Guidance and delivery system for high-energy pulsed laser light

A deflectable guidewire for a catheter includes a core extending from the proximal end of the guidewire toward the distal end thereof; a hollow sheath for encasing the core; the hollow core including a deflecting region at the distal end of the guidewire, the deflecting region having a first side and a second side, the first side being on an opposite side of a plane extending axially through the hollow means than the second side; the first side of the deflecting region being able to be compressed a greater distance than the second side of the deflecting region; the core being fixed to the hollow sheath at the distal side of the deflecting region; and a control handle mounted at the proximal end of the core and the hollow sheath for applying tension to the core with respect to the hollow sheath so that the tension causes the first side of the deflecting region to compress a greater amount than the second side, thus deflecting the deflecting region. In one embodiment, the core may include laser conducting optical fibers. Other features of the present invention include arranging the optical fibers in the catheter in such a way that bending stresses are distributed evenly throughout the fibers when the catheter is bent. One method of doing this is to arrange the fibers in rows or layers wherein each row or layer is twisted in a direction that is opposite to the direction of the adjacent rows or layers. It has also been found that flexibility is increased without unacceptable transmission losses by using optical fibers having a diameter of about fifty microns.

BACKGROUND OF THE INVENTION 
The present invention is directed to a system for delivering high energy 
laser light by means of an optical waveguide, and in one particular 
application is concerned with laser angioplasty and a means for guiding 
such a system. 
The use of laser energy to ablate atherosclerotic plaque that forms an 
obstruction in a blood vessel is presently being investigated as a viable 
alternative to coronary bypass surgery. This procedure, known as 
angioplasty, essentially involves insertion of a fiber optic waveguide 
into the vessel, and conduction of laser energy through the waveguide to 
direct it at the plaque once the distal end of the waveguide is positioned 
adjacent the obstruction. In certain uses, to enable the physician to 
ascertain the location of the waveguide as it is being moved through the 
vessel, additional waveguides for providing a source of illuminating light 
and for conducting the image from inside the vessel to the physician are 
fed together with the laser waveguide. 
Most of the experimentation and testing that has been done in this area has 
utilized continuous wave laser energy, such as that produced by Argon Ion, 
Nd:YAG or Carbon Dioxide lasers. The light produced by this type of laser 
is at a relatively low energy level. Ablation of the obstruction is 
achieved with these types of lasers by heating the plaque with constant 
laser power over a period of time until the temperature is great enough to 
destroy it. 
While the use of continuous wave laser energy has been found to be 
sufficient to ablate an obstruction, it is not without its drawbacks. Most 
significantly, the destruction of the lesion is uncontrolled and is 
accompanied by thermal injury to the vessel walls immediately adjacent the 
obstruction. In an effort to avoid such thermal injury and to provide 
better control of the tissue removal, the use of a different, higher level 
form of laser energy having a wavelength in the ultra-violet range (40-400 
nanometers) has been suggested. See, for example, International Patent 
Application PCT/US84/02000, published Jun. 20, 1985. One example of a 
laser for producing this higher level energy is known as the Excimer 
laser, which employs a laser medium such as argon-chloride having a 
wavelength of 193 nanometers, krypton-chloride (222 nm), krypton-fluoride 
(248 nm), xenon-chloride (308 nm) or xenon-fluorine (351 nm). The light 
produced by this type of laser appears in short bursts or pulses that 
typically last in the range of ten to hundreds of nanoseconds and have a 
high peak energy level, for example as much as 200 mJ. 
Although the destruction mechanism involving this form of energy is not 
completely understood, it has been observed that each single pulse of the 
Excimer laser produces an incision which destroys the target tissue 
without accompanying thermal injury to the surrounding area. This result 
has been theorized to be due to either or both of two phenomena. The 
delivery of the short duration, high energy pulses may vaporize the 
material so rapidly that heat transfer to the nonirradiated adjacent 
tissue is minimal. Alternatively, or in addition, ultraviolet photons 
absorbed in the organic material might disrupt molecular bonds to remove 
tissue by photochemical rather than thermal mechanisms. 
While the high peak energy provided by Excimer and other pulsed lasers has 
been shown to provide improved results with regard to the ablation of 
atherosclerotic plaque, this characteristic of the energy also presents a 
serious practical problem. Typically, to couple a large-diameter laser 
beam into a smaller diameter fiber, the fiber input end is ground and 
polished to an optical grade flat surface. Residual impurities from the 
polishing compound and small scratches on the surface absorb the laser 
energy. These small imperfections result in localized expansion at the 
surface of the fiber when the laser energy is absorbed. The high-energy 
Excimer laser pulses contribute to high shear stresses which destroy the 
integrity of the fiber surface. Continued application of the laser energy 
causes a deep crater to be formed inside the fiber. Thus, it is not 
possible to deliver a laser pulse having sufficient energy to ablate 
tissue in vivo using a conventional system designed for continuous wave 
laser energy. 
This problem associated with the delivery of high energy laser pulses is 
particularly exacerbated in the field of coronary angioplasty because of 
the small diameter optical fibers that must be used. For example, a 
coronary artery typically has an internal diameter of two millimeters or 
less. Accordingly, the total external diameter of the angioplasty system 
must be below two millimeters. If this system is composed of three 
separate optical fibers arranged adjacent one another, it will be 
appreciated that each individual fiber must be quite small in 
cross-sectional area. 
A critical parameter with regard to the destruction of an optical fiber is 
the density of the energy that is presented to the end of the fiber. In 
order to successfully deliver the laser energy, the energy density must be 
maintained below the destruction threshold of the fiber. Thus, it will be 
appreciated that fibers having a small cross-sectional area, such as those 
used in angioplasty, can conduct only a limited amount of energy if the 
density level is maintained below the threshold value. This limited amount 
of energy may not be sufficient to efficiently ablate the obstructing 
tissue or plaque without thermal damage. 
Even if the energy density is quite high, the small beam that results from 
the small diameter fiber may not have a sufficiently large target area 
that effective ablation of the lesion results. Only a small fragment of 
the lesion might be ablated, and thus not provide adequate relief from the 
blockage. A further problem with the use of a fiber optic waveguide to 
direct laser energy for purposes of ablating atherosclerotic plaque is 
that of perforation of the blood vessel. Such perforations can be caused 
by the waveguide itself contacting and perforating the vessel. Such 
perforations can also be caused by the laser beam, particularly if the 
waveguide is not aligned properly within the blood vessel. The perforation 
problems are related to the intrinsic stiffness of the glass fibers of the 
waveguide and poor control of laser energy, regardless of laser source or 
wavelength. 
Also related to the stiffness of the glass fibers is the ability to control 
the position of the fibers radially within the blood vessels. The 
conventional systems employing fiber optic waveguides within a blood 
vessel do not provide means for controlling radial movement within the 
blood vessel. 
One known attempt at developing an angioplasty catheter is disclosed in 
U.S. Pat. No. 4,747,405. The known catheter includes a center guidewire 
lumen, a guidewire therein, and a single optical fiber disposed at a side 
of the catheter for emitting laser energy. The catheter also has a blunt 
leading end that does not facilitate progress through a blood vessel. A 
particular problem that potentially results from the disclosed arrangement 
of the single optical fiber and guidewire is that large segments of the 
lesion may become loose in the blood stream and could possibly cause an 
emboli. As a result, the known catheter includes a dedicated channel to 
remove the loosened debris. 
Currently "over the wire" laser catheters are guided through coronary 
arteries with nondeflectable guidewires. In a laser angioplasty procedure, 
the physician first introduces the guidewire into the blood vessel where 
treatment is required. The catheter is then tracked over the guidewire and 
the laser is activated as required to treat the vessel. Because plaque is 
frequently built up eccentrically within the vessel, it is frequently 
desirable during a procedure to change the general direction of the 
guidewire, and thus the path of the catheter within the vessel. With a 
nondeflectable guidewire, it can be difficult to change the direction of 
the guidewire. 
