U.S. patent application number 17/346637 was filed with the patent office on 2021-11-11 for system and process of utilizing ultrasound energy for treating biological tissue.
This patent application is currently assigned to Ojai Retinal Technology, LLC. The applicant listed for this patent is Ojai Retinal Technology, LLC. Invention is credited to David B. Chang, Jeffrey K. Luttrull, Benjamin W. L. Margolis.
Application Number | 20210346714 17/346637 |
Document ID | / |
Family ID | 1000005738612 |
Filed Date | 2021-11-11 |
United States Patent
Application |
20210346714 |
Kind Code |
A1 |
Luttrull; Jeffrey K. ; et
al. |
November 11, 2021 |
SYSTEM AND PROCESS OF UTILIZING ULTRASOUND ENERGY FOR TREATING
BIOLOGICAL TISSUE
Abstract
A process for heat treating biological tissue includes providing
a plurality of energy emitters formed into an array. Treatment
energy in the form of ultrasound energy is generated from the
plurality of emitters and applied to target tissue. The treatment
energy has energy and application parameters selected so as to
raise the target tissue temperature sufficiently to create a
therapeutic effect while maintaining an average temperature of the
target tissue over several minutes at or below a predetermined
temperature so as not to destroy or permanently damage the target
tissue.
Inventors: |
Luttrull; Jeffrey K.; (Los
Angeles, CA) ; Chang; David B.; (Tustin, CA) ;
Margolis; Benjamin W. L.; (Oakland, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Ojai Retinal Technology, LLC |
Los Angeles |
CA |
US |
|
|
Assignee: |
Ojai Retinal Technology,
LLC
Los Angeles
CA
|
Family ID: |
1000005738612 |
Appl. No.: |
17/346637 |
Filed: |
June 14, 2021 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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16039779 |
Jul 19, 2018 |
11077318 |
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17346637 |
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15918487 |
Mar 12, 2018 |
10874873 |
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16039779 |
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15629002 |
Jun 21, 2017 |
10278863 |
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15918487 |
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15583096 |
May 1, 2017 |
10953241 |
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15629002 |
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15460821 |
Mar 16, 2017 |
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15583096 |
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15232320 |
Aug 9, 2016 |
9962291 |
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15460821 |
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15214726 |
Jul 20, 2016 |
10531908 |
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15232320 |
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15178842 |
Jun 10, 2016 |
9626445 |
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15214726 |
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14922885 |
Oct 26, 2015 |
9427602 |
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15178842 |
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14921890 |
Oct 23, 2015 |
9381116 |
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14922885 |
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14607959 |
Jan 28, 2015 |
9168174 |
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14921890 |
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13798523 |
Mar 13, 2013 |
10219947 |
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14607959 |
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13481124 |
May 25, 2012 |
9381115 |
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13798523 |
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15460821 |
Mar 16, 2017 |
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16039779 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 18/14 20130101;
A61N 5/067 20210801; A61N 2005/0608 20130101; A61N 1/403 20130101;
A61N 5/0603 20130101; A61N 2005/063 20130101; A61N 2005/0659
20130101; A61N 5/025 20130101; A61F 2009/00863 20130101; A61B
18/1206 20130101; A61N 2005/0609 20130101; A61N 5/045 20130101;
A61F 9/00817 20130101; A61N 5/0625 20130101; A61N 2005/0604
20130101; A61B 2018/00494 20130101; A61N 7/022 20130101; A61N
2007/0043 20130101; A61N 2005/0651 20130101; A61N 2005/0662
20130101; A61B 2018/00005 20130101; A61B 2018/0016 20130101; A61F
9/00821 20130101 |
International
Class: |
A61N 5/06 20060101
A61N005/06; A61F 9/008 20060101 A61F009/008; A61N 5/04 20060101
A61N005/04; A61B 18/14 20060101 A61B018/14; A61N 5/02 20060101
A61N005/02; A61N 7/02 20060101 A61N007/02; A61B 18/12 20060101
A61B018/12 |
Claims
1. A process for heat treating biological tissue, comprising the
steps of: providing a plurality of energy emitters formed into an
array; generating treatment energy comprising ultrasound energy
from the plurality of emitters; providing an initial treatment to
the target tissue by repeatedly applying the energy to the target
tissue in a pulsed manner for the exposure duration comprising less
than one second; halting application of the pulsed energy to the
target tissue for an interval of time comprising between three
seconds and three minutes; and providing a secondary treatment to
the target tissue after the interval of time, within a single
treatment session, by repeatedly reapplying the pulsed energy for
the exposure duration comprising less than one second; wherein the
treatment energy has energy and application parameters selected so
as to raise the target tissue temperature sufficiently to create a
therapeutic effect while maintaining an average temperature of the
target tissue over several minutes at or below a predetermined
temperature so as to not destroy or permanently damage the target
tissue.
2. The process of claim 1, wherein the selected energy and
application parameters comprise tissue application spot size or
area, average power or average power density, and exposure
duration.
3. The process of claim 1, wherein during an interval of time,
comprising less than one second, between applications of energy
applied to a first area of the target tissue, applying treatment
energy to a second area of the target tissue sufficiently spaced
apart from the first area of the target tissue to avoid thermal
tissue damage of the target tissue.
4. The process of claim 3, wherein repeatedly applying, in an
alternating manner during the same treatment session, the treatment
energy to each of the first and second areas of the target tissue
until a predetermined number of energy applications to each of the
first and second areas of the target tissue has been achieved.
5. The process of claim 3, including the step of introducing a
phase delay in the activation of the energy emitters of the array
to generate treatment energy in a phased manner using a
predetermined delay of activation in order to apply treatment
energy to each of the first and second areas of the target
tissue.
6. The process of claim 3, including the step of activating the
energy emitters of the array sequentially in order to apply
treatment energy to each of the first and second areas of the
target tissue.
7. The process of claim 1, wherein the treatment energy raises the
target tissue to up to eleven degrees Celsius at least during
application of the pulsed treatment energy thereto.
8. The process of claim 1, wherein the average target tissue
temperature is maintained at six degrees Celsius or less over
several minutes.
9. The process of claim 8, wherein the average target tissue
temperature is maintained at one degree Celsius or less over
several minutes.
10. The process of claim 1, wherein the applying step comprises the
step of stimulating heat shock protein activation in the target
tissue.
11. The process of claim 1, wherein the treatment energy and
application parameters are selected to have an average power
density of 100-590 watts/square centimeter of target tissue, a spot
size between 100-500 microns, and a train exposure duration of 500
milliseconds or less.
12. The process of claim 1, wherein the ultrasound energy has a
frequency between approximately 1 MHz and 5 MHz, a duty cycle
between approximately 2% to 10% and a power of between
approximately 0.46 and 28.6 watts.
Description
RELATED APPLICATIONS
[0001] This application is a divisional of U.S. application Ser.
No. 16/039,779 filed Jul. 19, 2018, which is a continuation-in-part
of U.S. application Ser. No. 15/918,487 filed Mar. 12, 2018, which
is a continuation-in-part of U.S. application Ser. No. 15/629,002
filed Jun. 21, 2017, Ser. No. 15/583,096 filed May 1, 2017, Ser.
No. 15/460,821 filed Mar. 16, 2017, Ser. No. 15/232,320 filed Aug.
9, 2016 (now U.S. Pat. No. 9,962,291), Ser. No. 15/214,726 filed
Jul. 20, 2016, Ser. No. 15/178,842 filed Jun. 10, 2016 (now U.S.
Pat. No. 9,626,445), Ser. No. 14/922,885 filed Oct. 26, 2015 (now
U.S. Pat. No. 9,427,602), Ser. No. 14/921,890 filed Oct. 23, 2015
(now U.S. Pat. No. 9,381,116), Ser. No. 14/607,959 filed Jan. 28,
2015 (now U.S. Pat. No. 9,168,174), Ser. No. 13/798,523 filed Mar.
13, 2013, and Ser. No. 13/481,124 filed May 25, 2012. This
application is also a continuation-in-part of U.S. application Ser.
No. 15/460,821, filed Mar. 16, 2017.
BACKGROUND OF THE INVENTION
[0002] The present invention is generally directed to systems and
processes for treating biological tissue, such as diseased
biological tissue. More particularly, the present invention is
directed to a process for heat treating biological tissue using
energy having parameters and applied such so as to create a
therapeutic effect to a target tissue without destroying or
permanently damaging the target tissue.
[0003] The inventors have discovered that there is a therapeutic
effect to biological tissue, and particularly damaged or diseased
biological tissue, by controllably elevating the tissue temperature
up to a predetermined temperature range while maintaining the
average temperature rise of the tissue over several minutes at or
below a predetermined level so as not to permanently damage the
target tissue. More particularly, the inventors have discovered
that electromagnetic radiation, such as in the form of various
wavelengths of light, can be applied to retinal tissue in a manner
that does not destroy or damage the retinal tissue while achieving
beneficial effects on eye diseases. The inventors have found that a
light beam can be generated and applied to the retinal tissue cells
such that it is therapeutic, yet sublethal to retinal tissue cells
and thus avoids damaging photocoagulation in the retinal tissue
which provides preventative and protective treatment of the retinal
tissue of the eye. The treatment typically entails applying a train
of laser micropulses to radiate a portion of a diseased retina for
a total duration of less than a second. Each micropulse is on the
order of tens to hundreds of microseconds long, with the
microseconds being separated by one to several milliseconds, which
raises the tissue temperature in a controlled manner.
[0004] It is believed that raising the tissue temperature in such a
controlled manner selectively stimulates heat shock protein
activation and/or production and facilitation of protein repair,
which serves as a mechanism for therapeutically treating the
tissue. It is believed that this micropulse train thermally
activates heat shock proteins (HSPs) in the targeted tissue. In the
case of retinal tissue, the process thermally activates HSPs in the
retinal pigment epithelium (RPE) layer immediately behind the
retinal layer containing the visually sensitive rods and cones, and
that these activated HSPs then reset the diseased retina to its
healthy condition by removing and repairing damaged proteins. This
then results in improved RPE function, improves retinal function
and autoregulation, restorative acute inflammation, reduced chronic
inflammation, and systematic immunodulation. These laser-triggered
effects then slow, stop or reverse retinal disease, improve visual
function and reduce the risk of visual loss. It is believed that
raising tissue temperature in such a controlled manner to
selectively stimulate heat shock protein activation has benefits in
other tissues as well.
[0005] HSPs are a family of proteins that are produced by cells in
response to exposure to stressful conditions. Production of high
levels of heat shock proteins can be triggered by exposure to
different kinds of environmental stress conditions, such as
infection, inflammation, exercise, exposure of the cell to toxins,
oxidants, heavy metals, starvation, hypoxia, water deprivation and
tissue trauma.
[0006] It is known that heat shock proteins play a role in
responding to a large number of abnormal conditions in body
tissues, including viral infection, inflammation, malignant
transformations, exposure to oxidizing agents, cytotoxins, and
anoxia. Several heat shock proteins function as intra-cellular
chaperones for other proteins and members of the HSP family are
expressed or activated at low to moderate levels because of their
essential role in protein maintenance and simply monitoring the
cell's proteins even under non-stressful conditions. These
activities are part of a cell's own repair system, called the
cellular stress response or the heat-shock response.
[0007] Heat shock proteins are found in nearly every cell and
tissue-type of multicellular organisms as well as in explanted
tissues and in cultured cells. The HSPs typically comprise 3%-10%
of a cell's proteins, although when under stress the percentage can
rise to 15%. The density of proteins of a mammalian cells has been
found to be in the range of (2-4).times.10.sup.18CM.sup.-3. Thus,
the aforementioned percentages mean that the density of HSPs is
normally (1-4).times.10.sup.17CM.sup.-3, while under stress the
density can rise to (3-6).times.10.sup.17CM.sup.-3.
[0008] Heat shock proteins are typically named according to their
molecular weight, and act in different ways. An especially
ubiquitous heat shock protein is Hsp70, a protein with a molecular
weight of 70 killodaltons. It plays a particularly significant role
in protecting proteins that are just being formed and in rescuing
damaged proteins. It contains a groove with an affinity for
neutral, hydrophobic amino acid residues that can interact with
peptides up to 7 residues in length. Hsp70 has peptide-binding and
ATPase domains that stabilize protein structures in unfolded and
assembly-competent states. The HSPs play a role in preventing
aggregation of misfolded proteins, many of which have exposed
hydrophobic portions, and a facilitating the refolding of proteins
into their proper conformations. Hsp70 accomplishes this by first
binding to the misfolded or fragmented protein, a binding that is
made energetically possible by a site that binds ATP and hydrolyzes
it into ADP.
[0009] Hsp70 heat shock proteins are a member of extracellular and
membrane bound heat-shock proteins which are involved in binding
antigens and presenting them to the immune system. Hsp70 has been
found to inhibit the activity of influenza A virus
ribonucleoprotein and to block the replication of the virus. Heat
shock proteins derived from tumors elicit specific protective
immunity. Experimental and clinical observations have shown that
heat shock proteins are involved in the regulation of autoimmune
arthritis, type 1 diabetes, mellitus, arterial sclerosis, multiple
sclerosis, and other autoimmune reactions.
[0010] Accordingly, it is believed that it is advantageous to be
able to selectively and controllably raise a target tissue
temperature up to a predetermined temperature range over a short
period of time, while maintaining the average temperature rise of
the tissue at a predetermined temperature over a longer period of
time. It is believed that this induces the heat shock response in
order to increase the number or activity of heat shock proteins in
body tissue in response to infection or other abnormalities.
However, this must be done in a controlled manner in order not to
damage or destroy the tissue or the area of the body being treated.
It would also be desirable to maximize the amount of heat shock
protein activation within the cells of a targeted tissue during a
single treatment session. The present invention fulfills these
needs, and provides other related advantages.
SUMMARY OF THE INVENTION
[0011] The present invention is directed to a process for heat
treating biological tissues by applying treatment energy to a
target tissue to therapeutically treat the target tissue. A first
treatment to the target tissue is performed by generating treatment
energy and repeatedly applying the treatment energy to the target
tissue over a period of time so as to controllably raise a
temperature of the target tissue to therapeutically treat the
target tissue without destroying or permanently damaging the target
tissue. The generated treatment energy may be pulsed or rapidly
applied in succession. The target tissue may comprise retinal
tissue.
[0012] The energy parameters are selected so as to raise a target
tissue temperature up to 11.degree. C. to achieve a therapeutic
effect, wherein the average temperature rise of the tissue over
several minutes is maintained at or below a predetermined level so
as not to permanently damage the target tissue. The energy
parameters may be selected so that the target tissue temperature is
raised between approximately 6.degree. C. to 11.degree. C. at least
during application of the energy to the target tissue. The average
temperature rise of the target tissue over several minutes is
maintained at 6.degree. C. or less, such as at approximately
1.degree. C. or less over several minutes.
[0013] The treatment energy and application parameters are selected
such so as to therapeutically treat the target tissue without
destroying or permanently damaging the target tissue. The selected
energy and application parameters may comprise tissue application
spot size or area, average power or average power density, and
exposure duration. Other parameters which may be selected include
wavelength or frequency and duty cycle. For example, the treatment
energy and application parameters may be selected to have an
average power density of 100-590 watts per square centimeter of
target tissue, a target tissue application spot size between
100-500 microns, and a train exposure duration of 500 milliseconds
or less.
[0014] The treatment energy may comprise a light beam, a microwave,
a radiofrequency or an ultrasound. A device may be inserted into a
cavity of the body in order to apply the treatment energy to the
tissue. The treatment energy may be applied to an exterior area of
a body which is adjacent to the target tissue, or has a blood
supply close to a surface of the exterior area of the body.
[0015] The treatment energy may comprise a radiofrequency between
approximately 3 to 6 megahertz (MHz). It may have a duty cycle of
between approximately 2.5% to 5%. It may have a pulsed train
duration of between approximately 0.2 to 0.4 seconds. The
radiofrequency may be generated with a device having a coil radii
of between approximately 2 and 6 mm and approximately 13 and 57 amp
turns.
[0016] The treatment energy may comprise a microwave frequency of
between 10 to 20 gigahertz (GHz). The microwave may have a pulse
train duration of approximately between 0.2 and 0.6 seconds. The
microwave may have a duty cycle of between approximately 2% and 5%.
The microwave may have an average power of between approximately 8
and 52 watts.
[0017] The treatment energy may comprise a pulsed light beam, such
as one or more laser light beams. The light beam may have a
wavelength of between approximately 570 nm to 1300 nm, and more
preferably between 600 nm and 1000 nm. The pulsed light beam may
have a power of between approximately 0.5 and 74 watts. The pulsed
light beam has a duty cycle of less than 10%, and preferably
between 2.5% and 5%. The pulsed light beam may have a pulse train
duration of approximately 0.1 and 0.6 seconds.
[0018] The treatment energy may comprise a pulsed ultrasound,
having a frequency of between approximately 1 and 5 MHz. The
ultrasound has a train duration of approximately 0.1 and 05
seconds. The ultrasound may have a duty cycle of between
approximately 2% and 10%. The ultrasound has a power of between
approximately 0.46 and 28.6 watts.
[0019] The process of the present invention may comprise the steps
of providing a plurality of energy emitters formed into an array.
Treatment energy is generated from the plurality of emitters. The
treatment energy is applied to the target tissue, wherein the
treatment energy has energy and application parameters selected so
as to raise the target tissue temperature sufficiently to create a
therapeutic effect while maintaining an average temperature of the
target tissue over several minutes at or below a predetermined
temperature so as not to destroy or permanently damage the target
tissue.
[0020] The first treatment comprises applying the treatment energy
to the target tissue for a period of less than ten seconds, and
more typically less than one second. The first treatment creates a
level of heat shock protein activation in the target tissue. The
application of the treatment energy to the target tissue is halted
for an interval of time that preferably exceeds the period of time
of the first treatment. The interval of time may comprise several
seconds to several minutes, such as three seconds to three minutes,
or preferably between ten seconds to ninety seconds. After the
interval of time and within a single treatment session, a second
treatment is performed to the target tissue by repeatedly
reapplying the treatment energy to the target tissue so as to
controllably raise the temperature of the target tissue to
therapeutically treat the target tissue without destroying or
permanently damaging the target tissue. The second treatment
increases the level of heat shock protein activation in the target
tissue such that it is at a level which is higher than the level
after the first treatment.
[0021] During an interval of time, typically comprising less than
one second, between applications of treatment energy applied to a
first area of the target tissue, the treatment energy may be
applied to a second area of the target tissue sufficiently spaced
apart from the first area of the target tissue to avoid thermal
tissue damage of the target tissue. The treatment energy is
repeatedly applied, in an alternating manner during the same
treatment session, to each of the first and second areas of the
target tissue until the predetermined number of energy applications
to each of the first and second areas of the target tissue has been
achieved.
[0022] When utilizing an array, a phase delay in the activation of
the energy emitters of the array may be introduced to generate
treatment energy in a phased manner using a predetermined delay of
activation in order to apply treatment energy to each of the first
and second areas of the target tissue. Alternatively, the energy
emitters of the array may be activated sequentially in order to
apply treatment energy to each of the first and second areas of the
target tissue.
[0023] Other features and advantages of the present invention will
become apparent from the following more detailed description, taken
in conjunction with the accompanying drawings, which illustrate, by
way of example, the principles of the invention.
