U.S. patent application number 14/267415 was filed with the patent office on 2014-08-28 for transesophageal ultrasound using a narrow probe.
This patent application is currently assigned to Imacor Inc. The applicant listed for this patent is Imacor Inc. Invention is credited to Harold M. Hastings, Scott L Roth.
Application Number | 20140243672 14/267415 |
Document ID | / |
Family ID | 34652330 |
Filed Date | 2014-08-28 |
United States Patent
Application |
20140243672 |
Kind Code |
A1 |
Roth; Scott L ; et
al. |
August 28, 2014 |
Transesophageal Ultrasound Using a Narrow Probe
Abstract
Transesophageal echocardiography is implemented using a
miniature transversely oriented transducer that is preferably small
enough to fit in a 7.5 mm diameter probe, and most preferably small
enough to fit in a 5 mm diameter probe. Signal processing
techniques improve the depth of penetration to the point where the
complete trans-gastric short axis view of the left ventricle can be
obtained, despite the fact that the transducer is so small. The
reduced diameter of the probe (as compared to prior art probes)
reduces risks to patients, reduces or eliminates the need for
anesthesia, and permits long term direct-visualization monitoring
of patients' cardiac function.
Inventors: |
Roth; Scott L; (East Hills,
NY) ; Hastings; Harold M.; (Sheffield, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Imacor Inc |
Garden City |
NY |
US |
|
|
Assignee: |
Imacor Inc
Garden City
NY
|
Family ID: |
34652330 |
Appl. No.: |
14/267415 |
Filed: |
May 1, 2014 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
12732937 |
Mar 26, 2010 |
|
|
|
14267415 |
|
|
|
|
10997059 |
Nov 24, 2004 |
7717850 |
|
|
12732937 |
|
|
|
|
60525330 |
Nov 26, 2003 |
|
|
|
Current U.S.
Class: |
600/447 |
Current CPC
Class: |
A61B 8/12 20130101; G01S
15/8934 20130101; G01S 15/8915 20130101; G01S 7/52033 20130101;
A61B 8/4488 20130101; G01S 7/52079 20130101; A61B 8/0883 20130101;
G01S 7/52036 20130101; G01S 15/8925 20130101; A61B 8/445 20130101;
G01S 15/8977 20130101 |
Class at
Publication: |
600/447 |
International
Class: |
A61B 8/00 20060101
A61B008/00; A61B 8/08 20060101 A61B008/08; A61B 8/12 20060101
A61B008/12 |
Claims
1. A system for imaging regions that include at least two types of
tissue, the system comprising: an ultrasound imaging system; and a
probe including (a) a housing having a distal end and a flexible
shaft, (b) an phased array ultrasound transducer mounted in the
distal end of the housing, the transducer being made of a stack of
piezo element spaced at a pitch between 125 and 128 .mu.m with a
kerf of less than 30 .mu.m, and (c) an interface that operatively
connects the ultrasound transducer to the ultrasound imaging system
such that the ultrasound imaging system drives the ultrasound
transducer and receives return signals from the ultrasound
transducer, wherein the distal end has an outer diameter of less
than 7.5 mm and the flexible shaft has an outer diameter of less
than 7.5 mm, wherein the transducer is operated at between 6 and
7.2 MHz, and wherein the ultrasound transducer is configured to
emit a fan-shaped beam with a sector width of 60.degree. or
less.
2. The system of claim 1, wherein the distal end has an outer
diameter of less than 6 mm.
3. The system of claim 1, wherein the distal end has an outer
diameter of about 5 mm.
4. The system of claim 1, wherein the transducer is operated at
6.16 MHz.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of application Ser. No.
12/732,937, filed Mar. 26, 2010, which is a divisional of
application Ser. No. 10/997,059, filed Nov. 24, 2004 (now U.S. Pat.
No. 7,717,850), which claims priority to U.S. provisional
application No. 60/525,330, filed Nov. 26, 2003.
BACKGROUND
[0002] In the medical field, monitoring heart function impacts
critical decisions that relate to patient care. One type of prior
art heart monitor is the intravascular/intracardiac ultrasound
transducer (such as the Accunav.TM. transducer). This type of
transducer, however, is not well suited for transesophageal
echocardiography because the transducer elements are oriented
longitudinally instead of transversely, which limits the types of
images that can be obtained. A second type of prior art heart
monitor is the transesophageal echocardiography (TEE) transducer,
which is transversely oriented. However, in order to produce
repeatedly usable images, the azimuthal aperture of these
transducers must be quite large (e.g., 10-15 mm in diameter for
adults), which requires a correspondingly large probe. Because of
this large probe, conventional TEE often requires anesthesia, can
significantly threaten the airway, and is not well suited for
long-term monitoring of the heart.
SUMMARY OF THE INVENTION
[0003] Transesophageal ultrasound imaging is implemented using a
miniature transversely oriented transducer that is preferably small
enough to fit in a 7.5 mm diameter probe, and most preferably small
enough to fit in a 5 mm diameter probe. Signal processing
techniques provide improved depth of penetration, despite the fact
that the transducer is so small.
BRIEF DESCRIPTION OF THE DRAWINGS
[0004] FIG. 1 is an overall block diagram of a system for
monitoring cardiac function by direct visualization of the
heart.
[0005] FIG. 2 is a more detailed view of the probe shown in the
FIG. 1 embodiment.
[0006] FIG. 3 is a schematic representation of a displayed image of
the trans-gastric short axis view (TGSAV) of the left
ventricle.
[0007] FIG. 4 depicts the positioning of the transducer, with
respect to the heart, to obtain the TGSAV.
[0008] FIG. 5 shows a plane that slices through the trans-gastric
short axis of the heart.
[0009] FIG. 6A shows an optional probe interface configuration.
[0010] FIG. 6B is a graph of gain characteristics for a TGC
amplifier.
[0011] FIGS. 7A, 7B, and 7C show a first preferred transducer
configuration.
[0012] FIGS. 8A and 8B show a second preferred transducer
configuration.
[0013] FIG. 9 shows the components of spatial resolution.
[0014] FIG. 10 shows the interaction between the shape of the
resolution voxel and the boundary.
[0015] FIG. 11 shows the sector width.
[0016] FIG. 12 is a schematic illustration of the paths of the
ultrasound beam as it is swept through the sector.
[0017] FIG. 13 is a schematic illustration of the samples that
correspond to a section of one of the beams of FIG. 12.
[0018] FIG. 14 is a flowchart of a processing algorithm that uses
frequency characteristics of the return signal.
[0019] FIG. 15 is a graph of a function that maps a gain factor
onto an energy ratio.
[0020] FIGS. 16A and 16B show two alternative transducer
designs.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0021] FIG. 1 is an overall block diagram of a system that may be
used for continuous long term monitoring of cardiac function by
direct visualization of the heart. An ultrasound system 200 is used
to monitor the heart 110 of the patient 100 by sending driving
signals into a probe 50 and processing the return signals received
from the probe into images, using the image processing algorithms
described below. The images generated by those algorithms are then
displayed on a monitor 210, in any conventional manner.