Laser angioplasty systems for the treatment of nontotal occlusions include 
systems that have a bundle of fibers placed concentrically around a 
guidewire lumen, and which are covered in a protective sheath. Small 
diameter catheters, e.g., 1.0 to 1.7 mm O.D., have a ring of active 
optical fibers that extends from the center guidewire lumen to the outer 
protective sheath. Any occlusive material confronted by the advancing 
catheter will be removed by the active fiber ring. 
These small diameter catheters contain fewer fibers than large diameter 
catheters, e.g., 1.8 mm O.D. and greater. With the increased number of 
optical fibers in the large diameter catheters, catheter flexibility is 
reduced, thus making the catheter more difficult to use. However, large 
arteries require a large diameter catheter. 
Efforts to guide catheters by means of ultrasonic transducers are taught in 
U.S. Pat. Nos. 4,576,177; 4,821,731; and 4,862,893; as well as in an 
article titled Intraluminal Ultrasound Guidance of Transverse Laser 
Coronary Atherectomy by M. A. Martinelli et al. 
OBJECTS AND BRIEF STATEMENT OF THE INVENTION 
Accordingly, it is a general object of the invention to provide a novel 
system for delivering high energy pulsed laser light using an optical 
waveguide. 
It is a more specific object of the invention to provide such a delivery 
system that is particularly well suited to deliver ultraviolet laser 
energy in vivo for the ablation of atherosclerotic plaque. In this regard, 
it is a particular object of the present invention to provide a highly 
efficient waveguide for use in such a delivery system. 
It is yet another object of the present invention to provide such a 
delivery system that is adapted to minimize the likelihood of perforating 
or otherwise damaging a blood vessel in which the system is being used. 
It is a further object of the present invention to provide such a system 
that includes a guide for facilitating the maneuvering of the optical 
waveguides through the blood vessel in which the system is being used. 
It is another object of the present invention to provide a device for 
controlling the radial movement of the optical waveguide within the blood 
vessel in which the system is being used. 
It is another object of the present invention to overcome the 
aforementioned problems associated with the use of nondeflectable 
guidewires. In this regard, one aspect of a delivery system embodying the 
present invention relates to a guidance system that facilitates guiding an 
optical fiber system through a blood vessel. In one embodiment, the 
guidance system comprises a guidewire in which the distal tip thereof may 
be deflected by a mechanism that is controlled by an operator at the 
proximal end thereof. Such a mechanism includes a wire coil surrounding a 
center core. At some position along the coil, preferably the distal end, 
the wire is shaped in a nonuniform manner such that the diameter or width 
of the wire in the coil is smaller on one side of the guidewire than it is 
on the other side of the guidewire. The center core is fixed to the wire 
coil at the distal end of the nonuniform portion. At the proximal end of 
the guidewire, a deflection handle is provided that is able to provide 
tension on the core with respect to the wire coil so as to cause the wire 
coil to compress unevenly at the nonuniform section thereof. The uneven 
compression causes the coil, and thus the distal end of the guidewire to 
deflect. 
In another embodiment, the above-described mechanism for deflecting a 
guidewire is used in conjunction with a fiber optic catheter as the core. 
An ultrasound system is also used to facilitate guiding the catheter. In 
such an embodiment, the wire coil, which includes an uneven portion 
therein, is positioned about a fiber optic bundle. Ultrasound sensors are 
provided at the distal end of the system for providing information to the 
operator. In a further variation of this embodiment, a radiopaque band, 
which functions as an x-ray opaque marker, is used with the ultrasound 
sensors. 
In another embodiment of the present invention, the flexibility of large 
diameter catheters is enhanced by the utilization of a specially designed 
catheter and guidewire. The catheter comprises a central lumen for 
accommodating a guidewire. A concentric lumen surrounds the guidewire 
lumen and includes a plurality of optical fibers for carrying laser energy 
to the distal end of the catheter. In order to increase the effective 
ablating area of the catheter, while minimizing ischemia, the catheter has 
an expanded distal tip portion that increases the effective outer diameter 
of the ring of active optical fibers. However, in order to reduce the 
number of optical fibers needed to fill the ring of optical fibers, and 
thus increase the flexibility of the catheter, the guidewire lumen is 
similarly expanded at the distal tip. The guidewire also has an enlarged 
distal tip that fits within the expanded portion of the guidewire lumen. 
Such an arrangement enables the guidewire to center the catheter with 
respect to an occlusion, keeps the distal tip of the catheter collinear to 
the occluded vessel, and displaces occlusive material in an outward radial 
direction as it is advanced. 
Other features of the present invention include arranging the optical 
fibers in the catheter in such a way that bending stresses are distributed 
evenly throughout the fibers when the catheter is bent. One method of 
doing this is to arrange the fibers in rows or layers wherein each row or 
layer is twisted in a direction that is opposite to the direction of the 
adjacent rows or layers. 
It has also been found that flexibility is increased without jeopardizing 
transmission losses by using optical fibers having a diameter of about 
fifty microns. 
In yet another embodiment of the present invention, an eccentric guidewire 
lumen includes a fountain pen-like tip that forms a leading point of the 
catheter. The optical fibers are also arranged in a manner that conforms 
to the shape of the tip. 
In another embodiment of the present invention, a balloon is used to lock 
an outer catheter to a vessel wall while an inner catheter advances 
through the vessel to ablate a lesion therein.

DESCRIPTION OF THE ILLUSTRATED EMBODIMENTS 
In the following specification, a laser delivery system is described with 
particular reference to the use of a high energy pulsed laser, such as an 
Excimer laser, in an angioplasty system, to facilitate an understanding of 
the invention and its uses. 
Referring now to FIG. 1, one embodiment of the delivery system for high 
energy pulsed laser light is illustrated in greater detail. The delivery 
system comprises two basic elements. One of these is the optical fiber 12, 
and the other is the energy coupler 30. A fiber that is particularly 
suitable for use in the delivery of high energy pulsed ultraviolet laser 
light is a multi-mode fiber which has a relatively large core, or active 
area, relative to the area of its cladding, i.e., the outer skin of the 
fiber. The core is made of substantially pure synthetic fused silica, 
i.e., amorphous silicon dioxide. This material preferably has a metallic 
impurity content of no more than 30 parts per million, to provide better 
conduction of the transmitted laser energy than that which is obtainable 
with natural fused quartz. The term "metallic impurity" includes both, 
metals per se and oxides thereof. 
Even with such a low level of metallic impurity, defects in the silica 
fiber can serve as linear and non-linear absorption sites for the photons. 
These defects can vary from oxygen vacancy to unbonded silicon atoms found 
in any silica glass. They can result in lowered transmittance of 
ultraviolet radiation. Increasing the intensity (or energy) of the laser 
light that is introduced into one end of a fiber exhibiting such defects 
will not necessarily result in proportionally increased output at the 
other end. Rather, the increased intensity level can reduce the threshold 
level at which bulk damage occurs to the silica glass, and thereby 
destroys the delivery system. 
In accordance with one aspect of the present invention, the transmittance 
of high energy UV laser light in a fiber made of synthetic silica is 
enhanced by lightly doping the silica with a material which functions to 
repair some of the inherent structural defects of the silica. The silica 
is preferably doped with an OH-- radical, to thereby form so-called "wet" 
silica. It is believed that defects in silica that affect UV light 
transmission comprise oxygen hole centers and unbonded silica atoms. It is 
theorized that the doping of the silica with the OH- radical functions to 
repair these defects by eliminating the oxygen holes or vacancies in one 
case and by bonding to the silicon to form the SiO.sub.2 double bond. It 
has been reported that pure silica having only about 5 parts per million 
(ppM) of an OH radical has an absorption coefficient which is 2-3 times 
greater than silica having about 1200 ppM of the radical. See J. H. 
Stathes et al, Physical Review B., Vol. 29, 12, 1984, pp. 70-79. Other 
investigations have reported that an optical absorption band appears in 
silica fibers having a low OH-- content as a result of the fiber drawing 
process. See Kaiser et al, J. Opt. Soc. Am., 63, 1973, p. 1141 and J. Opt. 
