Rotary encoder for intravascular ultrasound catheter

An ultrasound system and method for intravascular ultrasonic imaging includes an array of beacons that are fixed to direct ultrasonic energy toward an imaging transducer, with individual beacons being identifiable in order to determine the angular position of the imaging transducer. Based upon the data related to beacon identification, operation of the imaging device is adaptively adjusted in order to compensate for variations in angular velocity of the transducer. Adaptive compensation may be performed by adjusting the pulse repetition rate of transmitted ultrasonic energy, by adjusting the scan conversion algorithm or mapping reflected ultrasonic energy, or by varying control of the drive structure for rotating the transducer. The beacons are preferably piezoelectrically active, but passive beacons may also be used. Position identification may be performed by techniques including amplitude sensing, phase sensing, pulse length sensing, and frequency sensing. As an alternative to rotation of the transducer, ultrasonic signals may be formed at a proximal end of a probe and then conducting the energy to the distal end via a waveguide.

TECHNICAL FIELD 
The invention relates generally to devices and methods for imaging the 
interior of a vessel, such as a blood vessel, and more particularly to 
devices and methods for reducing image distortion due to nonuniform 
rotational velocity of a rotating catheter. 
BACKGROUND ART 
Within the medical field, ultrasound systems are used for various imaging 
and treatment purposes. For example, there is an increasing appreciation 
of the diagnostic value of obtaining cross sectional images of coronary 
arteries by the method of intravascular ultrasound (IVUS). There are 
currently two general types of IVUS catheter systems. In a first type, 
subsets of an array of ultrasound transducers are sequentially excited in 
a manner to electronically steer an ultrasonic beam. This approach is 
sometimes referred to as the synthetic aperture focusing technique (SAFT). 
U.S. Pat. Nos. 4,917,097 to Proudian et al. and 5,186,177 to O'Donnell et 
al. describe use of this approach. 
The second approach in the design of an IVUS catheter system is one in 
which the ultrasonic beam is redirected mechanically, rather than 
electrically. There are three subclasses of this mechanical approach. IVUS 
systems in the first subclass include either a rotating transducer or a 
rotating mirror in the distal end of the catheter that enters the vessel. 
A motor is coupled to the catheter at the proximal end that remains at the 
exterior of the bony. A drive shaft connects the proximal motor to the 
rotating distal transducer or mirror. U.S. Pat. Nos. 4,794,931 and 
5,000,185 to Yock teach this technique. In the second subclass, rotation 
is confined to the distal end. U.S. Pat. Nos. 5,176,141 to Bom et al. and 
5,240,003 to Lancee et al. teach incorporation of a micro motor at the 
distal end for rotating a transducer or a mirror. Alternatively, a 
fluid-driven turbine may be used to rotate the transducer or the mirror, 
as taught by U.S. Pat. No. 5,271,402 to Yeung et al., which is assigned to 
the assignee of the present invention. The third subclass is one in which 
the ultrasonic beam is generated at the proximal end of the catheter and 
is channeled to the distal end via a rotating waveguide, as taught by U.S. 
Pat. No. 5,284,148 to Dias et al., which is assigned to the assignee of 
the present invention. 
The mechanical-rotating approach of directing an ultrasonic beam from a 
distal end of a catheter is more prevalently used than the approach of 
electronically aiming the beam. The mechanical approach can be implemented 
using a single transducer, while the electronic approach requires an array 
of transducers to be contained in the distal end that must pass through a 
vessel, such as a blood vessel. However, one concern in the use of an IVUS 
imaging system in which mechanical rotation is required is that the 
rotational velocity of the rotating structure will be nonuniform. A 
nonuniform rotational velocity will distort the image that is formed. One 
cause of nonuniformity with respect to the rotational velocity is the 
existence of mechanical friction and binding of the catheter as it spins 
in the tortuous path of the coronary arteries. Although the proximal end 
of the catheter is rotating at the desired velocity, any binding of the 
catheter along its length will lead to a distal rotational velocity that 
is different than the desired velocity. Assuming a constant proximal 
velocity from a drive motor variations in the distal velocity are 
typically related to storage of energy in the drive shaft in the form of 
torsion. If the catheter rotates at a velocity that is greater or less 
than the desired rotational velocity, reflected ultrasonic energy that is 
received from a particular location will be portrayed in the resulting 
image as being from an incorrect location. 