[0022] FIG. 2 shows more details of the probe 50, which is
connected to the ultrasound system 200. At the distal end of the
probe 50 there is a housing 60, and the ultrasound transducer 10 is
located in the distal end 64 of the housing 60. The next portion is
the flexible shaft 62, which is positioned between the distal end
64 and the handle 56. This shaft 62 should be flexible enough so
that the distal end 64 can be positioned past the relevant
anatomical structures to the desired location, and the handle 56
facilitates the positioning of the distal end 64 by the operator.
Optionally, the handle 56 may contain a triggering mechanism 58
which the operator uses to bend the end of the housing 60 to a
desired anatomical position as described below.
[0023] At the other end of the handle 56 is a cable 54, which
terminates, at the proximal end of the probe 50, at connector 52.
This connector 52 is used to connect the probe 50 to the ultrasound
system 200 so that the ultrasound system 200 can operate the probe.
Signals for the ultrasound system 200 that drive the transducer 10
travel through the probe 50 via appropriate wiring and any
intermediate circuitry (not shown) to drive the transducer 10, and
return signals from the transducer 10 similarly travel back through
the probe 50 to the ultrasound system 200 where they are ultimately
processed into images. The images are then displayed on the monitor
210 in a manner well known to persons skilled in the relevant
art.
[0024] In the preferred embodiments, the housing 60 has an outer
diameter of less than 7.5 mm. The probe contains the ultrasound
transducer 10 and connecting wires, and the housing 60 can be
passed through the mouth or nose into the esophagus and
stomach.
[0025] The returned ultrasound signals are processed in the
ultrasound system 200 to generate an image of the heart.
Preferably, additional signal processing is used to significantly
improve image production, as described below. FIG. 3 shows a
displayed image of the trans-gastric short axis view (TGSAV) of the
left ventricle (LV), which is a preferred view that can be imaged
using the preferred embodiments. The illustrated image of the TGSAV
appears in a sector format, and it includes the myocardium 120 of
the LV which surrounds a region of blood 130 within the LV. The
image may be viewed in real time or recorded for later review,
analysis, and comparison. Optionally, quantitative analyses of
cardiac function may be implemented, including but not limited to
chamber and vessel dimensions and volumes, chamber function, blood
flow, filling, valvular structure and function, and pericardial
pathology.
[0026] Unlike conventional TEE systems, the relatively narrow
housing used in the preferred embodiments makes it possible to
leave the probe in position in the patient for prolonged periods of
time.
[0027] As best seen in FIGS. 4 and 5, the probe 50 is used to
introduce and position the transducer 10 into a desired location
within the patient's body. The orientation of the heart within the
chest cavity is such that the apex of the left ventricle is
positioned downward and to the left. This orientation results in
the inferior (bottom) wall of the left ventricle being positioned
just above the left hemidiaphragm, which is just above the fundus
of the stomach. During operation, the transducer 10 emits a
fan-shaped beam 90. Thus, positioning the transducer 10 in the
fundus of the stomach with the fan-shaped beam 90 aimed through the
left ventricle up at the heart can provide a trans-gastric short
axis view image of the heart 110. The plane of the fan-shaped beam
90 defines the image plane 95 shown in FIG. 5. That view is
particularly useful for monitoring the operation of the heart
because it enables medical personnel to directly visualize the left
ventricle, the main pumping chamber of the heart. Note that in
FIGS. 4 and 5, AO represents the Aorta, IVC represents the Inferior
Vena Cava, SVC represents the Superior Vena Cava, PA represents the
pulmonary artery, and LV represents the left ventricle.
[0028] Other transducer positions may also be used to obtain
different views of the heart, typically ranging from the
mid-esophagus down to the stomach, allowing the operator to
directly visualize most of the relevant cardiac anatomy. For
example, the transducer 10 may be positioned in the lower
esophagus, so as to obtain the conventional four chamber view.
Transducer positioning in the esophagus would typically be done
without fully flexing the probe tip, prior to advancing further
into the stomach. Within the esophagus, desired views of the heart
may be obtained by having the operator use a combination of some or
all of the following motions with respect to the probe: advance,
withdraw, rotate and slight flex.
[0029] For use in adults, the outer diameter of the housing 60 is
preferably less than about 7.5 mm, more preferably less than about
6 mm, and is most preferably about 5 mm. This is significantly
smaller than conventional TEE probes. This size reduction may
reduce or eliminate the need for anesthesia, and may help expand
the use of TEE for cardiac monitoring beyond its previous
specialized, short-term settings. When a 5 mm housing is used, the
housing is narrow enough to pass through the nose of the patient,
which advantageously eliminates the danger that the patient will
accidentally bite through the probe. Alternatively, it may be
passed through the mouth like conventional TEE probes. Note that
the 5 mm diameter of the housing is similar, for example, to
typical NG (naso-gastric) tubes that are currently successfully
used long-term without anesthesia in the same anatomical location.
It should therefore be possible to leave the probe in place for an
hour, two hours, or even six hours or more.
[0030] The housing wall is preferably made of the same materials
that are used for conventional TEE probe walls, and can therefore
withstand gastric secretions. The wiring in the probe that connects
the transducer to the rest of the system may be similar to that of
conventional TEE probes (adjusted, of course, for the number of
elements). The housing is preferably steerable so that it can be
inserted in a relatively straight position, and subsequently bent
into the proper position after it enters the stomach. The probe tip
may be deflected by various mechanisms including but not limited to
steering or pull wires. In alternative embodiments, the probe may
use an intrinsic deflecting mechanism such as a preformed element
including but not limited to pre-shaped materials. Optionally, the
probe (including the transducer housed therein) may be
disposable.
[0031] For imaging the TGSAV of the LV, the probe tip is preferably
ultimately "ante-flexed" (flexed towards the front of the patient)
approx. 70-110 degrees. This may be implemented, for example, by
building a triggerable ante-flex (e.g., on the order of 70 degrees)
into the probe through a combination of a pre-formed element, a
device to prevent flexing during insertion and a trigger to release
the preformed element from the insertion limit once the probe is in
the desired anatomic location. Optionally, a pull-wire may be used
for steering to provide the additional 0-40 degrees of flex after
the transducer is lowered to the appropriate depth. The triggerable
ante-flex component is preferably designed so that it will present
little resistance to returning to the unflexed position during
probe removal.
[0032] FIG. 6A shows an optional configuration that is similar to
the FIG. 1 embodiment, except that the circuitry that interfaces
with the probe 50 is relocated to an interface box 203. The rest of
the ultrasound system remains in the main processing unit 201,
which communicates with the interface box 203 via an appropriate
cable 205. The interface box 203 contains circuitry to amplify the
signals from the transducer 10 and/or digitize those signals. Using
such an interface box advantageously provides shorter signal paths
for those parts of the circuit that are most sensitive to
electrical noise (i.e., where the signals are small). The transmit
signals that drive the transducer 10 may also be generated within
the interface box 203 if desired.