BRIEF DESCRIPTION OF THE DRAWINGS
[0024] The accompanying drawings illustrate the invention. In such
drawings:
[0025] FIGS. 1A and 1B are graphs illustrating the average power of
a laser source compared to a source radius and pulse train duration
of the laser;
[0026] FIGS. 2A and 2B are graphs illustrating the time for the
temperature to decay depending upon the laser source radius and
wavelength;
[0027] FIGS. 3-6 are graphs illustrating the peak ampere turns for
various radiofrequencies, duty cycles, and coil radii;
[0028] FIG. 7 is a graph depicting the time for temperature rise to
decay compared to radiofrequency coil radius;
[0029] FIGS. 8 and 9 are graphs depicting the average microwave
power compared to microwave frequency and pulse train
durations;
[0030] FIG. 10 is a graph depicting the time for the temperature to
decay for various microwave frequencies;
[0031] FIG. 11 is a graph depicting the average ultrasound source
power compared to frequency and pulse train duration;
[0032] FIGS. 12 and 13 are graphs depicting the time for
temperature decay for various ultrasound frequencies;
[0033] FIG. 14 is a graph depicting the volume of focal heated
region compared to ultrasound frequency;
[0034] FIG. 15 is a graph comparing equations for temperature over
pulse durations for an ultrasound energy source;
[0035] FIGS. 16 and 17 are graphs illustrating the magnitude of the
logarithm of damage and HSP activation Arrhenius integrals as a
function of temperature and pulse duration;
[0036] FIG. 18 is a diagrammatic view of a light generating unit
that produces timed series of pulses, having a light pipe extending
therefrom, in accordance with the present invention;
[0037] FIG. 19 is a cross-sectional view of a photostimulation
delivery device delivering electromagnetic energy to target tissue,
in accordance with the present invention;
[0038] FIG. 20 is a diagrammatic view illustrating a system used to
generate a laser light beam, in accordance with the present
invention;
[0039] FIG. 21 is a diagrammatic view of optics used to generate a
laser light geometric pattern, in accordance with the present
invention;
[0040] FIG. 22 is a top plan view of an optical scanning mechanism,
used in accordance with the present invention;
[0041] FIG. 23 is a partially exploded view of the optical scanning
mechanism of FIG. 22, illustrating the various component parts
thereof;
[0042] FIG. 24 illustrates controlled offsets of exposure of an
exemplary geometric pattern grid of laser spots to treat the target
tissue, in accordance with an embodiment of the present
invention;
[0043] FIG. 25 is a diagrammatic view illustrating the use of a
geometric object in the form of a line controllably scanned to
treat an area of the target tissue;
[0044] FIG. 26 is a diagrammatic view similar to FIG. 25, but
illustrating the geometric line or bar rotated to treat the target
tissue;
[0045] FIG. 27 is a diagrammatic view illustrating an alternate
embodiment of the system used to generate laser light beams for
treating tissue, in accordance with the present invention;
[0046] FIG. 28 is a diagrammatic view illustrating yet another
embodiment of a system used to generate laser light beams to treat
tissue in accordance with the present invention;
[0047] FIG. 29 is a cross-sectional and diagrammatic view of an end
of an endoscope inserted into the nasal cavity and treating tissue
therein, in accordance with the present invention;
[0048] FIG. 30 is a diagrammatic and partially cross-sectioned view
of a bronchoscope extending through the trachea and into the
bronchus of a lung and providing treatment thereto, in accordance
with the present invention;
[0049] FIG. 31 is a diagrammatic view of a colonoscope providing
photostimulation to an intestinal or colon area of the body, in
accordance with the present invention;
[0050] FIG. 32 is a diagrammatic view of an endoscope inserted into
a stomach and providing treatment thereto, in accordance with the
present invention;
[0051] FIG. 33 is a partially sectioned perspective view of a
capsule endoscope, used in accordance with the present
invention;
[0052] FIG. 34 is a diagrammatic view of a pulsed high intensity
focused ultrasound for treating tissue internal the body, in
accordance with the present invention;
[0053] FIG. 35 is a diagrammatic view for delivering therapy to the
bloodstream of a patient, through an earlobe, in accordance with
the present invention;
[0054] FIG. 36 is a cross-sectional view of a stimulating therapy
device of the present invention used in delivering photostimulation
to the blood, via an earlobe, in accordance with the present
invention;
[0055] FIGS. 37A-37D are diagrammatic views illustrated in the
application of micropulsed energy to different treatment areas
during a predetermined interval of time, within a single treatment
session, and reapplying the energy to previously treated areas, in
accordance with the present invention;
[0056] FIGS. 38-40 are graphs depicting the relationship of
treatment power and time in accordance with the embodiments of the
present invention;
[0057] FIG. 41 is a graph depicting wavefront from two sources
separated by a distance;
[0058] FIG. 42 is a depiction of a square array of square antennas
or sources, which can be used in accordance with the present
invention;
[0059] FIG. 43 is a graph depicting the shape of radiation pattern
from a square antenna array;
[0060] FIG. 44 is a graph depicting a form of typical radiation
pattern along an X-axis from a far field array;
[0061] FIG. 45 is a graph depicting an envelope of the pattern of
FIG. 44;
[0062] FIG. 46 is another graph depicting the width of individual
lines of the pattern of FIG. 44;
[0063] FIG. 47 is a plot graph depicting the determinant of the
line separation;
[0064] FIG. 48 is a block diagram of components of a steerable
array system;
[0065] FIG. 49 is a plot graph showing induced tissue temperature
rise and drops;
[0066] FIGS. 50-52 are graphs depicting variables of three
different coils, in accordance with the present invention;
[0067] FIG. 53 is a graph depicting the plots of FIGS. 50-52
superimposed upon one another;
[0068] FIG. 54 is a block diagram for an induction array which can
be used in accordance with the present invention;
[0069] FIGS. 55A and 55B are graphs depicting the behavior of HSP
cellular system components over time following a sudden increase in
temperature;
[0070] FIGS. 56A-56H are graphs depicting the behavior of HSP
cellular system components in the first minute following a sudden
increase in temperature;
[0071] FIGS. 57A and 57B are graphs illustrating variation in the
activated concentrations of HSP and unactivated HSP in the
cytoplasmic reservoir over an interval of one minute, in accordance
with the present invention; and
[0072] FIG. 58 is a graph depicting the improvement ratios versus
interval between treatments, in accordance with the present
invention.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0073] As shown in the accompanying drawings, and as more fully
described herein, the present invention is directed to a system and
method for delivering a pulsed energy, such as ultrasound,
ultraviolet radiofrequency, microwave radiofrequency, one or more
light beams, and the like, having energy parameters selected to
cause a thermal time-course in tissue to raise the tissue
temperature over a short period of time to a sufficient level to
achieve a therapeutic effect while maintaining an average tissue
temperature over a prolonged period of time below a predetermined
level so as to avoid permanent tissue damage. It is believed that
the creation of the thermal time-course stimulates heat shock
protein activation or production and facilitates protein repair
without causing any damage.
[0074] The inventors have discovered that electromagnetic radiation
can be applied to retinal tissue in a manner that does not destroy
or damage the retinal tissue while achieving beneficial effects on
eye diseases. More particularly, a laser light beam can be
generated that is therapeutic, yet sublethal to retinal tissue
cells and thus avoids damaging photocoagulation in the retinal
tissue which provides preventative and protective treatment of the
retinal tissue of the eye. It is believed that this may be due, at
least in part, to the stimulation and activation of heat shock
proteins and the facilitation of protein repair in the retinal
tissue. This is disclosed in U.S. patent application Ser. No.
14/607,959 filed Jan. 28, 2015, Ser. No. 13/798,523 filed Mar. 13,
2013, and Ser. No. 13/481,124 filed May 25, 2012, the contents of
which are hereby incorporated by reference as if made in full.
[0075] Various parameters of the light beam must be taken into
account and selected so that the combination of the selected
parameters achieve the therapeutic effect while not permanently
damaging the tissue. These parameters include laser wavelength,
radius of the laser source or tissue application spot, laser power,
total pulse train duration, and duty cycle of the pulse train.
[0076] The selection of these parameters may be determined by
requiring that the Arrhenius integral for HSP activation be greater
than 1 or unity. Arrhenius integrals are used for analyzing the
impacts of actions on biological tissue. See, for instance, The CRC
Handbook of Thermal Engineering, ed. Frank Kreith, Springer Science
and Business Media (2000). At the same time, the selected
parameters must not permanently damage the tissue. Thus, the
Arrhenius integral for damage may also be used, wherein the solved
Arrhenius integral is less than 1 or unity. Alternatively, the
FDA/FCC constraints on energy deposition per unit gram of tissue
and temperature rise as measured over periods of minutes be
satisfied so as to avoid permanent tissue damage. The FDA/FCC
requirements on energy deposition and temperature rise are widely
used and can be referenced, for example, at
www.fda.gov/medicaldevices/deviceregulationandguidance/guidancedocument
s/ucm073817.htm#attacha for electromagnetic sources, and Anastosio
and P. LaRivero, ed., Emerging Imaging Technologies. CRC Press
(2012), for ultrasound sources. Generally speaking, tissue
temperature rises of between 6.degree. C. and 11.degree. C. can
create therapeutic effect, such as by activating heat shock
proteins, whereas maintaining the average tissue temperature over a
prolonged period of time, such as over several minutes, such as six
minutes, below a predetermined temperature, such as 6.degree. C.
and even 1.degree. C. or less in certain circumstances, will not
permanently damage the tissue.
[0077] The inventors have discovered that generating a
subthreshold, sublethal micropulse laser light beam which has a
wavelength greater than 532 nm and a duty cycle of less than 10% at
a predetermined intensity or power and a predetermined pulse length
or exposure time creates desirable retinal photostimulation without
any visible burn areas or tissue destruction. More particularly, a
laser light beam having a wavelength of between 570 nm-1300 nm, and
in a particularly preferred embodiment between 600 nm and 1100 nm,
having a duty cycle of approximately 2.5%-10% and a predetermined
average power or power intensity (such as between 100-590 watts per
square centimeter at the retina or approximately 1 watt per laser
spot for each treatment spot at the retina) and a predetermined
pulse train length or exposure time (such as between 100 and 600
milliseconds or less) creates a sublethal, "true subthreshold"
retinal photostimulation in which all areas of the retinal pigment
epithelium exposed to the laser irradiation are preserved and
available to contribute therapeutically. In other words, the
inventors have found that raising the retinal tissue at least up to
a therapeutic level but below a cellular or tissue lethal level
recreates the benefit of the halo effect of the prior art methods
without destroying, burning or otherwise damaging the retinal
tissue. This is referred to herein as subthreshold diode micropulse
laser treatment (SDM).
[0078] SDM does not produce laser-induced retinal damage
(photocoagulation), and has no known adverse treatment effect, and
has been reported to be an effective treatment in a number of
retinal disorders (including diabetic macular edema (DME)
proliferative diabetic retinopathy (PDR), macular edema due to
branch retinal vein occlusion (BRVO), central serous
chorioretinopathy (CSR), reversal of drug tolerance, and
prophylactic treatment of progressive degenerative retinopathies
such as dry age-related macular degeneration, Stargardts' disease,
cone dystrophies, and retinitis pigmentosa. The safety of SDM is
such that it may be used transfoveally in eyes with 20/20 visual
acuity to reduce the risk of visual loss due to early
fovea-involving DME.
[0079] A mechanism through which SDM might work is the generation
or activation of heat shock proteins (HSPs). Despite a near
infinite variety of possible cellular abnormalities, cells of all
types share a common and highly conserved mechanism of repair: heat
shock proteins (HSPs). HSPs are elicited almost immediately, in
seconds to minutes, by almost any type of cell stress or injury. In
the absence of lethal cell injury, HSPs are extremely effective at
repairing and returning the viable cell toward a more normal
functional state. Although HSPs are transient, generally peaking in
hours and persisting for a few days, their effects may be long
lasting. HSPs reduce inflammation, a common factor in many
disorders.
[0080] Laser treatment can induce HSP production or activation and
alter cytokine expression. The more sudden and severe the
non-lethal cellular stress (such as laser irradiation), the more
rapid and robust HSP activation. Thus, a burst of repetitive low
temperature thermal spikes at a very steep rate of change
(.about.7.degree. C. elevation with each 100 .mu.s micropulse, or
70,000.degree. C./sec) produced by each SDM exposure is especially
effective in stimulating activation of HSPs, particularly compared
to non-lethal exposure to subthreshold treatment with continuous
wave lasers, which can duplicate only the low average tissue
temperature rise.
[0081] Laser wavelengths below 550 nm produce increasingly
cytotoxic photochemical effects. The lower wavelength limit
realistically usable by the process of the present invention is
determined by the undesirable absorption by the visual pigments and
other absorbers, including blood, the lens of the eye, etc. At
approximately 570 nm, the sum of the optical densities of the long
wavelength sensitive and medium wavelength sensitive visual
pigments in the eye and the blood exceeds the optical density of
the melanin. The absorption is dominated by melanin between 570 nm
and 650 nm, where above 650 nm the absorption is practically all
due to the melanin in the RPE. However, at higher wavelengths, such
as above 1300 nm, there is a decrease in melanin absorption with
increasing absorption by the water in the vitreous of the eye. At
1300 nm, for instance, the melanin absorbance is only 0.048 of what
it is at 810 nm, and the radiation power due to this effect alone
would have to be increased by a factor of 20 compared to the power
at 810 nm to achieve the same temperature increase. Accordingly,
the present invention can be performed at a broad range of
wavelengths between 570 nm to 1300 nm, with the more preferable
range of wavelengths being 600 nm to 1100 nm, and an even more
preferable range of wavelengths of 650 nm to 900 nm, with the
particularly preferred operating wavelength at approximately 810
nm. At these wavelengths, the melanin absorption is dominant and
the heating primarily in the desired RPE and the wavelength is at a
safe distance from the wavelengths where appreciable absorption
occurs in the visual pigments as shorter wavelengths or water at
longer wavelengths, which will create undesirable heating of the
eye and other tissues. At 810 nm, SDM produces photothermal, rather
than photochemical, cellular stress. Thus, SDM is able to affect
the tissue without damaging it.
[0082] It has been found that the average required treatment power
between tissue reset and tissue damage can be calculated with the
wavelength used, the radiation train duration, preferably being
between 0.03 and 0.8 seconds and a retinal application spot by the
radiation being between 10 and 500 microns. A duty cycle of less
than 10% and preferably between 2.5% and 5% with a total pulse
duration of between 100 milliseconds and 600 milliseconds has been
found to be effective. The corresponding peak powers, during the
individual pulse, are obtained from the average powers by dividing
by the duty cycle. The average power can vary between 0.0000069 to
37.5 watts within a wavelength between 570 nm-1300 nm, a pulse
train duration between 30-800 milliseconds, and a treatment spot
between 10-700 microns.
[0083] The clinical benefits of SDM are thus primarily produced by
sub-morbid photothermal cellular HSP activation. In dysfunctional
cells, HSP stimulation by SDM results in normalized cytokine
expression, and consequently improved structure and function. The
therapeutic effects of this "low-intensity" laser/tissue
interaction are then amplified by "high-density" laser application,
recruiting all the dysfunctional cells in the targeted tissue area
by densely/confluently treating a large tissue area, including all
areas of pathology, thereby maximizing the treatment effect. These
principles define the treatment strategy of SDM described
herein.
[0084] Because normally functioning cells are not in need of
repair, HSP stimulation in normal cells would tend to have no
notable clinical effect. The "patho-selectivity" of near infrared
laser effects, such as SDM, affecting sick cells but not affecting
normal ones, on various cell types is consistent with clinical
observations of SDM. SDM has been reported to have a clinically
broad therapeutic range, unique among retinal laser modalities,
consistent with American National Standards Institute "Maximum
Permissible Exposure" predictions. While SDM may cause direct
photothermal effects such as entropic protein unfolding and
disaggregation, SDM appears optimized for clinically safe and
effective stimulation of HSP-mediated repair.
[0085] As noted above, while SDM stimulation of HSPs is
non-specific with regard to the disease process, the result of HSP
mediated repair is by its nature specific to the state of the
dysfunction. HSPs tend to fix what is wrong, whatever that might
be. Thus, the observed effectiveness of SDM in retinal conditions
as widely disparate as BRVO, DME, PDR, CSR, age-related and genetic
retinopathies, and drug-tolerant NAMD. Conceptually, this facility
can be considered a sort of "Reset to Default" mode of SDM action.
For the wide range of disorders in which cellular function is
critical, SDM normalizes cellular function by triggering a "reset"
(to the "factory default settings") via HSP-mediated cellular
repair.
[0086] The inventors have found that SDM treatment of patients
suffering from age-related macular degeneration (AMD) can slow the
progress or even stop the progression of AMD. Most of the patients
have seen significant improvement in dynamic functional log MAR
mesoptic visual acuity and mesoptic contrast visual acuity after
the SDM treatment. It is believed that SDM works by targeting,
preserving, and "normalizing" (moving toward normal) function of
the retinal pigment epithelium (RPE).
[0087] SDM has also been shown to stop or reverse the
manifestations of the diabetic retinopathy disease state without
treatment-associated damage or adverse effects, despite the
persistence of systemic diabetes mellitus. On this basis it is
hypothesized that SDM might work by inducing a return to more
normal cell function and cytokine expression in diabetes-affected
RPE cells, analogous to hitting the "reset" button of an electronic
device to restore the factory default settings. Based on the above
information and studies, SDM treatment may directly affect cytokine
expression via heat shock protein (HSP) activation in the targeted
tissue.
[0088] As heat shock proteins play a role in responding to a large
number of abnormal conditions in body tissue other than eye tissue,
it is believed that similar systems and methodologies can be
advantageously used in treating such abnormal conditions,
infections, etc. As such, the present invention is directed to the
controlled application of ultrasound or electromagnetic radiation
to treat abnormal conditions including inflammations, autoimmune
conditions, and cancers that are accessible by means of fiber
optics of endoscopes or surface probes as well as focused
electromagnetic/sound waves. For example, cancers on the surface of
the prostate that have the largest threat of metastasizing can be
accessed by means of fiber optics in a proctoscope. Colon tumors
can be accessed by an optical fiber system, like those used in
colonoscopy.
[0089] As indicated above, subthreshold diode micropulse laser
(SDM) photostimulation has been effective in stimulating direct
repair of slightly misfolded proteins in eye tissue. Besides HSP
activation, another way this may occur is because the spikes in
temperature caused by the micropulses in the form of a thermal
time-course allows diffusion of water inside proteins, and this
allows breakage of the peptide-peptide hydrogen bonds that prevent
the protein from returning to its native state. The diffusion of
water into proteins results in an increase in the number of
restraining hydrogen bonds by a factor on the order of a thousand.
Thus, it is believed that this process could be applied to other
tissues and diseases advantageously as well.
[0090] As explained above, the energy source to be applied to the
target tissue will have energy and operating parameters which must
be determined and selected so as to achieve the therapeutic effect
while not permanently damaging the tissue. Using a light beam
energy source, such as a laser light beam, as an example, the laser
wavelength, duty cycle and total pulse train duration parameters
must be taken into account. Other parameters which can be
considered include the radius of the laser source as well as the
average laser power. Adjusting or selecting one of these parameters
can have an effect on at least one other parameter.
[0091] FIGS. 1A and 1B illustrate graphs showing the average power
in watts as compared to the laser source radius (between 0.1 cm and
0.4 cm) and pulse train duration (between 0.1 and 0.6 seconds).
FIG. 1A shows a wavelength of 880 nm, whereas FIG. 1B has a
wavelength of 1000 nm. It can be seen in these figures that the
required power decreases monotonically as the radius of the source
decreases, as the total train duration increases, and as the
wavelength decreases. The preferred parameters for the radius of
the laser source is 1 mm-4 mm. For a wavelength of 880 nm, the
minimum value of power is 0.55 watts, with a radius of the laser
source being 1 mm, and the total pulse train duration being 600
milliseconds. The maximum value of power for the 880 nm wavelength
is 52.6 watts when the laser source radius is 4 mm and the total
pulse drain duration is 100 milliseconds. However, when selecting a
laser having a wavelength of 1000 nm, the minimum power value is
0.77 watts with a laser source radius of 1 mm and a total pulse
train duration of 600 milliseconds, and a maximum power value of
73.6 watts when the laser source radius is 4 mm and the total pulse
duration is 100 milliseconds. The corresponding peak powers, during
an individual pulse, are obtained from the average powers by
dividing by the duty cycle.