With reference to FIG. 1, in the ideal, a catheter 10 is positioned 
coaxially with a vessel 12 for which ultrasonic intravascular images are 
to be formed. The catheter has a diameter that is relatively small 
compared to the diameter of the vessel. In this ideal situation, the 
catheter is rotated at a constant angular velocity, .omega..sub.0. Thus, 
there is a one-to-one correspondence between the anticipated directions of 
ultrasound transmission and reception and the actual directions of 
transmission and reception. Electrical signals generated in response to a 
reception of reflected energy may be accurately collected, processed and 
displayed. 
Rather than a constant rotational velocity, FIG. 1 illustrates a "biphasic 
velocity profile," in which the catheter is rotating too quickly, 
.omega..sub.0 +.DELTA..omega..sub.1, for a portion of each revolution and 
too slowly, .omega..sub.0 -.DELTA..omega..sub.2, for another portion. By 
integrating the rotational velocity as a function of time, it is possible 
to obtain the angular position as a function of time. In FIG. 2, the 
anticipated rotational velocity is a constant, thereby producing a 
straight line 14 having a slope of .omega..sub.0. However, the actual 
biphasic velocity profile with an excessive rotational velocity over a 
first portion of each cycle and a diminished rotational velocity over a 
second portion is shown by plot 16. The difference between the anticipated 
rotational velocity 14 and the actual rotational velocity 16 is shown at 
plot 18, which is referred to as the angular error curve (AEC). The AEC 
determines the degree of error of an image. 
The combination of nonuniform rotational velocity and eccentric catheter 
placement can lead to the distortion of the shape of a vessel wall. As an 
example, FIG. 3 shows an actual vessel wall 20 compared to the image of 
the actual vessel wall when a catheter has a biphasic velocity profile and 
is located eccentrically as illustrated. The dotted lines in FIG. 3 
represent the ultrasound A lines, which are fired either too early or too 
late as a result of the incorrect velocity. For each ultrasound A line 
which is too early or too late, the angular position information to 
imaging equipment is incorrect. On the other hand, the range information 
to the imaging equipment is correct. Thus, the correct range information 
is rotated by the imaging equipment to the anticipated angular orientation 
in order to calculate the portrayal of the vessel wall range segment. The 
distorted image of the vessel wall is then constructed as the spline which 
connects all the "rotated" A lines. FIG. 3 illustrates an overestimation 
of the vessel area, as well as a distortion of the local curvature of the 
wall. Underestimations of the lumen area are also possible with different 
eccentric placements of the catheter. 
Previous attempts to measure the position, or orientation, of the 
transducer have included a fluoroscope marker, as taught by Scribner et 
al. in U.S. Pat. No. 5,054,492, a magnetic resonance imaging-based 
tracking system, as taught by Dumoulin et al. in U.S. Pat. No. 5,271,400 
and an external ultrasound-based system, as taught by Crowley in U.S. Pat. 
No. 5,131,397. While each of these systems operates sufficiently for its 
intended purpose, each system requires bulky and expensive additional 
equipment in order to perform the IVUS imaging of a vessel wall. Moreover, 
none of the systems recognizes the nonuniform rotational velocity problem 
or addresses a solution to the problem. 
U.S. Pat. No. 5,243,988 to Sieben et al. describes the use of markers, 
preferably periodic variations in wall thickness of the catheter sheath, 
as a means for rotary encoding, but concerns exist. First, the numbers 
identified in the patent with regard to the sheath thickness and the 
transducer frequency are such that the walls are likely to be too thin to 
be resolved. Second, the abrupt changes in wall thickness may cause 
distortion of the ultrasound beam, due to refraction, unless the speed of 
sound in the sheath is close to that in water, which is not true for most 
plastics that are used to construct conventional sheaths. The distortion 
of the beam would result in degradation of image quality. Moreover, abrupt 
changes in wall thickness reduce the number of angular positions that can 
be encoded. 
In a thesis entitled "Scanning Mechanisms for Intravascular Ultrasound 
Imaging: A Flexible Approach," Erasmus University, Rotterdam, 1993, by H. 
ten Hoff, non-uniformity of angular velocity was again addressed. 