[0033] Electrical noise may be further reduced using a variety of
techniques. For example, in one embodiment, the interface box 203
houses a preamplifier that serves as the first stage in the
amplification/processing chain, and separate power supplies are
used for the interface box and the main processing unit 201 to
reduce electronic noise pass-through. In another embodiment, the
interface box 203 houses a preamplifier that serves as the first
stage in the amplification/processing chain, and the preamplifier
operates on battery power. For both of these embodiments, time gain
compensation (TGC) is preferably implemented in that preamplifier.
TGC compensates for the fact that the return signals from distant
scatterers are weaker than those for nearby scatterers by
increasing the gain for signals with longer travel time. TGC may be
implemented using conventional techniques that are well known to
persons skilled in the relevant art. An example of suitable gain
vs. delay characteristics for TGC is shown in FIG. 6B, where the
x-axis represents the delay between transmission of the ultrasound
pulse and detection of the return signal, and corresponds to depth
as follows:
Depth(in cm)=0.077 cm/.mu.s.times.delay(in .mu.s).
[0034] Implementing TGC in the preamplifier facilitates efficient
digitization. The preamplifier may also provide amplitude
compandoring (a form of compression) to further facilitate
efficient digitization. Optionally, the preamplifier's output may
be digitized in the interface box, in which case only digital
signals would be sent from the interface box to the main processing
unit to further reduce electrical noise. These digital signals may
even be opto-isolated to eliminate all possible electrical
connections in the return path, to reduce electrical noise
pass-through further still.
[0035] The preferred embodiments described herein provide a good
quality image of the TGSAV of the LV from a transducer that is
small enough to fit in the narrow housing described above. FIGS.
7A-7C depict a first preferred transducer 10. FIG. 7A shows the
location of the transducer 10 in the distal end of the housing 60,
and also includes a top view 22 of the transducer 10 surrounded by
the wall of the housing 60 and a front cutaway view 24 of the
transducer 10.
[0036] As best seen in FIG. 7B, the azimuth axis (Y axis) is
horizontal, the elevation axis (Z axis) is vertical, and the X axis
projects out of the page towards the reader. When steered straight
forward by energizing the appropriate elements in the transducer,
the beam will go straight out along the X axis. The steering
signals can also send the beam out at angles with respect to the X
axis, in a manner well know to persons skilled in the relevant
arts.
[0037] The transducer 10 is preferably a phased array transducer
made of a stack of N piezo elements L.sub.1 . . . L.sub.N, an
acoustic backing 12, and a matching layer in the front (not shown),
in a manner well known to those skilled in the relevant art. As
understood by persons skilled in the relevant arts, the elements of
phased array transducers can preferably be driven individually and
independently, without generating excessive vibration in nearby
elements due to acoustic or electrical coupling. In addition, the
performance of each element is preferably as uniform as possible to
help form a more homogeneous beam. Optionally, apodization may be
incorporated into the transducer (i.e., tapering the power driving
transducer elements from a maximum at the middle to a minimum near
the ends in the azimuthal direction, and similarly tapering the
receive gain).
[0038] The preferred transducers use the same basic operating
principles as conventional TEE transducers to transmit a beam of
acoustic energy into the patient and to receive the return signal.
However, while the first preferred transducer 10 shown in FIGS.
7A-7C shares many characteristics with conventional TEE
transducers, the first preferred transducer 10 differs from
conventional transducers in the following ways:
TABLE-US-00001 TABLE 1 conventional TEE first preferred Feature
transducer transducer Size in the transverse (azimuthal) 10-15 mm
about 4-5 mm direction Number of elements 64 about 32-40 Size in
the elevation direction 2 mm about 4-5 mm Front face aspect ratio
about 1:5 about 1:1 (elevation:transverse) Operating frequency 5
MHz about 6-7.2 MHz
In FIG. 7A, the elevation is labeled E and the transverse aperture
is labeled A on the front cutaway view 24 of the transducer 10. The
location of the wall of the housing 60 with respect to the
transducer 10 can be seen in the top view 22.
[0039] FIG. 7C shows more details of the first preferred transducer
10. Note that although only eight elements are shown in all the
figures, the preferred transducer actually has between about 32-40
elements, spaced at a pitch P on the order of 130 .mu.m. Two
particularly preferred pitches are approximately 125 .mu.m (which
is convenient for manufacturing purposes) and approximately 128
.mu.m (0.6 wavelength at 7.2 MHz). When 32-40 elements are spaced
at a 125 .mu.m pitch, the resulting azimuth aperture A (sometimes
simply called the aperture) of the transducer 10 will be between 4
and 5 mm. The reduced element count advantageously reduces the wire
count (compared to conventional TEE transducers), which makes it
easier to fit all the required wires into the narrower housing. The
kerf K (i.e., the spacing between the elements) is preferably as
small as practical (e.g., about 25-30 .mu.m or less). Alternative
preferred transducers may have between about 24-48 elements, spaced
at a pitch between about 100-150 .mu.m.
[0040] A second preferred transducer 10' is shown in FIGS. 8A-8B.
This transducer 10' is similar to the first preferred transducer 10
described above in connection with FIGS. 7A-7C, except it is taller
in the elevation direction. Similar reference numbers are used in
both sets of figures to refer to corresponding features for both
transducers. Numerically, the second transducer differs from
conventional transducers in the following ways:
TABLE-US-00002 TABLE 2 conventional TEE second preferred Feature
transducer transducer Size in the transverse (azimuthal) 10-15 mm
about 4-5 mm direction Number of elements 64 about 32-40 Size in
the elevation direction 2 mm about 8-10 mm Front face aspect ratio
about 1:5 about 2:1 (elevation:transverse) Operating frequency 5
MHz about 6-7.2 MHz
[0041] In alternative embodiments, the transducer 10 may be built
with a size in the elevation direction that lies between the first
and second preferred transducers. For example, it may have a size
in the elevation direction of about 7.5 mm, and a corresponding
elevation:transverse aspect ratio of about 1.5:1.