Soc. Am. 63, 1974, p. 1765. Apparently, an increase in the OH-- content of 
silica reduces both types of absorption sites described above, and in 
accordance with the present invention this concept is applied to a system 
for delivering high peak energy ultraviolet laser pulses to thereby 
enhance the efficiency of the energy transmittance. Preferably, the silica 
that makes up the fibers contains about 200 to 2000 ppM of the OH-- 
radical, most preferably 1200 ppM. 
In another embodiment of the invention, the silica that is used to produce 
the fibers of the delivery system is doped with fluorine. Fluorine doped 
silica exhibits even lower attenuation than high OH silica. It appears 
that the fluorine functions to shift the absorption band gap in the 
SiO.sub.2 structure, to facilitate the transmittance of a large number of 
photons at low wavelengths. For multimode fibers having diameters in the 
range of 100 micrometers to 1500 micrometers, the silica preferably should 
contain between 0.25 and 2.0 wt % fluorine, most preferably 1.0 wt %. 
As a further feature of the invention, the silica can be doped with both 
the OH-- radical and fluorine. When both of these materials are used in 
combination, the OH radical content should range between 200 and 2000 ppM, 
and the fluorine should comprise between 0.5 and 3 wt % of the silica. 
In the context of the present invention, the fiber can be a single fiber or 
a bundle of fibers having a total diameter in the range of 100-2,000 
microns. A bundle of close-packed small-diameter fibers is preferred 
because they provide greater overall flexibility and thereby more easily 
accommodate the twists and tight turns that are required to feed the 
delivery system through body cavities. This is particularly desirable 
where a larger diameter waveguide is required to deliver a relatively 
large diameter beam with uniform intensity, such as in vascular 
angioplasty. This entire structure can be surrounded by a protective 
flexible jacket 28 made of a material which is not damaged by ultraviolet 
light. More particularly, when the fiber undergoes sharp bends, for 
example at the juncture of two arteries, light losses occur. These losses 
may be enough to melt some types of jacket materials such as silicone and 
nylon. However, UV light resistant materials, for example UV cured 
acrylate compound or TEFLON.RTM., can sustain high bending losses without 
degradation and are therefore more desirable for the jacket. 
In a preferred form of the invention, the protective jacket is incorporated 
as part of the fiber itself, rather than being a separate piece of 
structure which surrounds all of the fibers. As noted previously, every 
fiber comprises a core and a cladding which surrounds the core to maintain 
the transmitted light energy within the core. The cross-sectional area of 
the fiber might normally have a core/cladding ratio of 80/20 to provide 
suitable flexibility. Typically, both the core and the cladding are made 
of glass, with the cladding being appropriately modified (e.g., doped) to 
provide it with a lower index of refraction. In this conventional 
structure, the protective jacket comprises a third layer which surrounds 
the core and cladding. 
In accordance with one aspect of the invention, the conventional glass 
cladding is eliminated and the core of the fiber is directly surrounded by 
a coating of organic material. One specific preferred material is UV-cured 
acrylate. It has a lower index of refraction than silica, and thereby 
functions to maintain the laser energy within the core. It also serves to 
protect the silica glass, and hence eliminates the need for a third layer. 
This reduces the overall size of the fiber and hence enables the net 
cross-sectional area of the core to be increased for a delivery system 
having a given outer diameter. 
Further details regarding the composition of preferred coatings can be 
found in U.S. Pat. No. 4,511,209, the disclosure of which is incorporated 
herein by reference. 
By limiting the coating of the fiber to about 10% of the diameter of the 
fiber and by keeping the coating arranged substantially uniformly about 
the fiber, the fibers can be made as small as 50 microns in diameter. Such 
a fiber is approximately 16 times as flexible as a fiber having a diameter 
of 100 microns. Accordingly, the multi-fiber catheters described herein 
can be used with 50.mu. optical fibers for increased flexibility and 
limited transmission losses. 
A silica fiber of this construction can typically accommodate input energy 
up to a level around 30 Mj/mm.sup.2 produced by a commercially available 
Excimer laser with a 10-40 pulse. If the density of the energy is 
increased above this level, the input end of a conventional fiber having a 
planar, polished surface will be damaged or destroyed if the laser is 
applied directly to it. Unfortunately, this density level is about the 
minimum that is required to produce ablation of calcified plaque, thus 
providing no tolerance range if the intended use of the delivery system is 
for angioplasty. Accordingly, to enable a higher level of energy to be 
conducted in the fiber, an energy coupler 38 can be provided at the input 
end of the fiber. See FIG. 1. In this embodiment, the energy coupler 
comprises a section of fiber that has a larger cross-sectional area than 
the main portion of the fiber. This larger cross-sectional area gradually 
tapers to the nominal diameter of the fiber, to provide a funnel-shaped 
input section. 
Production of such a shape on the end of the fiber can be accomplished by 
appropriate design of the die through which the silica is drawn to produce 
the fiber. By interrupting the drawing of the fiber, a bulbous mass 
remains at one end of the fiber. This mass can be cut and polished to 
produce the funnel-shaped input section. 
In operation, the increased area of the funnel-shaped coupler decreases the 
input energy density for a given level of energy within the fiber. 
Accordingly, the area of the input end can be appropriately dimensioned to 
enable a sufficient amount of energy for ablation of tissue to be coupled 
into the fiber without damaging the input end. Once it has been coupled 
in, the density of the energy is increased by decreasing the 
cross-sectional area of the fiber within the tapered section, so that a 
greater amount of energy can be conducted within the fiber than would be 
possible without such a device. 
A second embodiment of an energy coupler is illustrated in FIG. 2. In this 
embodiment, the optical fiber has a uniform diameter along its length and 
terminates at a flat polished end. The end section of the fiber is encased 
within a ferrule 32 made of a suitable material such as brass, for 
example. An aluminum casing 33 having an annular ring 34 projecting from 
the inner wall thereof is threaded onto the ferrule. A TEFLON.RTM. O-ring 
35 disposed between the end of the annular ring and the ferrule provides a 
watertight seal between the casing and the ferrule. A second O-ring 36 is 
disposed on top of the annular ring and supports a glass plate 38 made of 
z-cut quartz, for example. This arrangement forms a fluid-tight cavity 40 
between the ferrule 32, the casing 33 and the glass plate 38. The glass 
plate can be held in place by means of a third O-ring 42 and a clamping 
ring 44 disposed on the top of the casing. The fluid tight cavity is 
filled with liquid which acts as a buffer to the input end of the fiber, 
enabling laser energy having a relatively high density to be coupled into 
the fiber without damage thereto. The liquid within the cavity can be 
distilled and deionized water, or it can be a transparent oil having an 
index of refraction that is matched to that of the fiber 12. 
A third embodiment of an energy coupler is illustrated in FIGS. 3A and 3B. 
In this embodiment, the input end of the fiber is provided with a fused 
semispherical lens 46. This lens can be formed by melting the material of 
the fiber itself with a micro-torch, to produce a high purity silica lens 
with no impurities or cracks. Alternatively, the lens 46 can be a 
separately ground lens that is attached to the flat end of the fiber. The 
fiber 12 can be tapered as shown in FIG. 1, or it can have a uniform 
diameter along its length. 
A second lens, preferably a plano-convex lens 47, focuses the input beam 
from the laser to a focal point 48. The input lens 46 on the fiber is 
axially aligned with the lens 47 and is located at a distance from the 
lens 47 which is greater than the focal length of that lens. Thus, the 
focused laser energy appears to be coming from a point source. The lens 46 
collimates this focused energy and couples it into the fiber. 
The input end of the fiber with the lens 46 and the focusing lens 47 are 
housed within a chamber 49. This chamber is provided with a vacuum port 50 
to enable the chamber to be evacuated of air. If air were present between 
the lenses 46 and 47, the highly concentrated energy at the focal point 48 
might cause a breakdown of nitrogen and oxygen gases that could 
contaminate the lens 46. In addition, the vacuum environment keeps out 
dust and other particles which could settle on the lens 46 and act as a 
heat sink, destroying the roundness of the lens. Alternatively, this 
chamber 49 can be filled with a liquid, such as water or oil, for example, 
which matches the index of refraction of the silica fiber. The higher 
index of refraction of the liquid reduces the dielectric shock when the 
pulse propagates from the liquid transmission medium to the fiber, 
relative to that which is experienced when air is the transmission medium. 
