Ứng Dụng Laser Trong Da Liễu Và Y Học (Lasers In Dermatology – 2018)

Lasers in Dermatology and Medicine

Keyvan Nouri
Editor
Lasers in Dermatology
and Medicine
Dermatologic Applications
Second Edition

ISBN 978-3-319-76116-9        ISBN 978-3-319-76118-3  (eBook)

Library of Congress Control Number: 2018950469
© Springer International Publishing AG, part of Springer Nature 2011, 2018
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Editor
Keyvan Nouri
Leonard M. Miller School of Medicine
University of Miami
Miami, FL
USA

v
This, Nouri’s book, is a thorough, recent, practical, and refreshing one that
puts “laser dermatology” into a broader perspective; it is a pleasure to update
my brief contribution for this edition. Almost immediately after the first laser
was created in 1960, a handful of visionary physicians recognized the poten-
tial for surgical applications, starting with the organ systems readily accessi-
ble to light. Lasers in laryngology, ophthalmology, and dermatology are so
fully adopted now that the standards of care have forever been changed. Now,
light is marching inside the body. Laser lithotripsy is widely practiced all over
the world. Know-how about lasers and biomedical optics is jumping between
medical specialties. Optical coherence tomography, a rapid form of live
microscopy invented for retinal imaging, is starting to impact dermatology
while making a larger splash for upper GI tract and coronary artery diagnostic
imaging. Dermatology was the first to figure out how to target individual
pigmented cells with laser pulses, a capability later adopted into ophthalmol-
ogy for glaucoma treatment. Recently, the various optical nanoparticles
developed for laser photo-thermal cancer therapy are being used in dermatol-
ogy for acne treatment.
How did we get such a wide, almost dazzling, variety of treatment lasers
in dermatology? (Because, we need them for different uses in various practice
settings; lasers are the most tissue-specific surgical tools in existence.) Do we
really need so many? (Well, we need most all of them. Only a few are inter-
changeable.) Are the mechanistic, clinical, safety, ethical, and practice-related
chapters of this book worthy of study? (Yes.) Can’t we just learn which but-
tons to push, in courses provided by the more reputable device manufacturers
just after a laser is purchased? (This approach is foolish beyond words, yet
such fools exist). Even more foolish are those who purchase a used laser and
start using it without any training whatsoever.
A great asset of this book is the breadth of its practical, clinical discus-
sions. There is no substitute for hands-on training, which cannot be obtained
even from this practical book. If you use lasers in practice, talk with your
colleagues and attend medical laser conferences in which you are free to ask
questions to faculty who are not trying to sell something. Many laser compa-
nies provide useful information, but are inherently biased. Laser companies
are restricted from discussing off-label indications. FDA clearance of a device
for a particular indication cannot be taken as assurance that it will work safely
and effectively enough to satisfy you and your patients, while lack of FDA
clearance for a specific indication cannot be taken as assurance that it will not
Foreword—And Forward!

vi
work safely and effectively. Some of the best uses for dermatological lasers
are not FDA-labeled indications, and probably never will be.
It is remarkable what lasers already can do for our patients, yet this field is
clearly still in its youth. What comes next? With the advent of fiber laser
technology, various industries and telecommunications now have extremely
powerful, efficient, wavelength-versatile lasers that operate reliably for
decades with little or no maintenance. Those have begun to make their way
into dermatology, and may ultimately do better what we do now, plus add
wholly new capabilities. Fractional lasers have taught us how amazingly tol-
erant skin is, to a large volume of micro-injury. Up to 30% of skin can be
killed or removed in random, full-thickness wounds that heal rapidly without
scarring. The caveat is that every little wound must be less than about 0.4 mm
wide. Given that, is it possible to “target” anything in the skin that can be
localized, regardless of its optical or thermal properties? If we knew where
various things are in the skin, can’t we just aim at them? Yes, we could!
Image-guided smart fractional lasers will be used to selectively treat struc-
tures and lesions not now addressed with lasers—and with that, we will have
software-programmable laser targeting. For example, all three cutaneous
glands—eccrine, sebaceous, and apocrine—are reasonable targets, as well as
nerves, lymphatics, sensory end organs, mast cells, antigen-presenting cells,
and other components of normal skin. Microscopy-driven ablative lasers may
even rival conventional microscopic margin-controlled tumor surgery, some
day. When laser microscopy and laser tissue ablation are finally married, sur-
gical oncology in general may be impacted. This new era is coming sooner
than you think.
I have been fortunate to play a role in launching many aspects of laser
dermatology, starting with some fundamental understanding of skin optics,
the concept of selective photothermolysis, lasers specifically designed for
dermatological use, permanent laser hair removal, scanning confocal laser
microscopy, and “fractional” laser treatments. Each of these arose from try-
ing to understand or solve one clinical problem, but now the panoply of clini-
cal laser applications far exceeds the initial effort. For example, fractional
lasers arose as a safer alternative to fully ablative laser skin resurfacing, a
safer way to induce skin remodeling. We had no idea that tissue so grossly
abnormal as a hypertrophic wound scar could be stimulated to normalize
itself this way. Fractional ablative lasers also offer a new way for delivery of
topical agents, including very high molecular weight macromolecules, parti-
cles, and even cells. The current widespread and diverse use of lasers in der-
matology attests not so much to new technology, as to the extreme value of
astute clinical observations made by dedicated dermatologists. Nouri’s text is
aimed exactly at achieving that. So please be a gourmet laser chef, not a
short-order cook. Contribute to an amazing and evolving part of
dermatology.
Thank you, Dr. Nouri and the many authors involved in this text, for your
excellent contribution.

