This video was created to support their published research. The authors did research using several lasers and slices of a sheep’s brain to try and determine the best parameter for treating TBI (Traumatic Brain Injury) with a desired fluency of 0.9 to 15 joules/cm2 at a depth of 2 cm. They state that getting the energy through the skull is especially difficult so they test multiple options so test the transfer rate. They started out using a continuous output split 980/810nm system (the only company that makes that type of split system, 80% of the power at 980nm and 20% of the power at 810nm, is LiteCure with their LightForce series). The result was less than 1/2% of the energy reached a depth of 2cm. Then they switched to pulsing and got an increase in the energy transfer. When they switched to a 810nm-only 15 watt system with pulsing the transfer rate increased to 16% of the output energy reached the target depth.
Here are some rough numbers to review the feasibility of using this system for treatment. If the duty cycle is 70%, the system will deliver 1.68 joules per second at a depth 2cm (15wattS*70%*16%). To get 5 joules/cm2 over 15 x 15 cm treatment area would require a total of 1125 joules at depth. This would take 23 minutes.
This research shows that only class 4 systems can delivery the level of power needed for this kind of therapy in a typical rushed doctor's office. A class 3b system with 1 watt would take 4 - 5 hours per treatment to get the same dosage.
The original research publication is titled " Treatments for traumatic brain injury with emphasis on transcranial near-infrared laser phototherapy"
video length: (9:18)
1Department of Zoology, The George S. Wise Faculty of Life Sciences, Tel-Aviv University, Tel-Aviv, Israel
2Department of Cell Biology and Immunology, The George S. Wise Faculty of Life Sciences, Tel-Aviv University, Tel-Aviv, Israel
Received 27 May 2013; revised 29 June 2013; accepted 16 July 2013
In this study, we investigated the hypothesis that photo- biostimulation by low-energy laser therapy (LLLT) applied to the bone marrow (BM) of myocardial in- farcted rats may attenuate the scarring processes that follow myocardial infarction (MI). Wistar rats under- went experimental MI. LLLT (Ga-Al-As diode laser) was applied to the BM of the exposed tibia at differ- ent time intervals post-MI (4 hrs, 48 hrs and 5 days). Sham-operated infarcted rats served as control. In- farct size was significantly reduced (55%) in the la- ser-treated rats as compared to the control non-treat- ed rats, at 2 weeks post-MI. A significant 3-fold in- crease was observed in the density of desmin immu- nopositive stained cells 14 days post-MI in the infarc- ted area of the laser-treated rats as compared to the non-laser-treated controls. The electron microscopy from the control infarcted rat hearts revealed a typi- cal interphase area between the intact myocardium and the infarcted area, with conspicuous fibroblasts with collagen deposition dispersed among them. In rats that were laser treated (to BM), the interphase zone demonstrated cells with different intracellular struc- tures. There was also a significant increase in the per- centage of c-kit positive cells and macrophages in the circulating blood of the laser treated rats as compar- ed to control non treated ones. In the majority of the cells clusters of myofibrils anchored to well-developed Z-lines and structures resembling the morphological characteristics of mature intact cardiomyocytes were evident. In conclusion, LLLT to the BM of rats post- MI induces cardiogenesis mainly at the borders of the infarcted area in the heart.
Keywords: Low-Level Laser Therapy; Myocardial Infarction; Macrophage; Desmin; Ultrastructure; c-Kit Positive Cells
Regenerative capacity and mitotic activity in the heart are confined mainly to the lower vertebrates . Amputation of ~20% of the zebrafish’s ventricular myocardium re- sulted in full regeneration without scarring . In am- phibians, heart injury was associated with increased cell proliferation of myocytes and enhanced regeneration . The adult mammalian heart was traditionally considered to be a post-mitotic organ with terminally differentiated cardiac myocytes. However, this dogma has recently been challenged by several studies and reviews [4-8]. These studies have suggested that cardiac myocytes are replaced throughout the lifespan even in the human heart, and that myocytes can regenerate from resident cardiac progenitor cells (CPC) as well as from bone marrow (BM). Studies in human infarcted hearts have shown evidence of cytoki- nesis of cells in the heart and evidence of cardiac stem cells that are activated in response to ischemic injury. This growth response is attenuated in chronic heart fail- ure . Some studies have reported that cardiac myocyt- es can be derived from BM; specifically, side population precursor cells following induction of myocardial infarc- tion (MI) by left anterior descending artery (LAD) liga- tion [10-12]. Contradicting these findings, other laborato- ries using genetic markers have reported that lineage ne- gative, c-kit+ BM cells did not differentiate into cardio- myocytes . It was also suggested that BM-derived stem cells may stimulate the small population of stem cells in the ischemic heart to proliferate and differentiate to enhance cardiac repair post-MI . In a recent study transient regenerative potential in the mouse heart was demonstrated during the neonatal period .
Low-level laser therapy (LLLT) has been found to modulate various biological processes [16,17], such as increasing mitochondrial respiration and ATP synthesis , facilitating wound healing and promoting the proc- ess of skeletal muscle regeneration and angiogenesis [19- 21]. In an experimental model of the infarcted heart in rats and dogs, it was demonstrated that LLLT application directly to the infarcted area in the heart at optimal power parameters significantly reduced scar tissue formation [22-24]. This phenomenon was partially attributed to a significant elevation in ATP content, heat shock proteins, vascular endothelial growth factor (VEGF), inducible ni- tric oxide (NO) synthase, and angiogenesis in the ischemic zone of the laser-irradiated rats, as compared to non- irradiated rats .
The effect of photobiostimulation on stem cells or pro- genitor cells has not been extensively studied. LLLT ap- plication to normal human neural progenitor cells signi- ficantly increases ATP production in these cells . LLLT delivery to MSCs and cardiac stem cells in vitro caused a significant enhancement in their proliferation rate [27,28]. LLLT has also been shown to increase the proliferation rate of adipose-derived stem cells in vitro . Recently, we demonstrated that LLLT application to autologous BM could induce mesenchymal stem cells (MSCs) in the BM to proliferate and cause their recruit- ment and specific homing in on the infarcted rat heart and not on other organs [30,31]. The laser treatment to the BM also caused a marked and statistically significant reduction of 79% in the scarring and ventricular dilata- tion followed MI as compared to infarcted non-laser- treated rats. The aim of the present study was to investi- gate the possibility that induction of stem cells in the BM of rats by LLLT could also affect cardiogenesis in the in- farcted rat heart.
2. MATERIALS AND METHODS
2.1. Experimental Procedures
A total of 21 Wistar male rats, weighing 200 - 250 gr, that underwent ligation of the LAD artery to induce MI, were used as described by us previously . All the ex- perimental procedures were approved by the animal care committee of Tel-Aviv University. Briefly, rats were anes- thetized with Avertin (1 ml/100 g body weight I.P.) and the lungs were ventilated. Thoractomy was performed by invasion of the intercostals muscles between the 5th and 6th rib to expose the heart. The LAD artery was occluded 2 mm from the origin with 5-0 polypropylene thread (Ethicon Inc., Cincinnati, OH). Following LAD artery occlusion the chest muscles and skin were sutured and the rats were ventilated until they woke up. The infarcted rats were divided randomly into two groups. In one group LLLT was applied directly to the BM 4 hrs, 48 hrs and 5 days post-MI (see below). The second group was non-laser-treated (the rat’s bone was exposed for the same duration as the laser-treated group but the laser was not turned on). Food and water were supplied ad libitum. Rats were sacrificed 14 days post-MI.
2.2. Laser Application
After induction of MI rats were randomly assigned to a laser-treated or control non-laser-treated group. A diode (Ga-Al-As) laser, wavelength 804 nm with a tunable po- wer output of maximum of 400 mW (Lasotronic Inc., Zug, Switzerland) for application to the BM was used. The laser device was equipped with a metal-backed glass fiber optic (1.5 mm diameter). An infrared viewer (Laso- tronic Inc. Zug, Switzerland) and infrared-sensitive de-tecting card (Newport, Inc., Irvine, CA) were used to de- termine the infrared irradiation area. Laser application was done by a 10 mm longitudinal cut in the skin above the medial aspect, and further delicate cleaning of the bone surface was carried out. The tip of the fiber optic (1.5 mm diameter) was placed perpendicularly to the center of the exposed medial aspect of the tibia and power den- sity of 10 mW/cm2 was applied to the BM. The laser was applied for a duration of 100 sec (energy density 1.0 J/cm2). Left or right exposed tibias were chosen at random for LLLT application. In sham-operated infarcted rats that served as control the tibias were exposed and the fi- ber optic was placed as described above but the laser beam was not turned on.
2.3. Histology and Electron Microscopy
A defined cross-section sample (2 mm thick) from the central part of the infarcted area was taken from all hearts for histology. Eight micron paraffin sections were pre- pared from the tissue samples of each heart. Infarct size was determined using Masson’s trichrome staining as described by us previously . Three observers, blinded to control or laser-treated rats, analyzed infarct size. Six microscopic slides from the infarcted area of each heart were chosen at random for determination of infarct size. Infarct size was expressed as the percentage of the total infarcted area relative to the total area of the left ventri- cle (LV) in each section, using image analysis software Sigma Scan Pro (Sigma, St. Louis, MO).
For electron microscopy three tissue samples from each of the control and laser-irradiated rat hearts were taken from the interphase zone between the infarcted and non-infarcted tissue by macroscopic examination. Fixa- tion was performed in 3.5% glutaraldehyde in 0.1 M ca- codylate buffer for 24 hrs followed by embedment in Epon-812. Semi-thin sections (1 micron) were prepared in order to localize the interphase zone. Thin sections were then prepared and stained with uranyl acetate and lead citrate followed by examination with a Jeol electron microscope.
The total number of cells immunostained for desmin (bone marrow cells or newly formed) in the infarcted area were determined using a desmin kit (Zytomed Laboratory, Ber- lin, Germany). The procedure was performed at room temperature with anti-mouse (dilution 1:25 - 1:50) primary antibody for 60 min. Following washing, slides were in- cubated with HRP secondary antibody for mouse for 30 min followed by DAB Chromogen system (Covance Inc., Dedham). Slides were rinsed again in wash buffer, stain- ed in Hematoxylin for nuclei detection, mounted and viewed using a Zeiss microscope equipped with a camera and video screen. The total number of desmin immuno- stained cells within the infarcted area was counted and their density expressed as the percentage of the total area of the infarct using SigmaPro software.
