Infrared laser catheter system

Laser energy produced by a laser operating in the mid-infrared region (approximately 2 micrometers) is delivered by an optical fiber in a catheter to a surgical site for biological tissue removal and repair. Disclosed laser sources which have an output wavelength in this region include: Holmium-doped Yttrium Aluminum Garnet (Ho:YAG), Holmium-doped Yttrium Lithium Fluoride (Ho:YLF), Holmium-doped Yttrium-Scandium-Gadolinium=Garnet (HO:YSGG), Erbium-doped YAG, Erbium-doped YLF and Thulium-doped YAG. Laser output energy is applied to a silica-based optical fiber which has been specially purified to reduce the hydroxyl-ion concentration to a low level. The catheter may be comprised of a single optical fiber or a plurality of optical fibers arranged to give overlapping output patterns for large area coverage. In a preferred application for the removal of atheroscleotic plaque, a Holmium-doped laser operating in the wavelength range of from about 1.9 to about 2.1 micrometers is preferred. For removal of such plaque by a Holmium-doped laser, it has been found that the threshold energy density should be greater than about 0.6 joules/mm.sup.2 per pulse, and that the pulse width should be substantially less than about 83 milliseconds, and that the repetition rate should be in the range of from about 1 to about 10 Hertz.

FIELD OF THE INVENTION 
This invention relates to laser catheters and optical fiber systems for 
generating and transmitting energy to a surgical site in a living body for 
the purposes of tissue removal or repair. 
BACKGROUND OF THE INVENTION 
While lasers have been used for many years for industrial purposes such as 
drilling and cutting materials, it is only recently that surgeons have 
begin to use lasers for surgical operations on living tissue. To this end, 
laser energy has been used to repair retinal tissue and to cauterize blood 
vessels in the stomach and colon. 
In many surgical situations, it is desirable to transmit laser energy down 
an optical fiber to the surgical location. If this can be done, the 
optical fiber can be included in a catheter which can be inserted into the 
body through a small opening, thus reducing the surgical trauma associated 
with the operation. In addition, the catheter can often be maneuvered to 
surgical sites which are so restricted that conventional scalpel surgery 
is difficult, if not impossible. For example, laser energy can be used to 
remove atherosclerotic plaque from the walls of the vasculature and to 
repair defects in small-diameter artery walls. 
A problem has been encountered with laser surgery in that prior art lasers 
which have been used for industrial purposes often have characteristics 
which are not well suited to percutaneous laser surgery. For example, a 
laser which is conventionally used for scientific purposes is an excimer 
laser which is a gas laser that operates with a gas mixture such as 
Argon-Fluorine, Krypton-Fluorine or Xenon-Fluorine. Another common 
industrial laser is the carbon dioxide or CO.sub.2 laser. 
Both the excimer laser and the CO.sub.2 laser have been used for surgical 
purposes with varying results. One problem with excimer lasers is that 
they produce output energy having a wavelength in the range 0.2-0.5 
micrometers. Blood hemoglobin and proteins have a relatively high 
absorption of energy in this wavelength range and, thus, the output beam 
of an excimer laser has a very short absorption length in these materials 
(the absorption length is the distance in the materials over which the 
laser beam can travel before most of the energy is absorbed). 
Consequently, the surgical site in which these lasers are to be used must 
be cleared of blood prior to the operation, otherwise most of the laser 
energy will be absorbed by intervening blood before it reaches the 
surgical area. While the removal of blood is possible if surgery is 
performed on an open area it is often difficult if surgery is to be 
performed via a catheter located in an artery or vein. 
An additional problem with excimer lasers is that the output energy pulse 
developed by the laser is very short, typically about ten nanoseconds. In 
order to develop reasonable average power, pulses with extremely high peak 
power must be used. When an attempt is made to channel such a high peak 
power output into an optical fiber, the high peak power destroys the 
fiber. Thus, excimer lasers have a practical power limit which is 
relatively low. Consequently, when these lasers are used for biological 
tissue removal, the operation is slow and time consuming. 
The CO.sub.2 generates output energy with a wavelength on the order of 10 
micrometers. At this wavelength, the absorption of blood hemoglobin is 
negligible but the absorption by water and tissue is relatively high. 
Scattering at this wavelength is also very low. Although these 
characteristics of the CO.sub.2 laser are favorable for surgical 
applications, the CO.sub.2 laser suffers from the same drawback as excimer 
lasers in that the absorption length is relatively short due to the high 
absorption of the laser energy in water. Thus, the surgical area must be 
cleared of blood prior to the operation. 
Unfortunately, the CO.sub.2 laser also suffers from a serious additional 
problem. Due to the long wavelength, the output energy from the carbon 
dioxide laser cannot be presently transmitted down any optical fibers 
which are suitable for use in percutaneous surgery (present fibers which 
can transmit energy from a CO.sub.2 laser are either composed of toxic 
materials, are soluble in water or are not readily bendable, or possess a 
combination of the previous problems) and, thus, the laser is only 
suitable for operations in which the laser energy can be either applied 
directly to the surgical area or applied by means of an optical system 
comprised of prisms or mirrors. 
Accordingly, it is an object of the present invention to provide a laser 
catheter system which uses laser energy of a wavelength that is strongly 
absorbed in water, in bodily tissues and atherosclerotic plaque. 