[0092] The volume of the tissue region to be heated is determined
by the wavelength, the absorption length in the relevant tissue,
and by the beam width. The total pulse duration and the average
laser power determine the total energy delivered to heat up the
tissue, and the duty cycle of the pulse train gives the associated
spike, or peak, power associated with the average laser power.
Preferably, the pulsed energy source energy parameters are selected
so that approximately 20 to 40 joules of energy is absorbed by each
cubic centimeter of the target tissue.
[0093] The absorption length is very small in the thin melanin
layer in the retinal pigmented epithelium. In other parts of the
body, the absorption length is not generally that small. In
wavelengths ranging from 400 nm to 2000 nm, the penetration depth
and skin is in the range of 0.5 mm to 3.5 mm. The penetration depth
into human mucous tissues is in the range of 0.5 mm to 6.8 mm.
Accordingly, the heated volume will be limited to the exterior or
interior surface where the radiation source is placed, with a depth
equal to the penetration depth, and a transverse dimension equal to
the transverse dimension of the radiation source. Since the light
beam energy source is used to treat diseased tissues near external
surfaces or near internal accessible surfaces, a source radii of
between 1 mm to 4 mm and operating a wavelength of 880 nm yields a
penetration depth of approximately 2.5 mm and a wavelength of 1000
nm yields a penetration depth of approximately 3.5 mm.
[0094] It has been determined that the target tissue can be heated
to up to approximately 11.degree. C. for a short period of time,
such as less than one second, to create the therapeutic effect of
the invention while maintaining the target tissue average
temperature to a lower temperature range, such as less than
6.degree. C. or even 1.degree. C. or less over a prolonged period
of time, such as several minutes. The selection of the duty cycle
and the total pulse train duration provide time intervals in which
the heat can dissipate. A duty cycle of less than 10%, and
preferably between 2.5% and 5%, with a total pulse duration of
between 100 milliseconds and 600 milliseconds has been found to be
effective. FIGS. 2A and 2B illustrate the time to decay from
10.degree. C. to 1.degree. C. for a laser source having a radius of
between 0.1 cm and 0.4 cm with the wavelength being 880 nm in FIG.
2A and 1000 nm in FIG. 2B. It can be seen that the time to decay is
less when using a wavelength of 880 nm, but either wavelength falls
within the acceptable requirements and operating parameters to
achieve the benefits of the present invention while not causing
permanent tissue damage.
[0095] It has been found that the average temperature rise of the
desired target region increasing at least 6.degree. C. and up to
11.degree. C., and preferably approximately 10.degree. C., during
the total irradiation period results in HSP activation. The control
of the target tissue temperature is determined by choosing source
and target parameters such that the Arrhenius integral for HSP
activation is larger than 1, while at the same time assuring
compliance with the conservative FDA/FCC requirements for avoiding
damage or a damage Arrhenius integral being less than 1.
[0096] In order to meet the conservative FDA/FCC constraints to
avoid permanent tissue damage, for light beams and other
electromagnetic radiation sources, the average temperature rise of
the target tissue over any six-minute period is 1.degree. C. or
less. FIGS. 2A and 2B above illustrate the typical decay times
required for the temperature in the heated target region to
decrease by thermal diffusion from a temperature rise of
approximately 10.degree. C. to 1.degree. C. as can be seen in FIG.
2A when the wavelength is 880 nm and the source diameter is 1
millimeter, the temperature decay time is 16 seconds. The
temperature decay time is 107 seconds when the source diameter is 4
mm. As shown in FIG. 2B, when the wavelength is 1000 nm, the
temperature decay time is 18 seconds when the source diameter is 1
mm and 136 seconds when the source diameter is 4 mm. This is well
within the time of the average temperature rise being maintained
over the course of several minutes, such as 6 minutes or less.
While the target tissue's temperature is raised, such as to
approximately 10.degree. C., very quickly, such as in a fraction of
a second during the application of the energy source to the tissue,
the relatively low duty cycle provides relatively long periods of
time between the pulses of energy applied to the tissue and the
relatively short pulse train duration ensure sufficient temperature
diffusion and decay within a relatively short period of time
comprising several minutes, such as 6 minutes or less, that there
is no permanent tissue damage.
[0097] The parameters differ for the individual energy sources,
including microwave, infrared lasers, radiofrequency and
ultrasound, because the absorption properties of tissues differ for
these different types of energy sources. The tissue water content
can vary from one tissue type to another, however, there is an
observed uniformity of the properties of tissues at normal or near
normal conditions which has allowed publication of tissue
parameters that are widely used by clinicians in designing
treatments. Below are tables illustrating the properties of
electromagnetic waves in biological media, with Table 1 relating to
muscle, skin and tissues with high water content, and Table 2
relating to fat, bone and tissues with low water content.
TABLE-US-00001 TABLE 1 Properties of Electromagnetic Waves in
Biological Media: Muscle, Skin, and Tissues with High Water Content
Wavelength Dielectric Conductivity Wavelength Depth of Reflection
Coefficient Frequency in Air Constant .sigma.H .lamda.H Penetration
Air-Muscle Interface Muscle-Fat Interface (MHz) (cm) H (mho/m) (cm)
(cm) r o r o 1 30000 2000 0.400 436 91.3 0.982 +179 10 3000 160
0.625 118 21.6 0.956 +178 27.12 1106 113 0.612 68.1 14.3 0.925 +177
0.651 -11.13 40.68 738 97.3 0.693 51.3 11.2 0.913 +176 0.652 -10.21
100 300 71.7 0.889 27 6.66 0.881 +175 0.650 -7.96 200 150 56.5 1.28
16.6 4.79 0.844 +175 0.612 -8.06 300 100 54 1.37 11.9 3.89 0.825
+175 0.592 -8.14 433 69.3 53 1.43 8.76 3.57 0.803 +175 0.562 -7.06
750 40 52 1.54 5.34 3.18 0.779 +176 0.532 -5.69 915 32.8 51 1.60
4.46 3.04 0.772 +177 0.519 -4.32 1500 20 49 1.77 2.81 2.42 0.761
+177 0.506 -3.66 2450 12.2 47 2.21 1.76 1.70 0.754 +177 0.500 -3.88
3000 10 46 2.26 1.45 1.61 0.751 +178 0.495 -3.20 5000 6 44 3.92
0.89 0.788 0.749 +177 0.502 -4.95 5800 5.17 43.3 4.73 0.775 0.720
0.746 +177 0.502 -4.29 8000 3.75 40 7.65 0.578 0.413 0.744 +176
0.513 -6.65 10000 3 39.9 10.3 0.464 0.343 0.743 +176 0.518
-5.95
TABLE-US-00002 TABLE 2 Properties of Electromagnetic Waves in
Biological Media: Fat, Bone, and Tissues with Low Water Content
Wavelength Dielectric Conductivity Wavelength Depth of Reflection
Coefficient Frequency in Air Constant .sigma.L, .lamda.L
Penetration Air-Fat Interface Fat-Muscle Interface (MHz) (cm) L
(mmho/m) (cm) (cm) r o r o 1 30000 10 3000 27.12 1106 20 10.9-43.2
241 159 0.660 +174 0.651 +169 40.68 738 14.6 12.6-52.8 187 118
0.617 +173 0.652 +170 100 300 7.45 19.1-75.9 106 60.4 0.511 +168
0.650 +172 200 150 5.95 25.8-94.2 59.7 39.2 0.458 +168 0.612 +172
300 100 5.7 31.6-107 41 32.1 0.438 +169 0.592 +172 433 69.3 5.6
37.9-118 28.8 26.2 0.427 +170 0.562 +173 750 40 5.6 49.8-138 16.8
23 0.415 +173 0.532 +174 915 32.8 5.6 55.6-147 13.7 17.7 0.417 +173
0.519 +176 1500 20 5.6 70.8-171 8.41 13.9 0.412 +174 0.506 +176
2450 12.2 5.5 96.4-213 5.21 11.2 0.406 +176 0.500 +176 3000 10 5.5
110-234 4.25 9.74 0.406 +176 0.495 +177 5000 6 5.5 162-309 2.63
6.67 0.393 +176 0.502 +175 5900 5.17 5.05 186-338 2.29 5.24 0.388
+176 0.502 +176 8000 3.75 4.7 255-431 1.73 4.61 0.371 +176 0.513
+173 - 10000 3 4.5 324-549 1.41 3.39 0.363 +175 0.518 +174, -
[0098] The absorption lengths of radiofrequency in body tissue are
long compared to body dimensions. Consequently, the heated region
is determined by the dimensions of the coil that is the source of
the radiofrequency energy rather than by absorption lengths. Long
distances r from a coil the magnetic (near) field from a coil drops
off as 1/r.sup.3. At smaller distances, the electric and magnetic
fields can be expressed in terms of the vector magnetic potential,
which in turn can be expressed in closed form in terms of elliptic
integrals of the first and second kind. The heating occurs only in
a region that is comparable in size to the dimensions of the coil
source itself. Accordingly, if it is desired to preferentially heat
a region characterized by a radius, the source coil will be chosen
to have a similar radius. The heating drops off very rapidly
outside of a hemispherical region of radius because of the
1/r.sup.3 drop off of the magnetic field. Since it is proposed to
use the radiofrequency the diseased tissue accessible only
externally or from inner cavities, it is reasonable to consider a
coil radii of between approximately 2 to 6 mm.
[0099] The radius of the source coil(s) as well as the number of
ampere turns (NI) in the source coils give the magnitude and
spatial extent of the magnetic field, and the radiofrequency is a
factor that relates the magnitude of the electric field to the
magnitude of the magnetic field. The heating is proportional to the
product of the conductivity and the square of the electric field.
For target tissues of interest that are near external or internal
surfaces, the conductivity is that of skin and mucous tissue. The
duty cycle of the pulse train as well as the total train duration
of a pulse train are factors which affect how much total energy is
delivered to the tissue.
[0100] Preferred parameters for a radiofrequency energy source have
been determined to be a coil radii between 2 and 6 mm,
radiofrequencies in the range of 3-6 MHz, total pulse train
durations of 0.2 to 0.4 seconds, and a duty cycle of between 2.5%
and 5%. FIGS. 3-6 show how the number of ampere turns varies as
these parameters are varied in order to give a temperature rise
that produces an Arrhenius integral of approximately one or unity
for HSP activation. With reference to FIG. 3, for an RF frequency
of 6 MHz, a pulse train duration of between 0.2 and 0.4 seconds, a
coil radius between 0.2 and 0.6 cm, and a duty cycle of 5%, the
peak ampere turns (NI) is 13 at the 0.6 cm coil radius and 20 at
the 0.2 cm coil radius. For a 3 MHz frequency, as illustrated in
FIG. 4, the peak ampere turns is 26 when the pulse train duration
is 0.4 seconds and the coil radius is 0.6 cm and the duty cycle is
5%. However, with the same 5% duty cycle, the peak ampere turns is
40 when the coil radius is 0.2 cm and the pulse train duration is
0.2 seconds. A duty cycle of 2.5% is used in FIGS. 5 and 6. This
yields, as illustrated in FIG. 5, 18 amp turns for a 6 MHz
radiofrequency having a coil radius of 0.6 cm and a pulse train
duration of 0.4 seconds, and 29 amp turns when the coil radius is
only 0.2 cm and the pulse train duration is 0.2 seconds. With
reference to FIG. 6, with a duty cycle of 2.5% and a radiofrequency
of 3 MHz, the peak ampere turns is 36 when the pulse train duration
is 0.4 seconds and the coil radius is 0.6 cm, and 57 amp turns when
the pulse train duration is 0.2 seconds and the coil radius is 0.2
cm.
[0101] The time, in seconds, for the temperature rise to decay from
approximately 10.degree. C. to approximately 1.degree. C. for coil
radii between 0.2 cm and 0.6 cm is illustrated for a radiofrequency
energy source in FIG. 7. The temperature decay time is
approximately 37 seconds when the radiofrequency coil radius is 0.2
cm, and approximately 233 seconds when the radiofrequency coil
radius is 0.5 cm. When the radiofrequency coil radius is 0.6 cm,
the decay time is approximately 336 seconds, which is still within
the acceptable range of decay time, but at an upper range
thereof.
[0102] Microwaves are another electromagnetic energy source which
can be utilized in accordance with the present invention. The
frequency of the microwave determines the tissue penetration
distance. The gain of a conical microwave horn is large compared to
the microwave wavelength, indicating under those circumstances that
the energy is radiated mostly in a narrow forward load. Typically,
a microwave source used in accordance with the present invention
has a linear dimension on the order of a centimeter or less, thus
the source is smaller than the wavelength, in which case the
microwave source can be approximated as a dipole antenna. Such
small microwave sources are easier to insert into internal body
cavities and can also be used to radiate external surfaces. In that
case, the heated region can be approximated by a hemisphere with a
radius equal to the absorption length of the microwave in the body
tissue being treated. As the microwaves are used to treat tissue
near external surfaces or surfaces accessible from internal
cavities, frequencies in the 10-20 GHz range are used, wherein the
corresponding penetration distances are only between approximately
2 and 4 mm.
[0103] The temperature rise of the tissue using a microwave energy
source is determined by the average power of the microwave and the
total pulse train duration. The duty cycle of the pulse train
determines the peak power in a single pulse in a train of pulses.
As the radius of the source is taken to be less than approximately
1 centimeter, and frequencies between 10 and 20 GHz are typically
used, a resulting pulse train duration of 0.2 and 0.6 seconds is
preferred.
[0104] The required power decreases monotonically as the train
duration increases and as the microwave frequency increases. For a
frequency of 10 GHz, the average power is 18 watts when the pulse
train duration is 0.6 seconds, and 52 watts when the pulse train
duration is 0.2 seconds. For a 20 GHz microwave frequency, an
average power of 8 watts is used when the pulse train is 0.6
seconds, and can be 26 watts when the pulse train duration is only
0.2 seconds. The corresponding peak power are obtained from the
average power simply by dividing by the duty cycle.
[0105] With reference now to FIG. 8, a graph depicts the average
microwave power in watts of a microwave having a frequency of 10
GHz and a pulse train duration from between 0.2 seconds and 0.6
seconds. FIG. 9 is a similar graph, but showing the average
microwave power for a microwave having a frequency of 20 GHz. Thus,
it will be seen that the average microwave source power varies as
the total train duration and microwave frequency vary. The
governing condition, however, is that the Arrhenius integral for
HSP activation in the heated region is approximately 1.
[0106] With reference to FIG. 10, a graph illustrates the time, in
seconds, for the temperature to decay from approximately 10.degree.
C. to 1.degree. C. compared to microwave frequencies between 58 MHz
and 20000 MHz. The minimum and maximum temperature decay for the
preferred range of microwave frequencies are 8 seconds when the
microwave frequency is 20 GHz, and 16 seconds when the microwave
frequency is 10 GHz.
[0107] Utilizing ultrasound as an energy source enables heating of
surface tissue, and tissues of varying depths in the body,
including rather deep tissue. The absorption length of ultrasound
in the body is rather long, as evidenced by its widespread use for
imaging. Accordingly, ultrasound can be focused on target regions
deep within the body, with the heating of a focused ultrasound beam
concentrated mainly in the approximately cylindrical focal region
of the beam. The heated region has a volume determined by the focal
waist of the airy disc and the length of the focal waist region,
that is the confocal parameter. Multiple beams from sources at
different angles can also be used, the heating occurring at the
overlapping focal regions.
[0108] For ultrasound, the relevant parameters for determining
tissue temperature are frequency of the ultrasound, total train
duration, and transducer power when the focal length and diameter
of the ultrasound transducer is given. The frequency, focal length,
and diameter determine the volume of the focal region where the
ultrasound energy is concentrated. It is the focal volume that
comprises the target volume of tissue for treatment. Transducers
having a diameter of approximately 5 cm and having a focal length
of approximately 10 cm are readily available. Favorable focal
dimensions are achieved when the ultrasound frequency is between 1
and 5 MHz, and the total train duration is 0.1 to 0.5 seconds. For
example, for a focal length of 10 cm and the transducer diameter of
5 cm, the focal volumes are 0.02 cc at 5 MHz and 2.36 cc at 1
MHz.
[0109] With reference now to FIG. 11, a graph illustrates the
average source power in watts compared to the frequency (between 1
MHz and 5 MHz), and the pulse train duration (between 0.1 and 0.5
seconds). A transducer focal length of 10 cm and a source diameter
of 5 cm have been assumed. The required power to give the Arrhenius
integral for HSP activation of approximately 1 decreases
monotonically as the frequency increases and as the total train
duration increases. Given the preferred parameters, the minimum
power for a frequency of 1 GHz and a pulse train duration of 0.5
seconds is 5.72 watts, whereas for the 1 GHz frequency and a pulse
train duration of 0.1 seconds the maximum power is 28.6 watts. For
a 5 GHz frequency, 0.046 watts is required for a pulse train
duration of 0.5 seconds, wherein 0.23 watts is required for a pulse
train duration of 0.1 seconds. The corresponding peak power during
an individual pulse is obtained simply by dividing by the duty
cycle.
[0110] FIG. 12 illustrates the time, in seconds, for the
temperature to diffuse or decay from 10.degree. C. to 6.degree. C.
when the ultrasound frequency is between 1 and 5 MHz. FIG. 13
illustrates the time, in seconds, to decay from approximately
10.degree. C. to approximately 1.degree. C. for ultrasound
frequencies from 1 to 5 MHz. For the preferred focal length of 10
cm and the transducer diameter of 5 cm, the maximum time for
temperature decay is 366 seconds when the ultrasound frequency is 1
MHz, and the minimum temperature decay is 15 seconds when the
microwave frequency is 5 MHz. As the FDA only requires the
temperature rise be less than 6.degree. C. for test times of
minutes, the 366 second decay time at 1 MHz to get to a rise of
1.degree. C. over the several minutes is allowable. As can be seen
in FIGS. 12 and 13, the decay times to a rise of 6.degree. C. are
much smaller, by a factor of approximately 70, than that of
1.degree. C.
[0111] FIG. 14 illustrates the volume of focal heated region, in
cubic centimeters, as compared to ultrasound frequencies from
between 1 and 5 MHz. Considering ultrasound frequencies in the
range of 1 to 5 MHz, the corresponding focal sizes for these
frequencies range from 3.7 mm to 0.6 mm, and the length of the
focal region ranges from 5.6 cm to 1.2 cm. The corresponding
treatment volumes range from between approximately 2.4 cc and 0.02
cc.
[0112] Examples of parameters giving a desired HSP activation
Arrhenius integral greater than 1 and damage Arrhenius integral
less than 1 is a total ultrasound power between 5.8-17 watts, a
pulse duration of 0.5 seconds, an interval between pulses of 5
seconds, with total number of pulses 10 within the total pulse
stream time of 50 seconds. The target treatment volume would be
approximately 1 mm on a side. Larger treatment volumes could be
treatable by an ultrasound system similar to a laser diffracted
optical system, by applying ultrasound in multiple simultaneously
applied adjacent but separated and spaced columns. The multiple
focused ultrasound beams converge on a very small treatment target
within the body, the convergence allowing for a minimal heating
except at the overlapping beams at the target. This area would be
heated and stimulate the activation of HSPs and facilitate protein
repair by transient high temperature spikes. However, given the
pulsating aspect of the invention as well as the relatively small
area being treated at any given time, the treatment is in
compliance with FDA/FCC requirements for long term (minutes)
average temperature rise <1K. An important distinction of the
invention from existing therapeutic heating treatments for pain and
muscle strain is that there are no high T spikes in existing
techniques, and these are required for efficiently activating HSPs
and facilitating protein repair to provide healing at the cellular
level.
[0113] The pulse train mode of energy delivery has a distinct
advantage over a single pulse or gradual mode of energy delivery,
as far as the activation of remedial HSPs and the facilitation of
protein repair is concerned. There are two considerations that
enter into this advantage:
[0114] First, a big advantage for HSP activation and protein repair
in an SDM energy delivery mode comes from producing a spike
temperature of the order of 10.degree. C. This large rise in
temperature has a big impact on the Arrhenius integrals that
describe quantitatively the number of HSPs that are activated and
the rate of water diffusion into the proteins that facilitates
protein repair. This is because the temperature enters into an
exponential that has a big amplification effect.