2.5. Flow Cytometry Analysis
Blood samples were taken 2 and 7 days post-IR injury for fluorescence-activated cell sorting (FACS) analysis. 100 μl of blood were mixed with different antibodies: anti-mouse CD117 (c-kit) PE (eBioscience San Diego, USA) and rat IgG2b isotype control PE (eBioscience San Diego, USA) and anti-rat macrophage marker PE (eBio- science San Diego, USA) and mouse IgG2a K isotype control PE (eBioscience San Diego, USA), were used for the FACS analysis according to the manufacturer’s guide- lines. Forty five min post incubation of the whole fresh blood with the relevant antibodies, 2 ml of Fix/Lyse so- lution (eBioscience, San Diego, USA) was added. After mixture the suspended cells were left for 60 min in the dark at room temperature. Centrifugation was performed for 10 min, supernatant was removed and washing of the pellet was performed with 2 ml of Flow Cytometry Stain- ing Buffer Solution (eBioscience, San Diego, California, USA). After another centrifugation for 10 minutes the supernatant was decanted. The pellet containing mono- nucleated cells was resuspended in 200 μl of flow stain buffer for FACS analysis.
2.6. Statistical Analysis
The SigmaStat 2.0 (Sigma, St. Luis, USA) software was used for statistical analysis. Tests were performed first for normality distribution, followed by parametric (stu- dent’s t-test) test.
Application of LLLT to the infarcted heart caused a sig- nificant (p = 0.049) reduction of 55% in infarct size as compared to control. The present of macrophages and c- kit positive cells in the blood was determined by FACS analysis (Figure 1). It was found that at 5 days post MI there was a statistical significant 2-fold higher concentra- tion of macrophages and significant 1.4-fold higher c-kit positive cells (mesenchymal cells) in the laser treated rats as compared to the infarcted non laser treated rats. Des- min immunostaining of histological sections of the in- farcted zone from laser-treated rats demonstrated a higher density of positively stained cells than in the non laser-treated ones (Figures 2-4). In the interphase zone, cells extending from the myocardium towards the in
Figure 1. Percent (out of total mononucleated cells) of macro- phages and c-kit positive cells in blood of control and laser treated rats (to the bone marrow) 5 days post MI as revealed by FACS analysis. The results are mean ± S.E.M of 15 rats at each group. Statistical significance *p < 0.05; **p < 0.01.
Figure 2. Representative desmin immunostained light micro- graphs of the infarcted zone of non-laser-treated rats (a, c) and laser-treated rats (to the bone marrow at 4 and 48 hrs and 5 days) (b, d) taken 2 weeks post-MI. Note that the zone in the control non-laser-treated rats contains mainly collageneous mate- rial with a few desmin immunopositive cells in the infarcted area (a, c); while in the laser-treated rats the zone displays posi- tive desmin staining in extended outgrowths (arrow) from the myocardium (MC) in (b), and in the cytoplasm of many cells in the infarcted area in (d). IF, Infarcted area. Bar = 50 μm.
farcted area showed higher immunostaining for desmin in the laser-treated rat hearts as compared to the control non-treated ones (Figure 2). The cell density of desmin immune-positive cells was also determined quantitatively in histological sections of both the infarcted laser-treated rats and infarcted non-laser-treated rats. The cell density was significantly (p < 0.01) 3-fold higher in the infarcted area of the laser-treated rats as compared to the non-la- ser-treated controls (Figure 4).
The electron micrographs of all samples taken from the control non-laser-treated infarcted rat hearts revealed a typical interphase area between intact and infarcted heart (Figure 5(a)). Adjacent to the non-ischemic intact myocardium there were conspicuous fibroblasts with col- lagen deposition dispersed among them (Figure 5(a)). In all samples taken from the laser-irradiated hearts the in- terphase zone between intact and infarcted area demon- strated different characteristics to those of the non-laser- treated infarcted rat hearts. Cells with newly-formed or- ganized contractile myofilaments dispersed in the cyto- plasm were detected in groups of several cells (Figure 5(b)). In these cells numerous mitochondria, clusters of ribosomes, and conspicuous clusters of contractile pro- teins were evident in the cytoplasm (Figures 6-8). Some cells contained dispersed contractile myofilaments in the cytoplasm that were still in an early stage of organization (Figure 6). The organization of newly-formed contractile myofilaments in the cytoplasm was observed in various
Figure 3. Representative desmin immunostained light micro- graphs of the interphase of the infarcted zone of laser-treated rats. Note that desmin positively stained cross-sections of myo- fibers (arrows) intermingled in the infarcted zone in (a). In (b) immunopositively stained cross-sections of myofibers (arrow) are visible in the infarcted area (IF). In (c) newly-formed car- diomyocytes (NC) are seen, with the desmin immunostaining mainly confined to the Z-line. Bar = 50 μm.
Figure 4. Density of desmin positively stained area (relative to total area) in the infarcted areas of control (non-laser-treated) and laser-treated (to the bone marrow) rats at 14 days post-MI. Results are mean+ S.E.M from 6 - 8 rats in each group. **p < 0.01.
Figure 5. Electron micrographs of typical interphase zone be- tween myocardium and infarcted area of control non-laser- treated (a) and laser-treated (b) to bone marrow rats. Note intact myocardium (MY) and adjacent fibroblast (FB) in the infarcted area surrounded by collagen (CL) deposition in (a). In (b) sev- eral newly-formed cardiomyocytes (marked with asterix) with conspicuous well-organized myofilaments (MF) in their cyto- plasm are evident adjacent to blood capillaries (CA). EN, En- dothelial cell.
degrees of maturation in those cells. In some cells the myofilaments were dispersed in the cytoplasm and in others they were organized in clusters anchored to well- developed Z-lines (Figure 7(a)). In certain cells the myo- filaments were organized parallel to the longitudinal di- rection of the cells, resembling the morphological char- acteristics of mature intact cardiomyocytes (Figure 7(b)). Some of the cells were also seen in a process of forma- tion of typical intercalated disc between them (Figure 9).
4. DISCUSSION AND CONCLUSION
The most significant outcome of this study was the ap- pearance of newly-formed cardiomyocytes following laser treatment to the BM, as indicated by light and electron microscopy. There was a 3-fold increase in the density of
Figure 6. Electron micrographs of most probably newly-formed cardiomyocytes at an early stage of organization of contractile myofilaments. Note myofilaments (MF) in the cytoplasm. M, Mitochondrion. Bar = 1 μm.
Figure 7. Electron micrographs of most probably newly-formed cardiomyocytes with early (a) and late (b) stages of the organi- zation of the contractile myofilaments in the cytoplasm. Note contractile myofilaments that are dispersed (DMF) in the cyto- plasm with a few organized in clusters anchored to Z-lines (Z) in (a). In (b) myofilaments (MF) are organized in parallel to the longitudinal axis of the cardiomyocyte, resembling their orga- nization in mature cardiomyocyte. N, Nucleus. Bar = 1 μm.
desmin immunostained cells in the infarcted rat hearts that had been laser treated. Desmin is a protein found in the cytoplasm of developing myocytes and cardiomyo- cytes . The significantly higher occurrence of des- min-positive cells in the infarcted area of the laser- treated hearts may indicate the synthesis of new contrac- tile proteins in the developing new cardiomyocytes, re- sembling the process that takes place during embryonic development. The ultrastructural features of the cells in the interphase between the intact myocardium and the
Figure 8. Electron micrographs of typical interphase zone be- tween myocardium and infarcted area of laser-treated infarcted rat heart. Note numerous mitochondria (M) in the cytoplasm of the cardiomyocytes in (a) and (b). Also note organized contrac- tile myofilament with well-developed Z-lines (Z), some dis- persed myofilaments and clusters of ribosomes (R). Bar = 1 μm.
Figure 9. Electron micrographs of typical intercalated disk formation in the interphase region of the infarcted heart of la- ser-treated rats. Formation of intercalated disks (ID) between cells (marked with asterix) is evident. Note that the most proba- bly newly-formed cardiomyocytes contain clusters of myofila- ments (MF) in the cytoplasm that are conspicuous in their obli- que or cross-sections (arrows). Bar = 1 μm.
infarcted myocardium of the laser-treated rats, as shown in this study, clearly resemble the characteristics of car- diomyocytes during embryonic development of the heart . Furthermore, the clusters of ribosomes and the nu- merous clusters of mitochondria in the cytoplasm of these cells may characterize cells that are active in the synthe- sis of proteins. It was previously demonstrated that direct LLLT to the infarcted hearts of rats, dogs and pigs caus- ed a significant reduction of scarring post-MI [23,24]. It was suggested that part of this reduction could be ex- plained by the regenerative response that takes place in the interphase zone .
The results of the present study indicate that the LLLT
applied to autologous BM attenuates the concentration of macrophages and MSC in the circulating blood. We have previously shown that LLLT application to the BM of infarcted rats caused a 2 fold enhancement in the rate of proliferation of MSC in the BM . Those cells that most probably leave the BM to the circulating blood in- deed show a significant elevation of their concentration (as reveled by the FACS analysis in the present paper) at 5 days post MI. Consequently these cells probably home in on the infarcted heart, and even migrate specifically to the infarcted area . These cells may induce cardiac stem cells to differentiate to newly-formed cardiomyo- cytes, as suggested previously by Hatzistergos et al. . Indeed, it was found that endogenous c-kit+ cardiac stem cells were increased by 20-fold in the rat infarcted heart compared to control, following transcardial injection of BM-derived MSCs . Such induction may be enabled due to paracrine secretion of various growth factors by the laser-stimulated MSC that originated from the BM. The possibility that paracrine secretion occurs in im- planted stem cells during cell therapy to the heart post- MI has been suggested previously . Another mecha- nism that may take place after homing of stem cells to the infarcted heart of the laser-stimulated rats is that these cells continue to proliferate in the appropriate mi-lieu of the interphase zone in the infarcted heart and then differentiate to cardiomyocytes .
Another possible mechanism that maybe associates with the reduction of infarct size is the significant increase in the concentration of macrophages in the circulation fol- lowing LLLT to the BM as revealed from the FACS analysis in the present study. These findings corroborate with studies indicating that macrophages activity in the infarcted area at early stages post MI cause reduction of scarring post MI [35,36]. Thus, it could be postulated that more macrophages that will eventually home in the infarcted area from the circulating blood in the laser treated rats will also contribute to the reduction of scar- ring.
Although the findings of the present study do not in- dicate the extent of regenerative capacity of the rat in- farcted heart post-laser-irradiation, they do reveal a shift from practically no cardiomyocytes in the tissue samples taken from the non-laser-treated hearts, to the presence of newly-formed cardiomyocytes in all the electron mi- croscope sections taken from the hearts of rats that are laser-treated to the BM.
In conclusion, to the best of our knowledge, this is the first study to demonstrate the appearance of newly-form- ed cardiomyocytes in the infarcted area following LLLT to autologous BM in the infarcted rat heart. The mecha- nisms associated with this phenomenon remain to be elu- cidated in further studies.
This study was partially supported by the Elizabeth and Nicholas Shle- zak Super-center for Cardiac Research and Medical Engineering. The authors wish to acknowledge N. Paz for editing the manuscript and V. Wexler for helping with preparation of the figures.