It is another object of the present invention to provide a laser catheter 
system which is capable of providing laser energy that can be transmitted 
through existing silica-based optical fibers. 
It is a further object of the present invention to provide a laser catheter 
system in which optical scattering is minimized and which has a 
medium-length absorption length to confine the energy to the area of 
interest. 
It is yet another object of the present invention to provide an optical 
catheter system with a laser that can be operated on either a pulsed mode 
or a continuous wave mode. 
It is still another object of the present invention to provide a laser 
catheter system which can be used for biological material removal and 
biological material repair. 
It is still another further object of the present invention to provide a 
laser catheter system which can be used for removal of atherosclerotic 
plaque within a living body. 
SUMMARY OF THE INVENTION 
The foregoing objects are achieved and the foregoing problems are solved in 
one illustrative embodiment of the invention in which a laser catheter 
system employs a laser source operating in the wavelength region of 
1.4-2.2 micrometers. Illustrative laser sources operating this region are 
Holmium-doped YAG, Holmium-doped YLF, Holmium doped YSGG, Erbium-doped 
YAG, Erbium doped YLF and Thulium-doped YAG lasers. 
In the inventive laser system, the above-noted lasers are used with a 
specially-treated silica fiber that has been purified to reduce the 
concentration of hydroxyl (OH--) ions. 
For biological tissue removal, the laser source may be operated in a pulsed 
mode with a relatively long pulse of approximately 0.2-5 milliseconds at 
an energy level of approximately 1-2 joules per pulse, for a spot size of 
the order of 1.5 millimeters in diameter. With this time duration and 
energy level, the peak power of the laser pulse is approximately 1 
kilowatt. This amount of power can easily be tolerated by the silica 
fiber, but is sufficient for rapid material removal. With a repetition 
rate in the range of 1-10 hertz, the average power delivered to a surgical 
site by such a laser will be under 10 watts. 
In particular, for removal of atherosclerotic plaque from a living body, 
particularly satisfactory results are obtained using a Holmium-doped laser 
source operating in a pulsed mode in a wavelength range of from about 1.90 
to about 2.10 micrometers, and at a threshold energy density of at least 
about 0.6 joules/mm.sup.2. The pulse width used should be substantially 
less than a thermal time constant for plaque, or substantially less than 
about 83 milliseconds. The repetition rate typically is about 2 Hertz. 
Alternatively, for biological tissue repair, the laser source can be 
operated in a low power continuous wave mode to repair, by coagulation, of 
tissue by a process similar to "spot welding".

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
The absorption and scattering characteristics versus output wavelength of a 
plurality of known laser systems are shown in FIG. 1. FIG. 1 has a 
logarithmic scale representing the absorption coefficient in units of 
cm.sup.-1 along the vertical axis and the incident energy wavelength in 
micrometers along the horizontal axis. 
From FIG. 1, it can be seen that excimer laser systems which utilize 
conventional gas mixtures, such as Argon-Fluorine, Krypton-Fluorine and 
Xenon Fluorine, and Argon gas lasers produce output energy which falls in 
the 0.2-0.5 micrometer wavelength region. In this region, the absorption 
of blood hemoglobin and proteins is very high. Consequently, the 
absorption length is very short (about 5-10 microns) and virtually no 
blood can be present between the fiber end and the surgical site during 
the operation. Thus, it is necessary to remove blood from the surgical 
area when these lasers are used for surgical purposes. 
In addition, for lasers such as Argon, the absorption of water reaches a 
minimum at 0.5 micrometers so that it is necessary to use a higher power 
laser than is desirable to achieve sufficient power in the surgical area 
for material cutting and removal. Also, due to the low absorption of the 
laser output in water and hemoglobin, the absorption length is very long 
(approximately 1 mm). In addition, scattering for these lasers is 
relatively high, causing difficulty in controlling the laser energy and a 
danger of tissue damage outside the surgical area due to scattering of the 
laser energy. 
At the other end of the wavelength spectrum shown in FIG. 1 are carbon 
monoxide and carbon dioxide lasers producing outputs at 5 and 10 
micrometers, respectively. At these wavelengths scattering is negligible 
and absorption by water and tissue is relatively high and thus both lasers 
have good surgical properties. Unfortunately, due to the high absorption 
of water, the absorption length is relatively short (about 20 microns). 
Further, silica-based optical fibers in present use which are suitable for 
percutaneous surgical use have a practical "cutoff" in transmission which 
occurs approximately at 2.3 micrometers, and, thus, the output energy from 
carbon monoxide and carbon dioxide lasers cannot be transmitted through 
such an optical fiber. 
In accordance with the invention, laser sources of interest are those that 
lie in the wavelength range of approximately 1.4-2.15 micrometers. As 
shown in FIG. 1, in this range, the energy absorption of water is 
relatively high whereas optical scattering is relatively low. Illustrative 
lasers which are useful with the present invention comprise Erbium-doped 
Yttrium Aluminum Garnet (YAG) with a wavelength of 1.55 micrometers, 
Erbium-doped Yttrium Lithium Fluoride (YLF) with a wavelength of 1.73 
micrometers, Thulium-doped YAG with a wavelength of 1.88 micrometers, 
Holmium-doped YLF with a wavelength of 2.06 micrometers, Holmium-doped YAG 
at a wavelength of 2.1 micrometers, and Holmium-doped 
Yttrium-Scandium-Gadolinium-Garnet (YSGG) at a wavelength of 2.088 
micrometers. The absorption of the laser energy produced by lasers in this 
latter group by water is moderately high and, consequently, the absorption 
by biological tissues of such energy will also be relatively high. 