Acoustic, capacitive, electromagnetic and optical techniques were 
considered as means for angle detection. Correction of the image was then 
performed in one of two ways. According to one method, an angle detection 
signal was used to measure or estimate the traversed angle between two 
successive directions of ultrasonic transmission, emitted at equal time 
intervals. The resulting information was then fed to display processing 
for correctly positioning the corresponding image-lines. H. ten Hoff 
concluded, however, that the tangential resolution became dependent upon 
the rotational angle, which diminished the image quality. The second 
method was one of using the angle detection signal to trigger the emission 
of ultrasonic pulses at equal traversed angle increments, so that 
image-lines were then generated periodically. The paper briefly referred 
to use of acoustic techniques to determine the angle detection signal, but 
focused upon optical determination because of a number of identified 
drawbacks to the acoustic solution. The identified drawbacks included low 
resolution of ultrasonic reflecting structures, multiple reflection, and 
shadowing. 
What is needed is an ultrasound device and method by which the angular 
orientation of a catheter tip can be tracked in real time in order to 
reliably and repeatedly identify the position and properties of a specific 
anatomic structure, such as calcified plaque, and to adaptively correct 
angular velocity profiles which potentially extend over more than one 
rotational cycle. 
SUMMARY OF THE INVENTION 
The above object has been met by an ultrasound device and method that 
utilize beacons that are positioned to either reflect (passive) or 
transmit (active) ultrasonic energy toward a rotating transducer. 
Typically, the beacons are in a field of view of the rotating ultrasound 
transducer used to image anatomic structures. The beacons should be 
configured and positioned to minimize their effect upon an image to be 
formed. Moreover, since the beacon positions are known, it is possible to 
boost the gain of those segments of rf A-lines corresponding to signals 
from structures lying beyond the beacon positions, with the boosting of 
the gain being relative to those segments of the rf A-lines corresponding 
to signals from the beacons. Gain can be adjusted to compensate for 
reduced signal strength due to partial reflection of sound from the 
beacons. In practice, beacons may be placed at the exterior or the 
interior of a sheath that houses the transducer. Each beacon preferably 
has a unique "acoustic signature," thereby allowing each beacon to be 
uniquely identified. This may be achieved by tailoring the acoustic 
properties of each beacon (e.g., the acoustic impedance of a passive 
beacon) and/or by varying with angular position the ultrasound signal used 
to interrogate the beacons. 
With regard to the operation of the beacons, three basic steps are involved 
in overcoming image distortions due to nonuniform rotational velocity. The 
three steps are: directing energy from the beacon; identifying individual 
beacons; and providing adjustments in response to detection of a variation 
in angular velocity. 
Passive beacons, i.e., beacons that redirect rather than transmit acoustic 
energy, are the preferred embodiment, since the use of passive beacons 
requires the least amount of modification to existing IVUS imaging 
systems. As one alternative, the passive beacons may be formed of a 
material that provides a change in acoustic impedance. In an active-beacon 
embodiment, the beacons may be a layer of piezoelectric material between 
two electrode layers. When an electrical signal is formed across the 
electrode layers, the piezoelectric layer transmits an ultrasonic signal 
to the rotating transducer, which then converts the ultrasonic signal back 
to an electrical signal for identification of the individual beacon by 
external imaging equipment. 
The second step involves individually identifying beacons in order to 
permit calculation of angular velocities of the rotating transducer. The 
second and third steps may be performed during a normal imaging sequence 
or during a pre-image calibration sequence. As previously noted, there are 
two approaches to uniquely identifying the beacons. These approaches may 
be used alone or in concert. In the first approach, the beacons themselves 
are physically different and are interrogated with identical ultrasonic 
signals. For example, by varying the reflection coefficient of the 
beacons, it is possible to encode angular position as a function of the 
amplitude of the reflected signal. In the second approach, the beacons are 
physically identical, but are distinguished from one another by varying 
the ultrasonic signal that they reflect (passive) or broadcast (active). 
One example of this second approach involves passive beacons and a 
variation in the number of cycles per burst (pulse length) transmitted by 
the rotating transducer, with the variation being a function of the 
anticipated angular position. By counting the number of cycles in the 
signal reflected from each beacon, the actual angular position can be 
determined and compared to the anticipated angular position in order to 
detect errors in the angular velocity. Angular velocity errors 
encompassing more than one revolution can be detected by monitoring the 
number of cycles returned from each beacon over a number of revolutions. 