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Dr. Hamblin discusses the use of low level laser therapy for all type of brain injuries. He is an expert in all type of light healing (see below). He has performed much of his research on rats. He claims several key points:
This video is mainly about TBI but the principals are universal. Dr Hamblin is associated with Thor laser so there is some potential for bias but he is also assocated with Harvard and the Wellman Centre. Introduction to Low Level Laser Therapy (LLLT) for Traumatic Brain Injury (TBI) by Mike Hamblin. Wellman Centre for Photomedicine, Harvard Medical School.
video length: (6:18)
This video is restricted for minors.
In this video Dr. Michael Hamblin is interviewed by youtuber Selfhacked. Dr. Hamblin discusses the use of red and near infrared light in low level light therapy, he talks about the different possibilities of LLLT as well as some of his personal experiences with LLLT. He explains the reasons behind some of the effects from LLLT noticed in patients over the years, including it's effect on calcium in cells, ion channels, and infrared light vs. ultraviolet light. Dr. Hamblin also notes the differences between LLLT, bright light therapy, and light from the sun that is used for theraputic purposes. The majority of the video is spent discussing the effect of LLLT on the brain specifically.
video length: (52:08)
This video shows before and after treatment footage of a patients with advanced Parkinsons. Treatment lasts about 2 weeks. Dr Riner is using the brain and neurostim setting on the brain, C5 Nerve Root and the Ulnar nerve in the elbow.
The TheraLazr is the prototype for the Avant LZ30 series of lasers.
video length: (2:03)
LLLT being used to treat a patient with TMJ. She experiences pain in her cheeks and neck particularly, and afterward claims to have no pain or discomfort. The doctor states that the LLLT has reduced inflammation, and relaxed the muscles that were causing the spasms.
video length: (2:43)
The laser info starts around 30 minutes into the presentation.
The presentation includes research done on rats for the following conditions:
The research was supported by Thor so it could be biased but their research indicates that 810nm provides better stimulation of the cells.
video length: (17:09)
In this video, you will see one of the best graphic representations of the photobiomodulation process including the release of ATP, NO, ROS.
This video was created by LightCure so at the end of the video, they try to make the point that class 4 laser are better than class 3 system. With the release of high-power class 3b (with power levels over 15 watts) system, this part of the video out of date.
video length: (3:08)
Dr. Burke shares some of his quite impressive results in patients who recieved LLLT, incuding:
video length: (13:46)
News report about the use of low level light therapy (LLLT) on a surgical wound that wouldn't heal after nearly a year in a patient with diabetes.
video length: (2:17)
Young man with mucositis recieved a bone marrow transplant and was given LLLT both pre-, and post-op. He was expected to have serious mouth sores, but thanks to the LLLT he had minimal sores that quickly went away.
video length: (1:35)
Chad Davis, DVM, gives a demonstration of the effecacy of LLLT on a horse.
video length: (4:39)
This short video shows 3 before and after videos of paralized dogs who recieved LLLT and were able to walk again.
video length: (1:25)
This video discusses the basics of Low Level Laser Therapy. You will learn a little bit about lasers and laser history, and what makes a cold laser a cold laser. It also talks about the difference between lasers and LED's and why the latter may be less effective for medical therapy.
You'll find information on the treatment parameters of LLLT which are:
It also goes over what indications the FDA has approved LLLT and infrared light for.
video length (12:13)
Dr. Michael Hamblin Talks about the differences between photodynamic therapy (PDT) and low level laser therapy (LLLT), and explains the basics of LLLT.
video length: (4:35)
Class IV K Laser is an excellent new treatment for helping muscle pulls, sprains, strains, and joint injuries. Cold Laser is a great treatment for speeding the recovery of knee pain, chondromalacia patella, patella tendonitis, and knee sprains. Lasers help increase cellular ATP, which is the cells energy source. Cells use the increased energy for healing and repair. Lasers are also excellent at increasing metabolic and repair process within the tissue. They decrease inflammation and pain around muscles, tendons, and joints. In addition, the speed recovery and healing of nerves, especially pain nerves.
video length: (1:51)
Dr. Adam Zuckerman talks about his use of LLLT in patients with diabetes.
video length: (1:46)
Veterinarian Ron Hirschberg tells of how he first experienced laser therapy on his own arthritis, and decided to invest in lasers for his practice. Since he started using laser therapy on the pets he works with he has seen many positive results, he explains that gross profits from NSAIDs decreased from 0.71% to 0.3%, and laser profit now makes up 4% of the practice's income. He will treat between 3 and as many as 13 patients with LLLT, 95% of which he says improve noticably faster than those without LLLT.
viedo length: (13:37)
In this video Spokane Chiropractor Dr. Patrick Dougherty gives a quick demonstration of how cold laser therapy is used as a chiropractic treatment to help with range of motion by affecting the nervous system. This can be used as an effective adjunct to chiropractic adjustments when the brain is having a difficult time holding on to the input that the adjustments provides to the brain.
video length: (4:34)
This short animation from gives a simple description of how LLLT effects individual cells and things like:
video length: (1:21)
Dr. Larry Lytle discusses LLLT, particularly for pain in an interview.
video length: (27:37)
Cold Laser Therapy can be very effective for those suffering from pain caused by auto accidents. Cold Laser Therapy is equally effective for those suffering from pain caused by work-related accidents. Athletes get great relief from sports-related injuries using Cold Laser Therapy. Cold Laser Therapy is a powerful therapy in the battle to relieve back pain, neck pain and joint pain.
Importantly, studies to date indicate that Cold Laser Therapy has no serious side effects when used by a trained healthcare professional. It is a non-invasive procedure requiring no surgical incision. There is no recovery time after a treatment. You do not have to take any medications relating to Cold Laser Therapy.
video length: (5:49)
Using LLLT to treat neck pain, procedure without wavelength or power, or any other laser parameters given.
video length: (1:14)
Dr. Fritz teaches Obstetrics & Gynecology at the Academy for Oriental Medicine in Austin.
After getting undergraduate and graduate degrees in Biology from the University of Virginia, Dr. Vanessa Fritz graduated from the National College of Natural Medicine in Portland, Oregon, with a doctorate in Naturopathic Medicine (ND) as well as a Master of Science in Oriental Medicine (MSOM).
video length: (9:32)
This video gives a somewhat in-depth list of dental issues that LLLT can be used to treat and how to treat them, along with a simple explaination of what LLLT is. At (16:48) there is a demonstration on how to use a Zolar laser for LLLT.
video length: (19:49)
LLLT is known to dramatically improve conditions associated with soft tissue inflammations, not only by reducing pain, but also by providing a significant therapeutic advantage resulting in inflammation reduction, as well as expediting the healing process.
video length: (3:25)
A man experiencing tooth sensitivity atests to pain reilief from LLLT.
video length: (0:54)
Dr. Bernard Filner MD Discusses his use of LLLT in pain management.
video length: (2:48)
Significant changes in blood flow or in the integrity of cerebral vessels are believed to cause cerebrovascular disease (CVD) and to contribute to dementias including Alzheimer’s disease . Stroke, the most serious form of CVD, is one of the leading causes of death and adult disability worldwide. Acute treatments for stroke, however, are severely limited. Neuroprotective drugs under development show promise at halting the ischemic cascade, but as yet, no such compound has received federal approval in the United States. One of the biggest limitations to this development is the lack of understanding of the mechanisms by which cerebral vessels react to factors such as ischemia, inflammation, blood pressure changes, metabolic demands, and trauma . In order to address these fundamental questions, functional brain imaging techniques such as fMRI and intrinsic signal optical imaging (ISOI) have emerged as tools to visualize and quantify cerebral hemodynamics.
In the neuroscience community, ISOI has long been used to study the organization and functional architecture of different cortical regions in animals and humans [3–5] (see other chapters in this book). Three sources of ISOI signals that affect the intensity of diffusely reflected light derive from characteristic physiologic changes in the cortex. For functional neuronal activation, these have been observed to occur over a range of timescales, including (1) light scattering changes, both fast (over 10 s of milliseconds) and slow (i.e., > ~0.5 s) (2) early (~0.5–2.5 s) absorption changes from alterations in chromophore redox status, i.e., the oxy/deoxy-hemoglobin ratio (known as the “initial dip” period), and (3), slower (~2–10 s) absorption changes due to blood volume increase (correlated with the fMRI BOLD signal). Light scattering changes have been attributed to interstitial volume changes resulting from cellular swelling, organelle swelling due to ion and water movement, capillary expansion, and neurotransmitter release [6,7]. The slower absorption factors have been demonstrated to correlate with the changes in metabolic demand and subsequent hemodynamic cascades following neuronal activation [4,8,9].
Using animal models of acute and chronic brain injury, ISOI has been used to quantify the acute hemodynamic events in response to stroke, including focal ischemia and cortical spreading depression (CSD) [10–21]. Researchers have also used ISOI to locate and quantify the spatial extent of the stroke injury, including ischemic core, penumbra, and healthy tissue zones [18,22]. CSD also plays a key role in migraine headache, and recent laser speckle imaging studies have revealed the neurovascular coupling mechanism to the transmission of headache pain [23,24].
To fully understand the underlying mechanisms in vascular changes associated with cerebrovascular diseases such as stroke, an optical imaging technique that has the capability to rapidly separate absorption from scattering effects can enhance the information content of traditional ISOI, enabling (1) more accurate quantitation of hemodynamic function, (2) isolation of the electro-chemical changes characterized by light scattering, and (3) longitudinal chronic injury studies of function where structural reorganization due to neovascularization can cause significant alterations in scattering [25,26].
Quantitative diffuse optical methods  such as spatially-resolved reflectance, diffuse optical spectroscopy (DOS), and tomography (DOT), and diffuse correlation spectroscopy (DCS) possess exquisite sensitivity to these functional and structural alterations associated with brain injury, and have been applied to the study of CSD [11,15,28]. DOS and DOT utilize the near-infrared spectral region (600–1000 nm) to separate and quantify the multispectral absorption (μa) and reduced scattering coefficients (μs′), providing quantitative determination of several important biological chromophores such as deoxy-hemoglobin (HbR), oxy-hemoglobin (HbO2), water (H2O), and lipids. Concentrations of these chromophores represent the direct metrics of tissue function such as blood volume fraction, tissue oxygenation, and edema. Additionally, the scattering coefficient contains important structural information about the size and density of scatterers and can be used to assess tissue composition (exctracellular matrix proteins, cell nuclei, mitochondria) as well as follow the process of tissue remodeling (wound healing, cancer progression). DOS utilizes a limited number of source-detector positions, e.g., 1–2, but often employs broadband content in temporal and spectral domains . In contrast, DOT typically utilizes a limited number of optical wavelengths (e.g., 2–6) and a narrow temporal bandwidth, but forms higher resolution images of subsurface structures by sampling a large number of source-detector “views.” To achieve maximal spatial resolution, the ideal DOT design would employ thousands of source-detector pairs and wavelengths. However, several engineering considerations including measurement time and instrument complexity currently limit the practicality of this approach.
