new ultrasound ppt

7/28/2019 New ultrasound ppt

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Ultrasound Imaging Ultrasonic waves Interaction with tissue

Tissue properties

Production and detection of ultrasound

Fields of ultrasound transducers Image formation

Resolution in ultrasound images

Electronic focusing

Real time imaging

Doppler ultrasound

Demonstration of ultrasoundtechnology, meet in IAHS lobby,Thursday 5:45 pm.

Download overheads at

http://www.jccresearch.ca/people/personalpage.asp?person=Patterson

Further reading (available electronically at McMaster)

PNT Wells, Ultrasonic imaging of the human body, Reports

on Progress in Physics62

671-722 (1999).


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The Ultrasound Spectrum

0.015 mm

0.1 MHZ

1.0 MHz

10 MHz

100 MHz

1 GHz

0.02 MHz (20 kHz, hearing threshold)

15 mm

1.5 mm

0.15 mm

1.5 mm

FREQUENCY

WAVE-

LENGTH

IN WATER

l f = 1500 ms -1In water

Medical imaging

Ultrasound microscopy

Therapeutic applications


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Description of Ultrasonic Waves

In general, both compressional and shear waves can

propagate in a solid, but soft tissue will support only

compressional waves.

While there are minor differences in the velocity of

ultrasound in soft tissues, the value 1540 meters persecond is assumed in image reconstruction. The variation

of velocity with frequency ( i.e. dispersion) can also be

ignored. A round trip of 1 cm takes 13 microseconds in

tissue.

The product of wavelength and frequency is equal to thewave velocity, so in soft tissue the wavelength of 10 MHz

ultrasound is 0.15 mm. The wavelength plays an

important part in determining the ultimate resolution that

can be achieved with an imaging system.

The energy transported by the wave per unit time perunit area is referred to as the intensity. The usual units

are Watts per square centimeter. The average intensity in

diagnostic applications is below 100 mW cm -2, but the

peak intensity can be much higher.


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Propagation of Ultrasonic Waves

Absorption

This occurs even when the wave propagates in a homogeneousmedium such as a tank of water. For simple fluids, the main

mechanism is viscous loss. For complex media, such as tissue.

relaxation processes in which acoustic energy is coupled to

changes in molecular conformation are important. Both

mechanisms are dependent on frequency - viscous losses vary

as f2 and relaxation losses somewhere between linear and

quadratic dependence.

Scattering

Scattering of the ultrasonic wave does not take place unless

the wave encounters a change in acoustic impedance. For ourpurposes, the acoustic impedance is Z = c where is the

density and c is the speed of sound. For water Z is 1.48 X 106

kg m-2 s -1 and most soft tissues are within a ten per cent of this

value. Note that the acoustic impedance of compact bone (e.g.

the skull) is about five times higher and that of air is lower by a

factor of 3,700.

The physical process of scattering depends in a complex way

on the size of the inhomogeneity and its acoustic impedance

relative to the surrounding medium.In general, the fraction of

incident energy scattered by the inhomogeneity will increase

with both of these quantities.


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Implications of Specular Reflection

Consider a soft tissue/bone interface. The intensity reflection

coefficient is (7.80 - 1.63)2 / (7.80 + 1.63)2 = 0.42. In other

words, the transmission is 58%. This means that if we were

trying to use ultrasound to image the brain, we would only get33% of the incident intensity into the brain, and only 33% of the

scattered intensity out of the brain.

100%58%

33%

BRAIN

SKULL

SKIN

This is one reason that

ultrasound is not used

to image the adult brain,but it is used in infants

where the skull is not

yet calcified. The reflection

coefficient at gas/tissue

interfaces is even larger,

so ultrasound is not useful

in imaging the lung. This alsoexplains why a coupling gel is

used during ultrasound exams.

The gel fills the space

between the transducer

and the skin and prevents

reflection by trapped air.