However, the absorption by water is not as high as the absorption of CO 
and CO.sub.2 laser energy. Thus, the absorption length will be longer for 
the lasers operating in the 1.4-2.2 micrometer range than for CO.sub.2 
lasers. Typically, the absorption length in the body for lasers operating 
in the 1.4-2.2 micrometer range is about 200 microns. Therefore, it is 
still possible to operate satisfactorily even with 10-30 microns of blood 
between the fiber end and the surgical site. 
Of particular interest is the absorption of the laser energy by 
atherosclerotic plaque, since an important use of laser catheter systems 
is angioplasty, particularly the clearing of blocked arteries. FIG. 2 is a 
plot of the absorption by plaque of electromagnetic energy versus 
wavelength for energy in the wavelength range of 0.2-2.2 micrometers. As 
shown in FIG. 2, the absorption by plaque of electromagnetic energy 
reaches a minimum in the 0.8-1 micrometer wavelength range and generally 
increases with increasing wavelength in the wavelength region of 1-2.2 
micrometers. In the wavelength range from 1.4-2.2 micrometers, the 
wavelength range produced by laser in the above-mentioned group, the 
absorption by plaque is at a relatively high value. 
A schematic diagram of a typical solid-state laser construction is shown in 
FIG. 3. The laser assembly consists of a laser crystal 1 and an excitation 
device such as a flashlamp 3. Typically, for the crystal compositions 
disclosed above, the laser crystal must be cooled to cryogenic 
temperatures to provide low laser-threshhold operation. Cryogenic cooling 
is typically provided by enclosing crystal 1 in a quartz or fused-silica 
jacket 4 through which liquid nitrogen is circulated. Liquid nitrogen 
enters jacket 4 by means of an inlet pipe 5 and leaves by means of an 
outlet pipe 6. The laser cavity is formed by a high reflectivity concave 
mirror 10 and a partial reflector 12. 
Generally, the crystal is excited by optical pumping which is, in turn, 
accomplished by irradiating the crystal with light from a flashlamp 3. A 
flashlamp which is typically used with the inventive laser compositions is 
a high-pressure Xenon flashlamp. Lamp 3 may also be surrounded by a quartz 
flow tube (not shown) through which coolant is pumped. 
Crystal 1 and flashlamp 3 are enclosed in a reflector 2 which concentrates 
the flashlamp energy into the laser crystal. To maximize energy transfer 
from lamp 3 to crystal 1, the inner surface of reflector 2 is coated with 
a material chosen to have high-reflectivity at the pumping wavelength of 
the laser crystal--illustratively, aluminum or silver. In order to provide 
thermal insulation to prevent condensation on the system optics, it may be 
necessary to evacuate the interior of reflector 2 or to provide a vacuum 
jacket around crystal 1. 
The construction of cryogenic solid-state lasers is conventional and 
described in a variety of sources; accordingly such construction will not 
be discussed further in detail herein. A more complete discussion of 
construction details of a typical cryogenic laser is set forth in an 
article entitled "TEM.sub.oo Mode Ho:YLF laser", N. P. Barnes, D. J. 
Gettemy, N.J. Levinos and J. E. Griggs, Society of Photo-Optical 
Instrumentation Engineers, Volume 190--LASL Conference on Optics 1979, pp 
297-304. 
FIG. 4 of the drawing is a plot of the illustrative pulse shape developed 
by a laser in the preferred group when used in the "material removal" 
mode. FIG. 4 shows light intensity along the vertical axis increasing in 
the downward direction versus time increasing towards the right. Although, 
as shown in FIG. 4, the laser source has been adjusted to produce an 
output pulse of relatively long time duration, most of the output energy 
is released within approximately 1 millisecond of the beginning of the 
pulse. It should also be noted, as illustrated in FIG. 4, that lasers in 
the preferred laser group exhibit a "spiking" phenomenon caused by 
internal relaxation oscillations in the laser crystal. The spiking 
behavior causes local increases in laser intensity which have a large 
magnitude, but a very short time duration. Similar "spiking" behavior has 
been found advantageous when lasers are used to drill metals and other 
materials for industrial purposes and it is believed that such "spiking" 
behavior enhances the laser usefulness for biological material removal. 
FIG. 5 is a schematic diagram of a laser/catheter system employing a solid 
state laser of the type shown in detail in FIG. 3. More specifically, the 
infrared output energy of laser 21 is focused by a conventional focusing 
lens onto the end of the optical fiber which is held in a conventional 
fiber optic connector 24. Fiber optic connector 24 is, in turn, connected 
to a tube 27 which houses a single optical fiber. Tube 27 is connected to 
a conventional two-lumen catheter 30 by means of a bifurcation fitting 28. 