In this chapter we present the basic principles of a new, noncontact quantitative optical imaging technology, modulated imaging (MI) [30–32], and provide examples of MI performance in 2 rat models of brain injury, cortical spreading depression (CSD) and stroke. MI enables both DOS and DOT concepts with high spatial (<1 mm) and temporal resolution (<1 s) in a simple, scan-free platform. MI is capable of both separating and spatially-resolving optical absorption and scattering parameters, allowing wide-field quantitative mapping of tissue optical properties. While compatible with time-modulation methods, MI alternatively uses spatially modulated illumination for imaging of tissue constituents. Periodic illumination patterns of various spatial frequencies are projected over a large area of a sample. The diffusely reflected image is modified from the illumination pattern due to the turbidity of the sample. Typically, sine-wave illumination patterns are used. The demodulation of these spatially modulated waves characterizes the modulation transfer function (MTF) of the material, and embodies the sample optical property information.
The MI instrument platform was introduced originally by Cuccia et al.  Based on this design, we have developed a custom multispectral near-infrared (NIR) MI spectroscopy system capable of imaging between 650 and 1000 nm. A diagram of this system is shown in Figure 12.1.
Broadband NIR illumination is provided by an intensity-stabilized 250 W quartz-tungsten-halogen (QTH) lamp (Oriel QTH Source with Light Intensity Controller, Newport Corporation-Oriel Instruments, Stratford, Connecticut). Light is collimated and refocused with a pair of aspheric F/#0.7 optical lens systems (Oriel Aspherab). A custom-sized 3.5 in square hybrid hot mirror (Reynard Corporation, i.e., R00670-00) was placed between the lenses to limit the illumination to wavelengths below 1000 nm. Light engine optics taken from a digital projector (NEC HT1000) serve to homogenize and direct the light onto a 0.7 in digital micromirror device (DMD Discovery™ 1100 with ALP Accessory Package, ViALUX, Germany). Grayscale spatial sinusoid patterns are projected at 400 Hz using the ViALUX software development toolkit, which generates the necessary pulse-width modulation of binary sub-frames to produce a specified grayscale bit-depth (1–8 bits). Finally, a fixed focal length (f = 100 mm) projection lens illuminates the tissue at a slight angle from normal with a 15 × 25 mm illumination field. Detection was performed at normal incidence using a CRI Nuance™ camera system, which combines a 12-bit CCD camera and a liquid crystal tunable filter (LCTF; λ = 650–1100 nm, Δλ = 10 nm). To avoid specular reflection, crossed linear polarizers are used in the illumination and detection arms. For this system, the former is a 1.5 in diameter NIR linear polarizer (Meadowlark Optics, VLM-200-IR-R) placed immediately after the projection lens, and the first stage of the Nuance LCTF serves as the latter. The DMD, CCD, and LCTF are controlled via USB by a laptop computer, and synchronized using LabVIEW software (LabVIEW 8, National Instruments), enabling fast acquisition of a series of patterns with various spatial frequencies.
A detailed description of SFD measurement, calibration, and diffusion modeling is provided by Cuccia . In this work, we modeled diffuse reflectance using a transport-based White Monte Carlo (WMC) method [33,34]. Previously, we have found that compared with Monte Carlo, (1) diffusion predictions over- and underestimate low- and high-frequency diffuse reflectance, respectively, and (2) the quantitative accuracy of diffusion degrades with decreasing albedo . Due to the moderate albedo of brain tissue (μs′/μa ~ 10–20), we chose to analyze all brain data with the WMC approach. This homogeneous tissue model is a significant simplification of the multilayered rat brain, and more work is necessary to accurately model this complex system. We discuss further the consequences of our simple model in Section 12.2.5.
In this chapter, we use two inversion methods to calculate the absorption and reduced scattering from measurements of diffuse reflectance. When high measurement precision is desired, we use a “sweep” in spatial frequency space, producing an overdetermined set of diffuse reflectance measurements, which can be fitted to our WMC forward model predictions using least-squares minimization. This method is performed for all spatially averaged region analysis of optical properties and chromophores. When increased acquisition and/or processing speed is desired, we alternatively use a rapid two-frequency lookup table method based on cubic spline interpolation . This data can be achieved with a minimal 3-phase, single frequency image set (by demodulating and averaging the images to obtain AC and DC amplitude maps, respectively). On typical personal computers this approach is capable of millions of inverse lookup calculations per second, and is therefore used to calculate all high-resolution images including time sequences. The signal-to-noise ratio (and thus the measurement precision) of either approach is limited by the data sampling, with the two-frequency method having a lower precision with the tradeoff of higher acquisition and processing speed.
The quantitative absorption coefficient is assumed to be a linear (Beer’s law) summation of individual chromophore absorption contributions:
where ci and ?i(λ) represent chromophore concentrations and molar extinction coefficients, respectively. Using reported extinction coefficients of HbO2/HbR35 and H2O,36 we can invert Equation 12.1 and calculate tissue chromophore concentration separately at each pixel by linear least-squares fitting to the multispectral absorption images. Total hemoglobin (HbT) and oxygen saturation (StO2) can then be calculated as HbT = HbR + HbO2 and StO2= HbO 2/(HbR + HbO2) * 100, respectively.
On a pixel-by-pixel basis, diffuse reflectance versus spatial frequency is fitted to the WMC forward model to extract the local absorption and reduced scattering optical property contrast. This process is repeated for each wavelength, resulting in multi-spectral absorption and scattering spectra at each pixel. The measured contrast from discrete absorbers and scatterers on millimeter and submillimeter spatial scales, however, will possess partial volume effects in all three spatial dimensions. This is due to the physical light transport length scales in tissue, limiting the true x-y resolution of optical property contrast to many detector pixels . This phenomenon is not unique to MI, but present in all planar reflectance imaging measurements of turbid media. Absorption and scattering are calculated using a homogeneous reflectance model, extracting a locally averaged sampling of optical property contrast. Based on simulations of the tissue MTF for varying optical properties , we expect the resulting image resolution to scale directly with the transport length, l* = (μa + μs′)− 1, and the spatial frequency of illumination. In this chapter, we place quantitative emphasis on average optical properties and chromophores measured over a field of view that is greater than l*. Spatial maps and videos of these parameters are displayed and referred to as “contrast maps,” with the caveat that high resolution features will exhibit degraded quantitative accuracy.
MI spectroscopy measurements were performed on an in vivo Wistar rat model with a thinned-skull preparation. All procedures were performed in accordance with approved IACUC protocol guidelines. The animals were anesthetized, placed in a stereotaxic frame, their skulls thinned and glass coverslip applied. This preparation is described in detail by Masino et al.  The resulting thinned skulls allowed direct imaging of the cortex over a 5 × 7 mm field-of-view (whisker barrel cortex, centered at the C2 location). In order to investigate the sensitivity of MI toward studying acute cortical injury, we induced cortical spreading depression (CSD) by applying 1 M KCl solution to the surface of the cortex through a perforated section of skull and dura, located approximately 3 mm above the camera’s imaging field.
For each of three animals, our MI measurement protocol was twofold. Prior to CSD induction, baseline spatial modulation data were acquired at 6 spatial frequencies (3-phase projections each) from 0 to 0.26 mm−1, at 10 nm intervals over the entire range between 650 and 980 nm. Depending on the wavelength, image acquisition times ranged from 200 ms to 4 s, with total spectral imaging time of approximately 30 s per spatial pattern. The entire measurement (34 wavelengths, 3 phases, 6 frequencies) was repeated three times for statistical averaging yielding an entire measurement time of approximately 30 min.
Next, rapid dynamic measurements were performed, beginning 1 min prior to K+Cl− administration. Here, a significantly reduced data set was chosen in order to achieve high temporal resolution. Two spatial frequencies (0 and 0.26 mm−1) were acquired with three phase projection images, as described in Section 12.2.2, at each of four wavelengths (680, 730, 780, and 830 nm). The resulting 12 images took in total 6 s, permitting a repetition rate of 10 measurements per minute. The animals were followed for a period of 10 min for rats 1 and 2, and a period of 30 min for rat 3.
All images in this study were smoothed by 2D convolution with a Gaussian filter function (FWHM = 3 pixels), and baseline repetitions were averaged prior to data processing. Additionally, time-series data were post-processed by smoothing slightly in time (Gaussian FWHM of 2 timepoints = 12 s).
Because of the differential absorption sensitivity at low and high frequencies, optimal optical property separation is achieved when a large range of frequencies is used . In Figure 12.2a, we depict this differential sensitivity using diffuse reflectance (MTF) predictions versus frequency, increasing μa by 100% from 0.02 (black line) to 0.04mm−1 (gray line). This is done for two values of μs′, 0.6 (solid lines) to 1.2mm−1 (dashed lines), simulating a 100% change in scattering. Notice that the low frequencies have a significant reflectance change due to absorption, while high frequency reflectance remains nearly unchanged. Conversely, reflectance changes due to scattering are observed at all spatial frequencies. In Figure 12.2b, we further visualize this by plotting the reflectance sensitivity to 1% changes in absorption and scattering. Whereas DC reflectance is equivalently sensitive to a fractional change in either absorption or scattering, at high spatial frequencies absorption contrast is lost while scattering contrast is retained. For instance, notice that at our maximum measurement frequency of 0.26 mm−1 the reflectance is roughly 24 times more sensitive to scattering compared to absorption (ΔRd = 0.56 μs′ versus 0.024 * 10−3 for μa). This plays an important role in Section 12.3.2 during our discussion of dynamic scattering measurement.
In realistic heterogeneous tissues, a tradeoff exists between maximizing the frequency range for optical property accuracy and obtaining similar sampling volumes. As tissue is a low-pass spatial filter, high frequencies are attenuated quickly with depth. Using diffusion-based forward modeling, we have estimated mean sampling depths at 650 nm using measured average background optical properties of brain tissue. This was done by predicting the depth sensitivity to contrast from a planar perturbation in absorption, given a background fluence profile from spatial frequencies 0 and 0.26 mm−1. Based on these results, we observe qualitatively similar depth sampling, with mean depth sampling ranging between 2.5 mm and 1.2 mm (for fx = 0 and 0.26 mm−1, respectively). In all cases maximal sensitivity was found in the first 1–2 mm, where cortical hemodynamic changes occur.
In Figure 12.3a we show a grayscale planar reflectance image of the cortical region of rat 1 at 650 nm. A dotted-line box denotes the region-of-interest (ROI) used for analysis, selected for its uniform illumination and the absence of cerebral bruising. The Monte Carlo-model fitting of spatial frequency data allows calculation of the absorption and reduced scattering coefficients. In Figure 12.3b we show the spatially averaged diffuse reflectance at 650 nm and the corresponding multi-frequency fit. Excellent agreement is observed between measurement data and the model-based fit, with derived μa and μs′ coefficients of 0.033 and 0.70 mm−1, respectively.