[0042] The transducer 10 preferably has the same transverse
orientation (with respect to the axis of the housing 60) as
conventional TEE transducers. When the transducer is positioned in
the stomach (as shown in FIG. 4), the image plane (azimuthal/radial
plane) generated by the transducer intersects the heart in the
conventional short axis cross-section), providing the trans-gastric
short axis view of the heart, as shown in FIGS. 3 and 5. The
transducer is preferably as wide as possible in the transverse
direction within the confines of the housing. Referring now to the
top view 22 in FIG. 7A, two examples of transducers that will fit
within a 5 mm housing are provided in the following table, along
with a third example that fits in a housing that is slightly larger
than 5 mm:
TABLE-US-00003 TABLE 3 first second third Parameter example example
example number of elements in the transducer 38 36 40 a (azimuthal
aperture) 4.75 mm 4.50 mm 5.00 mm b (thickness) 1.25 mm 2.00 mm
2.00 mm c (inner diameter of housing at 4.91 mm 4.92 mm 5.39 mm the
transducer) housing wall thickness 0.04 mm 0.04 mm 0.04 mm outer
diameter of housing 4.99 mm 5.00 mm 5.47 mm
Referring now to the top view 22 in FIG. 8A, the three examples in
Table 3 are also applicable for fitting the second preferred
transducer 10' within a 5-5.5 mm housing.
[0043] The above-describe embodiments assume that the housing is
round. However, other shaped housings may also be used to house the
transducer, including but not limited to ellipses, ovals, etc. In
such cases, references to the diameter of the housing, as used
herein, would refer to the diameter of the smallest circle that can
circumscribe the housing. To account for such variations in shape,
the housing may be specified by its outer perimeter. For example, a
5 mm round housing would have a perimeter of 5.pi. mm (i.e., about
16 mm). When a rectangular transducer is involved, using an oval or
elliptical housing can reduce the outer perimeter of the housing as
compared to a round housing. For example, an oval that is bounded
by a 6 mm.times.2 mm rectangle with its corners rounded to a radius
of 0.5 mm contains a 5 mm.times.2 mm rectangular region, which can
hold the third example transducer in Table 3. Allowing for a 0.04
mm housing wall thickness yields an outer perimeter of 15.4 mm,
which is the same outer perimeter as a 4.9 mm diameter circle. The
following table gives the outer perimeters that correspond to some
of the diameters discussed herein:
TABLE-US-00004 TABLE 4 outer diameter outer perimeter 2.5 mm 8 mm 4
13 5 16 6 19 7.5 24
[0044] Since the characteristics of the last one or two elements at
each end of the transducer may differ from the characteristics of
the remaining elements (due to differences in their surroundings),
the last two elements on each side may be "dummy" elements. In such
a case, the number of active elements that are driven and used to
receive would be the total number of element (shown in Table 3)
minus four. Optionally, the wires to these dummy elements may be
omitted, since no signals need to travel to or from the dummy
elements. Alternatively, the wires to may be included and the last
two elements may be driven, with the receive gain for those
elements severely apodized to compensate in part for their position
at the ends of the transducer.
[0045] Preferably, conventional beam-forming techniques are used to
generate and aim a beam of acoustic energy in the desired
directions. For example, focusing in the azimuthal direction may be
accomplished by phasing (i.e., timing the excitation of individual
elements L.sub.1 . . . L.sub.N in the array, and using appropriate
time delays in the returns of individual elements before summing
the respective returns into an ultrasound return signal). Focusing
in the elevation direction may be accomplished based on the
near-field and far-field properties of the sound signal, and will
depend upon the physical height of the elements in the elevation
direction and optional acoustic lenses.
[0046] Resolution adequate to determine LV size and function
depends upon a combination of resolution in azimuth, elevation, and
axis. This combination is referred to as "spatial resolution" and
is illustrated in FIG. 9. FIG. 9 shows the image plane 320 and a
scan line 310 that lies on the image plane 320. The axial direction
AX is defined by the scan lines 310, and the transducer (not shown)
is located far back along the AX axis. Out at the voxels being
imaged, the azimuthal direction AZ is perpendicular to the AX axis
within the image plane 320, and the elevation axis EL is
perpendicular to the image plane 320. In an ideal system, each
voxel would be a point. In real-world systems, however, the voxels
have a volume that is defined by the resolution in all three
directions AX, AZ and EL, as shown for voxel 330. Similarly, while
the image plane 320 is depicted as a thin plane, the real-world
image plane will have a thickness in the elevation direction EL
that is equal to the thickness of the voxel 330 in the elevation
direction.
[0047] The general formula for azimuthal and elevation resolutions
is:
.DELTA..theta..apprxeq.1.22.lamda./d,
where .DELTA..theta. denotes the beamwidth in radians, .lamda. the
wavelength (corresponding to the transducer center frequency) and d
the aperture in the given direction (azimuth or elevation). The
wavelength .lamda. and aperture d are measured in the same units
(e.g., .mu.m).
[0048] Axial resolution depends indirectly upon the wavelength
.lamda.. Although the inventors are not aware of any specific
formula for axial resolution, it is typically on the order of 16-64
times the wavelength. Thus, increasing the center frequency
increases all three components of spatial resolution. A center
frequency on the order of 5-10 MHz is high enough to provide
adequate resolution.
[0049] FIG. 10 illustrates the interplay between the three
components in determining the interaction between the shape of the
resolution voxel and the boundary orientation in detecting and
determining boundaries. It shows the same voxel 330 that appears in
FIG. 9, and also shows an illustrative piece 340 of the boundary
being imaged that coincides with that voxel. If the boundary
orientation is random with respect to the resolution voxel, one
suitable approach is to make the resolution voxel as cubical as
possible. In order to obtain that shape, the azimuth and elevation
resolutions for a given voxel should be approximately equal, which
occurs when the front face of the transducer is approximately
square, as it is for the first preferred transducer discussed above
in connection with FIGS. 7A-7C.
[0050] For the first preferred transducer, the elevation aperture
is approximately the same as the azimuth aperture. In other words,
the front face of the transducer has a elevation:transverse aspect
ratio that is approximately 1:1 (i.e., it is approximately square).
A square transducer with a width of 4-5 mm in the transverse
direction would therefore have an area of approximately 16-25
mm.sup.2.
[0051] The formulas for azimuthal and elevation resolution are:
.DELTA..theta..sub.AZ=1.22.times..lamda./d.sub.AZ
and
.DELTA..theta..sub.EL=1.22.times..lamda./d.sub.EL
where .DELTA..theta..sub.AZ and .DELTA..theta..sub.EL are the
azimuth and elevation resolutions, respectively, (both measured in
radians); and d.sub.AZ and d.sub.EL are the azimuth and elevation
apertures, respectively. These components may be combined into a
single equation for overall resolution as a function of area and
frequency, as follows:
.DELTA..theta..sub.OVERALL=1.5.times..lamda..sup.2/(d.sub.AZ.times.d.sub-
.EL)
[0052] As explained above, increasing the center frequency results
in increased resolution. However, increasing the center frequency
also reduces the penetration depth due to frequency-dependent
attenuation, which is governed by the approximate formula
a.apprxeq.0.5f.times.r
where a denotes the one-way attenuation in dB, f is the center
frequency in MHz, and r the depth in cm. Thus, one-way frequency
dependent attenuation will typically be about 0.5 dB MHz.sup.-1
cm.sup.-1 and typical round-trip frequency dependent attenuation
will typically be about 1 dB MHz.sup.-1 cm.sup.-1.