[0115] It is important that the temperature rise not remain at the
high value (10.degree. C. or more) for long, because then it would
violate the FDA and FCC requirements that over periods of minutes
the average temperature rise must be less than 1.degree. C. (or in
the case of ultrasound 6.degree. C.).
[0116] An SDM mode of energy delivery uniquely satisfies both of
these foregoing considerations by judicious choice of the power,
pulse time, pulse interval, and the volume of the target region to
be treated. The volume of the treatment region enters because the
temperature must decay from its high value of the order of
10.degree. C. fairly rapidly in order for the long term average
temperature rise not to exceed the long term FDA/FCC limit of
6.degree. C. for ultrasound frequencies and 1.degree. C. or less
for electromagnetic radiation energy sources.
[0117] For a region of linear dimension L, the time that it takes
the peak temperature to e-fold in tissue is roughly L.sup.2/16 D,
where D=0.00143 cm.sup.2/sec is the typical heat diffusion
coefficient. For example, if L=1 mm, the decay time is roughly 0.4
sec. Accordingly, for a region 1 mm on a side, a train consisting
of 10 pulses each of duration 0.5 seconds, with an interval between
pulses of 5 second can achieve the desired momentary high rise in
temperature while still not exceeding an average long term
temperature rise of 1.degree. C. This is demonstrated further
below.
[0118] The limitation of heated volume is the reason why RF
electromagnetic radiation is not as good of a choice for SDM-type
treatment of regions deep with the body as ultrasound. The long
skin depths (penetration distances) and Ohmic heating all along the
skin depth results in a large heated volume whose thermal inertia
does not allow both the attainment of a high spike temperature that
activates HSPs and facilitates protein repair, and the rapid
temperature decay that satisfies the long term FDA and FCC limit on
average temperature rise.
[0119] Ultrasound has already been used to therapeutically heat
regions of the body to ease pain and muscle strain. However, the
heating has not followed the SDM-type protocol and does not have
the temperature spikes that are responsible for the excitation of
HSPs.
[0120] Consider, then, a group of focused ultrasound beams that are
directed at a target region deep within the body. To simplify the
mathematics, suppose that the beams are replaced by a single source
with a spherical surface shape that is focused on the center of the
sphere. The absorption lengths of ultrasound can be fairly long.
Table 3 below shows typical absorption coefficients for ultrasound
at 1 MHz. The absorption coefficients are roughly proportional to
the frequency.
TABLE-US-00003 TABLE 3 Typical absorption coefficients for 1 MHz
ultrasound in body tissue: Body Tissue Attenuation Coefficient at 1
MHz (cm.sup.-1) Water 0.00046 Blood 0.0415 Fat 0.145 Liver
0.115-0.217 Kidney 0.23 Muscle 0.3-0.76 Bone 1.15
[0121] Assuming that the geometric variation of the incoming
radiation due to the focusing dominates any variation due to
attenuation, the intensity of the incoming ultrasound at a distance
r from the focus can be written approximately as:
l(r)=P/(4.pi.r.sup.2) [1]
where P denotes the total ultrasound power.
[0122] The temperature rise at the end of a short pulse of duration
t.sub.p at r is then
dT(t.sub.p)=P.alpha.t.sub.p/(4.pi.C.sub.vr.sup.2) [2]
where .alpha. is the absorption coefficient and C.sub.v is the
specific volume heat capacity. This will be the case until the r is
reached at which the heat diffusion length at t.sub.p becomes
comparable to r, or the diffraction limit of the focused beam is
reached. For smaller r, the temperature rise is essentially
independent of r. As an example, suppose the diffraction limit is
reached at a radial distance that is smaller than that determined
by heat diffusion. Then
r.sub.dif=(4Dt.sub.p).sup.1/2 [3]
where D is the heat diffusion coefficient, and for r<r.sub.dif,
the temperature rise at t.sub.p is
dT(r.sub.dif,t.sub.p)=3P.alpha./(8.pi.C.sub.vD) when r<r.sub.dif
[4]
[0123] Thus, at the end of the pulse, we can write for the
temperature rise:
dT.sub.p(r)={P.alpha.t.sub.p/(4.pi.C.sub.v}[(6/r.sub.dif.sup.2)U{r.sub.d-
if-r)+(1/r.sup.2)U(r-r.sub.dif)] [5]
[0124] On applying the Green's function for the heat diffusion
equation,
G(r,t)=(4.OMEGA.Dt).sup.-3/2 exp[-r.sup.2/(4Dt)] [6]
to this initial temperature distribution, we find that the
temperature dT(t) at the focal point r=0 at a time t is
dT(t)=[dT.sub.o/{(1/2).+-.(.pi..sup.1/2/6)}][(1/2)(t.sub.p/t).sup.3/2+(.-
pi..sup.1/2/6)(t.sub.p/t)] [7]
with
dT.sub.o=3P.alpha./(8.pi.C.sub.vD) [8]
[0125] A good approximation to eq. [7] is provided by:
dT(t).apprxeq.dT.sub.o(t.sub.p/t).sup.3/2 [9]
as can be seen in FIG. 15, which is a comparison of eqs. [7] and
[9] for dT(t)/dT.sub.o at the target treatment zone. The bottom
curve is the approximate expression of eq [9].
[0126] The Arrhenius integral for a train of N pulses can now be
evaluated with the temperature rise given by eq. [9]. In this
expression,
dT.sub.N(t)=.SIGMA.dT(t-nt.sub.l) [11]
where dT(t-nt.sub.l) is the expression of eq. [9] with t replaced
by t-nt.sub.l and with ti designating the interval between
pulses.
[0127] The Arrhenius integral can be evaluated approximately by
dividing the integration interval into the portion where the
temperature spikes occur and the portion where the temperature
spike is absent. The summation over the temperature spike
contribution can be simplified by applying Laplace's end point
formula to the integral over the temperature spike. In addition,
the integral over the portion when the spikes are absent can be
simplified by noting that the non-spike temperature rise very
rapidly reaches an asymptotic value, so that a good approximation
is obtained by replacing the varying time rise by its asymptotic
value. When these approximations are made, eq. [10] becomes:
.OMEGA.=AN[{t.sub.p(2k.sub.BT.sub.o.sup.2/(3EdT.sub.o)}exp[-(E/k.sub.B)l-
/(T.sub.o+dT.sub.o+dT.sub.N(Nt.sub.l))]+exp[-(E/k.sub.B)l/(T.sub.o+dT.sub.-
N(Nt.sub.l))]] [12]
where
dT.sub.N(Nt.sub.l).apprxeq.2.5dT.sub.o(t.sub.p/t.sub.l).sup.3/2
[13]
(The 2.5 in eq. [13] arises from the summation over n of
(N-n).sup.-3/2 and is the magnitude of the harmonic number (N,3/2)
for typical N of interest.)
[0128] It is interesting to compare this expression with that for
SDM applied to the retina. The first term is very similar to that
from the spike contribution in the retina case, except that the
effective spike interval is reduced by a factor of 3 for this 3D
converging beam case. The second term, involving dT.sub.N(Nt.sub.l)
is much smaller than in the retina case. There the background
temperature rise was comparable in magnitude to the spike
temperature rise. But here in the converging beam case, the
background temperature rise is much smaller by the ratio
(t.sub.p/t.sub.l).sup.3/2. This points up the importance of the
spike contribution to the activation or production of HSP's and the
facilitation of protein repair, as the background temperature rise
which is similar to the rise in a continuous ultrasound heating
case is insignificant compared to the spike contribution. At the
end of the pulse train, even this low background temperature rise
rapidly disappears by heat diffusion.
[0129] FIGS. 16 and 17 show the magnitude of the logarithm of the
Arrhenius integrals for damage and for HSP activation or production
as a function of dT.sub.o for a pulse duration t.sub.p=0.5 sec,
pulse interval t.sub.l=10 sec, and total number of pulses N=10.
Logarithm of Arrhenius integrals [eq. 12] for damage and for HSP
activation as a function of the temperature rise in degrees Kelvin
from a single pulse dT.sub.o, for a pulse duration t.sub.p=0.5
sec., pulse interval t.sub.l=10 sec., and a total number of
ultrasound pulses N=10. FIG. 16 shows the logarithm of the damage
integral with the Arrhenius constants A=8.71.times.10.sup.33
sec.sup.-1 and E=3.55.times.10.sup.-12 ergs. FIG. 17 shows the
logarithm of the HSP activation integral with the Arrhenius
constants A=1.24.times.10.sup.27 sec.sup.-1 and
E=2.66.times.10.sup.-12 ergs. The graphs in FIGS. 16 and 17 show
that .OMEGA..sub.damage does not exceed 1 until dT.sub.o exceeds
11.3 K, whereas .OMEGA..sub.hsp is greater than 1 over the whole
interval shown, the desired condition for cellular repair without
damage.
[0130] Equation [8] shows that when .alpha.=0.1 cm.sup.-1, a
dT.sub.o of 11.5 K can be achieved with a total ultrasound power of
5.8 watts. This is easily achievable. If .alpha. is increased by a
factor of 2 or 3, the resulting power is still easily achievable.
The volume of the region where the temperature rise is constant
(i.e. the volume corresponding to r=r.sub.d=(4Dt.sub.p).sup.1/2) is
0.00064 cc. This corresponds to a cube that is 0.86 mm on a
side.
[0131] This simple example demonstrates that focused ultrasound
should be usable to stimulate reparative HSP's deep in the body
with easily attainable equipment:
TABLE-US-00004 Total ultrasound power: 5.8 watts-17 watts Pulse
time 0.5 sec Pulse interval 5 sec Total train duration (N = 10) 50
sec
To expedite the treatment of larger internal volumes, a SAPRA
system can be used.
[0132] The pulsed energy source may be directed to an exterior of a
body which is adjacent to the target tissue or has a blood supply
close to the surface of the exterior of the body. Alternatively, a
device may be inserted into a cavity of a body to apply the pulsed
energy source to the target tissue. Whether the energy source is
applied outside of the body or inside of the body and what type of
device is utilized depends upon the energy source selected and used
to treat the target tissue.
[0133] Photostimulation, in accordance with the present invention,
can be effectively transmitted to an internal surface area or
tissue of the body utilizing an endoscope, such as a bronchoscope,
proctoscope, colonoscope or the like. Each of these consist
essentially of a flexible tube that itself contains one or more
internal tubes. Typically, one of the internal tubes comprises a
light pipe or multi-mode optical fiber which conducts light down
the scope to illuminate the region of interest and enable the
doctor to see what is at the illuminated end. Another internal tube
could consist of wires that carry an electrical current to enable
the doctor to cauterize the illuminated tissue. Yet another
internal tube might consist of a biopsy tool that would enable the
doctor to snip off and hold on to any of the illuminated
tissue.
[0134] In the present invention, one of these internal tubes is
used as an electromagnetic radiation pipe, such as a multi-mode
optical fiber, to transmit the SDM or other electromagnetic
radiation pulses that are fed into the scope at the end that the
doctor holds. With reference now to FIG. 18, a light generating
unit 10, such as a laser having a desired wavelength and/or
frequency is used to generate electromagnetic radiation, such as
laser light, in a controlled, pulsed manner to be delivered through
a light tube or pipe 12 to a distal end of the scope 14,
illustrated in FIG. 19, which is inserted into the body and the
laser light or other radiation 16 delivered to the target tissue 18
to be treated.
[0135] With reference now to FIG. 20, a schematic diagram is shown
of a system for generating electromagnetic energy radiation, such
as laser light, including SDM. The system, generally referred to by
the reference number 20, includes a laser console 22, such as for
example the 810 nm near infrared micropulsed diode laser in the
preferred embodiment. The laser generates a laser light beam which
is passed through optics, such as an optical lens or mask, or a
plurality of optical lenses and/or masks 24 as needed. The laser
projector optics 24 pass the shaped light beam to a delivery device
26, such as an endoscope, for projecting the laser beam light onto
the target tissue of the patient. It will be understood that the
box labeled 26 can represent both the laser beam projector or
delivery device as well as a viewing system/camera, such as an
endoscope, or comprise two different components in use. The viewing
system/camera 26 provides feedback to a display monitor 28, which
may also include the necessary computerized hardware, data input
and controls, etc. for manipulating the laser 22, the optics 24,
and/or the projection/viewing components 26.
[0136] With reference now to FIG. 21, in one embodiment, a
plurality of light beams are generated, each of which has
parameters selected so that a target tissue temperature may be
controllably raised to therapeutically treat the target tissue
without destroying or permanently damaging the target tissue. This
may be done, for example, by passing the laser light beam 30
through optics which diffract or otherwise generate a plurality of
laser light beams from the single laser light beam 30 having the
selected parameters. For example, the laser light beam 30 may be
passed through a collimator lens 32 and then through a mask 34. In
a particularly preferred embodiment, the mask 34 comprises a
diffraction grating. The mask/diffraction grating 34 produces a
geometric object, or more typically a geometric pattern of
simultaneously produced multiple laser spots or other geometric
objects. This is represented by the multiple laser light beams
labeled with reference number 36. Alternatively, the multiple laser
spots may be generated by a plurality of fiber optic waveguides.
Either method of generating laser spots allows for the creation of
a very large number of laser spots simultaneously over a very wide
treatment field. In fact, a very high number of laser spots,
perhaps numbering in the hundreds even thousands or more could be
simultaneously generated to cover a given area of the target
tissue, or possibly even the entirety of the target tissue. A wide
array of simultaneously applied small separated laser spot
applications may be desirable as such avoids certain disadvantages
and treatment risks known to be associated with large laser spot
applications.
[0137] Using optical features with a feature size on par with the
wavelength of the laser employed, for example using a diffraction
grating, it is possible to take advantage of quantum mechanical
effects which permits simultaneous application of a very large
number of laser spots for a very large target area. The individual
spots produced by such diffraction gratings are all of a similar
optical geometry to the input beam, with minimal power variation
for each spot. The result is a plurality of laser spots with
adequate irradiance to produce harmless yet effective treatment
application, simultaneously over a large target area. The present
invention also contemplates the use of other geometric objects and
patterns generated by other diffractive optical elements.
[0138] The laser light passing through the mask 34 diffracts,
producing a periodic pattern a distance away from the mask 34,
shown by the laser beams labeled 36 in FIG. 21. The single laser
beam 30 has thus been formed into hundreds or even thousands of
individual laser beams 36 so as to create the desired pattern of
spots or other geometric objects. These laser beams 36 may be
passed through additional lenses, collimators, etc. 38 and 40 in
order to convey the laser beams and form the desired pattern. Such
additional lenses, collimators, etc. 38 and 40 can further
transform and redirect the laser beams 36 as needed.
[0139] Arbitrary patterns can be constructed by controlling the
shape, spacing and pattern of the optical mask 34. The pattern and
exposure spots can be created and modified arbitrarily as desired
according to application requirements by experts in the field of
optical engineering. Photolithographic techniques, especially those
developed in the field of semiconductor manufacturing, can be used
to create the simultaneous geometric pattern of spots or other
objects.
[0140] The present invention can use a multitude of simultaneously
generated therapeutic light beams or spots, such as numbering in
the dozens or even hundreds, as the parameters and methodology of
the present invention create therapeutically effective yet
non-destructive and non-permanently damaging treatment. Although
hundreds or even thousands of simultaneous laser spots could be
generated and created and formed into patterns to be simultaneously
applied to the tissue, due to the requirements of not overheating
the tissue, there are constraints on the number of treatment spots
or beams which can be simultaneously used in accordance with the
present invention. Each individual laser beam or spot requires a
minimum average power over a train duration to be effective.
However, at the same time, tissue cannot exceed certain temperature
rises without becoming damaged. For example, using an 810 nm
wavelength laser, the number of simultaneous spots generated and
used could number from as few as 1 and up to approximately 100 when
a 0.04 (4%) duty cycle and a total train duration of 0.3 seconds
(300 milliseconds) is used. The water absorption increases as the
wavelength is increased. For shorter wavelengths, e.g., 577 nm, the
laser power can be lower. For example, at 577 nm, the power can be
lowered by a factor of 4 for the invention to be effective.
Accordingly, there can be as few as a single laser spot or up to
approximately 400 laser spots when using the 577 nm wavelength
laser light, while still not harming or damaging the tissue.
[0141] Typically, the system of the present invention incorporates
a guidance system to ensure complete and total retinal treatment
with retinal photostimulation. Fixation/tracking/registration
systems consisting of a fixation target, tracking mechanism, and
linked to system operation can be incorporated into the present
invention. In a particularly preferred embodiment, the geometric
pattern of simultaneous laser spots is sequentially offset so as to
achieve confluent and complete treatment of the surface.
[0142] This can be done in a controlled manner using an optical
scanning mechanism 50. FIGS. 22 and 23 illustrate an optical
scanning mechanism 50 in the form of a MEMS mirror, having a base
52 with electronically actuated controllers 54 and 56 which serve
to tilt and pan the mirror 58 as electricity is applied and removed
thereto. Applying electricity to the controller 54 and 56 causes
the mirror 58 to move, and thus the simultaneous pattern of laser
spots or other geometric objects reflected thereon to move
accordingly on the retina of the patient. This can be done, for
example, in an automated fashion using electronic software program
to adjust the optical scanning mechanism 50 until complete coverage
of the retina, or at least the portion of the retina desired to be
treated, is exposed to the phototherapy. The optical scanning
mechanism may also be a small beam diameter scanning galvo mirror
system, or similar system, such as that distributed by Thorlabs.
Such a system is capable of scanning the lasers in the desired
offsetting pattern.
[0143] The pattern of spots are offset at each exposure so as to
create space between the immediately previous exposure to allow
heat dissipation and prevent the possibility of heat damage or
tissue destruction. Thus, as illustrated in FIG. 24, the pattern,
illustrated for exemplary purposes as a grid of sixteen spots, is
offset each exposure such that the laser spots occupy a different
space than previous exposures. It will be understood that the
diagrammatic use of circles or empty dots as well as filled dots
are for diagrammatic purposes only to illustrate previous and
subsequent exposures of the pattern of spots to the area, in
accordance with the present invention. The spacing of the laser
spots prevents overheating and damage to the tissue. It will be
understood that this occurs until the entire target tissue to be
treated has received phototherapy, or until the desired effect is
attained. This can be done, for example, by applying electrostatic
torque to a micromachined mirror, as illustrated in FIGS. 22 and
23. By combining the use of small laser spots separated by exposure
free areas, prevents heat accumulation, and grids with a large
number of spots per side, it is possible to atraumatically and
invisibly treat large target areas with short exposure durations
far more rapidly than is possible with current technologies.
[0144] By rapidly and sequentially repeating redirection or
offsetting of the entire simultaneously applied grid array of spots
or geometric objects, complete coverage of the target, can be
achieved rapidly without thermal tissue injury. This offsetting can
be determined algorithmically to ensure the fastest treatment time
and least risk of damage due to thermal tissue, depending on laser
parameters and desired application.
[0145] The following has been modeled using the Fraunhoffer
Approximation. With a mask having a nine by nine square lattice,
with an aperture radius 9 .mu.m, an aperture spacing of 600 .mu.m,
using a 890 nm wavelength laser, with a mask-lens separation of 75
mm, and secondary mask size of 2.5 mm by 2.5 mm, the following
parameters will yield a grid having nineteen spots per side
separated by 133 .mu.m with a spot size radius of 6 .mu.m. The
number of exposures "m" required to treat (cover confluently with
small spot applications) given desired area side-length "A", given
output pattern spots per square side "n", separation between spots
"R", spot radius "r" and desired square side length to treat area
"A", can be given by the following formula:
m = A n .times. R .times. .times. floor .times. .times. ( R 2
.times. r ) 2 ##EQU00001##
[0146] With the foregoing setup, one can calculate the number of
operations m needed to treat different field areas of exposure. For
example, a 3 mm.times.3 mm area, which is useful for treatments,
would require 98 offsetting operations, requiring a treatment time
of approximately thirty seconds. Another example would be a 3
cm.times.3 cm area, representing the entire human retinal surface.