Analysis of multifrequency reflectance data separately at each pixel results in spatial maps of absorption and reduced scattering contrast. In Figure 12.3c, we plot the μa and μs′ maps recovered at 650 nm for rat 1. Note the strong absorption in the vein region, due to a large absorption by HbR at this wavelength. Below the images, we show histogram distributions of the corresponding quantitative maps above, indicating the degree of spatial variation in recovered optical properties. The mean and standard deviation for the pixel-wise μa and μs ′ were 0.030 ± 0.007 mm−1 and 0.63 ± 0.13 mm−1, respectively. These statistical results are in good agreement with the spatially averaged reflectance fit from Figure 12.3b, suggesting that our simple pixel-wise fitting approach yields optical properties similar to that calculated using a global analysis.
By mapping the absorption coefficient at multiple wavelengths, we can perform quantitative spectral imaging of tissue. In Figure 12.4, we summarize the baseline spectroscopy results for all three animals. In Figure 12.4a we show the μa (left) and μs′ (right) coefficients versus wavelength (circles) recovered from spatially averaged fitting. Data for rat 1 is shown in black (rat 2 in dark gray; rat 3 in light gray). Note the distinct spectral features in absorption, resulting from oxy- and deoxy-hemoglobin (HbO2, HbR), and water (H2O) absorption. The calculated scattering coefficient generally decays with increasing wavelength, and the results from a power law (μs ′ = A·λ(nm) −b, solid lines) fit are shown. A small residual coupling is observed between measured scattering and absorption spectral features. In particular, the scattering at the shortest and longest wavelengths appears to be underestimated by 5–10%, occurring where the corresponding absorption is highest (due to HbR and H2O, absorption features, respectively). Based on our experiments in layered tissue phantoms , we believe this effect is primarily due to frequency-dependent probing volumes in the presence of depth-heterogeneous structures.
Simultaneous linear fitting of the absorption to known extinction coefficients yields measures of chromophore concentration. Shown in Figure 12.4a, multispectral fitting (solid line) for rat 1 yields HbO2, HbR, H2O, HbT and StO2 values of 56.3 μM, 33.2 μM, 63.9%, 89.6 μM, and 56.3%, respectively. Tabulated results of chromophore values for all three animals are shown in Figure 12.4b. Lipid absorption near 930 nm was not apparent in the μa spectrum, and when included in the spectral analysis was not found to significantly affect the results. The small absorption “bump” at 900–910 nm is an artifact of imperfect phantom calibration due to the presence of a sharp, strong silicone absorption peak that is present in the phantom.
We note that the solution for chromophore concentration is well-determined when the number of wavelengths is at least equal to the number of chromophores. Therefore, as few as two wavelengths can be used to separate HbO2 and HbR (if a constant value of H2O is assumed). Repeating the above analysis with 780 and 830 nm only (assuming H2O = 65%) yields results for HbO2 and HbR within 10% of those from full spectral fitting. Repeating the above analyses using a simple diffusion-based model provided qualitatively similar results for absorption and scattering spectra, but in general was found to overestimate the absorption coefficient by 10–25%.
Absorption spectra at each pixel can be separately analyzed to yield spatial maps of local HbO2, HbR, and H2O distribution, shown in Figure 12.5. Notice the high concentration of HbR over the large superficial draining vessel (venous) regions, also reflected in the StO2 image, highlighting the effect of tissue oxygen extraction. Conversely, notice that the high albedo regions with less structural detail are highly oxygenated, with StO2 levels between 60 and 70%. Lastly, the H2O map reveals a relatively homogeneous distribution of water.
We performed measurements of CSD in each of the three rats, as described in Section 12.2.3. The results are presented as follows. We first present data for a single animal, choosing rat 3 for its long observation period of 30 minutes. Three ROIs are selected for analysis, and baseline MI spectroscopy results are reported for each of these regions. Next, the observed dynamic time courses of diffuse reflectance, optical properties, and chromophore concentrations are shown for each ROI. We then present the full spatio-temporal dynamic contrast data for rat 3 (2D + time) in the form of “snapshot” images.
Figure 12.6 summarizes the baseline spectroscopy measurements for rat 3. In Figure 12.6a, we show three regions of interest superimposed on the DC reflectance map, chosen to highlight three different characteristic temporal profiles observed within the field of view. In Figure 12.6b we show the baseline spectral fits for each of these regions, and in Figure 12.6c we tabulate the resulting calculated chromophore concentrations. In general, Region A (black) is a high albedo region lacking any large blood vessels, whereas Regions B (dark gray) and C (light gray) include high-absorption blood vessels and mild cerebral bruising from surgery. These differences are apparent in their recovered absorption spectra and fits, with on average 27% higher HbT, and 32% lower saturation in the vascular regions. Also, 7% higher H2O is found in Regions B and C, which may indicate increased edema due to bruising.
In Figures 12.7–12.9 (for regions A–C, respectively), we present the temporal dynamics of CSD in each ROI of rat 3 as measured by MI. In part (a) of each figure, we plot the multispectral diffuse reflectance changes at fx = 0 mm−1 (DC, top) and fx = 0.26 mm−1 (AC, bottom). In part (b), we plot the recovered Δμa (top) and Δμs′ (bottom) optical properties at each wavelength. While absolute values of diffuse reflectance and optical properties are measured separately at each time point, for visualization purposes all data are displayed as a change from that prior to KCl administration. Absolute optical property values at t = 0 (not shown) demonstrate excellent agreement (~5–10%) with full multifrequency baseline data.
Looking first at the reflectance time courses of Figure 12.7a (Region A), we see in general a series of three CSD events over the 30 minutes, with each transient event occurring for approximately 4.3 minutes. The first event occurs at minute 2.9 after KCl application, indicating an initial latency between the insult and the first resulting spreading depression wave. Reflectance contrast is present in both DC and AC frequency components, but with markedly different signatures. Generally, the DC time course shows a slow, gradual decay, punctuated by sharp, wavelength-dependent spikes/dips (for short/long wavelengths, respectively). Alternatively, the AC signature contains three sets of transient dips consistent across all wavelengths, with final values leveling off progressively lower than baseline. Discussed in detail in the following paragraph, we believe these AC changes are due primarily a result of optical scattering and may be related to neuronal depolarization. The corresponding derived optical properties in Figure 12.7b reflects this, with μs′ trends tracking directly with the measured AC reflectance. As expected, μa trends reveal similar wavelength-dependence of the DC reflectance (with opposite polarity), reflecting changes in HbO2 and HbR.
In Section 184.108.40.206 we noted that the diffuse reflectance at fx = 0.26 mm−1 is 23 times more sensitive to scattering changes compared to absorption. In this context, we propose that the observed magnitude of the CSD-induced AC reflectance changes can only be explained by changes in optical scattering. To concretely illustrate this point, we pick as an example the observed 780 nm AC diffuse reflectance dip in Figure 12.7a at t = 3.7 min of -0.003. Here, the corresponding change in reduced scattering in Figure 12.7b, Δμs′, is calculated to be −0.03 mm−1. In order for this change to instead be due to an absorption-only event, μa would need to increase by 121% from baseline (from 0.038 to 0.084 mm−1). This increase would also need to be accompanied by a drop in Rd (fx = 0 mm−1) of 0.12 (33%), whereas the actual observed DC reflectance only drops by 0.008 (<1%) and thus cannot explain the change. Secondly, we note that the three sets of AC reflectance dips occur consistently across all four wavelengths. While an approximate 120% increase in HbT could induce this decrease at high frequency, it would also require a large broad-wavelength decrease in the DC reflectance. We instead observe during these events that the DC increases at short wavelengths while the DC decreases at long wavelengths, suggesting primarily an exchange between HbO2 and HbR volume fractions, as opposed to a dramatic HbT change.
Regions A–C (Figures 12.7–12.9) were chosen to highlight three different time signatures observed in the field of view during the CSD dynamics. The most contrasting feature between all three regions is the measured AC reflectance and the derived scattering coefficient. In Region B (Figure 12.8), each CSD event appears to cause a biphasic scattering change, with a sharp increase and then decrease, whereas a monophasic dip was observed in Region A (Figure 12.7). Region C (Figure 12.9) appears even more complex with a triphasic rise-dip-rise temporal profile. We observe that Regions A to C are located with increasing proximity to the CSD induction point (3 mm above the imaging field).
Because fractional changes in scattering and absorption have an equal (and opposite) effect on DC reflectance (see Section 220.127.116.11), any scattering (i.e., pathlength) changes measured here could be misinterpreted as absorption events with traditional ISOI analyses (i.e., DC reflectance only). In our observations, the measured scattering change of up to −0.05 mm−1 would be interpreted as an increase in absorption of up to +0.005 mm−1, more than the maximum measured absorption change for wavelengths 730, 780, or 830 nm in any of the three regions. In order to account for differential pathlength changes, Kohl et al. proposed a multispectral model , which they used to differentiate dynamic scattering and absorption changes using ISOI. This approach improves ISOI accuracy, and has been generally adopted as the method of choice for quantitative functional imaging. For dynamic measurements, we see MI as an improvement over this approach as it alternatively uses frequency domain measurements at a single wavelength to derive absolute scattering and absorption coefficients. This potentially provides a simplified single-wavelength measurement apparatus for detection of scattering, and also avoids potential mis-estimation of background optical properties.
Light scattering changes induced by spreading depression have been reported previously, and a comprehensive review is provided by Somjen. With in vivo spatially resolved reflectance measurements, Kohl et al.  separated absorption from scattering and observed a biphasic scattering response similar to that of Region A. With simultaneous laser scattering and electrophysiological measurements, both Jarvis et al. and Tao et al. found a strong correlation between electrical and optical scattering changes [12,13,40]. Tao et al. noted spatial heterogeneity in the dynamic spreading depression (SD) waveform related to the proximity to the SD induction site, similar to our results.
Using linear spectral analysis of absorption at all four wavelengths, we calculated the time-dependent chromophore concentration for Regions A, B, and C, presented in Figure 12.10A,B,C, respectively. In each region, the calculated baseline concentrations of H2O were assumed to be constant. All three regions exhibit remarkably similar trends in HbR, HbO2, HbT, and StO2. This similarity is not clear in the DC traces of Figures 12.7–12.9, further highlighting the benefit of accurate separation of μa and μs′. Focusing on the first CSD event, there is a very consistent signature of: (1) a 2-minute latency post-KCl administration, (2) a 30-second period of decreasing StO2 (3) a dramatic spike in both StO2 (3–10%) and HbT (2–4 μM) with rise and decay times of approximately 1 minute each. For each region, the final StO2 is approximately 5–10% lower than baseline, while the HbT restores to baseline values. This process repeats again twice more, except that the phase (2) desaturation appears to be absent. Additionally, in the “vessel” Region 3, we observe a gradual increase in HbT over the 30 minutes, indicating chronic blood pooling.