[0053] The inventors have determined that a transducer center
frequency between about 6 and 7.2 MHz provides a good trade-off
between resolution and depth of penetration for TEE using a
transducer with a 4.75 mm azimuthal aperture. In the embodiments
described herein, that range of frequencies can typically provide
enough depth of penetration to image the far wall of the left
ventricle (in the TGSAV) so that the interior volume of the left
ventricle can be computed. (In most subjects, a 12 cm depth of
penetration is adequate to image the far wall. For many subjects, a
depth of penetration of about 9-10 cm will suffice).
[0054] When the transducer elements are spaced at a 125 .mu.m
pitch, using a transducer center frequency of 6.16 MHz is
particularly advantageous because it corresponds to a wavelength of
.lamda.=250 .mu.m. At that wavelength, the elements are spaced at a
pitch of 0.5.lamda., which is sometimes referred to as
"half-wavelength pitch". As is well known to those skilled in the
art, a half-wavelength pitch is excellent for eliminating grating
lobes while still minimizing element count for a given azimuth
aperture. Somewhat larger pitches, e.g. 0.6.lamda., still work
reasonably well in terms of eliminating grating lobes. Thus, for
transducers that can operate at a range of center frequencies,
acceptable performance can be maintained even if the frequency is
increased about 20% (i.e., to the point where the pitch becomes
about 0.6.lamda.).
[0055] As explained above, the formula for the angular resolution
is .theta..apprxeq..lamda./d. Referring to the tables above, the
first example of the first preferred transducer has a 38 element
transducer with a 125 .mu.m pitch, resulting in a 4.75 mm
transducer width (d=4750 .mu.m). It is preferably approximately
square and operates at a center frequency of 6.16 MHz (.lamda.=250
.mu.m). When those values for d and .lamda. are plugged into the
equation for resolution, the result is 0.apprxeq.0.053 radians,
which converts to approximately 3 degrees resolution in both
azimuth and elevation.
[0056] The increased size of the transducer in the elevation
direction helps improve the angular resolution of the system in the
elevation direction (as compared to a conventional TEE transducer
with a 2 mm elevation). This increased resolution in the elevation
direction helps compensate for losses in angular resolution in the
azimuth direction caused by shrinking the azimuthal aperture down
to about 4-5 mm.
[0057] The inventors have noticed that increasing the size of the
transducer in the elevation direction further, so that it is larger
than the size in the azimuthal direction provides improved
performance when imaging the far wall of the heart in the TGSAV.
This increase in transducer elevation causes the resolution voxel
to shrink in the elevation direction at distances that correspond
to the far wall of the LV, which results in increased resolution in
the elevation direction. The inventors believe that increasing
resolution in this direction is helpful at least in part because
the far wall is slanted about the Y axis with respect to the front
face of the transducers. (The Y axis is shown in FIG. 8B.)
Shrinking the size of the voxel in the elevation direction
therefore minimizes the variations of the components of return
signals arising from specular reflections that fall within a single
voxel.
[0058] The inventors have determined that the images of the TGSAV
are better when the transducer is more than 1.5 times as large in
the elevation direction as the transverse direction, and that the
best images of the TGSAV are obtained when the transducer is about
two times as large in the elevation direction as the transverse
direction, as it is for the second preferred transducer 10'
described above in connection with FIGS. 8A and 8B and Table 2.
[0059] Instead of the 90.degree. sector width that is typically
used in conventional TEE systems, the preferred embodiment uses a
smaller sector width (e.g., 60 degrees). Referring now to FIG. 11,
a 60.degree. sector 92 is shown emanating from the front face 14 of
the transducer 10. The effective azimuthal aperture at an angle
.theta. from the centerline CL can be obtained by multiplying the
(nominal) azimuthal aperture (at .theta.=0) by cos(.theta.).
Since)cos(30.degree.=0.866 and)cos(45.degree.=0.707, restricting
the sector width to 60.degree. (i.e., 30.degree. on each side of
the center line CL) causes a smaller degradation in worst-case
azimuthal aperture: azimuthal aperture is degraded by only 13.4%,
compared with 26.8% in the case of a 90.degree. sector width. For
example, the worst case aperture for a 4.75 mm wide transducer (5
mm housing diameter) in a 60.degree. sector would be about 4.11 mm.
The result is improved effective azimuthal aperture, which improves
the overall resolution obtainable with small transducers. If a
conventional 90.degree. sector were to be used, a 5.82 mm wide
transducer (6.1 mm housing diameter) would be needed in order to
provide the same worst case aperture.
[0060] After a beam of ultrasound energy is sent into the patient
using the transducer described above, the ultrasound return signal
is received, preferably by the same transducer. The transducer
converts the ultrasound return signal into an electrical return
signal. This process continues as the beam is swept through the
imaging sector. FIG. 12 is a schematic illustration of the path of
the ultrasound beam as it is swept through the sector, first along
line B.sub.1, then along line B.sub.2, and continuing on through
line B.sub.M. These scan lines B.sub.1 . . . B.sub.M correspond to
the fan-shaped beam 90 (shown in FIG. 4) and the sector 92 (shown
in FIG. 11). Although the illustration only includes a small number
(M) of scan lines, an actual system would have many more scan lines
that are much more densely packed, so as not to adversely impact
the azimuthal resolution.
[0061] The electrical return signal can be modeled as being an
amplitude-modulated signal, with the carrier frequency at the
center frequency, and with the modulation being caused in large
part from scatterer spacing and other tissue characteristics such
as the presence of connective tissue around heart muscle bundles.
The electrical return signal is demodulated and digitized (i.e.,
sampled) to form a demodulated and digitized return signal (DDRS).
A variety of conventional techniques that are well known to persons
skilled in the relevant arts may be used to form the DDRS. One
example is to digitize the electrical return signal and then
rectify the result (i.e., take the absolute value) to form a
rectified digitized ultrasound return signal. Another example is to
rectify the electrical return signal in analog form, and then
digitize the result to form the DDRS. Alternative demodulation
approaches may also be used to extract the modulation information
from the electrical return signal, including but not limited to
coherent demodulation, Hilbert transforms, and other demodulation
techniques that are well known to persons skilled in the relevant
arts.
[0062] FIG. 13 is a schematic illustration of the DDRS that
corresponds to a section of one of the ultrasound beams B.sub.1 . .
. B.sub.M of FIG. 12. Each sample is represented by a dot S0 . . .