Although the preferred embodiment employs a curved lens at the proximal 
input end of the fiber, it is possible to couple the energy into a fiber 
having a planar input surface. However, it is important to ensure that 
this surface is free of scratches and other imperfections. This can be 
accomplished by heating the end of the fiber with a micro-torch to cause 
the fiber material to melt and flow slightly, thereby removing the 
imperfections caused by polishing. 
The type of energy coupler shown in FIG. 3A serves to amplify the energy 
within the fiber. More particularly, the amplification factor is equal to 
the ratio of the diameter of the laser beam at the lens 47 to the diameter 
of the fiber. This ratio is also related to the magnification produced by 
the two lenses. Referring to FIG. 3B, the dimension FB is the focal length 
of the lens 47 and the dimension FA is the distance between the lens 47 
and the focal point 48. The magnification factor of these two lenses is 
defined as FB/FA. Since this factor must be equal to the laser energy 
amplification, the appropriate distance between the lenses 46 and 47, 
i.e., AB=FB FA, can be determined from the following relationship: 
##EQU1## 
where D.sub.L is the diameter of the laser beam and D.sub.F is the 
diameter of the fiber. 
Although illustrated as a separate element in the figures, it will be 
appreciated that the energy couplers could be incorporated into the 
structure of a laser, to provide an integrated laser and coupling system. 
Thus, with the combination of the lightly doped synthetic silica fiber and 
the energy coupler 30 that enables a greater level of energy to be 
conducted through the fiber, an amount of high energy laser light that is 
sufficient to produce an incision can be safely transmitted through an 
optical fiber waveguide without the risk of damage to the fiber. 
To further increase the peak energy that is delivered through the system, 
it is preferable to slightly increase the length of the pulses beyond the 
relatively short duration that is typically produced by commercial Excimer 
lasers and the like. For example, a pulse having a duration in the range 
of 10-3000 nsec, more preferably 30-1000 nsec, or 100-300 nsec, enables 
much higher peak energy to be applied with the same delivery system than a 
shorter pulse, yet is still sufficiently short to produce the desired 
cutting action. 
One example of a circuit for stretching the output pulses of a laser is the 
magnetic switch developed at the Jet Propulsion Laboratory by Drs. J. 
Ladunslager and T. Tacala. In this regard, it is not necessary that each 
lengthened or stretched pulse comprise a single, continuous pulse. Rather, 
it could comprise a burst of shorter length successive pulses which 
together provide an effective pulse length of the desired duration. With 
the increased energy that is provided by the lengthened pulses, the energy 
level within the fiber will likely be more than sufficient to enable the 
laser beam to ablate an obstruction. In fact, the beam could be expanded 
as it exits the fiber and still contain sufficient energy density to 
ablate tissue. By expanding the diameter of the laser beam, for example by 
means of an increasing taper or fusing a larger diameter fiber at the 
distal end of the fiber, a larger area of tissue is ablated to produce 
more favorable results towards obtaining better blood flow in a blood 
vessel while using a small diameter flexible fiber that can be easily 
propagated through the vessel. 
An additional method of coupling a substantially pure silica fiber delivery 
system to a high energy pulsed laser, an energy coupler as illustrated in 
FIG. 4 may be used. A pure silica fiber 102 is tapered at about a one 
degree (1.degree.) taper and is about 30 centimeters long. The fiber 102 
has an input end 104 with a diameter of about 15 mm. At the narrow end 106 
of the fiber 102 is a modular connector 106 that connects with a connector 
108 on the end of an optical waveguide 110. 
A source 112 of laser energy propagates laser energy into the tapered fiber 
102 such that, in a preferred embodiment the energy density at the 
connector 106 is about 50-100 milliJoules/mm.sup.2. 
In order to provide better control of an optical fiber during laser 
angioplasty, a guidance system may be employed. Referring now to FIG. 5, 
the guidance system includes a guidewire 70 and a sleeve 72. 
The sleeve 72 is preferably between 1 and 200 centimeters in length, and 
has a rounded tip 74 at its distal end. The tip 74 can be made of 
stainless steel and glued or welded to the sleeve, or it can be formed 
integrally on the sleeve 72. The diameter of the tip can vary from 1.2 to 
2.5 mm, depending upon the size of the blood vessel. The rounded tip 74 
serves as both a dilator to enlarge the blood vessel and as a device to 
blunt the tip of the optical fiber 12 so as to minimize trauma to the 
blood vessel. 
The sleeve 72 has at least two lumens 76, 78 therein. A first lumen 76 is 
designed to accept the guidewire 70 and is preferably within the range of 
twelve-thousandths (0.012) to thirty-eight thousandths (0.038) of an inch 
in diameter. The diameter of the first lumen may vary, depending on the 
diameter of the guidewire 70. 
The second sleeve lumen 78 is designed to enclose the optical fiber 12, or 
an array of fibers, if such is the case. The diameter of the second lumen 
78 may also vary according to the diameter of the optical fiber 12 or 
fibers being used. The distal end of the optical fiber 12 is bonded within 
the second lumen 78 by any suitable means well known to those skilled in 
the art of bonding. 
To use the guidance system, the guidewire 70 is threaded through the lumen 
of the blood vessel by means of an introducer catheter (not shown). The 
guidewire 70 is inserted up to the location of a total obstruction in the 
vessel, or in the case of a subtotal lesion, beyond the lesion. 
The sleeve 72 is then mounted onto the guidewire 70, with the guidewire 
extending through the first lumen 76 of the sleeve. The sleeve 72 and the 
optical fiber 12, which is bound thereto, are then advanced along the 
guidewire 70 until the sleeve 72 and the distal tip of the optical fiber 
12 are adjacent the lesion to be ablated. The combination of the guidewire 
70 and the sleeve 72 ensures that the optical fiber 12 remains in 
alignment with the blood vessel, thus avoiding perforation of the blood 
vessel by the tip of the optical fiber 12 during positioning of the fiber 
or by the laser beam during ablation. 
Alternatively, the guidewire 70 may initially be threaded through the first 
lumen 76 of the sleeve 72, prior to the insertion of the guidewire 70 into 
the blood vessel. Once the fiber 12 is adjacent the lesion, ablation of 
the lesion is conducted as described above. 
In a preferred embodiment, the second lumen 78 is eccentrically located 
within the sleeve 72. In such an arrangement, rotation of the optical 
fiber 12 while it is in the blood vessel causes rotation of the sleeve 72 
which causes the radial position of the optical fiber 12 to shift within 
the blood vessel. Accordingly, rotating the optical fiber 12 during lasing 
causes a larger lumen to be ablated within the blood vessel. 
Once the lasing is completed, the sleeve 72 and the optical fiber 12 can be 
withdrawn, leaving the guidewire 70 in place within the blood vessel. 
Angiographic dye can then be injected through a guiding catheter around 
the guidewire 70 to evaluate the results of the lasing operation. If the 
results are unsatisfactory, the entire procedure can be repeated, possibly 
using different laser parameters or fibers. 
A further, preferred embodiment of a catheter for coronary laser 
angioplasty, which operates in accordance with the foregoing principles, 
is illustrated in FIGS. 6-8. The catheter 80 is multi-compartmented. It 
includes a center lumen 82 which accommodates a guidewire 84. The 
guidewire is introduced through a suitable coupling device 86 at the 
proximal end of the catheter. The center lumen 82 is surrounded by a 
plurality of circumferentially disposed outer lumens 88. These lumens 88 
each house one or more substantially pure synthetic silica fibers 90. In 
the illustrated embodiment there are three outer lumens 88 each housing 
two fibers 90. It will be appreciated that a different number of lumens 
and/or fibers per lumen can be employed, as determined by the relative 
sizes of the fibers 90, the guidewire 84 and the diameter of the catheter. 