Illustratively, catheter 30 has two lumens passing axially therethrough to 
its distal end 34 so that an optical fiber can pass through one lumen and 
transmit laser energy from fiber optic connector 24 to lens tip 34. As 
previously mentioned, the optical fiber which passes through the catheter 
is specially purified to reduce the hydroxyl ion concentration to a low 
level, thus preventing the laser energy which is transmitted down the 
fiber from being highly absorbed in the fiber material. A fiber which is 
suitable for use with the illustrative embodiment is a fused-silica 
optical fiber part no. 822W manufactured by the Spectran Corporation 
located in Sturbridge, Mass. 
Advantageously, the mirrors and lenses (10, 12 and 22) which are used to 
form the IR laser cavity and focus the output energy beam are generally 
only reflective to energy with a wavelength falling within a narrow 
wavelength band and transparent to energy at other wavelengths. 
Consequently, the mirrors and lenses are transparent to visible light. An 
aiming laser 20 (for example, a conventional helium-neon laser) which 
generates visible light may be placed in series with IR laser 21 to 
generate a visible light beam. This light beam may be used to align 
mirrors 10 and 12 and to adjust focussing lens 22 so that the optical 
fiber system can be aligned prior to performing surgery. 
Also, the optical fibers used to transmit the IR energy from laser 21 to 
the surgical area can also be used to transmit the visible light from the 
aiming laser 20 to the surgical area. Thus, when the inventive system is 
used in performing surgery where the surgical area is visible to the 
surgeon, the light produced by laser 20 passes through the optical fiber 
in catheter 30 and can be used to aim the probe tip before laser 21 is 
turned on to perform the actual operation. 
The second lumen in catheter 30 is provided for transmission of a flushing 
fluid or to apply suction to the probe lens tip area to clear the area of 
blood during surgery. This latter lumen is connected through bifurcation 
fitting 28 to a second tube 29. Tube 29 may illustratively be terminated 
by a standard Luer-Lok fitting 26 which allows connection of the catheter 
to injectors and standard flow fittings. Solutions injected into the 
catheter through tube 29 pass through the lumen in catheter 30 and exit at 
the distal end via a small orifice 32. 
Probe tip 34 consists of a lens arrangement which forms the laser energy 
into a beam 36 which is used to perform the surgical operations. An 
enlarged view of the probe tip is shown in FIGS. 6 and 7. 
To ensure that the distal end of optical fiber 18 is spaced and oriented in 
a precise position with respect to the end of the probe, fiber 18 is 
mounted in a high-precision holder 58 which has a reduced diameter end 64 
that forms a shoulder 68. Shoulder 68, as will hereinafter be described, 
is used to hold the probe tip assembly together. Holder 58 has a 
precision-formed axial bore made up of two sections, including a 
large-diameter section 60 and a narrow-diameter section 63. Holder 58 may 
be made of glass, ceramic or other material capable of being formed to 
specified dimensions with precise tolerances. 
In order to attach holder 58 to the end of fiber 18, the fiber is first 
prepared as shown in FIG. 7. More particularly, prior to insertion of 
fiber 18 into holder 58, a portion of buffer sheath 61 is removed, 
exposing a length of optically-conductive core 65. Care is taken when 
stripping buffer sheath 61 from the fiber not to damage the layer of 
reflective cladding 67 located on the surface of core 65. After stripping, 
fiber 18 is inserted into holder 58 so that core 65 extends into the 
small-diameter bore 63 and sheath 61 extends into the large-diameter bore 
60. After fiber 18 has been inserted into holder 58, it may be fastened by 
epoxy cement to permanently affix the components. To complete the 
assembly, the end of fiber 18 which protrudes beyond surface 62 of holder 
58 may be finished flush with the surface by grinding the assembly or by 
carefully cleaving the fiber. 
Referring to FIG. 6, holder 58 (with fiber 18 fastened inside) is mounted 
within a glass tube 51 to shield the assembly. The front surface, 62, of 
holder 58 is spaced from the inner surface 142 of planar lens 144, which 
may be comprised of glass or sapphire, by means of a spacing ring 154. 
Ring 154 may illustratively be made of radiopaque material so that the 
catheter tip can be located inside the patient by means of a fluoroscope. 
Glass tubing 51 is bent over shoulder 68 of holder 58 to form a constricted 
end, 65, which holds the probe tip assembly together. A filler, 66, which 
may be made of a plastic such as TEFLON (trademark of the DuPont 
corporation for polytetrafluoroethylene) fills the annular space between 
catheter body 30 and end 65 of glass tube 51. The outer diameter of the 
entire assembly from catheter body 30 to glass tube 51 is substantially 
the same, providing a smooth, uniform surface along the entire length of 
the catheter as indicated in FIG. 6. 
FIG. 8 shows a schematic diagram of a wire-guided, four-fiber catheter for 
use with the present invention. The laser system is set up as previously 
described with the infrared laser 21 constructed in accordance with the 
above disclosure. A visible helium-neon aiming laser 20 may also be used 
in line with laser 21 for aiming purposes as discussed with the single 
fiber catheter. The output of the infrared laser 21 is directed towards a 
set of four mirrors 60-68 arranged at a 45.degree. angle with respect to 
the axis of beam 14. 