R. Rox Anderson
Foreword—And Forward!

vii
Laser technology is quickly evolving with the presence of newer lasers, along
with new indications, that are constantly being introduced. The use of lasers
has become a major discipline and is currently practiced in a variety of fields
of medicine today. This book specifically offers a comprehensive literature
covering the different ways lasers are being used in the field of dermatology.
The authors of Lasers in Dermatology and Medicine are well known in their
respective fields and have attempted to cover each topic in the most compre-
hensive, readable, and understandable format. Each chapter consists of an
introduction and summary boxes in bulleted formats with up-to-date informa-
tion highlighting the importance of each respective section, enabling the
reader to have an easy approach towards reading and understanding the vari-
ous topics on lasers. This book has been written with the sincere hope of the
editors and the authors to serve as a cornerstone of laser usage in dermatol-
ogy, ultimately leading to better patient care and treatments. Lasers in derma-
tology have clearly expanded. The areas or laser treatments include port wine
stains, vascular anomalies and lesions, pigmented lesions and tattoos, hair
removal and hair re-growth, acne, facial rejuvenation, psoriasis, hypopig-
mented lesions and vitiligo, and treatment of fat and cellulites, among others.
The lasers are also being used for treatment and diagnosis of skin cancers.
We anticipate that this book will be of interest to all the physicians in the
field of dermatology who use or are interested in using lasers in their practice.
We are extremely grateful to our contributing authors. This book will serve as
a potential study source for physicians that would like to expand their knowl-
edge in lasers and light devices.
Miami, FL, USA
Keyvan Nouri
Preface

ix
I would like to sincerely thank my family for their support and encourage-
ment throughout my life. Special thanks to Dr. William H. Eaglestein, Dr.
Lawrence A. Schachner (former Chairman of Dermatology at the University
of Miami School of Medicine), and Dr. Robert Kirsner (Chairman of the
Department of Dermatology and Cutaneous Surgery at the University of
Miami Miller School of Medicine). They have given me great support and
have served as mentors throughout my professional career. Their guidance
and encouragement over the years has been greatly appreciated. Dr. Dr. Perry
Robins, Dr. Robin Ashinoff, Dr. Vicki Levine, Dr. Seth Orlow, the late Dr.
Irvin Freedberg, Dr. Hideko Kamino, and the entire faculty and staff at
New  York University School of Medicine Department of Dermatology:
Thank you all for the wonderful learning and friendship during my surgery
fellowship.
I would like to thank the faculty and dermatology residents, and the staff
of the Department of Dermatology and Cutaneous Surgery at the University
of Miami Miller School of Medicine, for their teaching, expertise, and friend-
ship. Special acknowledgements to the Mohs and Laser Center staff at the
Sylvester Cancer Center for their dedication, hard work, and support on a
daily basis. I would also like to thank Dr. Ali Rajabi-Estarabadi, my research
fellow, for his diligence and hard work and the rest of the Mohs staff, includ-
ing Cathy Mamas, Juana Alonso, Gladys Quintero, Destini M. Adkins, and
Ileana P. Reyes.
Special thanks to my clinical research fellows in dermatologic surgery,
Sofia Iglesia, Ariel Eva Eber, Sebastian H.  Verne, Marina Perper, Robert
Magno, Alaleh Dormishian, and Samuel C. Smith, for all their hard work and
contributions to this book.
I would also like to acknowledge the publishing staff Mr. Grant Weston,
Ms. Tracy Marton, Mr. Leo Johnson, and the entire Springer Publishing team
for having done a superb job with the publication. It has been a pleasure
working with them and this excellent project to compile the textbook.
Lastly, I would like to sincerely thank all the authors of this textbook.
These individuals are world-renowned in their respective specialties and
without their time and energy, writing this book would have not been possi-
ble. These individuals have made this a comprehensive, up-to-date, and reli-
able source on Lasers in Dermatology and Medicine. I truly appreciate their
hard work and thank them for their contributions.
Keyvan Nouri
Acknowledgements

