-dimension & improving the spatial resolution in z

X-ray CT scanners typically run at 120 kV for generic scanning;

Recently high kV combined with 5 to 10 mm Al filtration leads to “hard” x-ray spectrum. 80- 100- 120- and 140-kV spectra are typical

Effective energy of 80-kV spectrum is about 40 keV, 140-kV is about 60 keV.

The range of effective x-ray energies in typical CT spectra is shown in Figure 10-19, overlaid on the mass attenuation coefficients for soft tissue.

Rayleigh scattering has the lowest interaction probability and the Compton scatter interaction has the highest interaction probability.

For 120- to 140-kV spectra, Compton scattering is 10-fold more likely than photoelectric effect in soft tissue.

For soft tissue, the CT image depicts physical properties for which Compton scattering is most dependent on electron density.

Compton scatter linear attenuation coefficient is proportional to:

Where:

μCompton = linear attenuation coefficient for Compton scattering interaction

ρ = mass density of the tissue in a voxel

N = Avogadro's number (6.023 × 1023)

Z = atomic number

A = atomic mass

For soft tissue, CT images depict physical properties for which:

Compton scattering is most dependent on electron density.

For hydrogen, the Z/A ratio is 1.

Implying hydrogenous tissues (i.e. adipose tissue) would have a higher μCompton

In reality:

lower density of:

adipose (ρ ≈ 0.94 g/cm3)

soft tissue (ρ ≈ 1)

dominate when it comes to the difference between soft tissue and adipose tissues.

Consequently, adipose tissue appears darker (has lower μCompton) than soft tissues such as liver or other organ parenchyma.

Primary Constituents of Soft Tissue

Element
Hydrogen 1
Carbon
Oxygen
Nitrogen

Basis for contrast between bone-, tissue-,

and

contrast-enhanced CT.

Where:

= Hounsfield units at position x,y,z

= average linear attenuation coefficient, voxel of tissue at location (x,y,z)

= linear attenuation coefficient of water

μ(x,y,z) = 0 for air, ratio becomes −1, HU = −1,000 for air.

HU range is defined only at these two points—water & air.

Adipose tissues range −80 to −30

Organ parenchyma runs +30 to +220

Bone +900 to + 2000

Metal > 3000

Iodinated contrast can run to HU = 3,095

Bladder (predominately water) should close to HU = 0

HU of adipose & soft tissues shift slightly with different tube voltages

CT Image Formation

Reconstruction

Converts the raw data to a series of CT images,

Reconstructed as a series of contiguous axial images.

Each individual CT image is 2D, those correspond to 3D cross section

Hence, the term volume element (voxel) is used.

Axial and coronal CT images

One picture element (pixel) indicated.

The 2D pixels in each image correspond to 3D voxels in patient.

Dimensions of in-plane (Δx, Δy)

&

Slice thickness in the z-axis (Δz)

Combine to form volume elements

or

voxel

In MDCT scanners, dimensions of the voxel are approximately equal on thin-slice reconstructions, where Δx = Δy ≈ Δz

When this happens the voxel is termed isotropic

Bow Tie Filter

Body areas scanned by a CT scanner are circular in cross section

Giving rise to an uncorrected x-ray fluence at the detector that is high at the edges and low in the center

A bow tie filter, is used on all whole body CT scanners and is located in the x-ray tube assembly.

This filter is designed to attenuate more toward the periphery of the field making signal levels at detector more homogeneous

Most CT Scans are Either: Head or Torso

Torso is typically broken up into:

Chest

Abdomen

Pelvis

These exams represent over ¾ of all CT procedures in a typical institution.

All of these body parts are either round or approximately round in shape.

The standard adult head is about 17 cm in diameter,

standard adult torso ranges from 24 up to 45 cm or greater,

CT scanners: minimum of two bow tie filters—a head and a body bow tie.

Head bow tie filter is also used in pediatric body imaging

B. Bow tie filter reduces dose to patient & no loss of image quality.

C. A beam shaping filter assembly from a commercial CT scanner is shown. This assembly has three different bow tie filters, as labeled.

(Courtesy J. Hsieh.)

A) The peripheral ray (marked P) passes through a shorter path of tissue to reach the exit point than the central ray (C) on a circular object, and thus the dose at the exit point for ray P is greater than for ray C.