The first mirror, 60, has a 25% reflective surface and directs 
approximately 1/4 of the energy to focusing lens 70. The second mirror of 
the set, 62, is a 33% reflector which directs 1/4 of the total energy to 
focusing lens 72. Mirror 64 is a 50% reflector which directs 1/4 of the 
total laser output to focusing lens 74. The last mirror in the set, mirror 
68, is a 100% reflector which directs the remaining 1/4 of the total 
energy to focusing lens 78. Mirrors 60-68 and lenses 70-78 are 
conventional devices. 
Focusing lenses 70-78 focus the output energy from IR laser 21 onto four 
fiber optic connectors 80-88. Connectors 80-88 are connected, 
respectively, to tubes 90-96 which are all connected, via a branch 
connector 102, to catheter 104. Each of tubes 90-96 contains a single 
optical fiber which transmits 1/4 of the total laser output energy through 
the catheter body to the catheter tip 108. An additional tube 98 is 
provided which is connected to branch fitting 102 and to a conventional 
Luer Lok connector, 100. This latter tube is connected to a central lumen 
in catheter body 104 through which flushing solutions may be injected or 
through which a guide wire may be inserted through the catheter for 
purposes of guiding the catheter to the surgical area. 
At catheter tip 108, the four optical fibers which pass through the 
catheter are arranged symmetrically so that the beams 110 overlap to 
illuminate a larger area. Tip 108 also has a hole on the center thereof, 
through which guidewire 112 can protrude to direct the catheter to the 
proper location. 
FIGS. 9 and 10 show detailed views of the illustrative four-fiber catheter 
tip. The four optical fibers 42 and the inner shaft 40 which holds the 
fibers, are held in a fiber holder 50 which is preferably formed from a 
radiopaque material such as stainless steel or platinum. Fiber holder 50 
is cylindrical and is provided with a central aperture, 54, which 
communicates with a lumen 34 of approximately the same size that passes 
through the center of the catheter body 104. Fiber holder 50 is provided 
with a plurality of longitudinally extending holes 56 which extend through 
the wall of holder 50 and receive, in a snug fit, the distal ends of the 
fiber cores 42. The distal face 58 of the combined fiber cores 42 and 
holder 50 is polished flat to butt flush against optically transparent cap 
52. 
Cap 52 is cylindrical and has the same outer diameter as catheter body 104 
so that the two pieces define a smooth and continuous diameter. Cap 52 may 
be formed of a transparent substance such as pyrex glass or sapphire and 
has an enlarged bore 62 extending in from its proximal end. Bore 62 
terminates at its end to form internal shoulder 60. A smaller diameter 
central aperture, 64, is formed in the distal end of cap 52 which aperture 
may have the same diameter as aperture 54 in fiber holder 50 and lumen 34 
in catheter body 104 to provide a smooth and continuous lumen which opens 
at the distal tip of the catheter. However, the aperture 64 in tip 52 may 
also be somewhat narrower than aperture 54 and lumen 34 as long as 
sufficient clearance is provided to accommodate a guidewire without 
adversely interfering with fluid flow and pressure measurements. 
Cap 52 is secured by an epoxy adhesive (placed on its inner surface 62) to 
the fiber holder 50 and also to the portion of the inner shaft 40 and 
fibers 42 which are disposed within the proximal end of the cap 52. The 
distal end of the catheter body 104 is heat shrunk around the inner shaft 
40 and fibers 42 to provide a smooth transition from cap 52 to catheter 
body 104. 
More complete construction details of a four fiber catheter suitable for 
use with the illustrative embodiment are given in co-pending U.S. patent 
application entitled "Wire Guided Laser Catheter", filed on May 22, 1985 
by Stephen J. Herman, Laurence A. Roth, Edward L. Sinofsky and Douglas W. 
Dickinson, Jr. 
FIG. 11 illustrates the output beam pattern developed by a four-fiber 
catheter, such as that described above, in which the four fibers are 
arranged in two diametrically-opposed pairs. The beam pattern from each of 
the four fiber ends is defined by a cone formed by the ray lines 70 in 
FIG. 11. The beam from each individual fiber 42 is emitted from the distal 
face of the fiber 42 and enters the distal segment 72 of cap 52 through 
the face defining the shoulder 60. The beam from each fiber is divergent 
and, in the illustrative embodiment, may have a half-angle in the range of 
6.degree.-16.degree., depending on the numerical aperture of the fibers 
which are used to construct the catheter. 
The diverging beam from each of the fibers 42 exits from the distal 
emission face 74 at the end of cap 52. FIGS. 11A, 11B and 11C illustrate 
the overall beam pattern (in cross-section) which is formed by the output 
of the four fibers as seen along image planes 11A, 11B and 11C in FIG. 11. 
At plane 11A, which is located at the emission face 74 of cap 52, the four 
beams in the illustrative embodiment are still separate. At plane 11B, the 
diverging beams have spread further and have begun to overlap. At the 
plane indicated as 11C, the beams have overlapped and define an envelop 73 
having an outer diameter which is slightly greater than the outer diameter 
of the catheter body 104. Preferably, at plane 11C, beams 70 will have 
overlapped to merge and cover a continuous pattern. Illustratively, such a 
merger will have occurred within a distance from the distal face 74 of tip 
52 which is approximately equal to the outer diameter of catheter 104 (a 
typical diameter is 1.5 millimeters). 