Furthermore, the guidewire lumen may be eccentrically arranged within the 
catheter. 
An enlarged cross-sectional side view of the distal end of the catheter is 
shown in FIG. 8. Here, each laser-energy conducting fiber 90 is fused at 
its end face 91 to a short section of a larger diameter fiber 92. For 
example, the fibers 90 might have a diameter of 200 microns throughout the 
length of the catheter 80, and the short end fibers 92 might have a 
diameter of 300 microns, or a 100 micron diameter fiber may have a short 
200 micron diameter fiber fused to its end. Each end fiber 92 can be about 
3 mm long, and can be made from the same silica material as the fibers 90. 
By virtue of the larger diameter fiber at the distal end, the laser beam 
can expand as it emerges from the fiber, thereby providing a larger area 
of coverage and subsequently a larger ablation area. Furthermore, the 
plural fibers located symmetrically around the guidewire provide uniform 
energy distribution over a larger area. 
The fibers 92 are held in place at the end of the catheter by means of a 
suitable epoxy 94. A gold marker ring 96 can be provided around the 
catheter at the distal end, to assist in locating the end of the fiber 
during a fluoroscopy and angiography procedure. 
Except at the very end of the catheter where the epoxy 94 is present, there 
is free space within each outer lumen 88 between the fiber 90 and the 
walls of the lumen. See also FIG. 22 for an illustration of the free 
space. If desired, this free space can be used to provide a saline 
solution, or other contrast media, to the site of the obstruction. The 
solution can be injected into the catheter through a suitable port 98 at 
the proximal end, and emerge through holes 100 in the side wall of the 
catheter at its distal end (see FIG. 8). 
In certain situations, there is a need to ablate a large area of a 
partially or fully occluded vessel in order to maintain an open lumen. The 
clinical efficacy of ablation in the large peripheral vessels is enhanced 
when the lumen created is larger than 2 mm. Long term reocclusion of large 
vessels with low pressure blood flow is reduced When the energy delivery 
system is used to ablate a hole without subsequent dilation of the vessel. 
Laser angioplasty systems for the treatment of nontotal occlusions include 
systems that have a bundle of fibers placed concentrically around a 
guidewire lumen, and which are covered in a protective sheath. Small 
diameter catheters, e.g., 1.0 to 1.7 mm O.D., have a ring of active 
optical fibers that extends from the center guidewire lumen to the outer 
protective sheath. Any occlusive material confronted by the advancing 
catheter will be removed by the active fiber ring. 
Traditional guidewire catheter systems are guided through the vessels with 
a nondeflectable guidewire. However, because plaque is frequently built up 
eccentrically within the vessel, it is frequently desirable during a 
procedure to change the general direction of the guidewire, and thus the 
path of the catheter within the vessel. With a nondeflectable guidewire, 
it is sometimes difficult to change the direction of the guidewire. For 
this reason, it is advantageous to have a deflectable guidewire, and in 
particular, a deflectable guidewire that can be manipulated by the 
operator from the proximal end thereof. 
Such a deflectable guidewire 206 is disclosed in FIGS. 9-14. The guidewire 
206 includes a core 208 extending substantially the entire length of the 
guidewire 206. A wire coil 210 is positioned about the core 208 at a 
position along the guidewire where flexibility or deflectability is 
desired, preferably at the distal end thereof. 
The coil includes three sections, a first section 210A, a middle section 
210B, and a third section 210C. The outside diameter of the coil 210 is 
between 0.012 and 0.038 inches, preferably about 0.018 inches. The first 
and second sections 210A and 210B are preferably manufactured from a 
length of radiopaque wire. The first section 210A of the wire coil is 
preferably about 2-3 cm. in length. At the proximal end of the first 
section 210A, the wire coil 210 is mechanically fixed to the core 208 at 
joint 214 through a method such as soldering or brazing. 
The wire comprising the middle section 210B of the wire coil is 
mechanically deformed, or otherwise has a nonuniform cross section. As 
seen in FIGS. 9 and 10, the wire is deformed or flattened on one side 208A 
of the core 208 so that when the wire is straight, gaps occur between the 
individual turns of the wire coil on the opposite side 208B of the core. 
Alternatively, as indicated in FIG. 12, the diameter D2 of the wire on one 
side of the core can be reduced by electropolishing or chemical etching. 
Or, the diameter D1 of the wire on one side of the core can be built up by 
selectively depositing material by a plating process. 
At the proximal end of the third section 210C, the wire coil 210 is fixed 
to an intermediate sheath 216 at joint 218 by a method such as soldering 
or brazing. The intermediate sheath 216 is a noncompressible conduit that 
retains and protects the core 208 and extends from the wire coil 210 to 
the proximal end of the guidewire. The intermediate sheathing is 
preferably semiflexible hypodermic tubing that can be made from any 
appropriate material. 
At the distal end of the guidewire is a blunt tip 212 for minimizing 
traumatic injury to the vessel in which it is being used. The distal end 
of the core 208 may be affixed to the distal blunt tip 212, as illustrated 
in FIG. 9. Alternatively, as illustrated in FIG. 13, it is not necessary 
that the core 208 extend all the way to the blunt tip 212. In that case, 
it is only necessary that the core 208 be fixed to the wire coil 210 at 
the distal end of the deformed second section of the wire coil 210B. If 
the core 208 does not extend to the blunt tip 212, it is preferable to 
have a substitute material 213 extend between the blunt tip 212 and the 
core 208 in order to prevent the coil 210 from being stretched and/or 
unwrapped. It may be preferable to use a fine wire or ribbon as the 
substitute material 213 so as to maintain maximum flexibility in the tip 
of the catheter. 
The core 208 is fabricated from a material having a high tensile strength, 
preferably from a single piece of stainless steel. The core 208 slides 
freely within the second and third sections of the coil 210B, 210C and 
within the intermediate sheath 216. As indicated above, and in FIG. 9, the 
core 208 is secured to the wire coil 210 at the joint 214 between the 
first and second sections of the wire coil 210A, 210B. At its proximal 
end, the core 208 is secured to a deflection handle 220. See FIG. 14A. 
The deflection handle 220 is designed to apply tension to the core 208, 
which in turn causes the second section 210B of the wire to deflect as 
this is the only area of the outer sheath that has spacing between the 
coil where a change in length can occur. See FIGS. 11 and 12. The 
deflection handle 220 includes a finger grip 222, and a thumb ring 224. 
The core 208 of the guidewire is fastened to a pin 226, which is in turn 
connected to the finger grip 222. The sheath 216 is fastened to a 
nonmoving part of the deflection handle 220 so that when the finger grip 
222 and the thumb ring 224 are brought together, as shown by the arrows in 
FIG. 14A, a tension is created in the core 208. Because of the unevenness 
of the second section 210B of the wire, the wire coil, and thus the entire 
guidewire deflects at the region of the second section 210B of the 
guidewire. 
By selecting a core material of high torque strength, the entire guidewire 
can be rotated by rotating the deflection handle, which is attached to the 
core 208. Such a rotation permits the user to deflect the guidewire in any 
radial direction. 
The concept of using a nonuniform wire to deflect a guidewire can also be 
applied to an energy transmitting guidewire or a catheter system. Turning 
attention to FIG. 15, the wire coil and intermediate sheathing arrangement 
can be applied to an optical fiber or optical fiber bundle. In the 
embodiment disclosed in FIG. 15, a wire coil 228 includes a first section 
228A and a second section 228B. The first section has a nonuniform cross 
section like the second section 210B of the FIG. 9 embodiment. The wire 
comprising the first section 228A of the wire coil is mechanically 
deformed, or otherwise has a nonuniform cross section. As seen in FIG. 15, 
the wire is deformed or flattened on one side so as to create gaps between 
the individual turns of the wire coil on the opposite side. Alternatively, 
the diameter of the wire on one side of the core can be reduced by 
electropolishing or chemical etching. Or, the diameter of the wire on one 
side of the core can be built up by selectively depositing material by a 
plating process. 