We show in Figure 12.11 the spatio-temporal evolution of both chromophore concentration and scattering changes from the first SD wave in rat 3. These are depicted in the form of a time derivative, i.e., (C(tn + 1) − C(tn))/(tn + 1 − tn), where C represents concentration/saturation/scattering values and tn represents time of acquisition for data point n. This visualization is appealing as it highlights the changes with high contrast . From left to right, we show HbO2, HbR, HbT, StO2, and μs′. Notice the wave in scattering which propagates from top right to bottom left, at a rate of approximately 3 mm/min. An increase, or “spike” in scattering is observed initially in the top right hand corner, in close proximity to the location of KCl administration. Note the large spikes in HbT and StO2 due to vascular activity from depression wave propagation through the measurement field. We observe a transient increase in saturation and blood volume. Over the longer time periods, however, we observe a slow, sustained trend toward hypoxia in the vein regions.
The spatio-temporal evolution of the scattering coefficient in Figure 12.11 reveals a spatially defined scattering wave (reduction in μs′) that precedes hemodynamic changes. The scattering drop is presumed to be a consequence of neuronal depolarization accompanying CSD. This observed wave pattern has been shown previously with reflectance ISOI and attributed to blood volume changes . Interestingly, the scattering depolarization wave is clearly followed in space and time by the increase in deoxyhemoglobin (HbR), decrease in saturation (StO2), and drop in oxyhemoglobin (HbO2); changes that are consistent with depolarization-induced neural tissue oxygen consumption.
In order to assess the sensitivity of MI to stroke, we conducted preliminary studies in a rat middle cerebral artery occlusion (MCAo) model, the most commonly involved artery in ischemic strokes. The left MCA was surgically cauterized using monopolar cautery or ligated to produce a permanent stroke. Figure 12.12 shows pre-versus post-MCAo results for a representative animal. Data were acquired at 5 wavelengt
Sue Hale, Physical Therapist. Hand therapist in Melbourne FL. Sue speaks of her experiences using the Microlight ML830® Cold Laser in her practice and personally. Sue uses the ML830® Laser daily for treating all types of injuries; from hand injuries, back pain and injuries, runner's injuries, knee injuries, and many other conditions.
video length: (1:17)
Django's 1st Low Level Laser Therapy (LLLT)
Django is now 13 months old and has had clicking and obvious pain in his right knee for the last 6 months. After a negative valley fever test & 2 knee x-rays showing no fracture or tumors,and being told by 2 vets, it's possibly a soft tissue (CCL/ACL) injury and that he needed a CT, MRI or arthroscopy to get a more definitive diagnosis. Django's owner decided to try LLLT to alleviate some of his pain.
video length: (0:32)
Django Post MPL Surgery Days 1-2
Django had surgery (medial imbrication, with a lateral release & anti-rotational sutures) on his left knee to correct a grade 2 medial patellar luxation August 20, 2015-. Here he is the day of surgery from check in through 48 hours later getting cold laser therapy.
video length: (1:27)
Here's Django approximately 9 weeks post-op, fully enjoying the newfound use of his leg.
video length: (1:41)
More videos of Django's journey can be found on his owner's Youtube channel, link below.
The young man featured in this video had a history of cognitive issues, including difficulty with reading and comprehension.
video length: (1:54)
This video covers the operation of the LZ30 family of lasers in Basic Mode.
video length: (3:47)
This video covers the Advanced Features of the LZ30 family of lasers.
video length: (6:17)
This video covers Safety and Regulatory Considerations for the use of the LZ30 family of lasers.
video length: (3:49)
Penetration depth test between a high powered class 4 970nm laser and a low powered class 3b LLLT 810nm laser, by shining through hand.
Result: 810nm laser passes through tissue better than 970nm Laser
video length: (3:09)
Demonstration of LLLT on a dog with inflammation, and pain in his paw.
video length: (3:36)
This study investigates the influence of gallium–arsenide (GaAs) laser photobiostimulation applied with different energy densities on skin wound healing by secondary intention in rats. Three circular wounds, 10 mm in diameter, were made on the dorsolateral region of 21 Wistar rats weighting 282.12 ± 36.08 g. The animals were equally randomized into three groups: Group SAL, saline solution 0.9%; Group L3, laser GaAs 3 J/cm2; Group L30, laser GaAs 30 J/cm2. Analyses of cells, blood vessels, collagen and elastic fibres, glycosaminoglycans and wound contraction were performed on the scar tissue from different wounds every 7 days for 21 days. On day 7, 14 and 21, L3 and L30 showed higher collagen and glycosaminoglycan levels compared to SAL (P < 0.05). At day 21, elastic fibres were predominant in L3 and L30 compared to SAL (P < 0.05). Type-III collagen fibres were predominant at day 7 in both groups. There was gradual reduction in these fibres and accumulation of type-I collagen over time, especially in L3 and L30 compared with SAL. Elevated density of blood vessels was seen in L30 on days 7 and 14 compared to the other groups (P < 0.05). On these same days, there was higher tissue cellularity in L3 compared with SAL (P < 0.05). The progression of wound closure during all time points investigated was higher in the L30 group (P < 0.05). Both energy densities investigated increased the tissue cellularity, vascular density, collagen and elastic fibres, and glycosaminoglycan synthesis, with the greater benefits for wound closure being found at the density of 30 J/cm2.
Laser photobiostimulation has been used as a non-invasive alternative to treat muscle injuries and skin wounds, and to control inflammatory processes and pain (Enwemeka et al. 2004; Reddy 2004). Although the use of laser light to accelerate the healing process was documented in the literature for the first time in 1971 (Mester et al. 1971; Shields & O'Kane 1994), and the efficacy of this therapeutic modality is proven, parameters about how it is used are still controversial (Tuner & Hode 1998; Moore et al. 2005). Parameters such as the type and source of laser light emission, number of applications, duration of treatment and mechanisms of action through which the laser light exerts its effects remain the focus of investigation in the ongoing search for efficient methodologies that justify and encourage the use of laser light in clinical practice. Several mechanisms have been proposed to explain the effects of laser light on biological tissues, including the absorption of light by the enzymes of the electron transport chain in the inner mitochondrial membrane, stimulation of the production of oxygen, and cell proliferation induced by photoactivation of the calcium channels (Shields & O'Kane 1994; Breitbart et al. 1996). Recent studies show that the main cells stimulated by laser light are macrophages and fibroblasts (Gonçalves et al. 2010a; Xavier et al. 2010). Macrophages are important cells responsible for releasing growth factors that stimulate proliferation, differentiation and synthesis of extracellular matrix components (Shields & O'Kane 1994; Reddy 2004; Gonçalves et al. 2010b). In in vitro experimental models examination of a wide range of wavelengths showed that wavelengths between 524 nm and 904 nm were related to decreased time of wound healing by stimulating fibroblast and keratinocyte differentiation, collagen production and skin neovascularization (Pogrel et al. 1997; Demidova-Rice et al. 2007).
Previous studies have shown that the gallium–arsenide laser (GaAs λ 660 nm) is able to stimulate skin wound healing in humans and laboratory animals with energy densities between 1 and 4 J/cm2 (Medrado et al. 2003; Pugliese et al. 2003; Reddy 2004). However, most of the work is restricted to investigating the effect of energy densities below 4 J/cm2, and reports on the effects of high energy densities in tissue repair are scarce and inconclusive. Thus, this study was designed to investigate the influence of laser photobiostimulation applied with different energy densities in a rat model of skin wound healing by secondary intention.
Twenty-one male Wistar rats (Rattus norvegicus), 10 week old and weighing 282.12 ± 36.08 g, obtained from the Biological Sciences Center, Federal University of Viçosa, Minas Gerais, Brazil, were used in this study. During the experiment, the animals were allocated to individual cages that were cleaned daily and maintained in an environment with controlled temperature (22 ± 2 °C), light (12 h light/dark cycles) and humidity (60–70%).
The experiment was conducted in accordance with International Ethical Standards for the Care and Use of Laboratory Animals and approved by the Ethics Committee for the Care and Use of Laboratory Animals of the Federal University of Viçosa (UFV; registration 005/2008).
Before the surgical wounds were made, the animals were anaesthetized using intramuscular ketamine (50 mg/kg) and xylazine (20 mg/kg). Then, trichotomy was performed on the dorsolateral region of the animals, and the area was defatted using ethyl ether (Merck®, Rio de Janeiro, Brazil) followed by the use of 70% ethanol and 10% povidone–iodine for anti-sepsis (Johnson Diversey®, Rio de Janeiro, Brazil). Three circular secondary intention wounds 10 mm in diameter were made in the dorsolateral region of the animals by removing the skin with a scalpel until the exposure of the muscle fascia. The standardized wound area was marked with a dermographic pencil and checked using an analogical pachymeter (Kingtools®, São Paulo, Brazil) (Gonçalves et al. 2013). After completion of the wounds, the animals were randomly divided into three groups with seven animals in each. Group saline (SAL, control): saline solution 0.9%; Group L3: GaAs laser (λ 660 nm, 3 J/cm2); Group L30: GaAs laser (λ 660 nm, 30 J/cm2). The laser device (Endophoton®, KLD, São Paulo, Brazil), which was previously calibrated by the manufacturer, presented an output of 20 mW, power density of 25.47 mW/cm2, visible radiation and a 0.79 cm2 circular beam. Laser light was applied transcutaneously at six equidistant points around the wound margin. The wounds were irradiated for 118.5 s in L3 to release 3 J/cm2 and 1185 s in L30 to release 30 J/cm2. The wounds were cleaned daily with 0.9% saline solution immediately before the laser application. The treatments were started immediately after the wound was made once a day for 21 days corresponding to the experiment duration.
The progress of wound closure was evaluated by measuring the wound area every 7 days in digitized images with the dimensions of 320 × 240 pixels (24 bits/pixel) obtained using a digital video camera (W320, Sony, Tokyo, Japan). The wound areas were calculated by computerized planimetry using the Image Pro-Plus image analysis software program, version 4.5, (Media Cybernetics®, Silver Spring, MA, USA), previously calibrated. Wound contraction index (WCI) was calculated using the following ratio: initial area of the wound (Ao) − area on the day of measurement (Ai)/initial area of the wound (Ao) × 100 (Gonçalves et al. 2013). The third wound was selected for this analysis because the tissue from this wound was collected on the final day of the experiment (21st).