For such a large treatment area, a much larger secondary mask size
of 25 mm by 25 mm could be used, yielding a treatment grid of 190
spots per side separated by 133 .mu.m with a spot size radius of 6
.mu.m. Since the secondary mask size was increased by the same
factor as the desired treatment area, the number of offsetting
operations of approximately 98, and thus treatment time of
approximately thirty seconds, is constant.
[0147] Of course, the number and size of spots produced in a
simultaneous pattern array can be easily and highly varied such
that the number of sequential offsetting operations required to
complete treatment can be easily adjusted depending on the
therapeutic requirements of the given application.
[0148] Furthermore, by virtue of the small apertures employed in
the diffraction grating or mask, quantum mechanical behavior may be
observed which allows for arbitrary distribution of the laser input
energy. This would allow for the generation of any arbitrary
geometric shapes or patterns, such as a plurality of spots in grid
pattern, lines, or any other desired pattern. Other methods of
generating geometric shapes or patterns, such as using multiple
fiber optical fibers or microlenses, could also be used in the
present invention. Time savings from the use of simultaneous
projection of geometric shapes or patterns permits the treatment
fields of novel size, such as the 1.2 cm{circumflex over ( )}2 area
to accomplish whole-retinal treatment, in a single clinical setting
or treatment session.
[0149] With reference now to FIG. 25, instead of a geometric
pattern of small laser spots, the present invention contemplates
use of other geometric objects or patterns. For example, a single
line 60 of laser light, formed by the continuously or by means of a
series of closely spaced spots, can be created. An offsetting
optical scanning mechanism can be used to sequentially scan the
line over an area, illustrated by the downward arrow in FIG.
25.
[0150] With reference now to FIG. 26, the same geometric object of
a line 60 can be rotated, as illustrated by the arrows, so as to
create a circular field of phototherapy. The potential negative of
this approach, however, is that the central area will be repeatedly
exposed, and could reach unacceptable temperatures. This could be
overcome, however, by increasing the time between exposures, or
creating a gap in the line such that the central area is not
exposed.
[0151] The field of photobiology reveals that different biologic
effects may be achieved by exposing target tissues to lasers of
different wavelengths. The same may also be achieved by
consecutively applying multiple lasers of either different or the
same wavelength in sequence with variable time periods of
separation and/or with different irradiant energies. The present
invention anticipates the use of multiple laser, light or radiant
wavelengths (or modes) applied simultaneously or in sequence to
maximize or customize the desired treatment effects. This method
also minimizes potential detrimental effects. The optical methods
and systems illustrated and described above provide simultaneous or
sequential application of multiple wavelengths.
[0152] FIG. 27 illustrates diagrammatically a system which couples
multiple treatment light sources into the pattern-generating
optical subassembly described above. Specifically, this system 20'
is similar to the system 20 described in FIG. 20 above. The primary
differences between the alternate system 20' and the earlier
described system 20 is the inclusion of a plurality of laser
consoles, the outputs of which are each fed into a fiber coupler
42. Each laser console may supply a laser light beam having
different parameters, such as of a different wavelength. The fiber
coupler produces a single output that is passed into the laser
projector optics 24 as described in the earlier system. The
coupling of the plurality of laser consoles 22 into a single
optical fiber is achieved with a fiber coupler 42 as is known in
the art. Other known mechanisms for combining multiple light
sources are available and may be used to replace the fiber coupler
described herein.
[0153] In this system 20' the multiple light sources 22 follow a
similar path as described in the earlier system 20, i.e.,
collimated, diffracted, recollimated, and directed to the projector
device and/or tissue. In this alternate system 20' the diffractive
element must function differently than described earlier depending
upon the wavelength of light passing through, which results in a
slightly varying pattern. The variation is linear with the
wavelength of the light source being diffracted. In general, the
difference in the diffraction angles is small enough that the
different, overlapping patterns may be directed along the same
optical path through the projector device 26 to the tissue for
treatment.
[0154] Since the resulting pattern will vary slightly for each
wavelength, a sequential offsetting to achieve complete coverage
will be different for each wavelength. This sequential offsetting
can be accomplished in two modes. In the first mode, all
wavelengths of light are applied simultaneously without identical
coverage. An offsetting steering pattern to achieve complete
coverage for one of the multiple wavelengths is used. Thus, while
the light of the selected wavelength achieves complete coverage of
the tissue, the application of the other wavelengths achieves
either incomplete or overlapping coverage of the tissue. The second
mode sequentially applies each light source of a varying wavelength
with the proper steering pattern to achieve complete coverage of
the tissue for that particular wavelength. This mode excludes the
possibility of simultaneous treatment using multiple wavelengths,
but allows the optical method to achieve identical coverage for
each wavelength. This avoids either incomplete or overlapping
coverage for any of the optical wavelengths.
[0155] These modes may also be mixed and matched. For example, two
wavelengths may be applied simultaneously with one wavelength
achieving complete coverage and the other achieving incomplete or
overlapping coverage, followed by a third wavelength applied
sequentially and achieving complete coverage.
[0156] FIG. 28 illustrates diagrammatically yet another alternate
embodiment of the inventive system 20''. This system 20'' is
configured generally the same as the system 20 depicted in FIG. 20.
The main difference resides in the inclusion of multiple
pattern-generating subassembly channels tuned to a specific
wavelength of the light source. Multiple laser consoles 22 are
arranged in parallel with each one leading directly into its own
laser projector optics 24. The laser projector optics of each
channel 44a, 44b, 44c comprise a collimator 32, mask or diffraction
grating 34 and recollimators 38, 40 as described in connection with
FIG. 21 above--the entire set of optics tuned for the specific
wavelength generated by the corresponding laser console 22. The
output from each set of optics 24 is then directed to a beam
splitter 46 for combination with the other wavelengths. It is known
by those skilled in the art that a beam splitter used in reverse
can be used to combine multiple beams of light into a single
output. The combined channel output from the final beam splitter
46c is then directed through the projector device 26.
[0157] In this system 20'' the optical elements for each channel
are tuned to produce the exact specified pattern for that channel's
wavelength. Consequently, when all channels are combined and
properly aligned a single steering pattern may be used to achieve
complete coverage of the tissue for all wavelengths. The system
20'' may use as many channels 44a, 44b, 44c, etc. and beam
splitters 46a, 46b, 46c, etc. as there are wavelengths of light
being used in the treatment.
[0158] Implementation of the system 20'' may take advantage of
different symmetries to reduce the number of alignment constraints.
For example, the proposed grid patterns are periodic in two
dimensions and steered in two dimensions to achieve complete
coverage. As a result, if the patterns for each channel are
identical as specified, the actual pattern of each channel would
not need to be aligned for the same steering pattern to achieve
complete coverage for all wavelengths. Each channel would only need
to be aligned optically to achieve an efficient combination.
[0159] In system 20'', each channel begins with a light source 22,
which could be from an optical fiber as in other embodiments of the
pattern-generating subassembly. This light source 22 is directed to
the optical assembly 24 for collimation, diffraction, recollimation
and directed into the beam splitter which combines the channel with
the main output.
[0160] It will be understood that the laser light generating
systems illustrated in FIGS. 20-28 are exemplary. Other devices and
systems can be utilized to generate a source of SDM laser light
which can be operably passed through to a projector device,
typically in the form of an endoscope having a light pipe or the
like. Also, other forms of electromagnetic radiation may also be
generated and used, including ultraviolet waves, microwaves, other
radiofrequency waves, and laser light at predetermined wavelengths.
Moreover, ultrasound waves may also be generated and used to create
a thermal time-course temperature spike in the target tissue
sufficient to activate or produce heat shock proteins in the cells
of the target tissue without damaging the target tissue itself. In
order to do so, typically, a pulsed source of ultrasound or
electromagnetic radiation energy is provided and applied to the
target tissue in a manner which raises the target tissue
temperature, such as between 6.degree. C. and 11.degree. C.,
transiently while only 6.degree. C. or 1.degree. C. or less for the
long term, such as over several minutes.
[0161] It is believed that stimulating HSP production in accordance
with the present invention can be effectively utilized in treating
a wide array of tissue abnormalities, ailments, and even
infections. For example, the viruses that cause colds primarily
affect a small port of the respiratory epithelium in the nasal
passages and nasopharynx. Similar to the retina, the respiratory
epithelium is a thin and clear tissue. With reference to FIG. 29, a
cross-sectional view of a human head 62 is shown with an endoscope
14 inserted into the nasal cavity 64 and energy 16, such as laser
light or the like, being directed to tissue 18 to be treated within
the nasal cavity 64. The tissue 18 to be treated could be within
the nasal cavity 64, including the nasal passages, and
nasopharynx.
[0162] To assure absorption of the laser energy, or other energy
source, the wavelength can be adjusted to an infrared (IR)
absorption peak of water, or an adjuvant dye can be used to serve
as a photosensitizer. In such a case, treatment would then consist
of drinking, or topically applying, the adjuvant, waiting a few
minutes for the adjuvant to permeate the surface tissue, and then
administering the laser light or other energy source 16 to the
target tissue 18 for a few seconds, such as via optical fibers in
an endoscope 14, as illustrated in FIG. 29. To provide comfort of
the patient, the endoscope 14 could be inserted after application
of a topical anesthetic. If necessary, the procedure could be
repeated periodically, such as in a day or so.
[0163] The treatment would stimulate the activation or production
of heat shock proteins and facilitate protein repair without
damaging the cells and tissues being treated. As discussed above,
certain heat shock proteins have been found to play an important
role in the immune response as well as the well-being of the
targeted cells and tissue. The source of energy could be
monochromatic laser light, such as 810 nm wavelength laser light,
administered in a manner similar to that described in the
above-referenced patent applications, but administered through an
endoscope or the like, as illustrated in FIG. 29. The adjuvant dye
would be selected so as to increase the laser light absorption.
While this comprises a particularly preferred method and embodiment
of performing the invention, it will be appreciated that other
types of energy and delivery means could be used to achieve the
same objectives in accordance with the present invention.
[0164] With reference now to FIG. 30, a similar situation exists
for the flu virus, where the primary target is the epithelium of
the upper respiratory tree, in segments that have diameters greater
than about 3.3 mm, namely, the upper six generations of the upper
respiratory tree. A thin layer of mucous separates the targeted
epithelial cells from the airway lumen, and it is in this layer
that the antigen-antibody interactions occur that result in
inactivation of the virus.
[0165] With continuing reference to FIG. 30, the flexible light
tube 12 of a bronchoscope 14 is inserted through the individual's
mouth 66 through the throat and trachea 68 and into a bronchus 70
of the respiratory tree. There the laser light or other energy
source 16 is administered and delivered to the tissue in this area
of the uppermost segments to treat the tissue and area in the same
manner described above with respect to FIG. 29. It is contemplated
that a wavelength of laser or other energy would be selected so as
to match an IR absorption peak of the water resident in the mucous
to heat the tissue and stimulate HSP activation or production and
facilitate protein repair, with its attendant benefits.
[0166] With reference now to FIG. 31, a colonoscope 14 could have
flexible optical tube 12 thereof inserted into the anus and rectum
72 and into either the large intestine 74 or small intestine 76 so
as to deliver the selected laser light or other energy source 16 to
the area and tissue to be treated, as illustrated. This could be
used to assist in treating colon cancer as well as other
gastrointestinal issues.
[0167] Typically, the procedure could be performed similar to a
colonoscopy in that the bowel would be cleared of all stool, and
the patient would lie on his/her side and the physician would
insert the long, thin light tube portion 12 of the colonoscope 14
into the rectum and move it into the area of the colon, large
intestine 74 or small intestine 76 to the area to be treated. The
physician could view through a monitor the pathway of the inserted
flexible member 12 and even view the tissue at the tip of the
colonoscope 14 within the intestine, so as to view the area to be
treated. Using one of the other fiber optic or light tubes, the tip
78 of the scope would be directed to the tissue to be treated and
the source of laser light or other radiation 16 would be delivered
through one of the light tubes of the colonoscope 14 to treat the
area of tissue to be treated, as described above, in order to
stimulate HSP activation or production in that tissue 18.
[0168] With reference now to FIG. 32, another example in which the
present invention can be advantageously used is what is frequently
referred to as "leaky gut" syndrome, a condition of the
gastrointestinal (GI) tract marked by inflammation and other
metabolic dysfunction. Since the GI tract is susceptible to
metabolic dysfunction similar to the retina, it is anticipated that
it will respond well to the treatment of the present invention.
This could be done by means of subthreshold, diode micropulse laser
(SDM) treatment, as discussed above, or by other energy sources and
means as discussed herein and known in the art.
[0169] With continuing reference to FIG. 32, the flexible light
tube 12 of an endoscope or the like is inserted through the
patient's mouth 66 through the throat and trachea area 68 and into
the stomach 80, where the tip or end 78 thereof is directed towards
the tissue 18 to be treated, and the laser light or other energy
source 16 is directed to the tissue 18. It will be appreciated by
those skilled in the art that a colonoscope could also be used and
inserted through the rectum 72 and into the stomach 80 or any
tissue between the stomach and the rectum.
[0170] If necessary, a chromophore pigment could be delivered to
the GI tissue orally to enable absorption of the radiation. If, for
instance, unfocused 810 nm radiation from a laser diode or LED were
to be used, the pigment would have an absorption peak at or near
810 nm. Alternatively, the wavelength of the energy source could be
adjusted to a slightly longer wavelength at an absorption peak of
water, so that no externally applied chromophore would be
required.
[0171] It is also contemplated by the present invention that a
capsule endoscope 82, such as that illustrated in FIG. 33, could be
used to administer the radiation and energy source in accordance
with the present invention. Such capsules are relatively small in
size, such as approximately one inch in length, so as to be
swallowed by the patient. As the capsule or pill 82 is swallowed
and enters into the stomach and passes through the GI tract, when
at the appropriate location, the capsule or pill 82 could receive
power and signals, such as via antenna 84, so as to activate the
source of energy 86, such as a laser diode and related circuitry,
with an appropriate lens 88 focusing the generated laser light or
radiation through a radiation-transparent cover 90 and onto the
tissue to be treated. It will be understood that the location of
the capsule endoscope 82 could be determined by a variety of means
such as external imaging, signal tracking, or even by means of a
miniature camera with lights through which the doctor would view
images of the GI tract through which the pill or capsule 82 was
passing through at the time. The capsule or pill 82 could be
supplied with its own power source, such as by virtue of a battery,
or could be powered externally via an antenna, such that the laser
diode 86 or other energy generating source create the desired
wavelength and pulsed energy source to treat the tissue and area to
be treated.
[0172] As in the treatment of the retina in previous applications,
the radiation would be pulsed to take advantage of the micropulse
temperature spikes and associated safety, and the power could be
adjusted so that the treatment would be completely harmless to the
tissue. This could involve adjusting the peak power, pulse times,
and repetition rate to give spike temperature rises on the order of
10.degree. C., while maintaining the long term rise in temperature
to be less than the FDA mandated limit of 1.degree. C. If the pill
form 82 of delivery is used, the device could be powered by a small
rechargeable battery or over wireless inductive excitation or the
like. The heated/stressed tissue would stimulate activation or
production of HSP and facilitate protein repair, and the attendant
benefits thereof.
[0173] From the foregoing examples, the technique of the present
invention is limited to the treatment of conditions at near body
surfaces or at internal surfaces easily accessible by means of
fiber optics or other optical delivery means. The reason that the
application of SDM to activate HSP activity is limited to near
surface or optically accessibly regions of the body is that the
absorption length of IR or visible radiation in the body is very
short. However, there are conditions deeper within tissue or the
body which could benefit from the present invention. Thus, the
present invention contemplates the use of ultrasound and/or radio
frequency (RF) and even shorter wavelength electromagnetic (EM)
radiation such as microwave which have relatively long absorption
lengths in body tissue. The use of pulsed ultrasound is preferable
to RF electromagnetic radiation to activate remedial HSP activity
in abnormal tissue that is inaccessible to surface SDM or the
like.
[0174] For deep tissue that is not near an internal orifice, a
light pipe may not be an effective means of delivering the pulsed
energy. In that case, pulsed low frequency electromagnetic energy
or preferably pulsed ultrasound can be used to cause a series of
temperature spikes in the target tissue.
[0175] Thus, in accordance with the present invention, a source of
pulsed ultrasound or electromagnetic radiation is applied to the
target tissue in order to stimulate HSP production or activation
and to facilitate protein repair in the living animal tissue. In
general, electromagnetic radiation may be ultraviolet waves,
microwaves, other radiofrequency waves, laser light at
predetermined wavelengths, etc. On the other hand, if
electromagnetic energy is to be used for deep tissue targets away
from natural orifices, absorption lengths restrict the wavelengths
to those of microwaves or radiofrequency waves, depending on the
depth of the target tissue. However, ultrasound is to be preferred
to long wavelength electromagnetic radiation for deep tissue
targets away from natural orifices.
[0176] The ultrasound or electromagnetic radiation is pulsed so as
to create a thermal time--course in the tissue that stimulates HSP
production or activation and facilitates protein repair without
causing damage to the cells and tissue being treated. The area
and/or volume of the treated tissue is also controlled and
minimized so that the temperature spikes are on the order of
several degrees, e.g. approximately 10.degree. C., while
maintaining the long-term rise in temperature to be less than the
FDA mandated limit, such as 1.degree. C. It has been found that if
too large of an area or volume of tissue is treated, the increased
temperature of the tissue cannot be diffused sufficiently quickly
enough to meet the FDA requirements. However, limiting the area
and/or volume of the treated tissue as well as creating a pulsed
source of energy accomplishes the goals of the present invention of
stimulating HSP activation or production by heating or otherwise
stressing the cells and tissue, while allowing the treated cells
and tissues to dissipate any excess heat generated to within
acceptable limits.
[0177] With reference now to FIG. 34, with ultrasound, a specific
region deep in the body can be specifically targeted by using one
or more beams that are each focused on the target site. The
pulsating heating will then be largely only in the targeted region
where the beams are focused and overlap. Pulsed ultrasound sources
can also be used for abnormalities at or near surfaces as well.
[0178] As illustrated in FIG. 34, an ultrasound transducer 92 or
the like generates a plurality of ultrasound beams 94 which are
coupled to the skin via an acoustic-impedance-matching gel, and
penetrate through the skin 96 and through undamaged tissue in front
of the focus of the beams 94 to a target organ 98, such as the
illustrated liver, and specifically to a target tissue 100 to be
treated where the ultrasound beams 94 are focused. As mentioned
above, the pulsating heating will then only be at the targeted,
focused region 100 where the focused beams 94 overlap. The tissue
in front of and behind the focused region 100 will not be heated or
affected appreciably.
[0179] The present invention contemplates not only the treatment of
surface or near surface tissue, such as using the laser light or
the like, deep tissue using, for example, focused ultrasound beams
or the like, but also treatment of blood diseases, such as sepsis.
As indicated above, focused ultrasound treatment could be used both
at surface as well as deep body tissue, and could also be applied
in this case in treating blood. However, it is also contemplated
that the SDM and similar treatment options which are typically
limited to surface or near surface treatment of epithelial cells
and the like be used in treating blood diseases at areas where the
blood is accessible through a relatively thin layer of tissue, such
as the earlobe.
[0180] With reference now to FIGS. 35 and 36, treatment of blood
disorders simply requires the transmission of SDM or other
electromagnetic radiation or ultrasound pulses to the earlobe 102,
where the SDM or other radiation source of energy could pass
through the earlobe tissue and into the blood which passes through
the earlobe. It would be appreciated that this approach could also
take place at other areas of the body where the blood flow is
relatively high and/or near the tissue surface, such as fingertips,
inside of the mouth or throat, etc.