Alternate embodiments are contemplated, wherein other characteristics of 
the ultrasonic signal are used to encode angular position, such as the use 
of the characteristics of amplitude, phase and frequency. The two 
approaches apply equally to utilization of passive beacons and active 
beacons. 
As yet another alternative to actually varying the characteristics of 
ultrasonic signals propagating between the beacons and the imaging 
transducer, identical ultrasonic signals may be received at the rotating 
transducer, but then the signals may be processed in a manner to determine 
beacon identity. As an example of this alternative, the ultrasonic signals 
received at the transducer from the beacons are converted into electrical 
signals and referenced to an internal electrical signal having a 
characteristic that varies with the anticipated angular position of the 
transducer. The reference signal may increase or decrease in frequency for 
a multiple of the anticipated period of a transducer revolution. By mixing 
the reference and beacon signals, the mixed signal may be monitored to 
detect differences between the actual mixed frequency and the frequency 
that is anticipated when the transducer is rotated at the constant 
selected angular velocity. 
Velocity error can then be determined using the acquired data from the 
operation of identifying individual beacons. For example, an angular error 
curve 18 of FIG. 2 may be calculated by comparing the anticipated angular 
position with the actual angular position as a function of time. 
The third operation identified above is the correction of a calculated 
error. In a first embodiment of adaptive adjustment, the repetition rate 
of ultrasound pulses from the rotating transducer is varied to achieve 
uniform angular increments between adjacent radio frequency A lines, 
regardless of rotational velocity. That is, the variations in rotational 
velocity remain uncorrected, but the pulse repetition rate varies to 
provide compensation. In a second embodiment, the scan conversion 
algorithm of external imaging equipment is affected such that while the 
angular sampling by the transducer may be nonuniform, the imaging 
equipment recognizes the nonuniformity and provides compensation at the 
imaging level. In the third embodiment, the rotational drive structure is 
controlled for the purpose of reducing the nonuniformity. For example, 
current to a drive motor may be increased when the transducer has a 
tendency to slow down and the current may be decreased when the transducer 
has a tendency of speeding up. 
An advantage of the invention is that the angular orientation of the 
transducer can be tracked throughout rotation without a significant 
increase in the cost of intravascular ultrasound imaging equipment. A more 
reliable cross sectional image of a blood vessel is thus achieved.

BEST MODE FOR CARRYING OUT THE INVENTION 
With reference to FIG. 4, in the ideal, ultrasonic energy is directed from 
a center 24 of a catheter for radiation in a rotary pattern. Ten lines 26, 
numbered 1-10, are shown in FIG. 4 as being equally spaced over a 
90.degree. sector of the catheter. All lines pass outwardly beyond a 
sheath 28 that houses the ultrasound transducer for generating the lines 
in response to an electrical signal. 
Line 1 contacts a first beacon 30. Line 10 is directed at a second beacon 
32. A third beacon 34 and a fourth beacon 36 are also shown in FIG. 4. 
The earliest arriving signals for lines 26 are representative of energy 
partially reflected back to the center 24 of the catheter. Reflected 
ultrasonic energy from the beacons is converted to an electrical signal by 
the transducer and is used to calibrate the rotational velocity. Later 
arriving signals on these lines, and all others, represent scattering from 
structures beyond the catheter and are used to form an image. As will be 
explained more fully below, a device and method are required to monitor 
early time segments of each line to detect whether or not a line is 
bearing on one of the plurality of beacons. This information can then be 
used to calibrate the rotational velocity of the catheter tip. 
Typically, the number of lines 26 within a 90.degree. sector greatly 
exceeds the number illustrated in FIG. 4. Likewise, there are preferably 
more than four beacons 30-36. The reduced numbers of lines and beacons are 
selected for the purpose of illustration. 
In the imaging of a vessel, a frame of data may be collected in 32,400 
.mu.sec. That is, a rotary pattern of ultrasound transmission is completed 
within this time period. This yields a frame rate of approximately 31 
frames/sec. In FIGS. 4 and 5, each frame includes 36 RF A-lines, so that 
reflected A-lines occur every ten degrees. In the ideal uniform rotational 
velocity of FIG. 4, this requires that the transducer be fired every 900 
.mu.sec. Assuming a master clock running at 1 MHz, the transducer is fired 
every 900th cycle of the master clock. Upon completion of a 90.degree. 
rotation, line 10 is fired and data collection begins for the second 
90.degree. sector. 