When the source is rotated 360 degrees around the object, this consideration leads to higher dose levels at the periphery

B) larger cylindrical objects no beam shaping filter

C) With an ideal bow tie filter, the attenuation thicknesses of the object are exactly compensated for, producing a nearly homogeneous dose distribution in the patient.

D) Bow tie filter designed for small cylinder is used for a larger cylinder (such as using a head bow tie for a body scan), dose will be too concentrated toward the center of the fan angle, higher doses will result centrally.

Image Noise Variance

Image noise at a given pixel in the CT image is the consequence of the noise from all of the projection data which intersect that pixel.

Noise variance (σ2) at a point is the propagation of noise variance from the individual projections (p1, p2, …, pN).

Simplified mathematical description of noise propagation is given by:

+ +…+

Detector Arrays

MDCT scanners use indirect (scintillating) solid-state detectors. Intensifying screens composed of rare earth crystals (such as Gd2O2S) packed into a binder, which holds the screen together.

To improve the detection efficiency of the scintillator material for CT imaging, the scintillation crystals (Gd2S2O and other materials as well) are sintered to increase physical density and light output.

Sintering: Heating phosphor crystals to just below melting point for long periods.

Densification occurs scintillating powder is converted into a high-density ceramic.

Phosphor is then scored with a saw or laser to create a number of individual detector elements in a detector module, for example, 64 × 64 detector elements.

An opaque filler is pressed into the space between detector elements to reduce optical cross talk between detectors. The entire fabrication includes a photodiode in contact with the ceramic detector

Electronics module has gain channels for each detector in the module and also contains the analog-to-digital converter, which converts the amplified electronic signal to a digital number.

Detector array modules are mounted onto an aluminum frame on the mechanical gantry assembly.

Aluminum frame forms an approximate radius of curvature from the position of the x-ray source. Each detector module can be removed and exchanged on the gantry for rapid repair, when necessary.

CT detector

Individual detector elements coupled to photodiodes layered on substrate electronics.

Spaces between dexel are scored out creating voids, which are filled with an optical filler material to reduce detector cross talk.

More expensive than radiography

Make use of sintered phosphors resulting in high-density ceramic detector arrays.

Increased density improves x-ray detection efficiency.

Detector array are mounted on electronics module includes power supplies for the amplifiers, amplification circuits for the detector module, and analog-to-digital converter systems.

Photograph of a detector module with electronics from a commercial CT scanner, courtesy J. Hsieh.

Beam Width and Slice Thickness

Slice thickness in single detector CT involved the relatively simple manipulation of the x-ray beam width using the collimator

Multiple detector array CT (MDCT) systems result in a divorce between slice thickness and x-ray beam width. The detector array width (T) determines the minimum CT slice width, while the overall x-ray beam thickness (>nT) is determined by the collimator

Cone beam CT (essentially MDCT) have a very large number of detector arrays: 64 to 128, 256, 320 detector channels

For 320 channels, T = 0.5 mm = collimated beam width of 160 mm

MDCT: slice thickness & beam width

With MDCT, slice thickness and x-ray beam width are decoupled.

Slice thickness is determined by the detector configuration

Beam width is determined by the collimator.

Detector thickness is T (measured at isocenter)

Number of detectors is n.

Example:

n = 64 slice (n)

T = 0.625 mm

collimated x-ray beam width is nT = 40 mm.

For contiguous images, the 20-cm section of abdomen discussed in the paragraph above could be acquired in just 5 scans (5 × 40 mm = 200 mm), and a total of 256 CT images, each 0.625 mm thick, would be acquired.

Compared to the single slice CT system, the scan time in this example went from 100s to 5 s, slice thickness went from 2 to 0.625 mm.

These capabilities underscore the advances that MDCT offers—faster scans and thinner slices.

Slice Thickness & Signal to Noise

T = 0.625 mm images are 8 times thinner than T = 5 mm images

T = 0.625 mm images provide more spatial resolution in the z-axis.

8-fold reduction in photons means that they are = 2.8 times noisier.

One can acquire thin T = 0.625 mm images with the same noise levels as the 5-mm images, by increasing the dose by a factor of 8…

Example

If 120 kV and 300 mA & pitch = 1 was used on a 5mm slice before

Increase the mA from 300 to 600 (2×),

Increase the scan time from 0.5 to 1.0 s (2×),

Increase pitch to 0.5 (2×).