xi
1 Laser-Tissue Interactions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .       1
Amanda Abramson Lloyd, Michael S. Graves, and Edward
Victor Ross
2 Laser Safety: Regulations, Standards
and Practice Guidelines. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .     37
Brienne D. Cressey, Ashley Keyes, and Murad Alam
3 Lasers for Treatment of Vascular Lesions . . . . . . . . . . . . . . . . .     49
Jayne Joo, Daniel Michael, and Suzanne Kilmer
4 Laser for Scars . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .     63
Voraphol Vejjabhinanta, Shalu S. Patel, and Keyvan Nouri
5 Laser Treatment of Leg Veins . . . . . . . . . . . . . . . . . . . . . . . . . . .     73
Julie K. Karen and Shields Callahan
6 Lasers and Lights for Treating Pigmented Lesions. . . . . . . . . .     83
Emmy M. Graber and Jeffrey S. Dover
7 Laser Treatment of Tattoos . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   109
Voraphol Vejjabhinanta, Caroline V. Caperton,
Christopher Wong, Rawat Charoensawad, and Keyvan Nouri
8 Laser for Hair Removal. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   121
Voraphol Vejjabhinanta, Keyvan Nouri, Anita Singh,
Ran Huo, Rawat Charoensawad, Isabella Camacho, and Ali
Rajabi-Estarabadi
9 Lasers for Resurfacing. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   137
Rungsima Wanitphakdeedecha and Tina S. Alster
10 Fractional Photothermolysis. . . . . . . . . . . . . . . . . . . . . . . . . . . .   165
Dieter Manstein, Hans-Joachim Laubac, Sofia Iglesia,
Alaleh Dormishian, Ali Rajabi-­Estarabadi, and Keyvan Nouri
11 Sub-surfacing Lasers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   197
Michael Howard Gold
12 Non-invasive Rejuvenation/Skin Tightening:
Light-Based Devices. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   213
Marina Perper, John Tsatalis, Ariel E. Eber, and Keyvan Nouri
Contents

xii
13 Laser and Light Therapies for Acne. . . . . . . . . . . . . . . . . . . . . .   227
Ali Rajabi-Estarabadi, Ariel E. Eber, and Keyvan Nouri
14 Lasers for Psoriasis and Hypopigmentation. . . . . . . . . . . . . . . .   237
Laura Jordan, Summer Moon, and James M. Spencer
15 Lasers for Adipose Tissue and Cellulite. . . . . . . . . . . . . . . . . . .   247
Molly Wanner and Mathew M. Avram
16 Photodynamic Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   261
Ariel E. Eber, Marina Perper, Sebastian H. Verne,
Robert Magno, Ibrahim Abdullah Omair ALOmair,
Mana ALHarbi, and Keyvan Nouri
17 Intense Pulsed Light (IPL). . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   275
Sanjana Iyengar, Keyvan Nouri, Peter Bjerring,
Kåre Christiansen, Robert A. Weiss, Girish S. Munavalli,
Sonal Choudhary, and Angel Leiva
18 Current Status of Light-Emitting Diode Phototherapy
in Dermatological Practice. . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   285
R. Glen Calderhead
19 Laser and Light for Wound Healing Stimulation . . . . . . . . . . .   339
Ehsan Azimi, Navid Bouzari, and Keyvan Nouri
20 Lasers in Hair Growth and Hair Transplantation. . . . . . . . . . .   351
Nicole E. Rogers, Marc R. Avram, Isabella Camacho,
and Ali Rajabi-Estarabadi
21 Photobiomodulation and Hair Growth. . . . . . . . . . . . . . . . . . . .   367
Molly B. Hirt and Ronda S. Farah
22 Reflectance Confocal Microscopy
in Oncological Dermatology. . . . . . . . . . . . . . . . . . . . . . . . . . . . .   375
Pablo Fernández-Crehuet Serrano,
Gonzalo Segurado-Miravalles, and Salvador González
23 Laser Clinical and Practice Pearls . . . . . . . . . . . . . . . . . . . . . . .   401
Hana Jeon, Lori A. Brightman, and Roy G. Geronemus
24 The Selection and Education of Laser Patients. . . . . . . . . . . . .   415
Murad Alam and Meghan Dubina
25 Anesthesia for Laser Surgery. . . . . . . . . . . . . . . . . . . . . . . . . . . .   427
Marina Perper, Ali Rajabi-Estarabadi, Ariel E. Eber,
Voraphol Vejjabhinanta, Ran Huo, Keyvan Nouri,
and John Tsatalis
26 Lasers in Skin of Color. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   437
Heather Woolery-Lloyd and Nkanyezi Ferguson
27 Laser Applications in Children. . . . . . . . . . . . . . . . . . . . . . . . . .   449
Jessica Cervantes, Sebastian H. Verne,
and Mercedes E. Gonzalez
Contents