Data collection for the ten A-lines is shown in FIG. 5. Since the 
transducer-beacon separation is fixed and is known, it can be expected 
that for those A-lines in which the transducer is bearing on a beacon, a 
reflected echo will be received at the transducer at a predetermined time, 
with the number of these fixed-time beacon echoes that are received within 
each frame being equal to the number of beacons. Thus, a "watch" period 
early in each A-line can be established, during which the system "looks" 
for an echo from a beacon. For example, in FIG. 5 the watch periods are 
shown within dashed lines, and watch periods 38 and 40 of lines 1 and 10, 
respectively, are shown as including echoes reflected from the first and 
second beacons 30 and 32 back to the transducer. Because the echoes from 
beacons 30 and 32 were received on those A-lines for which echoes are to 
be anticipated when the rotational velocity is fixed at the desired 
rotational velocity, velocity correction is not required. 
The reflected energy of A-line 1 during the watch period 38 identifies the 
angular position of the transmitter, since the position of the beacon is 
fixed and known. In the same manner, the angular position of the 
transducer during reception of the A-lines 10 is identifiable, since the 
beacon-reflected ultrasonic energy is received during watch period 40. 
Referring now to FIG. 6, ten A-lines 42 are transmitted and received during 
the same time sequence described above, but with the catheter rotating at 
an excessive speed. That is, the pulse repetition rate (PRR) is the same 
as that of FIG. 4, but the actual rotational velocity is greater than the 
anticipated rotational velocity. As a result, there is an angular 
under-sampling of the 90.degree. sector defined by the first and second 
beacons 30 and 32. The A-line 8 is extended along the radius intersecting 
the second beacon 32, while A-lines 9 and 10 extend into the second 
90.degree. sector of the rotational pattern. On the other hand, the scan 
conversion algorithm of the external imaging equipment continues to map 
the ten A-lines through a 90.degree. sector, rather than the actual 
115.7.degree. that the catheter has rotated during 8100 .mu.sec. 
The ten A-lines of FIG. 6 are represented in FIG. 7. The A-lines 1 includes 
transmitted energy 44 and reflected energy during a watch period 46. The 
watch period is defined by the known time required to receive reflected 
energy from a beacon that is spaced from a source by a fixed distance. 
With the catheter rotating at a speed greater than the anticipated speed, 
the watch period 48 of A-lines 10 does not include reflected energy from a 
beacon. Instead, transmitted energy 50 of A-lines 8 is reflected and 
received during a watch period 52. As will be explained more fully below, 
compensation for the difference between actual and anticipated rotational 
velocity is designed to reduce imaging distortions that would otherwise 
result. 
The rotational velocity represented by FIGS. 6 and 7 is excessive by a 
factor of 115.7.degree./90.degree., i.e. approximately 1.2856. One form of 
compensation would be to increase the PRR by the same factor. That is, the 
time between transmission of adjacent A-lines 42 must be decreased by the 
reciprocal of 1.2856. If n is the number of clock cycles between bursts, 
i.e. pulses, for the "corrected" PRR, then 90/115.7=n/900, so that n is 
equal to approximately 700. By decreasing the number of clock cycles 
between transmit pulses, it is possible to once again collect nine 
equally-spaced RF lines in a 90.degree. sector, despite the fact that the 
catheter is rotating at a faster rotational velocity than anticipated by 
the external imaging equipment. In another embodiment, the PRR remains 
uncorrected, and it is the scan conversion algorithm of the external 
imaging equipment that is affected in order to compensate for the 
difference between anticipated and actual rotational velocity. Changing 
the scan conversion algorithm in response to detected variations in 
angular velocity would be within the skill of a person within the art. 
However, the concern with this compensation approach is that resolution of 
imaged anatomical structures may be adversely affected in some cases. A 
third approach of providing compensation involves adaptively controlling 
mechanical drive of the catheter. For example, the electrical current to a 
rotational drive motor may be varied in response to detected velocity 
variations. While this approach will significantly reduce distortions, 
there are concerns that compensation will be hindered by the inertial 
effect of the catheter and other driven structure. 
Referring now to FIGS. 8 and 9, a catheter 54 is shown as having a 
transducer 56 mounted to a rotor 58. The transducer-and-rotor arrangement 
may be of any type known in the art. For example, the transducer 56 may be 
a piezoelectric layer sandwiched between electrode layers for conversion 
between electrical energy and ultrasonic energy. 