S143. Each sample corresponds to a point in 2D space based on the
direction of the beam and the time it took for the signal to travel
from the transducer to the point in question and back. For example,
if the return signal is digitized at 50 MHz, the time between
samples will be 0.02 .mu.s, which corresponds to a distance of
0.015 mm (based on the speed of sound in the body). Although the
illustration only includes only 144 samples, an actual system would
have many more samples in each scan line to provide the desired
resolution. For example, to obtain a depth of penetration of 12 cm
with the samples spaced 0.015 mm apart, 8000 samples would be
needed. Because the beam of ultrasound energy is swept about a
center point, polar coordinates are useful to organize the samples,
at least in this stage of the processing. In some embodiments, the
samples are analyzed entirely in polar coordinates, and only
converted into rectangular coordinates for viewing on a
conventional computer monitor. In other embodiments, the sample
space may be converted into rectangular coordinates at an earlier
stage of processing. The remaining explanation considers
coordinates along each scan line (constant .theta. in the (r,
.theta.) polar coordinate system, with r varying along the scan
line), and pixel data associated with the center of the pixel along
the scan line. Conversion to a sector image is well known in the
field of ultrasound imaging.
[0063] The samples of each scan line are preferably processed by
two different algorithms: one algorithm that analyzes intensity
characteristics of the samples, and one algorithm that analyzes
frequency characteristics of the samples.
[0064] For the first algorithm (i.e., the intensity algorithm) the
samples of the scan line is divided into a plurality of pixels,
with each pixel containing a plurality of samples. In the FIG. 13
example, each pixel contains 16 samples, as indicated by the boxes
labeled "WIAP j" (which stands for "Window for Intensity Algorithm
for Pixel j, where j is an integer from 0-8) that appear below the
corresponding samples. Pixel data generated by signal processing is
associated with the center position of the corresponding pixel. Of
course, other numbers of samples per pixels could also be used
instead of 16. In one preferred embodiment, for example, each pixel
contains eight samples. The intensity algorithm is preferably a
conventional image processing algorithm that converts the samples
into a conventional image. The intensity for any given pixel is
determined based on the amplitude of the samples that correspond to
that pixel, with higher intensities corresponding to larger
amplitudes. In the case of a 16 sample pixel, the average of those
16 samples would be used to determine the intensity at the pixel
(with higher average intensity values appearing brighter and lower
average intensity values appearing darker). Optionally, the
intensity level of the pixel (or the samples that make up that
pixel) may be compressed using conventional procedures such as
logarithmic compression.
[0065] The second algorithm (i.e., the frequency algorithm)
analyzes the frequency characteristics of the sample space and
determines the spatial frequencies in scatterer spacing. Examples
of suitable algorithms are described in U.S. Pat. No. 5,417,215
(hereinafter "the '215 patent"), which is incorporated herein by
reference. The article "Spectral Analysis of Demodulated Ultrasound
Returns: Detection of Scatterer Periodicity and Application to
Tissue Classification" by S. Roth, H. M. Hastings, et al.,
published in Ultrasonic Imaging 19 (1997) at pp. 266-277, is also
incorporated herein by reference.
[0066] The frequency algorithm provides a second result for each
pixel in the image (i.e., in addition to the result produced by the
intensity algorithm). Because most frequency analyzing algorithms
provides better results when a larger number of data samples is
used, and because each pixel only has a limited number of samples,
samples on either side of the pixel in question are preferably
combined with the samples in the pixel itself to increase the
number of samples. In the illustrated example, each pixel contains
16 samples, but the frequency algorithm for any given pixel
operates on 64 samples that are preferably centered in the pixel,
as indicated by the boxes labeled "Window for Freq. Algorithm for
Pixel k" (where k is an integer from 2-6) that appear below the
samples. In this case, for example, all the samples from pixels 2-4
plus half the samples from pixels 1 and 5 would be used to perform
the frequency analysis for pixel 3. Of course, other numbers of
samples could be used for the frequency analysis instead of 64.
Powers of 2, however, are preferable when a fast Fourier transform
(FFT) algorithm is used. Optionally, windowing techniques (such as
Hamming windows) may be used to weight the samples in the center
more heavily than the samples that are near the ends.
[0067] FIG. 14 is a flowchart of a suitable frequency algorithm. In
this algorithm, steps 1 and 2, taken together, attempt to discern
the material that the pixel in question is made of (and more
specifically, whether that pixel is blood or muscle) based on the
frequency characteristics of samples in the pixel and the samples
in the neighboring pixels.
[0068] In step 1, a Fourier analysis is performed on the samples to
determine the power distribution in the various frequency bands at
each pixel. The end result of the Fourier analysis of step 1 is a
set of amplitude coefficients for each of a plurality of different
frequencies, for each of the pixels (i.e., one set of coefficients
for the first pixel, a second set of coefficients for the second
pixel, etc.). The Fourier analysis may be implemented using any of
a variety of algorithms that are well known to persons skilled in
the relevant arts (e.g., a conventional FFT algorithm). Alternative
embodiments may use other frequency analysis tools to achieve
similar results, such as bandpass techniques (preferably
integer-based FIR recursive), wavelet techniques, etc. In step 2,
the ratio of power in a selected frequency band to the power in the
entire spectrum for each pixel is computed. Thus, for each pixel,
the following formula applies:
R=E.sub.BAND/E.sub.TOTAL
Where E.sub.BAND is the power in the selected frequency band,
E.sub.TOTAL is the total power in the portion of the spectrum, and
R is the ratio of those two powers. When a Fourier analysis is
used, the power in any given band equals the sum of the squares of
the amplitudes of the Fourier coefficients within the band. The
"selected frequency band" in this step is preferably selected so
that changes in the ratio R are correlated to differences in the
material that is being imaged (e.g., blood v. muscle).
Alternatively, it may be selected so that changes in the ratio R
are correlated to differences in S/N ratio, with larger Rs being
correlated with signal and smaller Rs being correlated with speckle
or electric noise. Optionally, different "selected frequency bands"
may be used for near returns and for distant returns. For example,
a wider frequency band may be used for signals that correspond to
distant structures. In other words, the band selection can be a
function of depth.
[0069] One suitable set of numeric values that results in a
correlation between R and the material being imaged will now be
discussed. Consider first the ultrasound return from a single
scatter at a depth of r mm. The return from this scatter arrives
after a time delay of t .mu.s, given by
t=r/v=r/(0.77 mm/.mu.s)=1.30 r .mu.s,
where the scaling factor of 0.77 mm/.mu.s represents a round trip
from the transducer to the scatterer and back (assuming the
velocity of sound in tissue is to be 1.54 mm/.mu.s).
[0070] The effects of scatterer periodicity upon the spectrum of
the demodulated ultrasound return may be calculated in the case of
separations large enough so that the ultrasound returns do not
overlap (i.e., separations .DELTA.r larger than .DELTA.r.sub.0=0.77
mm/.mu.s.times..DELTA.t). For example, in the case of an ideal one
cycle pulse, a 5 MHz center frequency, and a ideal wide-bandwidth
transducer,
.DELTA.t=1/f.sub.c=1/(5 MHz)=0.200 .mu.s,
and thus
.DELTA.r.sub.0=0.77 mm/.mu.s.times.0.200 .mu.s=0.154 mm.