A preferred application of the previously described laser catheter system 
of this invention is the removal of atherosclerotic plaque. The laser must 
be operated in a pulsed mode to remove such plaque. A continuous wave 
laser is unsuitable for plaque removal, since a continuous wave laser 
transmits insufficient energy to the surgical site at a time to vaporize 
the tissue before the energy is dissipated by thermal diffusion. As a 
consequence, typically, carbon deposits are formed in the affected area, 
and it is difficult to cleanly and accurately remove plaque from artery 
walls using a continuous wave laser. 
Although any of the lasers previously mentioned as operating in the 
preferred range of wavelengths would be suitable for the removal of 
atherosclerotic plaque, it has been found that Holmium doped lasers 
operating at a wavelength in the range of from about 1.90 to 2.10 
micrometers are particularly suited for such surgical applications. In the 
first place, Holmium-doped lasers operate in the preferred range of 
wavelengths of 1.4-2.2 micrometers, and they can be operated in a pulsed 
mode. Secondly, Holmium-doped lasers are generally very efficient, 
generating 10 watts of average power for every kilowatt of electrical 
power supplied to the lamp. Because of these power requirements, which are 
lower than for most conventional lasers, no open loop cooling is required. 
Furthermore, such lasers can be operated with a conventional 110 volt 
power source. As a consequence of the foregoing, Holmium-doped laser 
sources often are portable, and can be operated either in a hospital 
environment, or in a doctor's office, or even in the home. A laser source 
which is preferred is a Holmium-doped Yttrium-Scandium-Gadolinium Garnet 
(YSGG) laser operating at a wavelength of 2.088 micrometers. Such a laser 
can remove any tissue containing water. 
The pulse width of any laser which is used for plaque removal should be 
substantially less than the thermal time constant for that particular 
tissue. Typically, the pulse width should be roughly an order of magnitude 
less than the thermal time constant. For wavelengths in the range of 
1.4-2.2 micrometers, thermal diffusion, which is a function of absorption, 
is reduced to an insignificant factor when such pulse widths are used. For 
a given power or energy level delivered per pulse, pulses longer than the 
thermal time constant will be generally incapable of precisely removing 
tissue, because the energy delivered thereby will be dissipated through 
thermal diffusion. 
For all tissue, the thermal time constant is equal to the absorption length 
of that particular tissue type squared over four times its thermal 
diffusivity. For most tissue, the thermal diffusivity is equal to about 
1.times.10.sup.-3 cm.sup.2 /sec, while for atherosclerotic plaque, the 
thermal diffusivity is equal to about 1.2.times.10.sup.-3 cm.sup.2 /sec. 
The absorption length for most types of plaque is about 2.times.10.sup.-2 
cm. Thus, for plaque, the thermal time constant is approximately 
8.3.times.10.sup.-2 seconds. Therefore, typically, the pulse width should 
be of the order of about 8.3 milliseconds or less. 
The repetition rate is not critical, and typically is in the range of 1-10 
Hertz. The preferred rate, for use with a Holmium-doped laser, is about 2 
Hertz. 
The energy density delivered is defined as the energy per pulse divided by 
the spot size. The spot size is the area of the spot upon which the laser 
energy is incident. It has been predicted theoretically and verified 
experimentally that for plaque removal, the energy density of each pulse 
should exceed some threshold value. If the energy density is less than 
this threshold value, the energy will be dissipated by thermal diffusion 
before the tissue is vaporized, and no vaporization will occur. It has 
been found that most energy densities above the threshold value can be 
used, up to the operational limit of the laser and the associated fiber 
optic system. 
The threshold value of the energy density for plaque will vary, depending 
upon the particular tissue sample invoved. Generally, tissue samples with 
more calcification or less water content will vaporize at a higher 
threshold value, while tissue samples with less calcification, or a higher 
water content, will vaporize at a lower threshold value. 
A thermal model was developed for laser pulses whose pulse widths are small 
compared to the thermal time constant of the material with which the laser 
is to be used. Using this thermal model, one can develop predicted 
threshold energy densities per pulse for lasers operating at different 
wavelengths, as well as the extent of thermal damage or the extent of the 
denatured rim. 
In using this thermal model, the energy profile in the tissue can be 
described as a function with separable axial and radial components. The 
axial component can be described by the attentuation coefficient according 
to Beer's law and the radial component can be descried by a linearly 
expanding Gaussian distribution. The axial and radial components can be 
defined mathematically as follows: 
EQU J(r,z)=J(r).multidot.J(z) 
where 
EQU J(z)=J.sub.o e.sup.-.alpha.z 
and 
EQU J(r)=e.sup.-2r.spsp.2.sup./x.spsp.2, 
where the function J(r,z) is the spatial energy distribution, z is the 
axial distance into the tissue measured from the tissue-air or 
tissue-water interface, .alpha. is the combined absorption-scattering 
attenuation coefficient, r is the radial distance measured from the 
optical axis, and x is the 1/e.sup.2 beam width of the Gaussian-like 
radial light distribution. 