On the exterior of the catheter 54 is an array of eight passive beacons 60, 
62, 64, 66, 68, 70, 72 and 74. The only strict requirement of a beacon is 
that its position relative to the catheter be known. The externally 
mounted beacons are in the field of view of the ultrasound transducer 56. 
The beacons are mounted on an acoustic window 76 through which ultrasonic 
energy is transmitted and received. 
In the passive mode, the beacons 60-74 act as reflectors for ultrasound 
being transmitted by the catheter 54. It is important that the passive 
beacons be sufficiently reflective to appear as "bright spots" on an 
image, but not so bright that the spots shadow points in the field behind 
them. The passive beacons provide zones of varying acoustic impedance, so 
as to yield corresponding changes in reflectivity. Passive beacons may be 
formed of evaporated, sputtered or electroplated metal film, but this is 
not critical. Alternatively, an integrated impedance matching layer can be 
formed at the surface of the sheath. This might, for example, involve 
forming microgrooves in the sheath material and backfilling these 
microgrooves with a material that has a different acoustic impedance. The 
acoustic impedance and/or velocity of sound through this integrated 
impedance matching layer can be varied spatially by spatially varying the 
volume fraction of microgrooves in the sheath. 
The beacons 60-74 permit identification of the angular position of the 
catheter 54, but should not introduce artifacts that adversely affect 
imaging of anatomical structures. In software, the gain profiles of the 
A-lines that must "pass through" a beacon can be adjusted, so that 
shadowing of tissue behind the beacon can be overcome. Another alternative 
is to form the beacons to be acoustically transparent at the imaging 
frequency, but reflective at some other frequency. For example, a thin 
metal film may have a thickness that is a substantial fraction of a 
wavelength at a high frequency, but substantially acoustically transparent 
at a lower frequency. The catheter 54 may have two different transducers 
at the distal end. However, this technique is more easily implemented in 
use with an acoustic waveguide, wherein ultrasonic energy is generated at 
the proximal end and channeled to the distal end using waveguide 
principles. An intravascular imaging device having an acoustic waveguide 
is described in U.S. Pat. No. 5,284,148 to Dias et al., which is assigned 
to the assignee of the present invention. The acoustic waveguide works 
equally well with multiple frequencies. 
Another approach for minimizing artifacts produced by the beacons 60-74 is 
to use a single transducer which is rotated at one angle during imaging 
and at a second angle during calculation of angular velocity. At the 
imaging angle, the beacons are outside of the field of view of the single 
transducer. Alternatively, the single transducer may be repositioned 
during determination of angular velocity. In FIG. 8, a beacon 78 is shown 
as being mounted on the interior of the sheath wall 80. The rotor 58 and 
transducer 56 may be aligned with the beacon 78 during calculation of an 
angular error curve, and then aligned with the acoustic window 76 during 
an imaging procedure. In this embodiment, beacon 78 is one of an array of 
beacons and the beacons 64-74 are omitted. 
Referring now to FIG. 10, in an alternate embodiment, the distal end 82 of 
the device includes piezoelectrically active transducers. Thus, small 
transducers in the sheath wall generate ultrasonic energy in the direction 
of the rotating transducer 84. The rotating transducer is mounted on a 
drive shaft 86. The beacons include a continuous outer ground electrode 88 
and a continuous piezoelectric layer 90. The positions of the active 
beacons are defined by a pattern of inner electrodes 92, 94, 96 and 98. 
When an electrical signal is connected across ground electrode 88 and the 
patterned electrodes 92-98, the areas of the piezoelectric layer 90 
between the patterned electrodes and the ground electrode generate 
acoustic energy corresponding to the applied signal. In operation, the 
active beacons may be continuously transmitting as the imaging transducer 
84 rotates. The imaging transducer receives the signals from the four 
active beacons, and the external imaging equipment identifies the 
individual beacons in order to determine the angular position of the 
imaging transducer. Typically, the device includes more than four active 
beacons. 