[0071] The internal structure of cardiac muscle displays variations
on this and larger spatial scales. In contrast, scattering from
blood is characterized by full-developed speckle, including
variations on all, and especially much smaller spatial scales. As a
result, low frequencies are indicative of muscle, and high
frequencies are indicative of blood. This suggests defining the
upper limit of a low frequency band to be less than about 4 MHz,
corresponding to a minimal spatial scale .DELTA.r.sub.MIN of
.DELTA.r.sub.MIN=0.77 mm/.mu.s.times.1/(4 MHz)=0.77
mm/.mu.s.times.250 .mu.s=0.193 mm.
[0072] The inventors have performed tissue experiments that used a
signal digitized at 50 MHz (corresponding to sampling interval of
0.02 .mu.s), and computed the FFT in a 64 point window
(corresponding to 64.times.0.02 .mu.s=1.28 .mu.s, or 0.986 mm).
With that size window, the inventors selected a low frequency band
that included Fourier frequencies of between 2 and 5 cycles per
window (inclusive), which corresponds to frequencies between 2/1.28
MHz=1.56 MHz and 5/1.28 MHz=3.91 MHz.
[0073] Using the formula for R set forth above
(R=E.sub.BAND/E.sub.TOTAL) for this low frequency band, the ratio
of Fourier power in the low frequency band to the total Fourier
power is computed for each pixel. The end result of step 2 in FIG.
14 is a value of R for each pixel.
[0074] The inventors have found that, for the parameter values used
in this example, R-values of around 0.45 are significantly
correlated to the presence of muscle tissue at the pixel of
interest, and R-values of around 0.20 are significantly correlated
to blood or regions dominated by electronic noise. The remaining
part of the algorithm uses this information to improve the image by
increasing the intensity of the portions of the image that
correspond to muscle and decreasing the intensity of the portions
of the image that correspond to blood. Since blood is less
reflective than muscle, this difference enhances the contrast
between blood and muscle.
[0075] The inventors have determined that cardiac ultrasound images
are dramatically improved when the intensity of the areas with
R-values corresponding to muscle is increased to about 120% of its
original value, and when the intensity of the areas with R-values
corresponding to blood is decreased to between about 20% and 50% of
its original value. Thus, in step 3 of FIG. 14, a gain factor of
about 1.2 is assigned to those portions of the image with R values
of about 0.45, and a gain factor of between about 0.2 and 0.5 is
assigned to those portions of the image with R values of about
0.20. This gain factor is referred to herein and a "feature gain
factor" or FGF because the gain is feature dependant.
[0076] While most pixels in most images will have R-values that
permit the pixel to be classified as being either muscle or blood,
in some cases the classification is less clear. For example, pixels
that straddle a boundary between muscle and blood have less
predictable R values. In addition, although the R-values from blood
may average out to 0.20, any given pixel of blood may vary widely
from that R-value due to random statistical variations.
Accordingly, a monotonic, preferably smooth function may be used to
map R to FGF in some embodiments. FIG. 15 is an example of a
suitable function for this purpose. Optionally, additional
restrictions may be built into the mapping function, based on other
tissue characteristics.
[0077] Finally, in step 4 of FIG. 14, the results of the intensity
algorithm and the frequency algorithm are combined by multiplying
the intensity value for each pixel (obtained from the intensity
algorithm) by the FGF value for that pixel (obtained from the
frequency algorithm). The result is an enhanced image in which the
pixels that are probably muscle have been brightened while the
pixels that are probably blood have been dimmed. This enhanced
image is then displayed using conventional hardware and software
techniques (including, for example, using interpolation to convert
the polar coordinates to rectangular coordinates).
[0078] The actual choice of the Fourier frequency bands, R-values
and corresponding FGF values depends upon a variety of factors
including but not limited to transducer center frequency, sampling
rate, window size and any optional windowing techniques used in
signal processing, transducer bandwidth, the width of interrogating
pulse, etc. In one embodiment, for example, a transducer center
frequency of 7.5 MHz is used, the scan line is digitized at about
four times the center frequency (i.e., about 30 MHz), and the
distance between the samples is about 0.026 mm.
[0079] In alternative embodiments, other normalized (i.e.,
non-amplitude-dependent measures) may be used instead of dividing
E.sub.BAND by E.sub.TOTAL. For example, the ratio of power in a
first frequency band to the power in a second frequency band may be
used to compute R, as explained in the '215 patent (e.g., by
dividing E.sub.BAND1 by E.sub.BAND2). In alternative embodiments,
two or more Fourier analyses may be performed for each pixel, using
a corresponding number of lines of samples, where the center of
each line is contained within the pixel. For example, a two line
per pixel arrangement in which a first 1D Fourier analyses is
implemented along a line of samples in the radial direction, and a
second 1D Fourier analyses is implemented along a second line of
samples in the tangential direction. The results from those two
lines of samples are then merged (e.g., by averaging). In still
other embodiments, a 2D Fourier algorithm may be used instead of
the 1D algorithms described above.
[0080] Ordinarily, the above-described operations are performed on
the uncompressed image data. Under certain circumstances, however,
it may be possible to perform corresponding operations directly on
a compressed version of the image data.
[0081] Once the enhanced images have been generated, they may be
displayed using conventional hardware. The images may be
continuously updated and displayed for the entire time that the
probe is in position, so that the physician can visualize the
patient's heart in real time. In alternative embodiments, images
may be acquired and optionally stored periodically (e.g., by
capturing one or more complete heartbeats every two minutes).
Optionally, the ability to compare a prior heartbeat to the current
heartbeat may be provided by, for example, playing back a stored
video clip (or "loop") of an old heartbeat in one window, and
displaying the current image in a second window.
[0082] In contrast to conventional extended duration TEE using a
transducer with a 10-15 mm azimuthal aperture, which is ordinarily
done only under general anesthesia in the closely monitored
environment of an room, the smaller diameter of the preferred
embodiments described herein permits the preferred embodiments to
be used without general anesthesia, and in less closely monitored
environments. Optionally, the preferred embodiments may be used
with sedation or local anesthesia in place of the general
anesthesia that was used with conventional extended duration TEE.
It may even be possible to forgo the use of sedation or anesthesia
altogether. In such cases, the patient may optionally be medicated
with an analgesic.
[0083] Optionally, regions of high relevance as detected by the
feature gain factor may be highlighted, typically using
colorization, while preserving the intensity of the gray-scale
image, as explained in the '215 patent. Additional techniques for
image enhancing can be found in application Ser. No. 10/633,949,
filed Aug. 4, 2003, and entitled "Method and Apparatus for
Ultrasonic Imaging," which is incorporated herein by reference.