[0181] With reference again to FIGS. 35 and 36, an earlobe 102 is
shown adjacent to a clamp device 104 configured to transmit SDM
radiation or the like. This could be, for example, by means of one
or more laser diodes 106 which would transmit the desired frequency
at the desired pulse and pulse train to the earlobe 102. Power
could be provided, for example, by means of a lamp drive 108.
Alternatively, the lamp drive 108 could be the actual source of
laser light, which would be transmitted through the appropriate
optics and electronics to the earlobe 102. The clamp device 104
would merely be used to clamp onto the patient's earlobe and cause
that the radiation be constrained to the patient's earlobe 102.
This may be by means of mirrors, reflectors, diffusers, etc. This
could be controlled by a control computer 110, which would be
operated by a keyboard 112 or the like. The system may also include
a display and speakers 114, if needed, for example if the procedure
were to be performed by an operator at a distance from the
patient.
[0182] The proposed treatment with a train of electromagnetic or
ultrasound pulses has two major advantages over earlier treatments
that incorporate a single short or sustained (long) pulse. First,
the short (preferably subsecond) individual pulses in the train
activate cellular reset mechanisms like HSP activation with larger
reaction rate constants than those operating at longer (minute or
hour) time scales. Secondly, the repeated pulses in the treatment
provide large thermal spikes (on the order of 10,000) that allow
the cell's repair system to more rapidly surmount the activation
energy barrier that separates a dysfunctional cellular state from
the desired functional state. The net result is a "lowered
therapeutic threshold" in the sense that a lower applied average
power and total applied energy can be used to achieve the desired
treatment goal.
[0183] Power limitations in current micropulsed diode lasers
require fairly long exposure duration. The longer the exposure, the
more important the center-spot heat dissipating ability toward the
unexposed tissue at the margins of the laser spot. Thus, the
micropulsed laser light beam of an 810 nm diode laser should have
an exposure envelope duration of 500 milliseconds or less, and
preferably approximately 300 milliseconds. Of course, if
micropulsed diode lasers become more powerful, the exposure
duration should be lessened accordingly.
[0184] Aside from power limitations, another parameter of the
present invention is the duty cycle, or the frequency of the train
of micropulses, or the length of the thermal relaxation time
between consecutive pulses. It has been found that the use of a 10%
duty cycle or higher adjusted to deliver micropulsed laser at
similar irradiance at similar MPE levels significantly increase the
risk of lethal cell injury. However, duty cycles of less than 10%,
and preferably 5% or less demonstrate adequate thermal rise and
treatment at the level of the MPE cell to stimulate a biological
response, but remain below the level expected to produce lethal
cell injury. The lower the duty cycle, however, the exposure
envelope duration increases, and in some instances can exceed 500
milliseconds.
[0185] Each micropulse lasts a fraction of a millisecond, typically
between 50 microseconds to 100 microseconds in duration. Thus, for
the exposure envelope duration of 300-500 milliseconds, and at a
duty cycle of less than 5%, there is a significant amount of wasted
time between micropulses to allow the thermal relaxation time
between consecutive pulses. Typically, a delay of between 1 and 3
milliseconds, and preferably approximately 2 milliseconds, of
thermal relaxation time is needed between consecutive pulses. For
adequate treatment, the cells are typically exposed or hit between
50-200 times, and preferably between 75-150 at each location, and
with the 1-3 milliseconds of relaxation or interval time, the total
time in accordance with the embodiments described above to treat a
given area which is being exposed to the laser spots is usually
less than one second, such as between 100 milliseconds and 600
milliseconds on average. The thermal relaxation time is required so
as not to overheat the cells within that location or spot and so as
to prevent the cells from being damaged or destroyed. While time
periods of 100-600 milliseconds do not seem long, given the small
size of the laser spots and the need to treat a relatively large
area of the target tissue, treating the entire target tissue take a
significant amount of time, particularly for a patient who is
undergoing treatment.
[0186] Other pulsed energy sources, including microwave, radio
frequency and ultrasound is also preferably pulsed in nature and
have duty cycles and/or pulse trains and thus lag time or intervals
between micropulse energy applications to the target tissue.
Moreover, the target tissue previously treated with the micropulse
of the energy must be allowed to dissipate the heat created by the
energy application in order not to exceed a predetermined upper
temperature level which could permanently damage or even destroy
the cells of the target tissue. Typically, the area or volume of
target tissue to be treated is much larger than the area or volume
of target tissue which is treated at any given moment by the energy
sources, even if multiple beams of energy are created and applied
to the target tissue.
[0187] Accordingly, the present invention may utilize the interval
between consecutive applications to the same location to apply
energy to a second treatment area, or additional areas, of the
target tissue that is spaced apart from the first treatment area.
The pulsed energy is returned to the first treatment location, or
previous treatment locations, within the predetermined interval of
time so as to provide sufficient thermal relaxation time between
consecutive pulses, yet also sufficiently treat the cells in those
locations or areas properly by sufficiently increasing the
temperature of those cells over time by repeatedly applying the
energy to that location in order to achieve the desired therapeutic
benefits of the invention.
[0188] It is important to return to a previously treated location
within a predetermined amount of time to allow the area to cool
down sufficiently during that time, but also to treat it within the
necessary window of time. In the case of the laser light pulsed
energy applications, the laser light is returned to the previously
treated location within one to three milliseconds, and preferably
approximately two milliseconds, as one cannot wait one or two
seconds and then return to a previously treated area that has not
yet received the full treatment necessary, as the treatment will
not be as effective or perhaps not effective at all. However,
during that interval of time, typically approximately 2
milliseconds, at least one other area, and typically multiple
areas, can be treated with a laser light application as the laser
light pulses are typically 50 seconds to 100 microseconds in
duration. This is referred to herein as microshifting. The number
of additional areas which can be treated is limited only by the
micopulse duration and the ability to controllably move the light
beams from one area to another.
[0189] Currently, approximately four additional areas which are
sufficiently spaced apart from one another can be treated during
the thermal relaxation intervals beginning with a first treatment
area when using laser light. Thus, multiple areas can be treated,
at least partially, during the 200-500 millisecond exposure
envelope for the first area. Thus, in a single interval of time,
instead of only 100 simultaneous light spots being applied to a
treatment area, approximately 500 light spots can be applied during
that interval of time in different treatment areas. This would be
the case, for example, for a laser light beam having a wavelength
of 810 nm. For shorter wavelengths, such as 572 nm, even a greater
number of individual locations can be exposed to the laser beams to
create light spots. Thus, instead of a maximum of approximately 400
simultaneous spots, approximately 2,000 spots could be covered
during the interval between micropulse treatments to a given area
or location. Typically each location has between 50-200, and more
typically between 75-150, light applications applied thereto over
the course of the exposure envelope duration (typically 200-500
milliseconds) to achieve the desired treatment. In accordance with
an embodiment of the present invention, the laser light would be
reapplied to previously treated areas in sequence during the
relaxation time intervals for each area or location. This would
occur repeatedly until a predetermined number of laser light
applications to each area to be treated have been achieved.
[0190] Similarly, the one or more beams of microwave,
radiofrequency and/or ultrasound could be applied to second, or
additional treatment areas of the target tissue that is spaced
apart from the first treatment area, and after a predetermined
interval of time returning, if necessary, to the first treatment
area of the target tissue to reapply the pulsed energy thereto. The
pulsed energy could be reapplied to a previously treated area in
sequence during the relaxation time intervals for each area or
location until a desired number of applications has been achieved
to each treatment area. The treatment areas must be separated by at
least a predetermined minimum distance to enable thermal relaxation
and dissipation and avoid thermal tissue damage. The pulsed energy
parameters including wavelength or frequency, duty cycle and pulse
train duration are selected so as to raise the target tissue
temperature up to 11.degree. C., such as between approximately
6.degree.-11.degree. C., during application of the pulsed energy
source to the target tissue to achieve a therapeutic effect, such
as by stimulating HSP production within the cells. However, the
cells of the target tissue must be given a period of time to
dissipate the heat such that the average temperature rise of the
tissue over several minutes is maintained at or below a
predetermined level, such as 6.degree. C. or less, or even
1.degree. C. or less, over several minutes so as not to permanently
damage the target tissue.
[0191] This is diagrammatically illustrated in FIGS. 37A-37D. FIG.
37A illustrates with solid circles a first area having energy
beams, such as laser light beams, applied thereto as a first
application. The beams are controllably offset or microshifted to a
second exposure area, followed by a third exposure area and a
fourth exposure area, as illustrated in FIG. 37B, until the
locations in the first exposure area need to be re-treated by
having beams applied thereto again within the thermal relaxation
time interval. The locations within the first exposure area would
then have energy beams reapplied thereto, as illustrated in FIG.
37C. Secondary or subsequent exposures would occur in each exposure
area, as illustrated in FIG. 37D by the increasingly shaded dots or
circles until the desired number of exposures or hits or
applications of energy to the target tissue area has been achieved
to therapeutically treat these areas, diagrammatically illustrated
by the blackened circles in exposure area 1 in FIG. 37D. When a
first or previous exposure area has been completed treated, this
enables the system to add an additional exposure area, which
process is repeated until the entire area to be treated has been
fully treated. It should be understood that the use of solid
circles, broken line circles, partially shaded circles, and fully
shaded circles are for explanatory purposes only, as in fact the
exposure of the energy or laser light in accordance with the
present invention is invisible and non-detectable to both the human
eye as well as known detection devices and techniques, including
ophthalmoscopically and angiographically.
[0192] Adjacent exposure areas must be separated by at least a
predetermined minimum distance to avoid thermal tissue damage. Such
distance is at least 0.5 diameter away from the immediately
preceding treated location or area, and more preferably between 1-2
diameters away. Such spacing relates to the actually treated
locations in a previous exposure area. It is contemplated by the
present invention that a relatively large area may actually include
multiple exposure areas therein which are offset in a different
manner than that illustrated in FIG. 37. For example, the exposure
areas could comprise the thin lines illustrated in FIGS. 25 and 26,
which would be repeatedly exposed in sequence until all of the
necessary areas were fully exposed and treated. In accordance with
the present invention, the time required to treat that area to be
treated is significantly reduced, such as by a factor of 4 or 5
times, such that a single treatment session takes much less time
for the medical provider and the patient need not be in discomfort
for as long of a period of time.
[0193] In accordance with this embodiment of the invention of
applying one or more treatment beams at once, and moving the
treatment beams to a series of new locations, then bringing the
beams back to re-treat the same location or area repeatedly has
been found to also require less power compared to the methodology
of keeping the beams in the same locations or area during the
entire exposure envelope duration. With reference to FIGS. 38-40,
there is a linear relationship between the pulse length and the
power necessary, but there is a logarithmic relationship between
the heat generated.
[0194] With reference to FIG. 38, a graph is provided wherein the
x-axis represents the Log of the average power in watts of a laser
and the y-axis represents the treatment time, in seconds. The lower
curve is for panmacular treatment and the upper curve is for
panretinal treatment. This would be for a laser light beam having a
micropulse time of 50 microseconds, a period of 2 milliseconds of
time between pulses, and duration of train on a spot of 300
milliseconds. The areas of each retinal spot are 100 microns, and
the laser power for these 100 micron retinal spots is 0.74 watts.
The panmacular area is 0.55.sup.2, requiring 7,000 panmacular spots
total, and the panretinal area is 3.30.sup.2, requiring 42,000
laser spots for full coverage. Each RPE spot requires a minimum
energy in order for its reset mechanism to be adequately activated,
in accordance with the present invention, namely, 38.85 joules for
panmacular and 233.1 joules for panretinal. As would be expected,
the shorter the treatment time, the larger the required average
power. However, there is an upper limit on the allowable average
power, which limits how short the treatment time can be.
[0195] As mentioned above, there are not only power constraints
with respect to the laser light available and used, but also the
amount of power that can be applied to the eye without damaging eye
tissue. For example, temperature rise in the lens of the eye is
limited, such as between 4.degree. C. so as not to overheat and
damage the lens, such as causing cataracts. Thus, an average power
of 7.52 watts could elevate the lens temperature to approximately
4.degree. C. This limitation in power increases the minimum
treatment time.
[0196] However, with reference to FIG. 39, the total power per
pulse required is less in the microshift case of repeatedly and
sequentially moving the laser spots and returning to prior treated
locations, so that the total energy delivered and the total average
power during the treatment time is the same. FIGS. 39 and 40 show
how the total power depends on treatment time. This is displayed in
FIG. 39 for panmacular treatment, and in FIG. 40 for panretinal
treatment. The upper, solid line or curve represents the embodiment
where there are no microshifts taking advantage of the thermal
relaxation time interval, such as described and illustrated in FIG.
24, whereas the lower dashed line represents the situation for such
microshifts, as described and illustrated in FIG. 37. FIGS. 39 and
40 show that for a given treatment time, the peak total power is
less with microshifts than without microshifts. This means that
less power is required for a given treatment time using the
microshifting embodiment of the present invention. Alternatively,
the allowable peak power can be advantageously used, reducing the
overall treatment time.
[0197] Thus, in accordance with FIGS. 38-40, a log power of 1.0 (10
watts) would require a total treatment time of 20 seconds using the
microshifting embodiment of the present invention, as described
herein. It would take more than 2 minutes of time without the
microshifts, and instead leaving the micropulsed light beams in the
same location or area during the entire treatment envelope
duration. There is a minimum treatment time according to the
wattage. However, this treatment time with microshifting is much
less than without microshifting. As the laser power required is
much less with the microshifting, it is possible to increase the
power in some instances in order to reduce the treatment time for a
given desired retinal treatment area. The product of the treatment
time and the average power is fixed for a given treatment area in
order to achieve the therapeutic treatment in accordance with the
present invention. This could be implemented, for example, by
applying a higher number of therapeutic laser light beams or spots
simultaneously at a reduced power. Of course, since the parameters
of the laser light are selected to be therapeutically effective yet
not destructive or permanently damaging to the cells, no guidance
or tracking beams are required, only the treatment beams as all
areas can be treated in accordance with the present invention.
[0198] Although the present invention is described for use in
connection with a micropulsed laser, theoretically a continuous
wave laser could potentially be used instead of a micropulsed
laser. However, with the continuous wave laser, there is concern of
overheating as the laser is moved from location to location in that
the laser does not stop and there could be heat leakage and
overheating between treatment areas. Thus, while it is
theoretically possible to use a continuous wave laser, in practice
it is not ideal and the micropulsed laser is preferred.
[0199] While the information provided in connection with graphs
38-40 is derived from observations and calculations of laser light
beams as the energy source applied to retinal eye tissue, it is
believed that applying such pulsed light beams to other tissue will
achieve similar results in that moving the treatment beams to a
series of new locations, then bringing the beams back to re-treat
the same location or area repeatedly will not only save time but
also require less power compared to the methodology of keeping the
beams in the same location or area during the entire exposure
envelope duration. Similarly, it is believed that such power
conservation will also be achieved with other sources of pulsed
energy, including coherent and non-coherent light, microwave,
radiofrequency and ultrasound energy sources.
[0200] In accordance with the microshifting technique described
above, the shifting or steering of the pattern of light beams may
be done by use of an optical scanning mechanism, such as that
illustrated and described in connection with FIGS. 22 and 23. For
situations where the wavelength of the illumination or energy is
much less than the distance to the volume to be illuminated or
exposed, the steering can be accomplished by using phased arrays.
The illumination or energy in this case is said to be the "far
field". Phased arrays can be used for the microwave and ultrasound
illumination application or even for the laser light beam
source.
[0201] Steering for microwave, ultrasound and even for laser energy
sources may be done by use of multiple sources which provide an
"array". The basic idea for steering the illumination radiation
pattern of an array is constructive (and destructive) interference
between the radiation from the individual members of the array of
sources. With reference to FIG. 41, to illustrate this, it is only
necessary to consider two adjacent members of the array. FIG. 41
depicts the wavefront originating from two adjacent sources.
[0202] It is evident that for a wavefront that is depicted at an
angle .theta. with respect to the distance a between the two
sources, the amplitude of the wave from the source on the left is
proportional to exp[i.omega.t] whereas the amplitude of the wave
from the source on the right is proportional to exp[i.omega.t-ka
sin .theta.-.PHI.], where .omega. is the angular frequency of the
radiation, and k=2.pi./.lamda..
[0203] For constructive interference, these two waves should be "in
phase", i.e.
.PHI.constructive=ka sin .theta.+2n.pi. [14]
[0204] For destructive interference, these two waves should be "out
of phase", i.e.
.PHI.destructive=ka sin .theta.+(2n+1).pi. [15]
[0205] Accordingly, the illumination will be large in the
directions .theta. given by
sin .theta.=(1/ka)[.PHI.constructive-2n.pi.] [16]
[0206] In other words, the radiation can be steered to different
desired directions .theta. simply by choosing different delays
.PHI..
[0207] The delays can be introduced electronically into the
circuits for exciting the radiation sources. The means for doing
this have also been well discussed in the published literature:
analog delay circuits are available as well as digital delay
circuits.
[0208] Radiation patterns for microwave, ultrasound, and laser
sources are quite well-directed. If we estimate the divergence of
the radiation beam from a source of transverse dimension 2b by the
Airy disc expression
.THETA.1/2=0.6.lamda./b [17]
[0209] Then at a target distance D from the source, the half-width
w of the illuminated region is roughly
w=0.6.lamda.D/b [18]
[0210] If we require the separation of the illuminated regions to
be 2w, then the separation of the source s is roughly 3w:
a=1.8.lamda.D/b [19]
[0211] This can be a small separation if the source size is chosen
to be much larger than the radiation wavelength.
[0212] For example, for ultrasound, suppose we have a 5 MHz source
with a transverse dimension of 1 cm, and suppose the desired target
distance is 10 cm. Then the separation distance is a .apprxeq.0.5
cm.
[0213] As another example, a commercially available microwave
standard gain horn source, operating at 140-220 GHz has transverse
dimensions of 13.9 mm by 10.8 mm and a depth dimension of 32.2 mm.
For 200 GHz, the wavelength is 0.15 cm, and for a target distance
of 10 cm, the target width given by the equation [18] is
1.2.times.0.15.times.10/0.6=3 cm. For the spacing a of the horns,
eq. [19] then gives 9 cm.
[0214] Next, apply eqs. [17]-[19] to obtain rough estimates for a
steerable array of 810 nm laser radiation. Suppose b=2.times.810
nm, and suppose D=1 mm. Then eqs. [17]-[19] give .THETA.1/2=0.3,
w=0.3 mm, and a=0.9 mm.
[0215] For the radiofrequency application, however, the wavelength
of the radiofrequency radiation is typically much larger than a
human body dimensions. In that case, the treatment volume is said
to be in the "near field" of the radiofrequency source. Phased
arrays are not useful in near field, and a different method of
steering is required.
[0216] For radio frequency treatment, the wavelength of the
radiation is much larger than body dimensions. Thus, for 3-6 MHz,
the wavelengths range from 10,000 cm to 5000 cm. Accordingly, the
target region in the body is in the "near field" of the source,
i.e. the target distance and dimensions are much less than the
wavelength of the RF radiation. This means that the relevant
treatment fields are not radiation fields (as they were in the case
of microwave, ultrasound, and laser treatments), but are instead
induction fields.
[0217] The induction field from an RF coil is only large over
dimensions comparable to the coil dimension. The induction magnetic
fields drop off rapidly as 1/r3 for distances larger than this.
Accordingly, for a coil at the surface of the body, we can picture
the treatment volume as roughly a hemisphere with radius equal to
that of the coil.
[0218] For coils with radii between 2 and 6 mm, the treatment
volumes for these coils are rather close to the surface (distances
comparable to the coil dimensions). Larger coils can be used for
deeper targets. In keeping with the spacing criteria discussed
earlier, the spacing between the coils in a surface array would be
chosen to be comparable to the individual coil dimensions.
[0219] For the laser or other light beam and ultrasound sources,
the wavelengths are much less than the distances from the sources
to the target tissue. For these sources, then, the intensity
distributions from the arrays can be calculated in the "far field"
approximation. However, for the RF sources, the wavelength is much
larger than the distances between the sources to the target tissue.