For each group, 35 histological sections 8 μm thick stained with Fast green and Sirius red were used to quantify the levels of collagen and total protein in scar tissue using a previously described spectrophotometric method (López-De León & Rojkind 1985). In this method, the maximal absorbance to the Sirius red (540 nm) and Fast green (605 nm) dyes, correspond to the amount of collagen and non-collagen proteins respectively. For each section used in the collagen analysis, a corresponding serial section was obtained, which was used in the analysis of glycosaminoglycans. The tissue content of glycosaminoglycans was determined according to a modified procedure described by Corne et al. (1974). Sections were transferred immediately to 10 ml of 0.1% (w/v) Alcian blue 8GX solution (0.16 M sucrose solution buffered with 0.05 ml sodium acetate at pH 5). After successive rinses in 10 ml of 0.25 M sucrose solution, dye adhered to the tissue was extracted with 10 ml of 0.5 M magnesium chloride, and the absorbance of the resultant solution was analysed in a spectrophotometer at 580 nm.
Tissue fragments were collected from the different wounds every 7 days. Each fragment contained tissue removed from the centre of the wound and part of the uninjured adjacent tissue that had not received laser radiation. The fragments were put into Karnovsky's solution for 24 h and processed for paraffin embedding. Semiserial 4-μm-thick vertical uniform random (VUR) sections were obtained using a rotating microtome (Leica Multicut 2045®, Reichert-Jung Products, Jena, Germany). One of every 20 sections was used to avoid repeating analysis of the same histological area. Sections mounted on histology slides were stained with haematoxylin and eosin for visualization of cells and blood vessels (Karu 2003), Verhoeff's method for elastic fibres (Verhoeff 1908) and Sirius red dye (Sirius red F3B, Mobay Chemical Co., Union, NJ, USA) for marking collagen fibres observed under polarizing microscopy (Junqueira et al. 1979). Analysis of collagen was based on the birefringence properties of the collagen fibres, because under polarization, the thick collagen fibres (type I) appear in shades of bright colour ranging from red to yellow, whereas thin reticular fibres (type III) are shown in bright green (Gonçalves et al. 2010a).
The slides were visualized, and the images captured using a BX-60® light microscope (Olympus, São Paulo, Brazil) connected with a digital camera (QColor-3®, Olympus, São Paulo, Brazil). For each wound and staining method, 10 histological sections were analysed. For each section, five images were obtained randomly with a 20× objective lens, and the cells and blood vessels were quantified in the histological area. Under each image was applied an unbiased two-dimensional test area (At) of 69 × 103 μm2 at tissue level, so that the total histological area investigated was 24 × 106 μm2. The proportion of the histological area occupied by type-I and type-III collagen fibres was determined using the Quantum® software program (Department of Soil Science, Federal University of Viçosa, Viçosa, Brazil) (Gonçalves et al. 2010a).
The volume density of cells (Vv [cells], %), blood vessels (Vv [bvs], %) and elastic fibres (Vv [elf], %) was estimated as:
where ΣPp [cells; bvs; elf] denotes the total number of points on the cells, blood vessels or elastic fibres, and ΣPt is the total points of the test system (ΣPt = 200).
The length density of blood vessels (Lv [bvs], mm/mm3) and elastic fibres (Lv [elf], mm/mm3) was estimated as:
where ΣQ−[bvs] denotes the total number of blood vessel or elastic fibre profiles counted in the At, and ΣP [tissue] is the total number of points on the tissue (Brüel et al., 2005).
The surface area density of blood vessels (Sv [bvs], mm2/mm3) was estimated as:
where ΣI [bvs] denotes the total number of intersections between the cycloid arcs (here 44) and the blood vessel surface area, and l is the length of the cycloid arcs. The Image Pro-Plus 4.5® image analysis software (Media Cybernetics) was used in the stereological analysis.
The data were expressed as mean and standard deviation (mean ± SD). The normalcy of the data distribution was verified using the Shapiro–Wilk test. All variables investigated were subjected to the Kruskal–Wallis test for multiple comparisons. Statistical significance was established at P < 0.05. The analysis was performed using the software Sigma Stat 3.0® (Systat Software Inc., Chicago, IL, USA).
There were no significant differences in total collagen and glycosaminoglycan content in the uninjured tissues from the different groups (Table 1). At all investigated time points, the groups exposed to laser photobiostimulation had higher collagen content in the scar tissue compared with SAL (P < 0.05). At day 7, the content of glycosaminoglycans was higher in both groups exposed to laser irradiation in relation to SAL group. A similar result was observed at day 14, but only the group L30 was significantly different compared with SAL. At the end of the experiment, the content of glycosaminoglycans was significantly higher in L3 compared with the other groups.
The analysis of collagen fibres in the uninjured tissue showed no difference in the proportion of type-I and type-III fibres between the groups. On days 14 and 21, the groups receiving laser irradiation had higher proportion of type-I collagen fibres compared with SAL, with the best results in L30 (P < 0.05). At day 21, this variable was similar in L3 and L30. Animals in L3 and L30 had a higher proportion of type-III fibres compared with SAL on days 7 and 14, with the best results in L3 (P < 0.05). At day 21, the content of type-III fibres was similar in all groups (Figure 1).
The analysis of elastic fibres in the uninjured tissue showed no difference in the proportion of volume and length of elastic fibres between the groups. On day 21, the groups receiving laser irradiation had a higher proportion of volume (Vv) and length (Lv) of elastic fibres (elf) compared with SAL (P < 0.05) (Figure 2).
The extent of scar tissue occupied by blood vessels is shown in Table 2. There were no significant differences in volume, length or surface densities of blood vessels in the unharmed tissues (day 0). At day 7, all these parameters were significantly higher in both groups that received laser light compared with SAL, with better results in L30 (P < 0.05). On days 14 and 21, similar results were observed in L30 compared with other groups (P < 0.05).
The results of tissue cellularity are shown in Table 3. The unharmed tissue presented similar cellularity in all groups. On days 7 and 21, the groups L3 and L30 had higher cellularity in the granulation tissue compared with SAL (P < 0.05). At day 14, there was a higher volume density of cells in L3 compared with the other groups (P < 0.05).
Figure 3 colour shows photomicrographs of skin histological sections collected in both groups investigated. The uninjured skin showed similar cellularity and blood vessel density in all groups. On days 7, 14 and 21, there was increased cell distribution in all groups, with higher cellularity in L3 and L30 compared with the SAL (Figure 3 and Table 2). On days 7 and 14, increased density of blood vessels was observed mainly in the group L30 compared with the other groups. At day 21, there was a higher density of cells and blood vessels in both groups that received laser light compared with SAL.
At all times investigated, the group L30 showed a significant reduction in the wound area compared with other groups (P < 0.05). At day 7, the rate of wound closure was higher in the groups receiving laser irradiation compared with SAL (P < 0.05). A high rate of wound closure was identified in SAL at the end of the experiment (day 21). Total closure of the wound was achieved in L30 by day 21, a feature not found in the other groups (Table 4 and Figure 4).
The present study investigated the effect of different energy densities of the GaAs laser on skin wound healing. Using design-based stereology and spectrophotometric methods, the results indicated that the laser photobiostimulation was able to modify the morphology of the scar tissue in a time-dependent way leading to more efficient healing.
It is widely recognized that for healing to occur properly, synthesis of extracellular matrix is required, especially collagen, a protein that provides structural support for cell proliferation and neoangiogenesis (Liu et al. 2008; Gonçalves et al. 2010a,b2010b). The results of this study showed that both groups that received laser irradiation had a higher total collagen content at all time points analysed. These findings corroborate the results found by Medrado et al. (2003) and Gonçalves et al. (2010a,b2010b), which observed a significant increase in the collagen content in scar tissue 7 days after laser irradiation of skin wounds in rats. Collagen synthesis is an event directly related to the biomechanical properties of the scar tissue. In this context, the greatest collagen content gives the scar tissue greater resistance to mechanical stresses, a characteristic essential to the maintenance of tissue integrity and to reduced susceptibility to further injury (Karu 2003; Gonçalves et al. 2010a,b2010b).
Considering the different collagen types, both irradiated groups had a higher proportion of type-I and type-III collagen fibres than the control group. Both energy densities investigated were effective in stimulating the maturation of collagen in scar tissue, and the best results were found in group L30. Although laser irradiation has influenced the total levels of collagen, it is essential to identify the types of collagen produced in scar tissue. Traditionally, the assessment of type-I and type-III fibrillar collagens has provided an important indicator of the progression of the healing process (Karu 2003; Gonçalves et al. 2010a,b2010b). In the earlier stages of cutaneous wound healing the synthesis of type-III collagen predominates and is then gradually replaced by type-I collagen fibres, thicker, resilient and the type of collagen that predominate in normal tissue (unharmed). Thus, determining the proportion of type-I collagen fibres in relation to type-III fibres allows us to evaluate the level of remodelling and maturation of scar tissue, which in turn indicates how much this tissue approximates to the tissue when it is unharmed (Reddy 2004; Mendez et al. 2004; Gonçalves et al. 2010b). Considering these characteristics, it is widely recognized that therapeutic approaches that stimulate the synthesis of type-I collagen, leading to increased collagen maturation, are potentially useful strategies in the treatment of skin injuries (Medrado et al. 2003; Pugliese et al. 2003; Gonçalves et al. 2010a,b2010b).
An additional result shown in the present study was the influence of the laser photobiostimulation on the glycosaminoglycan content in irradiated tissue. This finding indicates a transient modification of some structural polysaccharides of the extracellular matrix during the healing of skin wounds. It is believed that this event is possibly related to the development of a structural and functional support able to stimulate the cell migration and differentiation (Pierce et al. 1991; Hodde 2002; Lai et al. 2006). It is known that the content and distribution of polysaccharides molecules are important to the hydration (attraction of water molecules – solvation water) and nutrition of the granulation tissue during the development of a vascular network that would allow the progression of tissue repair (Pierce et al. 1991; Hodde 2002; Lai et al. 2006). Although the quantity and quality of non-protein and protein components of the stromal tissue are important in tissue repair, currently there is not sufficient evidence as to how the laser irradiation modulates the synthesis and secretion of polysaccharide molecules to stimulate the healing of skin wounds. As the analysis of these molecules performed in this study is not as sensitive and specific as some molecular biology techniques, we cannot yet establish how much the induction of synthesis of polysaccharides contributes to the mechanism through which the laser photobiostimulation improves the healing process. Thus, further studies are needed in this area.