2 x 2 x 2 = 8

Heightened concerns about the radiation dose levels in CT indicate that using thicker CT slices for interpretation is a good way to reduce noise while keeping dose levels low.

MDCT Z-Axis Dose Profile

Heel effect is seen.

Beam profile slopes off “penumbra”

Penumbra edge is placed outside the active detector arrays.

X-ray beam not striking active detectors = reduction of geometrical dose efficiency.

Dose efficiency as a function of detector arrays.

As detector arrays become wider, fraction of beam represented by the penumbra is reduced, giving rise to very acceptable geometrical efficiencies for systems with more-numerous arrays.

4 array efficiency = 70%

128 array efficiency = 98%

Modes of CT Acquisition

Scanned Projection Radiograph

Basic Axial/Sequential Acquisition

Helical (Spiral) Acquisition

Scanned Projection Radiograph

Once the patient is on the table and the table is moved into the gantry bore, the technologist performs a preliminary scan called the CT radiograph. Also known as:

Scout view,

topogram,

scanogram,

localizer;

Scanned Projection Radiograph

CT radiograph is acquired with the CT x-ray tube and detector arrays stationary, the patient is translated through the gantry, and a digital radiographic image is generated from this line-scan data.

CT systems can scan anterior-posterior (AP), posterior-anterior (PA), or lateral. It is routine to use one CT radiograph for patient alignment; however, some institutions use the lateral localizer as well to assure patient centering.

The PA CT radiograph is preferred over the AP to reduce breast dose in women and girls.

Scanned Projection Radiograph

CT radiograph is used to set up the CT scan geometry

Place guidelines to define each end of the CT scan.

CT scan parameters are set

These parameters include:

kV

mA

gantry rotation time(s)

type of scan (helical or axial)

direction of scan

pitch

detector configuration

reconstruction kernels(s)

mA modulation parameters

Bounding lines for abdominal pelvis CT scan are shown.

Basic Axial/Sequential Acquisition

Axial (sequential) CT scanning (basic step-and-shoot mode)

x-ray beam is not “on” while the patient is being translated between acquisition cycles.

For 64 (n) detector arrays with width T = 0.625 mm, the table is moved a distance nT between axial acquisitions (about 40 mm in this example)

Table advance between each scan is designed to produce contiguous CT images in z-axis direction.

Table increment result in contiguous images,

x-ray beam is slightly wider than the beam width (nT)

X-ray beam overlap between locations increases radiation dose to patient

Helical (Spiral) Acquisition

Table moves at a constant speed

Gantry rotates around the patient.

Geometry results in the x-ray source forming a helix around the patient,

Advantage of helical scanning is speed—

Eliminating the start/stop motion of the table as in axial CT, there are no inertial constraints to the procedure.

Like threads on a screw, pitch describes the relative advancement of CT table per rotation of the gantry.

Helical (Spiral) Acquisition

The pitch of the helical scan is defined as:

Where:

Ftable = table feed distance per 360-degree rotation of the gantry

nT = nominal collimated beam width

n = number of detector elements

T = thickness of each element

Pitch Considerations

Pitch can range between 0.75 and 1.5;

Pitch = 1.0 corresponds to contiguous axial CT.

Pitch < 1.0 results in overs-canning hence higher dose to patient

Pitch > 1.0 represents under-scanning, resulting in lower dose to patient.

Pitch > 1

Helical Hell

acquired data at beginning and end of a helical scan do not have sufficient angular coverage to reconstruct artifact-free CT images.

Consequently, along the z-direction, helical scans start ½ nT before an axial scan would start, and stop ½ nT after an axial scan would end.

For example,  nT = 20 mm scan length of 200 mm,

axial scan would scan only the desired 200 mm,

helical scan require s220 mm acquisition to reconstruct 200 mm of image data.

In this example, 10% radiation dose to the patient is wasted.

Helical Heaven

Newer helical systems, have adaptive beam collimation

Thus shielding wasted dose at outer edges of helical scan (collimated out)

Dose thus saved as percentage of the overall dose increases for scans of smaller scan length, and for CT scanners with larger collimated beam widths (nT).