For these sources, the intensity distribution be calculated in the
"near field" approximation. For microwaves, at high frequencies,
the wavelengths are much less than the distance between the sources
and target tissue; however, at low microwave frequencies, the
wavelengths can be larger than the distance between the sources and
the target tissues. (Thus, at 1 and 100 GHz, the wavelengths are 30
cm and 3 mm, respectively). Accordingly, at high microwave
frequencies, the "far field" approximation applies, while at low
microwave frequencies, the "near field" approximation applies.
[0220] In the far field approximation, the expressions treat
kR>>1, where k=2.pi./.lamda. is the wavenumber, .lamda. is
the wavelength, and R is a typical distance between the source and
target: In this approximation, the energy is "radiated" from the
source to the target. In the near field approximation, the
expressions treat kR<<1: In this approximation, the fields
are not radiation fields, but are "induction" fields. The array
behaviors are markedly different in the two approximations.
[0221] For far field "radiation" arrays, the following is taken
into account. With reference now to FIG. 42, a square array of
square antennas or radiation sources is shown. Each antenna has a
side of length 2a, and the shortest distance between the centers of
adjacent antennas is 2d. There are a total of N antennas along a
line in the x direction and N antennas along the y-direction, for a
total of N.sup.2 antennas.
[0222] On using the far field approximation, we find for the
intensity I.sub.p at a distant observation point P:
I.sub.p/I.sub.o={4k.sup.2a.sup.4/(.pi..sup.2R.sub.o.sup.2)}Sin
c.sup.2{k.alpha.a)Sin
c.sup.2(k.beta.a){Sin(Nk.alpha.d)/Sin(k.alpha.d)}.sup.2{Sin(Nk.beta.d)/Si-
n(k.beta.d)}.sup.2 [20]
[0223] In this expression, it is assumed that the observation point
is located a distance R.sub.o from the antenna array and that the
intensity from a single antenna is I.sub.o. In addition, .alpha.
and .beta. are the deflection angles in the x and y directions,
respectively, and
Sin c(v)=Sin(v)/v [21]
[0224] Equation [20] can also be written in terms of the
coordinates X and Y along the x and y directions in the observation
plane by using the approximate relations
.alpha.=X/R.sub.o [22]
.beta.=Y/R.sub.o [23]
[0225] From eq. [20], we can see what the specific form of the
radiation pattern from the array is. FIG. 43 is a plot of a typical
radiation pattern from a square array. (Anomalies in the plot
appear due to the plotting routine employed. Because of plotting
inaccuracies, there is randomness in the height of some of the
peaks which should not be present, and not all of the peaks are
actually shown.) The X and Z dimensions are shown in centimeters,
but these dimensions can be changed easily in the equations
below.
[0226] FIG. 44 is the form of a typical radiation pattern along the
X-axis for a typical radiation pattern from a "far field" array.
The pattern results from the individual features shown in FIGS.
45-47.
[0227] Specifically, it is plot of
Sin
c.sup.2{k(X/R.sub.o)a){Sin(NkX/R.sub.o)d)/Sin(k(X/R.sub.o)d)}.sup.2
[0228] The envelope of the pattern is determined by the Sin
c.sup.2{k (X/R.sub.o)a) function. This is shown in FIG. 45. The
width of the individual lines is determined by the
Sin.sup.2(NkX/R.sub.o)d) function. This is shown in FIG. 46.
Finally, the separation of the lines is determined by the
Sin.sup.2(kX/R.sub.o)d) in the denominator. The lines occur every
time the function has a zero, i.e. whenever the argument of the
function is some multiple of .pi.. The function is plotted in FIG.
47.
[0229] With continuing reference to FIGS. 42-47, the widths of the
individual lines and the envelope are determined by the half-widths
of the Sin.sup.2(Nk.alpha.d) and Sin c.sup.2(k.alpha.a) functions,
respectively, and the spacing between the lines is determined by
the zeros of the Sin.sup.2((k.alpha.d) function.
Thus, we can write directly:
Width of envelope: .DELTA..sub.env.alpha.=.xi.(.lamda./a) [24]
Spacing between lines (spots): .DELTA..sub.sep .alpha.=.lamda./(2d)
[25]
Width of a single line (spot): .DELTA..sub.line
.alpha.=.xi.(.lamda./Nd) [26]
Number of lines (spots) along X-axis: N=2.xi.(d/a) [27a]
Number of spots in square pattern N.sup.2=4.xi..sup.2(d/a).sup.2
[27b]
[0230] In these expressions .xi. is a fraction on the order of 1/2
that describes where the corresponding Sin c or Sin function is
about half-max. (If it is desired to observe only where these
functions are larger and more uniform in magnitude, then .xi. can
be chosen smaller.)
[0231] A far field array, such as that illustrated in FIG. 42, can
be selectively and controllably steered. The position of the peaks
can be changed by introducing a phase delay in the excitation of
the antennas. Thus, the direction in the X direction can be changed
by introducing a phase delay .PHI..sub.n in the nth antenna in the
X-direction, that is proportional to n. To change the direction
from .alpha.=0 to an arbitrary .alpha..sub.o, the phase delay of
the nth antenna in the X direction is
.PHI..sub.n=-.alpha..sub.onkd. [28a]
[0232] In a similar manner the maximum peak direction in the Y
direction can be shifted from .beta.=0 to an arbitrary
.beta..sub.o. To change the direction from .beta.=0 to an arbitrary
.beta..sub.o, an additional phase delay is introduced to the mth
antenna in the Y direction is
.PHI..sub.m=-.beta..sub.omkd [28b]
[0233] With reference now to FIG. 48, a block diagram of its system
for exciting the antennas in the array, such as that illustrated in
FIG. 42, to irradiate a target tissue is shown. The array system of
FIG. 48 is applicable for the light beam, ultrasound and high
frequency microwave arrays. The computer controller provides the
desired power excitation and phase delays for steering the array.
The computer-controlled oscillator source activates the antennas
with appropriate phase delays to steer the antenna array peaks, as
described above.
[0234] A near field (induction) array, and particularly the
steering of such near field arrays, for low frequency microwaves
and RF differs markedly from the far field arrays discussed
above.
[0235] As an example, consider the near field (induction electric
field) from a circular coil carrying an alternating current I. If
the coil lies in the X-Y plane with its axis along the Z-direction,
then the vector potential A is in the azimuthal direction, and is
given by
A.PHI.=(.mu.l/.pi.k)(a/.rho.).sup.1/2[{1-(k.sup.2/2)}K(k.sup.2)-E(k.sup.-
2)] [29]
with
k.sup.2=4a.rho.[(a+.rho.).sup.2+Z.sup.2].sup.-1 [30]
Here
[0236] .mu. is the magnetic permeability of free space
[0237] a is the radius of the current carrying coil
[0238] .rho.=(X.sup.2+Y.sup.2).sup.1/2
[0239] E is the complete elliptic integral of the second kind
[0240] K is the complete elliptic integral of the first kind
[0241] The induction electric field is also in the azimuthal
direction, and is given by
E.sub..PHI.=-i.omega.A.sub..PHI. [31]
where
[0242] .omega. is the angular frequency of the alternating current
I.
[0243] The objective of the induction field is to heat the tissue
to activate heat shock proteins. The heating is achieved by
dielectric or Ohmic heating: Accordingly, the temperature rise in
the tissue is proportional to Im(.di-elect
cons.)(.omega.A.sub..PHI.).sup.2.
[0244] FIG. 49 is a plot of (.pi.A.sub..PHI./.mu.l).sup.2 vs the
dimensionless variables x'=X/a and z'=Z/a for a coil of radius a in
the Z=0 plane with its center at X=Y=0. In the figure, x' ranges
from -3 to +3, and z' ranges from 0 to 1.
[0245] With continuing reference to FIG. 49, it is shown that the
induced tissue temperature rise drops off rapidly as the axial
distance from the coil increases. The tissue between the coil and
about an axial distance equal to the radius of the coil divided by
2 can be expected to experience a temperature rise. Thus, if it is
desired to heat a tissue that is 5 cm from the surface, where the
coil sits, the coil should be approximately 10 cm in diameter. FIG.
49 also shows that the main heating will occur in a circular ring
equal in radius to the coil radius.
[0246] To illustrate the latter point, FIGS. 50-52 show
(.pi.A.sub..PHI./.mu.l).sup.2 vs the dimensionless variables x'=X/a
at z'=0.5 for three different coils of the same radius a. FIG. 50
is for a coil with its center at X=-a. FIG. 51 is for a coil with
its center at X=0, and FIG. 52 is for a coil with its center at
X=+a. FIG. 50 illustrates (.pi.A.sub..PHI./.mu.l).sup.2 vs the
dimensionless variable x'=X/a at z'=Z/a=0.5 for a coil of radius a
in the Z=0 plane with its center at X=-a, Y=0. FIG. 51 illustrates
(.pi.A.sub..PHI./.mu.l).sup.2 vs the dimensionless variable x'=X/a
at z'=Z/a=0.5 for a coil of radius a in the Z=0 plane with its
center at X=Y=0. FIG. 52 illustrates (.pi.A.sub..PHI./.mu.l).sup.2
vs the dimensionless variable x'=X/a at z'=Z/a=0.5 for a coil of
radius a in the Z=0 plane with its center at X=+a and Y=0.
[0247] FIG. 53 shows the plots of FIGS. 50-52 superimposed, where
(.pi.A.sub..PHI.)/.mu.l).sup.2 vs the dimensionless variable x'=X/a
at z'=Z/a=0.5 for three different locations of a coil of radius a
in the z'=0 plane. The left-most curve is for a coil with its
center at X=-a and Y=0; the middle curve is for a coil with its
center at X=Y=0; and the right-most curve is for a coil with its
center at X=a and Y=0.
[0248] With continuing reference to FIGS. 49-53, with respect to
depth of treatment, if a tissue at a distance Z.sub.o needs to be
treated by induction heating, a coil of radius 2Z.sub.o should be
used. It will treat all tissue between the surface and Z.sub.o. For
steering, for induction fields, the way to treat different
transverse positions is not to "steer" an array by phase delay, but
rather to activate individual coils sequentially. Each activated
coil will treat the region below it, primarily in a circular strip
beneath its circumference.
[0249] With reference now to FIG. 54, a block diagram for an
induction array (near field) for RF sources and low-frequency
microwave sources is shown. Here, the computer-controlled powered
oscillating current source selects the coils sequentially in order
to treat different transverse tissue positions. Thus, coils 1-N are
powered sequentially in order to steer the induction fields. Thus,
for the different types of radiation or energy, a different
steering mechanism or system is utilized in order to treat the
desired tissue at a desired depth.
[0250] As mentioned above, the controlled manner of applying energy
to the target tissue is intended to raise the temperature of the
target tissue to therapeutically treat the target tissue without
destroying or permanently damaging the target tissue. It is
believed that such heating activates HSPs and that the thermally
activated HSPs work to reset the diseased tissue to a healthy
condition, such as by removing and/or repairing damaged proteins.
It is believed by the inventors that maximizing such HSP activation
improves the therapeutic effect on the targeted tissue. As such,
understanding the behavior and activation of HSPs and HSP system
species, their generation and activation, temperature ranges for
activating HSPs and time frames of the HSP activation or generation
and deactivation can be utilized to optimize the heat treatment of
the biological target tissue.
[0251] As mentioned above, the target tissue is heated by the
pulsed energy for a short period of time, such as ten seconds or
less, and typically less than one second, such as between 100
milliseconds and 600 milliseconds. The time that the energy is
actually applied to the target tissue is typically much less than
this in order to provide intervals of time for heat relaxation so
that the target tissue does not overheat and become damaged or
destroyed. For example, as mentioned above, laser light pulses may
last on the order of microseconds with several milliseconds of
intervals of relaxed time.
[0252] Thus, understanding the sub-second behaviors of HSPs can be
important to the present invention. The thermal activation of the
HSPs in SDM is typically described by an associated Arrhenius
integral,
.OMEGA.=.intg.dt A exp[-E/k.sub.BT(t)] [28]
where the integral is over the treatment time and
[0253] A is the Arrhenius rate constant for HSP activation
[0254] E is the activation energy
[0255] T(t) is the temperature of the thin RPE layer, including the
laser-induced temperature rise
[0256] The laser-induced temperature rise--and therefore the
activation Arrhenius integral--depends on both the treatment
parameters (e.g., laser power, duty cycle, total train duration)
and on the RPE properties (e.g., absorption coefficients, density
of HSPs). It has been found clinically that effective SDM treatment
is obtained when the Arrhenius integrals is of the order of
unity.
[0257] The Arrhenius integral formalism only takes into account a
forward reaction, i.e. only the HSP activation reaction): It does
not take into account any reverse reactions in which activated HSPs
are returned to their inactivated states. For the typical subsecond
durations of SDM treatments, this appears to be quite adequate.
However, for longer periods of time (e.g. a minute or longer), this
formalism is not a good approximation: At these longer times, a
whole series of reactions occurs resulting in much smaller
effective HSP activation rates. This is the case during the
proposed minute or so intervals between SDM applications in the
present invention disclosure.
[0258] In the published literature, the production and destruction
of heat shock proteins (HSPs) in cells over longer durations is
usually described by a collection of 9-13 simultaneous mass-balance
differential equations that describe the behavior of the various
molecular species involved in the life cycle of an HSP molecule.
These simultaneous equations are usually solved by computer to show
the behavior in time of the HSPs and the other species after the
temperature has been suddenly raised.
[0259] These equations are all conservation equations based on the
reactions of the various molecular species involved in the activity
of HSPs. To describe the behavior of the HSPs in the minute or so
intervals between repeated applications of SDM, we shall use the
equations described in M. Rybinski, Z. Szymanska, S. Lasota, A.
Gambin (2013) Modeling the efficacy of hyperthermia treatment.
Journal of the Royal Society Interface 10, No. 88, 20130527
(Rybinski et al (2013)). The species considered in Rybinski et al
(2013) are shown in Table 4.
TABLE-US-00005 TABLE 4 HSP system species in Rybinski et al (2013)
description: HSP ubiquitous heat shock protein of molecular weight
70 Da (in free, activated state) HSF heat shock (transcription)
factor that has no DNA binding capability HSF.sub.3 (trimer) heat
shock factor capable of binding to DNA, formed from HSF HSE heat
shock element, a DNA site that initiates transcription of HSP when
bound to HSF.sub.3 mRNA messenger RNA molecule for producing HSP S
substrate for HSP binding: a damaged protein P properly folded
protein HSP.cndot.HSE a complex of HSP bound to HSF (unactivated
HSPs) HSF.sub.3.cndot.HSE a complex of HSF.sub.3 bound to HSE, that
induces transcription and the creation of a new HSP mRNA molecule
HSP.S a complex of HSP attached to damaged protein (HSP actively
repairing the protein)
[0260] The coupled simultaneous mass conservation equations for
these 10 species are summarized below as eqs. [29]-[38]:
d .function. [ HSP ] / d .times. t = ( I 1 + k 1 .times. 0 )
.function. [ HSPS ] + I 2 .function. [ HSPHSF ] + k 4 .function. [
m .times. R .times. NA ] - k 1 .function. [ S ] .function. [ HSP ]
- k 2 .function. [ HSP ] .function. [ HSF ] - I 3 .function. [ HSP
] .function. [ HSF 3 ] - k 9 .function. [ HSP ] [ 29 ] d .times. {
HSF ] / d .times. t = I 2 .function. [ HSPHSF ] + 2 .times. 1 3
.function. [ HSP ] .function. [ HSF 3 ] + k 6 .function. [ HSPHS
.times. F ] .function. [ S ] - k 2 .function. [ HSP ] .function. [
HSF ] - 3 .times. k 3 .function. [ HSF ] 3 - I 6 .function. [ HSPS
] .function. [ HSF ] [ 30 ] d .function. [ S ] / d .times. t = k 1
.times. 1 .times. { [ P ] + I 1 .function. [ HSPS ] + I 6
.function. [ SPS ] .function. [ HSF ] - k 1 .function. [ S ]
.function. [ HSP ] - k 6 .function. [ HSPHS .times. F ] .function.
[ S ] [ 31 ] d .function. [ HSPHS .times. F ] / d .times. t = k 2
.function. [ HSP ] .function. [ HSF ] + I 6 .function. [ HSPS ]
.function. [ HSF ] + I 3 .function. [ HSP ] .function. [ HSF 3 ] -
I 2 .function. [ HSPHSF ] - k 6 .function. [ HSPHS .times. F ]
.function. [ S ] [ 32 ] d .function. [ HSPS ] / d .times. t = k 1
.function. [ S ] .function. [ HSP ] + k 6 .function. [ HSPHS
.times. F ] .function. [ S ] - ( I 1 + k 1 .times. 0 ) .function. [
HSPS ] - I 6 .function. [ HSPS ] .function. [ H .times. SF ] [ 33 ]
d .function. [ HSF 3 ] / d .times. t = k 3 .function. [ HSF ] 3 + I
7 .function. [ HSF 3 ] .function. [ HSE ] - I 3 .function. [ HSP ]
.function. [ HSF 3 ] - k 7 .function. [ HSF 3 ] .function. [ H
.times. SE ] [ 34 ] .times. d .function. [ HSE ] / d .times. t = I
7 .function. [ HSF 3 ] .function. [ HSE ] - k 7 .function. [ HSF 3
] .function. [ H .times. SE ] [ 35 ] .times. d .function. [ HSF 3
.times. HSE ] / d .times. t = k 7 .function. [ HSF 3 ] .function. [
HSE ] - I 7 .function. [ HSF 3 ] .function. [ H .times. SE ] [ 36 ]
.times. d .function. [ mRNA ] / d .times. t = k 8 .function. [ HSF
3 .times. HSE ] - k 5 .function. [ mRNA ] [ 37 ] .times. d
.function. [ P ] / d .times. t = k 1 .times. 0 .function. [ HSPS ]
- k 1 .times. 1 .function. [ P ] [ 38 ] ##EQU00002##
[0261] In these expressions, [ ] denotes the cellular concentration
of the quantity inside the bracket. For Rybinski et al (2013), the
initial concentrations at the equilibrium temperature of 310K are
given in Table 5.
TABLE-US-00006 TABLE 5 Initial values of species at 310K for a
typical cell in arbitrary units [Rybinski et al (2013)]. The
arbitrary units are chosen by Rybinski et al for computational
convenience: to make the quantities of interest in the range of
0.01-10. [HSP(0)] 0.308649 [HSF(0)] 0.150836 [S(0)] 0.113457
[HSPHSF(0)] 2.58799 [HSPS(0)] 1.12631 [HSF.sub.3(0)] 0.0444747
[HSE(0)] 0.957419 [HSF.sub.3HSE(0)] 0.0425809 [mRNA(0)] 0.114641
[P(0)] 8.76023
[0262] The Rybinski et al (2013) rate constants are shown in Table
6.
TABLE-US-00007 TABLE 6 Rybinski et al (2013) rate constants giving
rates in min.sup.-1 for the arbitrary concentration units of the
previous table. l.sub.1 = 0.0175 k.sub.1 = 1.47 l.sub.2 = 0.0175
k.sub.2 = 1.47 l.sub.3 = 0.020125 k.sub.3 = 0.0805 k.sub.4 = 0.1225
k.sub.5 = 0.0455 k.sub.6 = 0.0805 l.sub.6 = 0.00126 k.sub.7 =
0.1225 l.sub.7 = 0.1225 k.sub.8 = 0.1225 k.sub.9 = 0.0455 k.sub.10
= 0.049 k.sub.11 = 0.00563271
[0263] The initial concentration values of Table 5 and the rate
constants of Table 6 were determined by Rybinski et al (2013) to
correspond to experimental data on overall HSP system behavior when
the temperature was increased on the order of 5.degree. C. for
several (e.g. 350) minutes.