In addition to the increased collagen and glycosaminoglycan content, the laser-treated groups also had a higher tissue area occupied by capillaries, with the best results in the group that received the highest energy density. Furthermore, this study confirmed previous findings (Moore et al. 2005; Houreld et al. 2010) that the laser radiation, in both low and high doses, stimulates the tissue cellularity and increases the synthesis of granulation tissue, which are aspects involved in tissue repair. These data are similar to those described by Corazza et al. (2007) and Gonçalves et al. (2010a). These authors showed the efficiency of high-energy dosages in the induction of fibroblast proliferation and neoangiogenesis. However, these findings are in contrast to previous studies that show better results in these variables with the use of low doses of energy, especially 1–4 J/cm2 (Tuner & Hode 1998; Medrado et al. 2003; Reddy 2004). A complex mechanism has been described through which the laser light stimulates the tissue repair. Studies with models of soft-tissue injuries have provided evidence that the photobiostimulation laser induces the synthesis and secretion of mitogens (Posten et al. 2005; Houreld et al. 2010; Xavier et al. 2010) such as vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF) and tumour necrosis factor alpha (TNF-α) by macrophages, neutrophils, endothelial cells and fibroblasts, which stimulate the reorganization and repair of damaged tissue through the induction of proliferation, cell differentiation and neoangiogenesis (Posten et al. 2005; Houreld et al. 2010; Xavier et al. 2010).
There is sufficient evidence that the synthesis and differentiation of parenchymal and stromal components of the tissue determine the progression of the reorganization of injured tissue and the quality of the neoformed tissue (Karu 2003; Posten et al. 2005; Corazza et al. 2007; Liu et al. 2008). Thus, therapeutic interventions that stimulate the production of cellular and molecular components of the granulation tissue have been effective in promoting faster closure of wounds in soft tissues (Gonçalves et al. 2010a,b2010b; Xavier et al. 2010). In the present study, the group that received a higher dose of laser radiation (L30) showed more rapid progression of wound closure compared with other groups. These data are similar to those found by Enwemeka et al. (2004) and Moore et al. (2005), which showed the influence of various parameters of laser photobiostimulation on the tissue repair, including reduction in the wound area mainly with moderate energy densities between 19 and 24 J/cm2. In contrast, in these same studies, densities below 8.25 J/cm2 did not improve the injuries' closing time, findings that are contrary to the results of Medrado et al. (2003), Pugliese et al. (2003) and Mendez et al. (2004) that demonstrated a higher closing speed of the injured tissue at low energy densities (2–4 J/cm2), while high doses led to a delay in tissue recovery.
The findings of the present study suggest that laser photobiostimulation can modulate the process of skin wound healing in a time-dependent way. The higher energy density investigated was more effective in modifying the morphology of the parenchyma and stroma of the scar tissue and led to a faster healing. Considering the findings of this study in relation to the contradictory results of previous investigations, it is evident that additional studies are required to investigate the effects of photobiostimulation lasers with different energy densities on biological tissues, especially in relation to ultrastructural and metabolic changes of injured tissues.
Demonstration of LLLT being used on a patient with a sports injury.
video length: (6:57)
This video gives a short description of how LLLT works, and shows the most basic procedure for using a laser for therapy. It also explains some of the differences between differnt types of lasers, and talks about penetratin depth.
video length: (2:22)
This is a compilation of news reports from the early 2000's about LLLT being used in a Canadian clinic primarily for atheletes.
video length: (13:47)
Therapist discuses the basics of LLLT while giving a basic demonstration of LLLT.
video length: (1:56)
LLLT used on sports injury.
video length: (0:41)
Dentist describes his experience using LLLT on patients.
video length: (3:57)
2 IA steroid injections + Hyaluronic Acid, systemic NSAID's, and chiropractic were administered, and joint support supplement was added to Rio's feed - with no improvement noted. Laser therapy was tried as a last resort prior to euthanasia. 15,000 total Joules were applied to Rio's left elbow and left shoulder. Over a year and a half later Rio is going strong
video length: (2:55)
Informational video on LLLT, that turns into an ad at around (2:45) for terraquant.
video length: (7:06)
See a flexor withdrawal reflex cold laser therapy demonstration for chiropractic care in this free health care video.
video length: (4:28)
LLLT works by using light to stimulate, regulate and accelerate cell function in the area being treated in order to heal, restore and improve damaged tissue.
LLLT can be treated on any part of the body and restores injured tissue to return to a normal level in both structure and function, which alleviates symptoms that include swelling, redness, damaged skin, and pain.
video length: (1:47)
This video gives a simple description of LLLT, however the laser is used through clothing for demo purposes this would not be the case in actual LLLT.
video length: (2:22)
Two videos of a tiger recieving LLLT on a wound and on it's eye, tiger does't seem to be in pain, but the videos simply don't show enough to make any real claims.
video length: (0:40)
Tiger's eye gets THOR LLLT / Low Level Laser Therapy treatment
video length: (0:39)
Grizzly Bear recieving LLLT for it's osteoarthritis of the hip.
video length: (0:54)
Action News reports on the use of LLLT for relief from Carpal Tunnel Syndrome. LLLT provides relief from pain and inflammation, migraines, arthritis, fibromyalgia, back pain, neck pain, muscle injuries, ligament injuries, and many more.
video length: (3:02)
Anna Johnson is the British 3 Day Olympic team equine physio. Here she demonstrates the THOR laser system for pain relief and tissue healing.
video length: (4:51)
In this video three physical therapists talk about using LLLT for pain management and for smoking cessation.
video length: (3:24)
This is essentially an ad for thor, and the physical therapy clinic mentioned, but it still contains a small amount of information on LLLT
video length: (2:00)
LLLT also known as cold laser was featured on Channel 7 News with Dr Roberta Chow using low level laser to treat back pain
video length: (2:06)
This video shows North Dallas and Plano chiropractor, Dr. Khayal, demonstrating the use of cold laser therapy, sometimes known as "cold light laser". Premier Health Chiropractic employs cold laser therapy as a part of its patients healing and recovery process, along with other chiropractic technologies such as the ProAdjuster and spinal decompression.
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Video of a LLLT procedure on a dental pateint.
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Fox News 10 reports on the Microlight ML830® Cold Laser used for all types of pain relief.
video length: (4:58)
Dr. Sam Lam talks about a laser therapy device for hairloss, it is basically an ad, so be mindful of that.
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this video talks about the benificial effects of LLLT on all cells that are injured or ailing:
It also states that LLLT has no effect on cells that don't need energy.
video length: (7:26)
Lasers as a medicinal tool have been researched ever since their discovery forty years ago. LASER is the acronym for Light Amplification by Stimulated Emission of Radiation. Albert Einstein was the first person to discover the presence of laser; however it was Theodre Maiman who invented the first working ruby laser. He managed to produce a red colored laser from ruby crystal which was so intense that it could bore through several layers of stacked razor metals. The laser produced was powerful but short-lived. Since then much research has taken place in this area.
Low level lasers are characterized by low intensity levels and were used by Endre Mester, a Hungarian scientist in his medical research and he presented the paper on utility of Low level lasers in medicine in 1969. This resulted in astounding discoveries -- laser beams relieved patients from pain, hastened recuperation, and drastically cut down marks and blemishes.
LLLT- The Science behind It
The photons, carriers of electromagnetic radiations, an inherent component of all wavelengths of light activate the multifunctional nucleotide, ATP. ATP (Adenosine Triphosphate) contain high energy phosphate bonds which transport energy to cells for biochemical processes including muscular contraction and enzymatic metabolism, thereby restoring the normal enzymatic balance and energy levels. This, in a radical but natural manner, accelerates the healing process.
LLLT is characterized by typical LASER attributes of coherence, polarization and monochromaticity. They are narrow and shiny beams which can penetrate and be easily assimilated by the body. Low level LASERs have a specific wavelength range. The frequency of light is given by the movement of light moving upward and downward.
How It Works
Low level laser therapy works in a similar manner to the photosynthesis in plants where the sun's energy is used by the plants which initiate crucial cellular processes hastening the cell production and rejuvenating processes of the plant cells.
In the similar way, the photons of LLT reach the human body. While the body can be compared to the plant, the low level laser light is similar to the sun light. Once absorbed, the LASER activates the cell metabolism and cell reconstruction. The Low Level Laser Therapy rays are capable of boring 3 inches deep into the body. These rays insert bio photons into the damaged and the living cells. These cells start producing ATP improving their function, strengthening the body resistance by producing collagen, enzymes thereby improving the synthesis of various hormones. These substances are basic for the healthy functioning of the body cells. Hence the tissues are healed and pain disappears.
With photons as the driving force, the Low Level Laser is the silent healer of wounds, pains and dermatological disorders. It is established beyond doubt that unlike other drug or therapy, Low Level Laser Therapy has no peripheral or undesirable secondary effect. A laser is critical in revitalizing the impaired or injured cells by improving the resistance or immunity. Low Level Laser Therapy will go a long way in medical history and it has come to a stay.
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Advertisement that at least shows LLLT being used on patients.
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This video is not exactly the autority on LLLT, it does show some results in the use of LLLT in hairloss at (6:24); however it has a shameless hairloss clinic plug at the end, so take it with a grain of salt.
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Welcome to the laser-therapy.us research tool. This tool is a searchable collection of technical publications, books, videos and other resources about the use of lasers for photobiomodulation. This tool includes almost the entire U.S. library of medicine research papers on LLLT, videos from Youtube associated with therapy lasers and the tables of contents from laser therapy books. This allows users to search for a keyword or condition and see resources about using lasers to treat that condition. All the resources include links to the original source so we are not making any statement about the use of lasers for treating non-FDA cleared application, we are simple summarizing what others have said. Where every possible, we have included a link to the orginal publication.
Here are some of our favorite queries:
This tool uses a broad match query so:
The results of the search are sorted based on 3 quality factors on a scale of 1 to 10 with 10 being the best score. Originally all the resources were given a 5-5-5 until they could be individually evaluated. These scores are purely opinion and are only used to simplify the rank of the results from more valuable to least valuable. This should not be considered a critique of any work. This system was created to help researchers (including ourselves) find the most usable resources for any cold laser therapy research. The resources are assigned values based on the following 3 factors:
Over the past few years of working with research, we found that a majority of the published resources are lacking in one of these three ranking factors.
The original goal of this research tool was to tie published resources to the protocols in the laser-therapy.us library. This connection allows users to trace each protocol back to a list of resources so the protocol can be researched and improved.
When many of the first research papers were published, the most power laser available for therapy were less than 100mW and many systems had to be pulsed to keep the laser from burning out too quickly. Today, system are available that will deliver up to 60,000mW of continuous output. Because of these power limitation, many early studies were limited to extremely low dosages by today’s standards. It takes a 50mW system 17 minutes to deliver 50 joules at the surface of the skin. If this was spread over a large area of damage or was treating a deeper problem, the actual dosages were much less than 1J/cm2. Today, we know that these dosages typically produce very little or no results.
About 80% of the resources in this database are in the near infrared wavelength. There is also some interest in the red wavelength (600 to 660nm) . Other wavelengths like blue, purple, and green have very little scientific research behind them and have not gotten much traction in the core therapy market with the exception of some fringe consumer products.
This research tool is free to use but we make no claims about the accuracy of the information. It is an aggregation of existing published resources and it is up to the user to determine if the source of the resources has any value. The information provided through this web site should not be used for diagnosing or treating a health problem or disease. If you have or suspect you may have a health problem, you should consult your local health care provider.