[0084] The preferred embodiments described above advantageously
permit non-invasive, intermediate and long-term monitoring of
cardiac function using a small transducer that fits into a housing
approximately 5 mm in diameter, thereby reducing or eliminating the
need for anesthesia. The preferred embodiments described above
combine a plurality of techniques to produce images that are
comparable to or better than images that were conventionally
obtained by much larger transducers. The images produced by the
preferred embodiments described above are repeatably and reliably
usable for monitoring heart function, with adequate penetration
depth to see the far wall of the left ventricle (10-12 cm) and
adequate resolution to determine LV size and function from an image
of the endocardial wall in real time, despite the use of a smaller
transducer. Thus, in contrast to prior art systems which provide a
depth of penetration that is less than 15 times the azimuthal
aperture of the transducer (e.g., obtaining 10 cm penetration using
a 10 mm transducer) the preferred embodiments can provide
penetration that is greater than 15 times the azimuthal aperture of
the transducer, or even greater than 20 times the azimuthal
aperture of the transducer (e.g., obtaining 10 cm penetration using
a 4.75 mm transducer).
[0085] The preferred embodiments described above use a probe that
is much narrower than conventional TEE probes, and may be used to
monitor heart function over an extended period of time and to
obtain an understanding of the patients' hemodynamic status. Such
information may be useful in choosing treatments and improving
outcome in many situations (including but not limited to critical
medical problems such as hypotension, pulmonary edema and heart
failure).
[0086] The above-described embodiments permit direct visualization
of cardiac function, which permits evaluation of a patient's
hemodynamic status including intravascular volume (normal, low or
high), cardiac contractility (how well the left ventricle pumps),
cardiac ischemia (inadequacy of blood flow to the heart muscle) and
cardiac tamponade (fluid in the pericardial sac limiting heart
function). For example, information about intravascular volume
status can be derived from directly visualizing the size of the
left ventricle and monitoring changes in size with treatments over
time. Information about contractility can be obtained by directly
visualizing the contraction (pumping) of the left ventricle, either
using qualitative visual estimates or quantitatively. Information
about ischemia is available during direct visualization of the left
ventricle since ischemia results in abnormal motion of the walls of
the left ventricle (wall motion abnormality). Information about
possible cardiac tamponade or pericardial effusion (fluid in the
heart sac) is available when using ultrasound to directly visualize
the heart.
[0087] The narrowness of the probe may enable the above-described
embodiments to provide this information for longer periods of time,
outside the operating room, and/or without anesthesia. The
above-described embodiments also lend themselves to use in settings
where interventional cardiac procedures are performed such as the
cardiac catheterization and electrophysiology laboratories, both
for monitoring the effects of physicians' interventions on cardiac
and hemodynamic function and for guiding the placement of devices.
For example, they may be used to help the physician correctly place
the pacing leads to achieve the desired result. The above-described
embodiments may also be used in non-cardiac applications in which a
narrower probe is needed or beneficial.
[0088] The above-described embodiments are not limited to
ultrasound imaging modes, and may be used in alternative ultrasound
modes (e.g., pulsed wave Doppler, continuous wave Doppler, and
color flow imaging Doppler modes). These alternative modes may be
performed using the same transducer as the above-described imaging
modes and may yield information which can be combined with images,
optionally in real-time. For example, color flow Doppler
information may be obtained during imaging of the mitral valve
(between the left atrium and left ventricle) while maintaining
transducer position in the mid to lower esophagus. Such an
application would permit evaluation of leakage of the mitral valve
(mitral regurgitation or insufficiency).
[0089] If desired, the preferred embodiments described above may be
scaled down for neonatal or pediatric use. In such cases, a
transducer that is between about 2.5 and 4 mm in the azimuthal
direction is preferable, with the elevation dimension scaled down
proportionally. Because less depth of penetration is required for
neonatal and pediatric patients, the operating frequency may be
increased. This makes .lamda. smaller, which permits the use of a
smaller transducer element spacing (pitch), and a correspondingly
larger number of elements per mm in the transducer. When such a
transducer is combined with the above-described techniques, the
performance should meet or surpass the performance of conventional
7.5 mm TEE probes for neonatal and pediatric uses.
[0090] The embodiments described herein may also be used in
non-cardiac applications. For example, the probe could be inserted
into the esophagus to monitor the esophagus itself, lymph nodes,
lungs, the aorta, or other anatomy of the patient. Alternatively,
the probe could be inserted into another orifice (or even an
incision) to monitor other portions of a patient's anatomy.
[0091] If desired, the center frequency may be lowered (e.g., down
to about 4.5 MHz) to provide additional depth of penetration when
needed (e.g., for very large patients). Although this will also
reduce the resolution, the result may be acceptable when very large
structures are being imaged. Alternatively, the transducer size and
housing diameter may be scaled up in size (e.g., to about 7 mm) if
the reduced resolution results in unusable images.
[0092] Numerous alternative and optional features may be
substituted and added to the above-described embodiment. One
optional feature is digital beamforming using significant
oversampling. For example, if the transducer is operated at 7 MHz,
and the return is digitized at 30.times.frequency, 30.times.7
MHz=210 MHz digitization would be required. That data could then be
downsampled by a factor of five to reduce the number of data points
to a 42 MHz sample. Such downsampling would reduce the noise floor
due to front-end noise by a factor of .apprxeq.5, (i.e., over 2
bits in power). Similarly, downsampling by a factor of 7 would
reduce the noise floor by a factor of 17.
[0093] FIG. 16A depicts the front face of an alternative 2D
transducer 500, which includes a 2D array of active elements 510.
The concepts described herein can also be implemented using this
type of transducer by making appropriate adjustments that will be
apparent to persons skilled in the relevant arts.
[0094] FIG. 16B depicts the front face of another alternative 2D
transducer design that is referred to as a "sparse 2D transducer."
The sparse 2D transducer 600 has a column 610 of "transmit"
elements 611, used for transmitting the ultrasound, and a row 620
of receive elements 621 used for receiving the ultrasound signal.
As shown, there is one element 630 common to both the column 610 of
transmit elements and the row 620 of receive elements. This common
element 630 may be used for transmission, reception, or both. This
transducer design reduces electronic noise by using separate
transmit and receive elements, which eliminates the need for
electronic transmit/receive switches at the elements. The concepts
described herein can also be implemented using this type of
transducer by making appropriate adjustments that will be apparent
to persons skilled in the relevant arts.
[0095] Alternative embodiments of the invention may use fewer
techniques and/or implement those techniques to a lesser extent,
and still maintain the ability to produce an acceptable image. For
example, depending on the other components in the system, it may be
possible to obtain an acceptable image using a 75.degree. sector
width, or even using a 90.degree. sector width. It may also be
possible to obtain an acceptable image using a transducer with an
elevation:transverse aspect ratio of about 2:3 in place of the
preferred 1:1 or 2:1 aspect ratios. Another alternative would be to
use some or all of the above-described techniques with a transducer
that is slightly larger than the preferred embodiments described
above, yet still smaller than conventional 10 mm TEE transducer.
Numerous other modifications to the above-described embodiments
will be apparent to those skilled in the art, and are also included
within the purview of the invention.
* * * * *