The amount of energy, .DELTA.J, deposited into a volume element is equal to 
the energy profile divided by the volume as a function of the radial 
distance measured from the optical axis or 
##EQU1## 
Besides the attenuation coefficient, the next most important parameter is a 
description of the 1/e.sup.2 radius, x. Beam spreading occurs due to 
multiple scattering. Measurements have shown that the beam spreading is 
linear in almost all cases, although the slope varies with different 
tissue samples. Because absorption dominates over scattering in 
mid-infrared wavelengths, spreading in this range should be quite small. 
Therefore, it is assumed that the radius expands with the same divergence 
as the incident beam. 
The threshold energy density is the sum of the energy required to raise the 
tissue temperature to 150.degree. C. from the ambient temperature of about 
37.degree. C. plus the energy required to overcome the heat of 
vaporization (2260 J/g). Thus, 
##EQU2## 
where p is the tissue density (1.2 g/cm.sup.3), c is the specific heat 
(3.6 J/g.degree.C.), and L.sub.v is the heat of vaporization (2260 J/g). 
It can be seen that the total energy density required to vaporize a volume 
element is 3200 J/cm.sup.3. When the deposited energy density is between 
488 and 3200 J/cm.sup.3, the tissue temperature remains at 150.degree. C. 
For purposes of comparison, the vaporization threshold is defined as being 
reached if enough energy is supplied to vaporize tissue out to the edge of 
a 1 millimeter hole. This definition means that both the central peak and 
the wings of the Gaussian energy distribution are sufficient to vaporize 
tissue. The thermal zone of damage is defined as the region in which the 
temperature has reached above 60.degree. C., and can vary in the axial and 
radial directions. Energy densities in terms of joules/mm.sup.2 for 
different wavelengths can be determined based on the assumption that the 
depth of the hole is equal to the absorption length of the radiation at 
that wavelength. 
Based upon the foregoing analysis, the following Table I sets forth the 
predicted threshold energy densities in joules per mm.sup.2 and the 
predicted sizes of the zones of axial and radial thermal damage in 
micrometers for the lasers previously indicated to operate within the 
preferred rang of wavelengths. All of these predicted energy densities are 
for atherosclerotic plaque having a high water content and little or no 
calcification. 
TABLE I 
______________________________________ 
Threshold 
Wave- Absorption 
Energy Thermal 
Length Coeff Density Damage (.mu.m) 
Laser (.mu.m) (cm.sup.-1) 
(J/mm.sup.2) 
Axial Radial 
______________________________________ 
Ho:YAG 2.1 50 .76 500 500 
Ho:YSGG 2.088 50 .76 500 500 
Er:YAG 1.55 2700 .0095 12 12 
Ho:YLF 2.06 50 .76 100 300 
Tm:YAG 1.88 50 .76 100 300 
Er:YLF 1.73 15 5.7 500 300 
______________________________________ 
Energy densities required to vaporize atherosclerotic plaque using a 
Holmium-doped laser also have been determined empirically from the 
experimental data. These threshold energy densities were determined using 
an apparatus 120 shown schematically in FIG. 12. Apparatus 120 includes a 
laser source 122, focusing lens and fiber optic connector 124, fiber optic 
126, and a tank 128 containing the tissue sample 130. Laser source 122 and 
connector and associated lens 124 are both conventional, as previously 
described with respect to FIGS. 3-11. Fiber optic 126 may contain a single 
optical fiber, or it may contain a bundle of optical fibers. Typically, in 
conducting these experiments, a single fiber was used. 
Tank 128 contains a saline solution in which the tissue sample 130 is 
immersed. Each optical fiber included within fiber optic 126 typically is 
a 100 micron fiber. Also immersed in the saline solution in tank 128 and 
covering tissue sample 130 is a sapphire plate 132. The end of fiber optic 
126 is placed in direct contact with the upper surface of plate 132, while 
the lower surface of plate 132 is in direct contact with the tissue 
sample. In conducting these experiments, typically the sapphire plate 132 
had a thickness of 1.52 millimeters. The sapphire plate 132 is used to 
permit a controlled spreading of the beam of light emitted from the end of 
fiber optic 126 to produce the desired spot size on the tissue sample. 
Tissue sample 130 used in each of these experiments comprises a layer of a 
fatty fibrous plaque 134 disposed on top of a section of an aorta wall 
136. Two different tissue samples were used. Experiments 1-9 were 
performed on Sample No. 2 while experiments 10-13 were performed as Sample 
No. 1. Sample No. 1 exhibited greater calcification and lower water 
content than Sample No. 2. Sample No. 2 exhibited virtually no 
calcification and a high water content. Prior to use, the tissue samples 
had been removed from the body, and had been carefully frozen and stored 
at a temperature of -70.degree. C. The tissue samples were later thawed 
for use in these experiments over a lengthy period of time to avoid damage 
to the tissue and to avoid destruction of the cells during the freezing 
and thawing processes. 
In each experiment, a Holmium doped YSGG laser was operated in a pulsed 
mode at a wavelength of 2.088 micrometers. The pulse width used was 400 
microseconds, and two pulses were emitted for each test with a spacing of 
one half second, or at a rate of 2 Hertz. The spot sizes and energies per 
pulse were varied to determine the threshold energy density. The results 
of these experiments are set forth below in Table II. 