The signals of the active beacons may be identical, with the external 
imaging equipment being operated to distinguish the beacons using one of 
the methods described below. Alternatively, the characteristics of beacon 
signals may be varied, such that the source of a beacon signal may be 
identified by analysis of the signal. For example, the amplitudes of the 
electrical signals to the patterned electrodes 92-98 may increase from 
electrode to electrode. As a result, the amplitudes of the generated 
beacon signals will increase correspondingly. Alternatively, the beacons 
may vary with respect to transmission frequency. However, this typically 
requires that the piezoelectric material vary in thickness from beacon to 
beacon, since the resonant frequency of a piezoelectric layer changes with 
the thickness of the layer. Another signal characteristic that can be 
varied is phase. Delays in the application of an excitation signal from 
one electrode 92-98 to the next electrode may be used as a method of 
providing phase-shifted, beacon-identifying signals. 
The outer covering 100 of the distal end 82 may be formed of nylon, while 
an inner shell 102 may be formed of polyethylene. The patterned electrodes 
92, 94, 96 and 98 may be formed of gold on the surface of the 
piezoelectric layer 90. The continuous ground electrode 88 may also be 
gold. A PVDF copolymer may be used as the piezoelectric layer 90. However, 
none of these materials is critical to the invention. 
Using techniques known in the art, the regions of the continuous 
piezoelectric layer 90 associated with the patterned electrodes 92-98 are 
made active by poling, while regions adjacent to the patterned electrodes 
remain piezoelectrically passive. Wiring to the electrodes may extend from 
the proximal end of the device, but it is possible to house a signal 
generator within the distal end 82 for excitation of the active beacons. 
In another embodiment, the active beacons associated with the patterned 
electrodes 92-98 are used as receptors, rather than transmitters. In this 
embodiment, the angular position of the imaging transducer 84 is 
determined by monitoring which active beacon is receiving acoustic energy 
from the imaging transducer. 
Referring now to FIGS. 4 and 11, identification of individual beacons 30, 
32, 34 and 36 may be accomplished by a frequency detection method. An 
ultrasonic signal is transmitted from the center 24 of the catheter. A 
beacon return burst 106, 108, 110 and 112 is directed toward the center 24 
from each of the four passive beacons. The return bursts are received by 
an imaging transducer in sequence as the imaging transducer is rotated. A 
reference signal 114, referred to herein as a "chirped" signal, is mixed 
with the return bursts. The reference signal sweeps in frequency from a 
starting frequency to a peak frequency in the time of one period of 
nominal rotation, i.e. the anticipated time for completing one transducer 
rotation. The mixing of the return bursts with the reference signal 114 is 
carried out to provide four mixed signals 116, 118, 120 and 122. The mixed 
signals increase in frequency. A graph of the change in frequency to the 
change in angular position of the imaging transducer will yield a straight 
line when the rotational velocity of the transducer is constant. 
If the imaging transducer has an actual rotational velocity that exceeds 
the anticipated rotational velocity between the first beacon 30 and the 
second beacon 32, as represented in FIG. 6, and if the actual rotational 
velocity is below the anticipated velocity with travel from the second 
beacon to the fourth beacon 36, operation will be as shown in FIG. 12. The 
return bursts 124, 126, 128 and 130 are identical to those of FIG. 11, 
other than with regard to timing. The excessive speed causes the return 
bursts 126 from the second beacon 32 to be too closely spaced to the first 
return burst 124. In comparison, the time between other return bursts is 
excessive. The reference signal 114 increases in frequency in the same 
manner described above. The mixed signals 132, 134, 136 and 138 formed by 
referencing the four return bursts with the chirped signal are not 
proportional with regard to increases in frequency. Consequently, a plot 
of frequency change to angular position of the imaging transducer will not 
be a straight line. The difference between the actual plot and the 
anticipated straight-line plot is an angular error curve (AEC) that can be 
employed to provide compensation using one of the approaches described 
above, i.e. PRR adjustment, A-lines mapping adjustment, or drive 
adjustment. Other compensation techniques may also be utilized. 
The "interrogation" of return bursts 106-112 and 124-130 by means of a 
reference signal 114 in order to individually identify beacons has been 
described with regard to correcting errors encountered over a single 
revolution of the catheter. However, the technique may be adapted to 
detect variations in angular velocity that are not restricted to a single 
rotation. The duration of the frequency sweep of the reference signal 114 
may be extended beyond the period of nominal rotation. 