[0264] Note that the initial concentration of HSPs is
100.times.0.308649/(8.76023+0.113457+1.12631)}=3.09% of the total
number of proteins present in the cell.
[0265] Although the rate constants of Table 6 are used by Rybinski
et al for T=310+5+315K, it is likely that very similar rate
constants exist at other temperatures. In this connection, the
qualitative behavior of the simulations is similar for a large
range of parameters. For convenience, we shall assume that the
values of the rate constants in Table 6 are a good approximation
for the values at the equilibrium temperature of T=310K.
[0266] The behavior of the different components in the Rybinski et
al cell is displayed in FIGS. 55A and 55B for 350 minutes for the
situation where the temperature is suddenly increased 5K at t=0
from an ambient 310K.
[0267] With continuing reference to FIG. 55, the behavior of HSP
cellular system components during 350 minutes following a sudden
increase in temperature from 37.degree. C. to 42.degree. C. is
shown.
[0268] Here, the concentrations of the components are presented in
computationally convenient arbitrary units. S denotes denatured or
damaged proteins that are as yet unaffected by HSPs; HSP denotes
free (activated) heat shock proteins; HSP:S denotes activated HSPs
that are attached to the damaged proteins and performing repair;
HSP:HSF denotes (inactive) HSPs that are attached to heat shock
factor monomers; HSF denotes a monomer of heat shock factor;
HSF.sub.3 denotes a trimer of heat shock factor that can penetrate
the nuclear membrane to interact with a heat shock element on the
DNA molecule; HSE:HSF.sub.3 denotes a trimer of heat shock factor
attached to a heat shock element on the DNA molecule that initiates
transcription of a new mRNA molecule; mRNA denotes the messenger
RNA molecule that results from the HSE:HSF.sub.3, and that leads to
the production of a new (activated) HSP molecule in the cell's
cytoplasm.
[0269] FIG. 55 shows that initially the concentration of activated
HSPs is the result of release of HSPs sequestered in the molecules
HSPHSF in the cytoplasm, with the creation of new HSPs from the
cell nucleus via mRNA not occurring until 60 minutes after the
temperature rise occurs. FIG. 55 also shows that the activated HSPs
are very rapidly attached to damaged proteins to begin their repair
work. For the cell depicted, the sudden rise in temperature also
results in a temporary rise in damaged protein concentration, with
the peak in the damaged protein concentration occurring about 30
minutes after the temperature increase.
[0270] FIG. 55 shows what the Rybinski et al equations predict for
the variation of the 10 different species over a period of 350
minutes. However, the present invention is concerned with SDM
application is on the variation of the species over the much
shorter O(minute) interval between two applications of SDM at any
single retinal locus. It will be understood that the preferred
embodiment of SDM in the form of laser light treatment is analyzed
and described, but it is applicable to other sources of energy as
well.
[0271] With reference now to FIGS. 56A-56H, the behavior of HSP
cellular system components during the first minute following a
sudden increase in temperature from 37.degree. C. to 42.degree. C.
using the Rybinski et al. (2013) equations with the initial values
and rate constants of Tables 5 and 6 are shown. The abscissa
denotes time in minutes, and the ordinate shows concentration in
the same arbitrary units as in FIG. 56.
[0272] FIG. 56 shows that the nuclear source of HSPs plays
virtually no role during a 1 minute period, and that the main
source of new HSPs in the cytoplasm arises from the release of
sequestered HSPs from the reservoir of HSPHSF molecules. It also
shows that a good fraction of the newly activated HSPs attach
themselves to damaged proteins to begin the repair process.
[0273] The initial concentrations in Table 5 are not the
equilibrium values of the species, i.e. they do not give d[ . . .
]/dt=0, as evidenced by the curves in FIGS. 55 and 56. The
equilibrium values that give d[ . . . ]/dt=0 corresponding to the
rate constants of Table 6 are found to be those listed in Table
7.
TABLE-US-00008 TABLE 7 Equilibrium values of species in arbitrary
units [Rybinski et al (2013)] corresponding to the rate constants
of Table 6. The arbitrary units are those chosen by Rybinski et al
for computational convenience: to make the quantities of interest
in the range of 0.01-10. [HSP(equil)] 0.315343 [HSF(equil)]
0.255145 [S(equil)] 0.542375 [HSPHSF(equil)] 1.982248 [HSPS(equil)]
5.05777 [HSF.sub.3(equil)] 0.210688 [HSE(equil)] 0.206488
[HSF.sub.3HSE(equil)] 0.643504 [mRNA(equil)] 0.1171274 [P(equil)]
4.39986
[0274] Note that the equilibrium concentration of HSPs is
100.times.{0.31 5343/(4.39986+5.05777+0.542375)}=3.15% of the total
number of proteins present in the cell. This is comparable, but
less than the anticipated 5%-10% total number of proteins found by
other researchers. However, we have not attempted to adjust
percentage upwards expecting that the general behavior will not be
appreciably changed as indicated by other researchers.
[0275] The inventors have found that a first treatment to the
target tissue may be performed by repeatedly applying the pulsed
energy (e.g., SDM) to the target tissue over a period of time so as
to controllably raise a temperature of the target tissue to
therapeutically treat the target tissue without destroying or
permanently damaging the target tissue. A "treatment" comprises the
total number of applications of the pulsed energy to the target
tissue over a given period of time, such as dozens or even hundreds
of light or other energy applications to the target tissue over a
short period of time, such as a period of less than ten seconds,
and more typically a period of less than one second, such as 100
milliseconds to 600 milliseconds. This "treatment" controllably
raises the temperature of the target tissue to activate the heat
shock proteins and related components.
[0276] What has been found, however, is that if the application of
the pulsed energy to the target tissue is halted for an interval of
time, such as an interval of time that exceeds the first period of
time comprising the "first treatment", which may comprise several
seconds to several minutes, such as three seconds to three minutes
or more preferably ten seconds to ninety seconds, and then a second
treatment is performed on the target tissue after the interval of
time within a single treatment session or office visit, wherein the
second treatment also entails repeatedly reapplying the pulsed
energy to the target tissue so as to controllably raise the
temperature of the target tissue to therapeutically treat the
target tissue without destroying or permanently damaging the target
tissue, the amount of activated HSPs and related components in the
cells of the target tissue is increased resulting in a more
effective overall treatment of the biological tissue. In other
words, the first treatment creates a level of heat shock protein
activation of the target tissue, and the second treatment increases
the level of heat shock protein activation in the target tissue
above the level due to the first treatment. Thus, performing
multiple treatments to the target tissue of the patient within a
single treatment session or office visit enhances the overall
treatment of the biological tissue so long as the second or
additional treatments are performed after an interval of time which
does not exceed several minute but which is of sufficient length so
as to allow temperature relaxation so as not to damage or destroy
the target tissue.
[0277] This technique may be referred to herein as "stair-stepping"
in that the levels of activated HSP production increase with the
subsequent treatment or treatments within the same office visit
treatment session. This "stair-stepping" technique may be described
by a combination of the Arrhenius integral approach for subsecond
phenomena with the Rybinski et al. (2013) treatment of intervals
between repeated subsecond applications of the SDM or other pulsed
energy.
[0278] For the proposed stair-stepping SDM (repetitive SDM
applications) proposed in this invention disclosure, there are some
important differences from the situation depicted in FIG. 55:
[0279] SDM can be applied prophylactically to a healthy cell, but
oftentimes SDM will be applied to a diseased cell. In that case,
the initial concentration of damaged proteins [S(0)] can be larger
than given in Table 7. We shall not attempt to account for this,
assuming that the qualitative behavior will not be changed. [0280]
The duration of a single SDM application is only subseconds, rather
than the minutes shown in FIG. 55. The Rybinski et al rate
constants are much smaller than the Arrhenius constants: the latter
give Arrhenius integrals of the order of unity for subsecond
durations, whereas the Rybinski et al rate constants are too small
to do that. This is an example of the different effective rate
constants that exist when the time scales of interest are
different: The Rybinski et al rate constants apply to phenomena
occurring over minutes, whereas the Arrhenius rate constants apply
to subsecond phenomena.
[0281] Accordingly, to analyze what happens in the proposed
stair-stepping SDM technique for improving the efficacy of SDM, we
shall combine the Arrhenius integral treatment appropriate for the
subsecond phenomena with the Rybinski et al (2013) treatment
appropriate for the phenomena occurring over the order of a minute
interval between repeated SDM applications: [0282] SDM subsecond
application described by Arrhenius integral formalism [0283]
Interval of O(minute) between SDM applications described by
Rybinski et al (2013) equations
[0284] Specifically, we consider two successive applications of
SDM, each SDM micropulse train having a subsecond duration. [0285]
For the short subsecond time scale, we assume that the unactivated
HSP's that are the source of the activated (free) HSP's are all
contained in the HSPHSF molecules in the cytoplasm. Accordingly,
the first SDM application is taken to reduce the cytoplasmic
reservoir of unactivated HSPs in the initial HSPHSF molecule
population from [0286] [HSPHSF(equil)] to
[HSPHSF(equil)]exp[-.OMEGA.], [0287] and to increase the initial
HSP molecular population from [0288] [HSP(equil)] to
[HSP(equil)]+[HSPHSF(equil)](1-exp[-.OMEGA.]) [0289] as well as to
increase the initial HSF molecular population from [0290]
[HSF(equil)] to [HSF(equil)]+[HSPHSF(equil)](1-exp[-.OMEGA.])
[0291] The equilibrium concentrations of all of the other species
will be assumed to remain the same after the first SDM application
[0292] The Rybinski et al equations are then used to calculate what
happens to [HSP] and [HSPHSF] in the interval .DELTA.t=O(minute)
between the first SDM application and the second SDM application,
with the initial values of HSP, HSF and HSPHSF after the first SDM
application taken to be [0293]
[HSP(SDM1)]=[HSP(equil)]+[HSPHSF(equil)](1-exp[-.OMEGA.]) [0294]
[HSF(SDM1)]=[HSF(equil)]+[HSPHSF(equil)](1-exp[-.OMEGA.]) [0295]
and [0296] [HSPHSF(SDM1)]=[HSPHSF(equil)]exp[-.OMEGA.] [0297] For
the second application of SDM after the interval .DELTA.t, the
values of [HSP], [HSF] and {HSPHSF] after the SDM will be taken to
be [0298] [HSP(SDM2)]=[HSP(At)]+[HSPHSF(At)](1-exp[-.OMEGA.])
[0299] [HSF(SDM2)]=[HSF(At)]+[HSPHSF(At)](1-exp[-.OMEGA.]) [0300]
and [0301] [HSPHSF(SDM2)]=[HSPHSF(.DELTA.t)]exp[-.OMEGA.] [0302]
where [HSP(.DELTA.t)], [HSF(.DELTA.t)], and [HSPHSF(.DELTA.t)] are
the values determined from the Rybinski et al (2013) equations at
the time .DELTA.t. [0303] Our present interest is in comparing
[HSP[SDM2)] with [HSP[SDM1)], to see if the repeated application of
SDM at an interval .DELTA.t following the first application of SDM
has resulted in more activated (free) HSP's in the cytoplasm. The
ratio .beta.(.DELTA.t,
.OMEGA.)=[HSP(SDM2)]/[HSP(SDM1)]={[{[HSP(.DELTA.t)]+[HSPHSF(.DELTA.t)](1--
exp[-.OMEGA.])}/{[HSP(0)]+[HSPHSF(0)](1-exp[-.OMEGA.])} [0304]
provides a direct measure of the improvement in the degree of HSP
activation for a repeated application of SDM after an interval
.DELTA.t from the first SDM application.
[0305] The HSP and HSPHSF concentrations can vary quite a bit in
the interval .DELTA.t between SDM applications.
[0306] FIGS. 57A and 57B illustrate the variation in the activated
concentrations [HSP] and the unactivated HSP in the cytoplasmic
reservoir [HSPHSF] during an interval .DELTA.t=1 minute between SDM
applications when the SDM Arrhenius integral .OMEGA.=1 and the
equilibrium concentrations are as given in Table 7.
[0307] Although only a single repetition (one-step) is treated
here, it is apparent that the procedure could be repeated to
provide a multiple stair-stepping events as a means of improving
the efficacy of SDM, or other therapeutic method involving
activation of tissue HSPs.
[0308] Effects of varying the magnitude of the Arrhenius integral
.OMEGA. and interval .DELTA.t between two distinct treatments
separated by an interval of time are shown by the following
examples and results.
[0309] Nine examples generated with the procedure described above
are presented in the following. All of the examples are of a
treatment consisting of two SDM treatments, with the second
occurring at a time .DELTA.t following the first, and they explore:
[0310] The effect of different magnitude Arrhenius integrals
.OMEGA. in the SDM treatments [Three different .OMEGA.'s are
considered: .OMEGA.=0.2, 0.5 and 1.0] [0311] The impact of varying
the interval .DELTA.t between the two SDM treatments [Three
different .DELTA.t's are considered: .DELTA.t=15 sec., 30 sec., and
60 sec.
[0312] As indicated above, the activation Arrhenius integral
.OMEGA. depends on both the treatment parameters (e.g., laser
power, duty cycle, total train duration) and on the RPE properties
(e.g., absorption coefficients, density of HSPs).
[0313] Table 8 below shows the effect of different .OMEGA.
(.OMEGA.=0.2, 0.5, 1) on the HSP content of a cell when the
interval between the two SDM treatments is .DELTA.t=1 minute. Here
the cell is taken to have the Rybinski et al (2013) equilibrium
concentrations for the ten species involved, given in Table 7.
[0314] Table 8 shows four HSP concentrations (in the Rybinski et al
arbitrary units) each corresponding to four different times: [0315]
Before the first SDM treatment: [HSP(equil)] [0316] Immediately
after the first SDM application: [HSP(SDM1)] [0317] At the end of
the interval .DELTA.t following the first SDM treatment:
[HSP(.DELTA.t)] [0318] Immediately after the second SDM treatment
at .DELTA.t: [HSP(SDM2)] [0319] Also shown is the improvement
factor over a single treatment: .beta.=[HSP(SDM2)]/[HSP(SDM1)]
TABLE-US-00009 [0319] TABLE 8 HSP concentrations at the four times
just described in the text: Effect of varying the SDM .OMEGA. for
two SDM applications on a cell when the treatments are separated by
.DELTA.t = 0.25 minutes = 15 seconds. [HSP.sub.(equil)]
[HSP.sub.(SDM1)] [HSP(.DELTA.t)] [HSP.sub.(SDM2)] .beta. .OMEGA. =
0.2 0.315 0.67 0.54 0.95 1.27 .OMEGA. = 0.5 0.315 1.10 0.77 1.34
1.22 .OMEGA. = 1.0 0.315 1.57 0.93 1.71 1.09
[0320] Table 9 is the same as Table 8, except that it is for an
interval between SDM treatments of .DELTA.t=0.5 minutes=30
seconds.
TABLE-US-00010 TABLE 9 HSP concentrations at the four times
described in the text: Effect of varying the SDM .OMEGA. for two
SDM treatments on a cell when the treatments are separated by
.DELTA.t = 0.5 minutes = 30 seconds. [HSP.sub.(equil)]
[HSP.sub.(SDM1)] [HSP(.DELTA.t)] [HSP.sub.(SDM2)] .beta. .OMEGA. =
0.2 0.315 0.67 0.44 0.77 1.14 .OMEGA. = 0.5 0.315 1.10 0.58 1.18
1.08 .OMEGA. = 1.0 0.315 1.57 0.67 1.59 1.01
[0321] Table 10 is the same as the Tables 8 and 9, except that the
treatments are separated by one minute, or sixty seconds.
TABLE-US-00011 TABLE 10 HSP concentrations at the four times just
described in the text: Effect of varying the SDM .OMEGA. for two
SDM treatments on a normal (healthy) cell when the treatments are
separated by .DELTA.t = 1 minute = 60 seconds. [HSP.sub.(equil)]
[HSP.sub.(SDM1)] [HSP(.DELTA.t)] [HSP.sub.(SDM2)] .beta. .OMEGA. =
0.2 0.315 0.67 0.30 0.64 0.95 .OMEGA. = 0.5 0.315 1.10 0.37 1.06
0.96 .OMEGA. = 1.0 0.315 1.57 0.48 1.51 0.96
[0322] Tables 8-10 show that: [0323] The first treatment of SDM
increases [HSP] by a large factor for all three .OMEGA.'s, although
the increase is larger the larger .OMEGA.. Although not displayed
explicitly in the tables, the increase in [HSP] comes at the
expense of the cytoplasmic reservoir of sequestered (unactivated)
HSP's: [HSPHSF(SDM1)] is much smaller than [HSPHSF(equil)] [0324]
[HSP] decreases appreciably in the interval .DELTA.t between the
two SDM treatments, with the decrease being larger the larger
.DELTA.t is. (The decrease in [HSP] is accompanied by an increase
in both [HSPHSF]--as shown in FIG. 44 and in [HSPS] during the
interval .DELTA.t--indicating a rapid replenishment of the
cytoplasmic reservoir of unactivated HSP's and a rapid attachment
of HSP's to the damaged proteins.) [0325] For .DELTA.t less than 60
seconds, there is an improvement in the number of activated (free)
HSP's in the cytoplasm for two SDM treatments rather than a single
treatment. [0326] The improvement increases as .DELTA.t becomes
smaller. [0327] For .DELTA.t becoming as large as 60 seconds,
however, the ratio .beta.=[HSP(SDM2)]/[HSP(SDM1)] becomes less than
unity, indicating no improvement in two SDM treatments compared to
a single SDM treatment although this result can vary depending on
energy source parameters and tissue type that is treated. [0328]
The improvement for .DELTA.t<60 seconds is larger the smaller
the SDM Arrhenius integral .OMEGA. is.
[0329] The results for the improvement ratio
.beta.=[HSP(SDM2)]/[HSP(SDM1)] are summarized in FIG. 45, where the
improvement ratio .beta.=[HSP(SDM2)]/[HSP(SDM1)] vs. interval
between SDM treatments .DELTA.t (in seconds) for three values of
the SDM Arrhenius integral .OMEGA., and for the three values of the
interval .DELTA.t=15 sec, 30 sec, and 60 sec. The uppermost curve
is for .OMEGA.=0.2; the middle curve is for .OMEGA.=0.5; and the
bottom curve is for .OMEGA.=1.0. These results are for the Rybinski
et al (2013) rate constants of Table 6 and the equilibrium species
concentrations of Table 4.
[0330] It should be appreciated that results of Tables 8-10 and
FIG. 58 are for the Rybinski et al. (2013) rate constants of Table
6 and the equilibrium concentrations of Table 7. The actual
concentrations and rate constants in a cell may differ from these
values, and thus the number results in Tables 8-10 and FIG. 58
should be taken as representative rather than absolute. However,
they are not anticipated to be significantly different. Thus,
performing multiple intra-sessional treatments on a single target
tissue location or area, such as a single retinal locus, with the
second and subsequent treatments following the first after an
interval anywhere from three seconds to three minutes, and
preferably ten seconds to ninety seconds, should increase the
activation of HSPs and related components and thus the efficacy of
the overall treatment of the target tissue. The resulting
"stair-stepping" effect achieves incremental increases in the
number of heat shock proteins that are activated, enhancing the
therapeutic effect of the treatment. However, if the interval of
time between the first and subsequent treatments is too great, then
the "stair-stepping" effect is lessened or not achieved.
[0331] The technique of the present invention is especially useful
when the treatment parameters or tissue characteristics are such
that the associated Arrhenius integral for activation is low, and
when the interval between repeated applications is small, such as
less than ninety seconds, and preferably less than a minute.
Accordingly, such multiple treatments must be performed within the
same treatment session, such as in a single office visit, where
distinct treatments can have a window of interval of time between
them so as to achieve the benefits of the technique of the present
invention.
[0332] Although several embodiments have been described in detail
for purposes of illustration, various modifications may be made
without departing from the scope and spirit of the invention.
Accordingly, the invention is not to be limited, except as by the
appended claims.
* * * * *
References