TABLE II 
__________________________________________________________________________ 
Energy/ 
Delivery 
Fluence, Ablation Thermal, 
Exper- 
Pulse, Area, (Joules/Area, 
OD, Damage 
iment 
(Joules) 
(mm.sup.2) 
mm.sup.2) 
(mm.sup.2) 
(mm) (mm) 
__________________________________________________________________________ 
1 122 .times. 10.sup.-3 
4.012 .times. 10.sup.-1 
0.304 
N/A N/A N/A 
2 222.5 .times. 10.sup.-3 
4.012 .times. 10.sup.-1 
0.555 
N/A N/A N/A 
3 310 .times. 10.sup.-3 
4.012 .times. 10.sup.-1 
0.773 
3.848 .times. 10.sup.-1 
100 .times. 10.sup.-6 
50 .times. 10.sup.-6 
4 816.2 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
25.98 
6.158 .times. 10.sup.-2 
280 .times. 10.sup.-6 
240 .times. 10.sup.-6 
5 475 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
15.13 
4.337 .times. 10.sup.-2 
235 .times. 10.sup.-6 
240 .times. 10.sup.-6 
6 268.7 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
8.55 
5.41 .times. 10.sup.-3 
262.5 .times. 10.sup.-6 
100 .times. 10.sup.-6 
7 875 .times. 10.sup.-3 
5.3 .times. 10.sup.-1 
1.65 
5.03 .times. 10.sup.-1 
800 .times. 10.sup.-6 
35 .times. 10.sup.-6 
8 760 .times. 10.sup.-3 
5.3 .times. 10.sup.-1 
1.434 
6.36 .times. 10.sup.-1 
900 .times. 10.sup.-6 
35 .times. 10.sup.-6 
9 655 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
20.85 
1.018 .times. 10.sup.-1 
360 .times. 10.sup.-6 
200 .times. 10.sup.-6 
10 185 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
5.89 
6.157 .times. 10.sup.-2 
280 .times. 10.sup.-6 
35 .times. 10.sup.-6 
11 139 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
4.43 
4.869 .times. 10.sup.-2 
249 .times. 10.sup.-6 
10 .times. 10.sup.-6 
12 271 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
8.63 
4.869 .times. 10.sup.-2 
249 .times. 10.sup.-6 
25 .times. 10.sup.-6 
13 74.5 .times. 10.sup.-3 
3.14 .times. 10.sup.-2 
2.37 
N/A N/A N/A 
__________________________________________________________________________ 
The value for the energy per pulse in Table II was measured in a 
conventional manner. The delivery area in square millimeters was 
calculated from the known, controlled spreading of the light beam as it 
passed from fiber optic 126 through sapphire plate 132. The fluence, in 
joules per square millimeter was calculated by dividing the energy per 
pulse by the delivery area. The ablation area is the area of the hole 
formed in the tissue sample. The extent of thermal damage is the distance 
beyond the rim of the hole where thermal damage was observed, while the OD 
is the outside diameter of the hole formed. The ablation area, OD and 
extent of thermal damage were all measured microscopically from 
histological slides prepared from the specimens using conventional 
histological techniques. 
The immersion of a tissue sample in a saline solution, as shown in FIG. 12, 
closely models the conditions inside the body. Although all of these 
experiments were conducted in vitro, as shown, one could expect nearly 
identical results in vivo. 
It can be seen from Table II that, for fatty fibrous plaque, having a high 
water content and virtually no calcification, the threshold energy density 
per pulse is at least about 0.6 J/mm.sup.2 for a Holmium-doped YSGG laser 
operating at a wavelength of 2.088 micrometers. It is expected that 
similar experiments using a Ho:YAG laser and a Ho:YLF laser would show 
that the energy threshold for such lasers is also at least about 0.6 
J/mm.sup.2 per pulse at their respective wavelengths of operation, based 
upon the fact that identical threshold energy densities are predicted by 
the theoretical model, as shown in Table I. 
For the other lasers described herein as being suitable, the theoretically 
predicted threshold energy densities per pulse for vaporization of fatty 
fibrous plaque having virtually no calcification and a high water content 
are as follows as found in Table I: Erbium YAG, 9.5 mJ/mm.sup.2 ; Erbium 
YLF, 5.7 J/mm.sup.2 ; and Thulmium YAG, 0.76 J/mm.sup.2. The precise 
threshold energy density depends upon the wavelength, and the degree of 
calcification and water content of the tissue sample. As indicated, the 
greater the calcification and the less the water content, the higher the 
threshold energy density required for vaporization of the tissue. 
The data set forth in Tables I and II also illustrate additional advantages 
of using a Holmium-doped laser within the wavelength ranges of 1.90 
micrometers to 2.1 micrometers. It can be seen that the zone of thermal 
damage is minimal, far less than one would expect from the prior art. 
Furthermore, virtually no carbonization was observed in each of the 
experiments. Such superior results were not achieved in tests with any 
other existing laser source which could be delivered via a non-toxic, 
durable, flexible fiber optic to a vein or artery within the body. 
In view of the above description, it is likely that modifications and 
improvements may occur to those skilled in the art within the scope of 
this invention. Thus, the above description is intended to be exemplary 
only, the scope of the invention being described by the following claims 
and their equivalents.