Another method of individually identifying the four beacons 30, 32, 34 and 
36 of FIG. 4 is described with reference to FIG. 13. In this embodiment, 
the method is based on broadcasting bursts 26 having a varying number of 
cycles. The number of cycles in a burst varies as a function of angular 
position. For example, a 3.5 French catheter has a transducer-to-sheath 
inner wall spacing of approximately 450 .mu.m. The wavelength in water at 
30 MHz is approximately 50 .mu.m. Thus, a maximum of approximately 18 
cycles can be transmitted, reflected at a beacon, and received by the 
imaging transducer. As an example, it will be presumed that there are 16 
broadcasted bursts, with the first burst having three cycles and each 
successive burst increasing by an increment of one cycle, so that the last 
burst of the sequence has a total of 16 cycles. If the angular velocity of 
the catheter is uniform and correct, the imaging transducer will be in the 
anticipated angular position with each return burst. In FIG. 13, the first 
return burst 140 includes the three cycles of the first broadcasted burst. 
A seven-cycle return burst 142 will be directed from the second beacon 32, 
an eleven-cycle return burst 144 will be directed from the third beacon 
34, and a fifteen-cycle return burst 146 will be directed from the fourth 
beacon 36. On the other hand, if the catheter is rotating too rapidly or 
too slowly, return bursts will have an insufficient number or an excessive 
number of cycles, respectively. The cycle count for each return burst can 
be monitored with a zero-crossing detector working in conjunction with a 
threshold detector. In order to implement this technique, it would be 
preferable to operate the imaging transducer at a frequency greater than 
30 MHz, so that the number of cycles per angular increment can be 
increased. Angular velocity changes that occur over more than one 
revolution of the catheter can be corrected by monitoring the number of 
cycles returned from each beacon over more than one revolution period. 
The cycle-count approach to correcting non-uniform angular velocity may be 
applied by the system of FIG. 14. Signals from the transducer 148 during 
an AEC calibration operation are passed to a receiver 150. A time gain 
control (TGC) 152 is set to attenuate signals beyond the desired beacon 
range. A beacon calibration module 154 passes the A-lines to an AEC module 
156 that calculates the angular error curve and outputs the calculations 
to a line rate correction module 158. The line rate correction module is 
capable of providing compensation in any of the three approaches described 
above. Thus, depending upon the position of a switch 160, the module 158 
will either (a) vary current to a drive motor 162 in order to reduce the 
velocity non-uniformities, (b) vary the scan converter algorithm 164 for 
mapping A-lines received at a signal processing module for accurate 
display 168 of the nonuniformly transmitted and received A-lines, or (c) 
vary the PRR as determined at a line rate clock 170, so that bursts are 
transmitted 172 at uniform spacing even though the angular velocity is 
nonuniform. 
Other embodiments of the invention involving phase shifting have been 
contemplated. For example, passive or active beacons may be spaced apart 
from the imaging transducer by varying distances, so that the arrival time 
of return bursts may be used to identify individual beacons. This would 
effectively track one wall of the sheath. In an alternative embodiment, 
the wall thickness could be varied gradually going from a minimum to a 
maximum in one revolution. The variation in wall thickness could be 
measured using a higher frequency transducer mounted near the imaging 
transducer. Measured wall thickness would then uniquely identify 
individual lines. 
In addition to acoustical encoding, optical, electrical and/or magnetic 
encoding may be employed. One advantage of the acoustical approach is that 
much of the required hardware is already in place in existing 
intravascular image systems. Nevertheless, other approaches are possible. 
Optical encoding may be performed by substituting light for the ultrasound 
and allowing the light to reflect off of reflectors located about the 
inner circumference of the catheter sheath. Electrical encoding may be 
performed by use of a commutator-type arrangement, wherein a signal is 
alternatingly switched on and off as a brush momentarily contacts 
conductive fingers placed on the inner circumference of the catheter 
sheath. Magnetic encoding may be performed by fixing a small magnetic 
member to the catheter so as to spin with the transducer. A single 
conductive loop with multiple "teeth," or a series of conductive loops, 
may be placed in the catheter wall about its circumference. The changing 
flux as the magnet is spun induces a voltage in the loop or loops, which 
can be picked up and used to determine angular velocity. 
While the acoustical approach has been described and illustrated as being 
implemented with devices in which the transducer is located at the distal 
end of the probe, this is not critical. The techniques apply equally to 
embodiments in which ultrasonic energy is conducted to the distal end via 
an acoustic waveguide, as taught in U.S. Pat. No. 5,284,184 to Dias et al. 
The rotary pattern of ultrasound transmission can then be formed by 
rotating the end of the waveguide or by rotating a reflector.