SKIN

TRANSDUCER

COUPLING GEL


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ttenuation

ogether, absorption and scattering result in attenuation of the

trasound wave. The intensity falls off exponentially with

stance. We could express this as exp (-x) using a linear

ttenuation coefficient , but with ultrasound it is conventionalo express the attenuation coefficient in dB cm-1.

igression on the decibel

he decibel scale is a way to represent the ratio of two intensities.

the two intensities are I1 and I2, we can express the ratio as

dB ratio = 10 log10 (I1 / I2)

o if the attenuation coefficient is, for example, 1 dB cm-1, after

ansmission through 10 cm of tissue the intensity will be reduced

y 10 dB, or a factor of 10. After 20 cm it would be reduced by

0 dB or a factor of 100.

requency dependence

he attenuation coefficient is a function of ultrasound frequency.

s shown on the next diagram, the coefficient is roughly

roportional to frequency for a variety of tissues. In fact, a good

ule-of-thumb is that the attenuation coefficient for soft tissue is

dB cm-1 MHz-1. Note the implications: at 1 MHz, transmission

rough 10 cm of tissue reduces intensity by a factor of 10, but at

0 MHz, transmission through 10 cm of tissue reduces intensity

y a factor of 10,000,000,000! This frequency-dependent

ttenuation imposes limits on the performance of imaging systems.


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Generation and detection of ultrasound

Both rely on the piezoelectric effect, first discovered in the late

1800s. Certain natural crystals (e.g. quartz) undergo a change

in physical size when an electric field is imposed on the crystal.

+ -

This change can launch an acoustic wave in the surroundingmedium. These materials also demonstrate a reciprocal

piezoelectric effect: a voltage is generated across the crystal

which is proportional to the applied pressure. This pressure

can result from an acoustic wave incident on the crystal. The

esulting voltage can be detected and amplified, so we have a

means of detecting acoustic waves as well as generating them.

n medical applications, the same transducer is used togenerate acoustic waves and to detect the waves scattered

by tissue. Synthetic materials demonstrate a much more

efficient conversion of electrical to acoustic energy. Most

medical transducers are fabricated from a ceramic material:

ead zirconate titanate, or PZT.

A transducer has a natural resonant frequency correspondingo a wavelength (in the material) that is twice the transducer

hickness. For PZT, a 1 mm thick transducer resonates at 2

MHz. While resonance contributes to efficient energy conversion,

t may be detrimental in other ways.


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Most ultrasound imaging is based on the localization of echoes

or scattered waves from structures within the tissue. Intuitively,

we would expect that it is easier to figure out where a short pulse

ame from because longer pulses would lead to temporal overlap

n the echoes and greater ambiguity about their source. Hence

is generally desirable for the source to emit a short pulse, buthis is not what we get from a resonating transducer:

PO

WER

PRESSURE

FREQUENCYTIME

The emitted pulse lasts a long time and looks like a pure tone.

This corresponds to a narrow frequency spectrum. If thetransducer is damped so that it does not resonate, the pulse is

much shorter, and the spectrum is correspondingly broad.

PRESSURE

POWE

R

TIMEFREQUENCY

BANDWIDTH


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his diagram shows a single element transducer with electrical

onnections and mechanical damping. Note also the matching

ayer. This is fabricated of material with an acoustic impedance

etween that of the PZT crystal (very high) and tissue. Its

hickness is one quarter wavelength. It can be shown that this

maximizes power transfer from the PZT to the tissue in the same

ay anti-reflective coatings work on glasses and other optics.

ote also, that the PZT crystal is shaped like a spherical shell.

his results in a focussed wave. Similar results could be achieved

ith an acoustic lens

s discussed later, single element transducers are no longer

sed in medical imaging. Multi-element arrays offer potential

or electronic focussing and real-time imaging.


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Acoustic fields of ultrasonic transducers

A very small transducer (compared to the ultrasound wavelength)

would act like a point source (and detector). This would be

useless for imaging because we would not be able to direct the

ultrasound pulse to a specific region, nor would we be able to tellwhich direction echoes were coming from.