An intermediate sheathing 230 is fixed to the proximal end of the wire coil 
228 via a joint that is made by a process such as soldering or brazing. A 
fiber optic core 234 extends from the distal end of the wire coil 228 
through the proximal end, and is ultimately connected to a source 227 of 
laser energy. See FIG. 14B. At its distal end, the fiber optic core 234 is 
fixed to the distal end of the wire coil 228 by means of epoxy 236 or some 
other suitable method. The distal end of the fiber optic core 234 is of 
course exposed so that laser energy may be transmitted forwardly from the 
catheter. 
At the proximal end of the intermediate sheathing and fiber optic core 234, 
a deflection handle is mounted. See FIG. 14B. The deflection handle is 
substantially identical to the deflection handle illustrated in FIG. 14A. 
Accordingly, a detailed description thereof will not be repeated herein. 
By appropriate control of the deflection handle, the wire coil, and thus 
the fiber optic core 234, can be deflected and thus guided within a 
vessel. 
FIGS. 16-18 disclose a catheter system substantially identical to the 
system disclosed in FIG. 15, except that the system disclosed in FIGS. 
16-18 includes an ultrasound system for facilitating guidance of the 
catheter during an ablating procedure. The core 238 comprises a plurality 
of small stranded optical fibers 240, a central optical fiber 242, and 
wires 244 for transmitting signals to and from ultrasound sensors at the 
distal end of the catheter. The small optical fibers 240 are approximately 
50 to 100 microns in diameter and the large central optical fiber 242 is 
approximately 150-200 microns in diameter. The actual sizes of the wires 
may vary in different applications. Alternatively, the central optical 
fiber 242 could be the same diameter as the outer fibers. In one 
embodiment, the central optical fiber 242 is 100 microns and the 
surrounding fibers 240 are 80 microns. The outside diameter of the 
catheter is preferably between 1.0 mm to 2.0 mm. 
At the distal end of the catheter, four ultrasound sensors 246 are 
symmetrically arranged about the optical fibers. The ultrasound sensors 
are fixed in place with epoxy 248, which also is used to fill in the gaps 
between each sensor 246. The ultrasound sensors comprise thin film 
ultrasound emitter/detectors, and are arranged so as to emit and detect 
signals in a radial direction with respect to the catheter. The ultrasound 
sensors are of the same type used in high-resolution, intraluminal 
imaging. However, in the present case, only a low pixel, low resolution 
system is used to merely delineate the blood vessel wall 239 without 
showing in great detail the diseased, or occluded, section 241 inside the 
vessel. 
Each of the ultrasound sensors 246 looks in a radial direction with respect 
to the catheter and sends a signal to the operator that is representative 
of the distance between the sensor and the wall of the vessel. In this 
manner, the operator can easily determine when the catheter is too close 
to the wall. The purpose of the ultrasound sensors is to detect the 
location of the wall so that during the ablating procedure, the 
laser-carrying optical fibers do not inadvertently contact the vessel 
wall. 
Because the ultrasound system is a low resolution system, the image seen by 
the operator is an outer circle representing the vessel wall and an inner 
circle representing the catheter. See FIGS. 19A and 19B. FIG. 19A 
illustrates the image seen when the catheter is in the center of a vessel. 
FIG. 19B illustrates the image seen when the catheter is adjacent a vessel 
wall. 
The catheter may have a marker band 247 fixed to the distal end of the 
catheter. The marker band 247 is preferably made from a radiopaque 
material, such as gold, so that it can be detected by an external x-ray, 
or comparable detecting system, in order to facilitate the guiding of the 
catheter to the lesion in the vessel. Thus, the catheter is first guided 
to the lesion using the external detecting system and the radiopaque 
marker. Then, control during the ablating process can be had with the use 
of the ultrasound system described above. Instead of using a radiopaque 
marker band, the deflectable coil itself may be made from a radiopaque 
material. 
The deflectability of the catheter facilitates the guiding of the catheter 
to the lesion to be ablated. However, during the ablating process, the tip 
of the catheter can be selectively deflected to permit a hole to be 
ablated that is larger in diameter than the diameter of the catheter. 
To increase the strength and torque of the core 238 of the catheter, each 
row of optical fibers may be stranded, i.e., wound or twisted, in a 
direction that is opposite to the direction in which each adjacent row is 
wound. Although the catheter in FIGS. 16-18 includes a solid fiber optic 
core 238, the concept of stranding the fibers in opposite directions can 
also be applied to a catheter that has a center guidewire lumen. FIG. 20 
illustrates a cutaway view of a catheter similar to the catheter of FIGS. 
16-18, except that it has a center guidewire lumen 282. There are shown 
only three layers of optical fibers stranded in accordance with the 
present invention. A preferred embodiment may have five or six layers. A 
first layer 280 of optical fibers is wound in a first (e.g., 
counterclockwise) direction about the plastic center guide lumen 282. A 
second layer 284 of optical fibers is wound over the first layer in an 
opposite direction, and a third layer 286 of optical fibers is wound over 
the second layer in the same direction as the first layer 280. A plastic 
outer cover 288 covers the third layer 286. 
As illustrated in FIGS. 21 through 24, the alternatingly stranded fibers 
perform better than coaxial parallel fibers. FIG. 21 is a simplified 
illustration of a catheter bent in 180.degree., and which has coaxial 
parallel fibers. A single exemplary fiber 290 is shown on the outer side 
of the bend, and a single exemplary fiber 292 is shown on the inner side 
of the bend. The outer fiber 290 tends to be stretched in tension, while 
the inner fiber 292 tends to be compressed. Both fibers are urged toward 
the center of the catheter as illustrated in FIG. 21. To accommodate these 
uneven forces acting on the fibers, such coaxial fiber catheters are 
generally designed, as illustrated in FIG. 22, with some clearance 293 
between the covering tube 294 and the fibers 296. The clearance is not an 
effective use of catheter space. 
In contrast to the design illustrated in FIGS. 21-22, a simplified 
illustration of fiber optics alternatingly stranded in accordance with the 
present invention is shown in FIGS. 23-24. As can be seen in FIG. 23, in a 
stranded catheter, the tensile and compressive forces caused by bending 
the catheter are evenly distributed among the fibers. As a result, there 
is little or no tendency for the catheter to flatten, as is illustrated in 
FIG. 21. In addition, as is illustrated in FIG. 24, there is no need to 
provide for clearance between the fibers 297 and the outer cover 299 if 
the fibers are stranded in accordance with the present invention. As a 
result, the overall diameter of the catheter can be made smaller. Another 
advantage of the stranded arrangement is that the catheter can be twisted 
more easily in that the stranded catheters will more easily transmit 
torque than the coaxial fibers. For example, if the assembly of FIG. 20 is 
twisted clockwise, the third layer 286 will tend to reduce in diameter and 
increase in length, while the second layer 284 will have the opposite 
effect, tending to increase in diameter and decrease in length. If the 
assembly is twisted counter-clockwise, the second layer 284 will tend to 
reduce in diameter and increase in length while the first layer 280 will 
tend to increase in diameter and decrease in length. 
Although FIG. 20 illustrates fibers wrapped in alternating directions, 
alternatively the fibers may be otherwise arranged within the catheter so 
that when the catheter is bent, the bending stresses are absorbed by the 
fibers in a substantially uniform manner. This is accomplished by 
arranging the fibers such that each fiber is bent a substantially equal 
amount during bending of the catheter. 
When performing angioplasty operations in large diameter vessels it is 
preferable to use catheters of a large diameter which have a large 
effective ablating area. In traditional large diameter catheters having a 
central guidewire lumen, there is a large number of optical fibers in the 
lumen between the outer sheath and the central guidewire lumen. Such a 
large number of fibers reduces the flexibility of the catheter. 
The outer diameter of the effective ablating area of the catheter can be 
increased, without decreasing the flexibility of the catheter, by 
increasing the outer diameter of both the outer sheath and the central 
guidewire lumen at the distal end of the catheter. The increased size of 
the outer sheath and guidewire lumen increases the diameter of the ring of 
active fiber optics, without requiring additional fibers. Thus, the 
flexibility of the catheter is not decreased. By increasing the diameter 
of the central guidewire lumen only at the distal end of the catheter, the 
likelihood of ischemia is decreased. 