Some manufacturers achieve adaptive beam collimation using set of motorized collimators in x-ray tube head assembly

others use a cam-shaped collimator, which rotate to achieve same effect

cone beam acquisition: whole organ imaging

320-detector (160mm wide) x-ray beam at scanner isocenter

Axial CT capable of acquiring entire data set along z, without table translation, due to very large detector array and cone angle.

Other full cone beam systems used in radiation oncology and dental CT use flat panel detectors that extend up to 300 mm in width.

third-generation cardiac scanners

Retrospective gating: Heart is imaged continuously with rapidly rotating CT scanner

Electrocardiogram (ECG) data recorded in synchrony with the CT data acquisition.

Heart: longer in z-dimension than many MDCT scanners can image (nT), procedure requires data acquisition at several table positions to image entire length of heart

If cardiac cycle is not synchrony with gantry rotation speed, after several cardiac cycles, enough data would be acquired over different phases of the cardiac cycle to reconstruct the heart.

Retrospective reconstruction, projection data are acquired over the entire cardiac cycle.

CT images can be synthesized using the ECG gating information over the entire cardiac cycle, and a rendered “beating heart” image can be produced

Interest in cardiac CT imaging is in identifying blockages in the coronary arteries,

For the evaluation of coronary stenoses, a temporal depiction of the heart is not necessary and only the end-diastole images are necessary.

Dynamic information achieved from retrospective cardiac imaging is visually dramatic, it represents a high-dose procedure.

Prospective cardiac gating

ECG triggers x-ray tube so projection data are only acquired during quiescent part of cardiac cycle, end-diastole.

This reduces dose to patient considerably and produces one cardiac CT image that is useful for assessing coronary artery stenosis and other anatomical anomalies in the heart.

Duel Source CT

Developed specifically for cardiac imaging.

Two full x-ray imaging systems—two x-ray tubes and two sets of detector arrays;

Tube B imaging chain has a smaller detector array resulting in a smaller field of view,

90-degree juxtaposition of two imaging chains, a 90-degree rotation of the gantry results in 180 degrees acquisition of projection data

Sufficient for true CT reconstruction at center of FOV.

330-ms for 360-degree gantry rotation, the

90-degree rotation requires 330/4 = 83 ms to acquire data for CT reconstruction (within a fraction of the typical cardiac cycle)

Dual Energy CT

Hounsfield predicted the use of CT for dual energy decomposition methods to separate density from the elemental composition of materials.

DSCT scanner was designed primarily for cardiac imaging; however, the availability of two complete x-ray systems on the same gantry clearly lends itself to the acquisition of a low and high kV data set during the same CT acquisition.

Single source CT systems rapidly switch the kV to achieve dual energy acquisition in a single acquisition. It is more difficult to switch the tube current (mA) with high temporal frequency since the tube current is determined by the x-ray tube filament temperature, which cannot be modulated at kilohertz rates due to thermal inertia of the filament.

Low kV (80 kV) needs higher output levels than the high kV (140 kV) setting. However, at the same mA, the x-ray tube is far more efficient at x-ray production at high energies.

Air kerma at isocenter at same mAs is about 3.3 times higher at the highest kV setting (135 to 140 kV) compared to the lowest kV setting (80 to 90 kV).

Single tube CT scanners can increase temporal pulse width at lower kV setting to increase relative signal at lower tube potentials

Trade-offs for dual energy CT between the dual source and single source approaches.

Dual source system has a smaller Tube B FOV, which limits the FOV of the dual energy reconstruction.

Both x-ray systems are simultaneously on a DSCT system, scatter from one impacts the other requiring correction techniques, which result in more noise.

On the other hand, DSCT systems have two distinct x-ray imaging chains and so additional added filtration can be added to the higher kV source, which increases energy separation and improves the energy subtraction.

A) Higher Z materials are below the line of identity. The Zs of hydrogen, aluminum, and calcium are 1, 13, and 20, respectively. The HU of calcium assumes 10% density. 

B. Water’s effective Z is 7.42. Materials with lower effective Z than water are above the line of identity. Materials with a higher effective Z are below the line of identify.