Lets consider a larger transducer as shown below. We can

conceptually divide the transducer into many small elements, all

vibrating in phase. Each emits a spherical wave, but at some

arbitrary point in the tissue, these waves do not arrive in phase

because the distance to each element is different. In order tocalculate the resulting field we must consider the possibility of

interference and sum all of the waves taking into account their

relative phase. This is exactly what you do to calculate the

diffraction pattern from a single slit in optics, and the same sort

of integral appears in the acoustic problem.

P

AXIS OF SYMMETRY


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This diagram shows the amplitude of the pressure field when

the disk transducer vibrates at a single frequency. Close to thetransducer in the near field there is a complex interference

pattern. In the far field we see a well-defined peak on the axis

of the transducer. The peak broadens and decreases in size

as we move farther from the transducer.

Mathematically, we can show that the amplitude of the acoustic

pressure wave is proportional to

2 J1(k a sin ) where k = 2/

k a sin

J1 is first order

Bessel function


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We can do the same calculation for a focused transducer with

the geometry shown below:

a

z

Focal plane

P

r

In the focal plane the pressure amplitude is proportional to

2 J1 (kr / 2f) where f = z / 2a is called

the f - number or f#(kr / 2f)

Note that the amplitude peak becomes narrower as we decrease

(i.e. increase frequency) or as we decrease the f- number.

If the transducer emits a pulse (as it usually does) rather than a

continuous wave, the calculation is more complex because the

pulse contains energy at many frequencies. The generalfeatures of the field are the same if we consider the frequency to

correspond to the average of the power spectrum. The sharp

maxima and minima of the interference pattern are not so

evident for pulsed excitation. Note that a focused transducer can

be used as a detector and that it will have the same directional

sensitivity.


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..so enough physics, how do you make images?

To obtain a complete image it is necessary to a fill in the box

by getting sufficient A scans. In the early days of ultrasound

imaging this was accomplished by moving a single element

transducer over the surface of the patient. Before considering

modern methods, it is instructive to examine the resolution

attainable with the older method.


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Typical ultrasound image note the high resolution at 12 MHz,

the speckle in the image, and the characteristic appearance

of the cyst in the thyroid gland.

Image courtesy of

GE Healthcare


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he lateral resolution depends on the focusing properties of the

ransducer. It can be improved by increasing the frequency or

y decreasing the f#. The first option is limited by the clinicalequirements for penetration. In general, systems use the highest

requency possible. For a small organ like the thyroid, a higher

requency can be used than is possible with a large organ like the

ver. The second option is limited by the fact that stronger

ocusing (i.e. lower f#) will result in a reduced depth-of-field.

his means that lateral resolution is rapidly degraded at depths

ut of the focal plane.


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Electronic focusing can be achieved with annular arrays or with

linear arrays of transducer elements. A lens can be used with a

linear array to provide some (fixed) focusing in the plane

perpendicular to the array.


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Real time ultrasound imaging

A major advantage of modern systems is the ability to acquire

mages in real time. Practically, this means that an image must

be obtained in a time comparable to persistence of human vision.

This requires about 30 images (or frames) per second. For a

rame rate F, there is a relationship between the number of Acans in the image, L, and the maximum depth to be imaged, d.

n order to avoid ambiguity in echo location we require

1/F = 2 L d/c where c is sound velocity

Clearly it is not possible to translate a transducer quickly enough

o get real time images. Early scanners overcame this problem bysing a rotating or rocking scanner:

Surface

Shaded area is the imaged region. Because of its shape, thisis sometimes called a sector scan.


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Another way to do this would be to have an array of transducer

excited sequentially, so that each is used to get an A scan. The

problem with this approach is that a single element has very poor

ateral resolution. This can be overcome by using groups of

array elements to steer and focus the beam as shown below:


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The implementation of these ideas requires sophisticated

electronics and fabrication techniques. Advances in these

technologies have allowed real time systems to become

small, portable, and cheap compared to other imaging methods.


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Example (from GEs clinical image library) of color Doppler

study of carotid stenosis.

nstead of imaging flow velocity, another mode is power Dopple

where the integral of the power spectrum is displayed. This

ncreases sensitivity and allows imaging of organ perfusion


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new ultrasound ppt

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