In FIGS. 25 and 26, such a catheter system is disclosed. Referring to FIG. 
25, a catheter 250 has an enlarged diameter at its distal end. The outer 
sheath 252 has an increased diameter at the distal end thereof. At the 
center of the catheter 250 is a guidewire lumen 254, which has a similarly 
enlarged diameter at its distal end. Between the guidewire lumen 254 and 
the outer sheath are one or more lumens 256 for optical fiber. 
Within lumen 256 is a radiopaque marker band 251 for facilitating guiding 
the catheter through a body lumen. The marker 251, which may be made from 
gold, is detectable by an external system, such as an x-ray system. A 
polyurethane bushing 255 is placed within the guidewire lumen 254 to 
support the enlarged expanded portion 253 of the catheter. 
In a preferred embodiment, the internal diameter of the expanded portion 
253 of the guidewire lumen is approximately 0.041 inches. The expanded 
portion of the optical fiber lumen is ring shaped, and has an outer 
diameter of approximately 0.085 inches and an inner diameter of 
approximately 0.056 inches. About 360 50.mu. fibers are enclosed within 
the lumen 256. The overall outer diameter of the expanded portion of the 
catheter is approximately 0.095 inches. As illustrated in FIG. 25, the 
front face of the catheter may be angled 25.degree. from a plane normal to 
the axial direction of the catheter. 
The guidewire 258 disclosed in FIG. 26 is designed for use with the 
catheter of FIG. 25, and is referred to as a "carrotwire" guidewire. The 
main portion of the guidewire comprises a length of stainless steel shaft 
260, preferably about 0.018" in diameter, that extends from the proximal 
end of the guidewire. Fixed to the distal tip of the shaft 260 is a 
secondary shaft 262 of reduced diameter, e.g., 0.008", and having a length 
of about 19 cm. The diameter of the last four centimeters of the secondary 
shaft 262 is further tapered to a diameter of about 0.002". At the distal 
tip of the secondary shaft 262 is a flat strip 266 of stainless steel 
having a width of about 0.002". 
The secondary shaft 262 is wrapped with a stainless steel wire coil 264. 
The first 13 cm. of the wire coil 264 adjacent the main stainless steel 
shaft is wrapped in substantial contact with the secondary shaft 262 so 
that the overall diameter of the coil is about 0.018", i.e., about the 
same diameter as the main stainless steel shaft. At the distal end of the 
13 cm. length, the diameter of the coil is expanded to approximately 
0.039". At this point of expansion, a radiopaque (e.g., gold) solder 
marker 267 is fixed to the coil so that the distal end of the guidewire 
258 can he easily detected by an external x-ray system. The expanded 
portion of the wire coil extends for about 2 cm., at which point the 
diameter of the coil is gradually reduced to about 0.018" at the point 
where the secondary shaft is fixed to the strip 266 of stainless steel. At 
the distal tip of the guidewire, a platinum coil of wire 268, which covers 
the strip 266 of stainless steel 266 is affixed to the stainless steel 
coil of wire 264 with a soldered joint 270. The distal tip of the platinum 
coil of wire 268 is soldered to the tip of the stainless steel strip 266. 
The tip of the guidewire is blunt so as to minimize traumatic injury to 
the vessel in which it is being used. Note that the expanded portion of 
the wire coil 264 is adapted to fit within the expanded portion of the 
center guidewire lumen 254 of FIG. 25. 
Because both the catheter 250 and the guidewire 258 are expanded at their 
distal end, the catheter 250 and the guidewire 258 coact so as to center 
the distal tip of the catheter 250 within the vessel, to keep the distal 
tip of the catheter 250 collinear to the axis of the host vessel, and to 
mechanically displace occlusive material in an outward radial direction 
such that the occlusive material will be presented to the active fiber 
ring of the catheter 250. 
Turning attention to FIGS. 27-29, an over-the-wire catheter 300 having an 
eccentric body of optical fibers 302 is disclosed. The catheter 302 
includes an outer sheath 304 that encloses the plurality of optical fibers 
302 and a guidewire lumen 306 that forms a fountain pen-like point 308 at 
the distal end thereof. The point 308 has tapered walls 310 that form a 
gradual transition from the tip to the outer sheath 304 of the catheter. 
The point 308 may be made integral with the outer sheath 304. 
As illustrated in FIG. 28, the optical fibers 302 themselves are of an 
unequal length, with the fibers closest to the guidewire 306 being longer 
than those farther away from the guidewire. The varying length of the 
optical fibers is optimal for situations where a lesion 312 is located on 
one portion of the vessel 314. As seen in FIG. 30, the guidewire lumen can 
pass over the lesion 312 and the graduated length of the optical fibers 
302 tends to conform to the shape of the lesion. 
The following systems address the problems of making large diameter holes 
in blood vessels using a small overall diameter delivery system. Turning 
attention now to FIG. 31, a blood vessel 114 is shown with a lesion 116 
therein. A catheter 118 includes a plurality of optical fibers 120 evenly 
distributed about a concentrically mounted inflatable balloon 122. The 
optical fibers 120 may be 50-400 microns. A guidewire 124, larger than 
0.012 inches in diameter, is concentrically located in a center lumen 
within the balloon 122. If desired, a metallic (gold) marker 126 may be 
located adjacent the distal end of the catheter, such that the catheter 
118 may be located by an x-ray or fluoroscopy system. 
In operation, the guidewire 124 is inserted through the lumen or blood 
vessel 114 until the guidewire passes through the lesion 116 that is to be 
ablated. The catheter 118 is then conveyed along the guidewire 124 until 
the distal end of the catheter 118 contacts the lesion 116. The fibers 120 
deliver an initial dose of high energy laser pulses to ablate the inner 
portion of the lesion 116. Subsequently, the balloon 122 is inflated to a 
predetermined pressure which then forces the fibers 120 into an array of a 
larger diameter. A subsequent delivery of high energy pulsed laser is then 
delivered to ablate the outer peripheral portions of the lesion 116. The 
balloon 122 may be additionally inflated if necessary to obtain an array 
of fibers 120 of a still larger diameter. 
With reference to FIG. 32, an additional preferred embodiment of the 
present invention is disclosed. A catheter 128 has three balloons 130 
located in an equally spaced arrangement at the distal end of the catheter 
128. A first lumen 132 is positioned concentrically among the balloons 
130. In the first lumen 132 is disposed a fiber optic instrument 133. If 
desired, a guidewire may be retained in a second lumen adjacent the first 
lumen 132. Additional lumens may also be included at the center of the 
balloons 130 to accommodate other instruments to assist with illumination 
or flushing, for example. 
In operation, the lumens at the center of the catheter may be positioned or 
tilted by selectively inflating and deflating the three balloons 130. In 
one mode of operation, the balloons may be inflated sequentially and 
continuously so as to selectively revolve the fibers in a circular pattern 
along the perimeter of the catheter. The fibers may be centered by 
inflating all of the balloons. 
In another embodiment, two, four or any other number of balloons may be 
used instead of three. 
As in the embodiment shown in FIG. 31, the embodiment shown in FIG. 32 may 
also include a marker, such as a gold band, at the distal end of the 
catheter. 
Another preferred embodiment of the present invention is illustrated in 
FIG. 33. In this embodiment a guidewire 150 has a balloon 152 at the end 
thereof. Markers 154 are mounted on the guidewire 150 adjacent the balloon 
152. The guidewire 150 is disposed through a guidewire lumen 164 in the 
center of a catheter 160. The catheter 160 further includes a plurality of 
fibers 162 arranged in an annular pattern about the guidewire lumen 164. 
If desired, a marker 166, such as a gold band, can be placed around the 
distal end of the catheter 160 to facilitate detection of the catheter. 