Z:

carbon = 6

nitrogen = 7

oxygen = 8

Zeff:

Teflon (CF2) = 8.4

polyethylene = 5.4

PMMA = 6.5

Muscle = 7.6

2 approaches: reconstructing dual energy CT images

Dual energy image (raw data) each projection angle can be subjected to logarithmic weighted subtraction techniques

The dual energy subtracted projection data can be used to reconstruct dual energy CT images

Alternatively, low kV and high kV data sets, reconstructed separately. Subtraction techniques or other digital image manipulation can be applied to reconstructed CT images.

In clinical practice, the latter technique is most widely used.

Note that the HU values in CT are already linear with respect to the linear attenuation coefficient, and so no additional logarithms are necessary.

Dual energy image: a linear combination of low- & high-energy CT images.

CT Angiography

CT has higher contrast resolution than traditional angiography,

Requires only venous access as opposed to more invasive arterial access for fluoroscopy-based angiography.

CTA is really just contrast-enhanced CT where the images have been processed to highlight the vascular anatomy using image processing techniques

Maximum intensity projection procedures produce excellent CTA images.

CT Perfusion

Used to evaluate vascular perfusion and other physiological parameters related to blood flow to a specific organ.

Used to evaluate stroke or vasospasm in the brain,

Used for abdominal organ imaging, especially in the oncologic setting.

CT perfusion is a high radiation dose study

CT scanner acquires images of an organ repeatedly in real time to quantify the flow of iodine-based contrast agent through that organ.

Typical head CT perfusion protocols call for the acquisition of about 40 CT images, each using a 1-s acquisition, resulting in about 40 s of scanning in the same area of tissue

CT scan starts prior to the intravenous injection of contrast, scanner produces a number of CT images prior to the arrival of contrast bolus to the organ of interest.

This is necessary because the mathematical models (Patlak, etc.) that are used to compute the parametric maps require an input function (HU versus time), typically the artery supplying blood to the organ of interest.

CT angiography (CTA)

Used with increasing frequency instead of fluoroscopy-based angiography.

For diagnostic procedures, the venous access for contrast agent injection is far less invasive than penetrating the (high-pressure) arterial system for iodine injection.

CTA images of the arteries in the leg and the head are shown, with false color rendering.

Shuttle Mode CT

CT scanner repeatedly images a volume of tissue wider than the detector array (nT).

Table rocks back and forth a prescribed distance, during the temporal acquisition procedure

Allows CT perfusion analysis over a larger section of tissue.

The temporal resolution is lower in shuttle mode, but adequate physiological information can in general be derived.

Image data being acquired during table acceleration and deceleration, implying that reconstruction algorithm needs to accommodate for the changing pitch.

CT perfusion, an extension of CT angiography, relatively long time sequence of repeated CT images is acquired

(A). Typically ~40 full CT data sets are acquired for brain perfusion studies, requiring about 40 seconds.

(B) physiological metrics such as vascular perfusion, time to peak enhancement, and blood volume can be computed using software designed for this purpose.

Dose Modulation

Some body regions in the patient are rounder than other regions. An elliptical cross section of a patient is shown in Figure 10-44A.

With fixed x-ray tube current and hence constant x-ray flux, the detectors will receive more primary x-rays when the gantry is in position “a” compared to gantry position “b”, due to the increased attenuation path in position b.

Noise variance in projection data propagates into reconstructed CT image and is more optimal when the projection data variance levels are nearly matched (Equation 10-3).

Hence, the higher signal levels (higher dose and lower signal variance) in position “a” are essentially wasted when combined with those in position “b” during CT image reconstruction.

When mA is modulated, mA is reduced in position “a” compared to position “b”

mA modulation: lower radiation dose can be used to achieve same image quality.

Dose Modulation

Tube Current Modulation

Also is active along the z-dimension of the patient.

As the scanner images patient's thorax, abdomen, and pelvis, the tube current is modulated to optimize the image quality through entire extent (along z) of the patient scan.

Overall mAs per slice along the z-axis of the patient is shown

Oscillatory nature of the angular mAs modulation is also shown.

(Courtesy SE McKenney.)

Tube current modulation: uses different radiation levels for regions in the body with different overall attenuation, and most tube current modulation schemes result in a more homogeneous distribution of CT noise in the patient images along the z-axis.

Here are mAs values used for the various body regions are indicated.

Tube Current Modulation