In operation, the guidewire 150 with the balloon 152 at one end thereof is 
inserted through the blood vessel until it is located beyond the lesion 
158. The balloon 152 is then inflated, thus centering the guidewire 150 
within the vessel. Such centering of the guidewire 150 minimizes the 
likelihood that the catheter 160 will contact the walls of the blood 
vessel 156 during the ablation process. A stop 168 may be placed on the 
guidewire 150 so as to prevent the catheter 160 from contacting and thus 
possibly rupturing the balloon 152. 
The fibers 162 in the catheter preferably range from about 50 to 300 
microns. The guidewire lumen 164 is large enough to allow free movement 
over a standard guidewire ranging from 0.014 to 0.038 inches in diameter. 
With reference to FIG. 34, another preferred embodiment of the present 
invention is disclosed within a blood vessel 170 having a lesion 172 
therein. A catheter 174 includes a guidewire and flushing lumen 178 
located adjacent an inner edge of the catheter. An optical fiber lumen 176 
has a width that extends from the guidewire and flushing lumen 178, 
through the center of the catheter to the edge of the catheter 
diametrically opposite the guidewire and flushing lumen 178. A plurality 
of optical fibers 180 made of substantially pure silica are disposed 
within the optical fiber lumen 176, and a guidewire 182 is disposed within 
the guidewire and flushing lumen 178. A balloon 184 is bonded at one point 
to the outer surface of the catheter 174. 
To operate, an initial ablation is performed with the balloon 184 in an 
uninflated condition. The ablation is effected by delivering a high energy 
pulsed laser, such as from an Excimer laser, through the optical fibers 
180 while advancing and rotating the fibers about the guidewire 182. The 
rotation enables the fibers to ablate a circular area about the size of 
the catheter outer diameter. 
In the next stage, the outer balloon is inflated, thus urging the fibers 
into a larger diameter. A second ablation is performed with the balloon 
inflated and while advancing and rotating the catheter so as to ablate an 
annular area having a diameter larger than the diameter of the circular 
area ablated during the first stage. During the second stage, the balloon 
is inflated to a predetermined size in order to ensure that the second 
annular area is contiguous to the circle ablated during the first stage. 
The Excimer energy is only capable of ablating a lesion in a forward 
direction and does not require that the blood between the lesion and the 
optical fibers be displaced with a solution, such as saline, that is 
transparent to the laser. In addition, the laser transmitted through the 
optical fibers is capable of ablating a lesion in a blood field. 
Therefore, there is no necessity to block the blood flow around the 
catheter during the ablation process. 
The tip of the catheter should have a round, conical, distal tip to 
minimize the trauma to the vessel wall. The optical fibers 180 comprise a 
highly flexible array of fibers that may range in size from 50 to 400 
microns in diameter, with a possible enlarged fiber tip output diameter to 
create a large ablation area. The fibers which are immobilized at the 
distal end of the catheter by epoxy are polished to form the round distal 
tip. 
The guidewire used to advance the catheter over is commonly used in 
interventional radiology and cardiology and is preferably larger than 
0.012 inch. 
If desired, markers 186, such as gold bands can be placed at the distal of 
the catheter to facilitate monitoring of the catheter by x-ray or 
fluoroscopy. 
With reference to FIG. 35, a catheter 174' is disclosed that is 
substantially the same as the catheter 174 of FIG. 34, described above. 
The significant difference between the catheter 174 of FIG. 34 and the 
catheter 174' of FIG. 35 is that the balloon 184' of FIG. 35 includes a 
recess 188' that enables blood to flow around the balloon 184' during its 
use. 
The off-center balloon catheters of the present invention function 
differently from other balloon catheters that are used in interventional 
cardiology and radiology. The off-center balloon catheters of the present 
invention are not dilation devices, such as those used in Percutaneous 
Transluminal Angioplasty (PTA) and Percutaneous Transluminal Coronary 
Angioplasty (PTCA). They are not intended to laterally dilate the blood 
vessels in which they are used or to press a catheter against the wall of 
a vessel. They also are not intended to stop the flow of blood and replace 
it with saline, as is commonly done to improve visualization during 
angioscopy. Rather, they function as positioning devices which enable the 
optical fibers, and hence the high energy laser, to be precisely 
positioned within the blood vessel. In this regard, they are not intended 
to contact the walls of the blood vessel tightly, so that translation 
and/or rotation of the catheter is possible while the balloons are 
inflated, to thereby provide a mobile angioplasty operation. 
With reference to FIG. 36, another embodiment of the present invention is 
illustrated within a blood vessel 190 having a lesion 192 therein. A 
catheter 194 has a smooth, rounded leading edge 196, which may be 
lubricous, e.g., coated with a water repellant chemical, for easy 
advancing and/or rotation through a blood vessel. 
In the center of the catheter 194 is a guidewire lumen 198, within which 
lumen a guidewire 200 is slidably disposed. Surrounding the central 
guidewire lumen 198 is an annular array of optical fibers 202 that may be 
used to deliver high energy pulsed laser for ablating the lesion 192. A 
balloon 204 is mounted to a location on the outer periphery of the 
catheter 194. 
In operation, the guidewire 200 is first threaded through the blood vessel 
190 until it is adjacent the lesion 192. The catheter 194 is then conveyed 
along the guidewire 200 with the balloon 204 deflated until the catheter 
194 is adjacent the lesion 192. At that point, a channel is ablated in the 
lesion 192 with laser energy delivered by the optical fibers 202. The size 
of the channel is substantially equal to the outside diameter of the 
catheter 194. The balloon 204 is then inflated, thus shifting the optical 
fibers 202 to a higher orbit. The lesion 192 is again ablated by the 
optical fibers 202 while advancing and/or rotating the catheter, thus 
forming a larger opening within the lesion 192. The balloon 204 may be 
further inflated, and the process repeated, as necessary. 
As a result of the embodiment disclosed in FIG. 36, the lesion 192 is 
uniformly ablated by a plurality of optical fibers 202. As a result, the 
likelihood of large segments of the lesion 192 being released into the 
blood flow, where they may cause emboli down stream is substantially 
reduced. Because the lesion 192 is broken into small segments, there is no 
need to have a dedicated channel in the catheter to remove debris. The 
system disclosed in FIG. 36 generally creates particles less than 10 g in 
diameter. 
A further refinement of the catheter system disclosed in FIG. 36 is 
disclosed in FIG. 37. An outer catheter 316 includes an outer lumen or 
passage (not shown) for retaining means for inflating the balloon 318. An 
internal catheter 320 includes a plurality of optical fibers 322 and is 
slidable in an axial direction with respect to the outer catheter 316. The 
internal catheter 320 further includes a central lumen for a guidewire 
324. 
In use, the outer catheter 316 and internal catheter 320 are advanced 
together along the guidewire 324 to a lesion 326 in a vessel 328. At the 
location of the lesion 326, the balloon 318 is inflated to lock the outer 
catheter 316 in place within the vessel 328. The internal catheter 320 can 
then be advanced as desired while ablating the lesion. 
FIG. 38 relates to a method of transmitting high energy pulsed laser beams 
through a substantially pure silica fiber. FIG. 38 depicts a pulse that is 
made up of a train or a group of time shifted subpulses. A high energy 
long pulse I is a super positioning of numerous subpulses A-H and has a 
shorter diffusivity of tissue. In other words, a high energy pulse 
duration can be stretched in time when ablating a tissue without causing 
undesired thermal damage to adjacent tissue due to the thermal diffusion 
through the tissue during the ablation process. The estimated time needed 
to move a high energy laser pulse from an ablation zone tissue is about 
one millisecond. 
U.S. Pat. No. 4,677,636 relates to such laser pulses, and the subject 
matter thereof is incorporated herein by reference. 
It will be appreciated by those of ordinary skill in the art that the 
present invention can be embodied in other specific forms without 
departing from the spirit or essential characteristics thereof. The 
presently disclosed embodiments are therefore considered in all respects 
to be illustrative and not restrictive. The scope of the invention is 
indicated by the appended claims rather than the foregoing description, 
and all changes that come within the meaning and range of equivalents 
thereof are intended to be embraced therein.