xiii
28 Dressing/Wound Care for Laser Treatment. . . . . . . . . . . . . . . .   467
Ariel E. Eber, Vincent M. Hsu, Stephen C. Davis,
and Keyvan Nouri
29 Prevention and Treatment of Laser Complications. . . . . . . . . .   475
Rachael L. Moore, Juan-Carlos Martinez, Ken K. Lee,
Yun Ehrlich, Brian Simmons, and Keyvan Nouri
30 Ethical Issues. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   493
Abel Torres, Tejas Desai, Alpesh Desai, William T. Kirby,
and Maren C. Locke
31 Medicolegal Issues (Documentation/Informed
Consent/Non-physician Operators). . . . . . . . . . . . . . . . . . . . . . .   499
Abel Torres, Tejas Desai, Sailesh Konda, Alpesh Desai,
and William T. Kirby
32 Photography of Dermatological Laser Treatment. . . . . . . . . . .   505
Shraddha Desai and Ashish C. Bhatia
33 Online Resources of Dermatologic Laser Therapies. . . . . . . . .   517
Shraddha Desai, Elizabeth Uhlenhake, and Ashish C. Bhatia
34 Starting a Laser Practice. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .   523
Vic A. Narurkar
35 Research and Future Directions . . . . . . . . . . . . . . . . . . . . . . . . .   527
Yakir S. Levin, Fernanda Hidemi Sakamoto,
and R. Rox Anderson
Index��������������������������������������������������������������������������������������������������������  541
Contents

xv
Murad  Alam  Department of Dermatology, Otolaryngology, and Surgery,
Northwestern University, Chicago, IL, USA
Department of Dermatology, Northwestern University, Chicago, IL, USA
Mana  ALHarbi  Department of Dermatology, College of Medicine, Al
Imam Muhammad Ibn Saud Islamic University (IMSIU), Riyadh, Kingdom
of Saudi Arabia
Ibrahim Abdullah Omair ALOmair  Department of Dermatology, College
of Medicine, Al Imam Muhammad Ibn Saud Islamic University (IMSIU),
Riyadh, Kingdom of Saudi Arabia
Tina S. Alster  Georgetown University Medical Center, Washington Institute
of Dermatologic Laser Surgery, Washington, DC, USA
Marc  R.  Avram  Department of Dermatology, Massachusetts General
Hospital, Harvard Medical School, Boston, MA, USA
Mathew  M.  Avram  Department of Dermatology, Massachusetts General
Hospital, Harvard Medical School, Boston, MA, USA
Ehsan  Azimi  Department of Dermatology, Cutaneous Biology Research
Center, Massachusetts General Hospital, Charlestown, MA, USA
Ashish C. Bhatia  Department of Dermatology, Northwestern University –
Feinberg School of Medicine, Chicago, IL, USA
Oak Dermatology, Schaumburg, IL, USA
Peter Bjerring  Department of Dermatology, Swansea University, Swansea,
Wales, UK
Navid Bouzari  South Shore Skin Center, Plymouth, MA, USA
Lori A. Brightman  Laser & Skin Surgery Center of New York, New York,
NY, USA
R.  Glen  Calderhead  Research Division, VP Medicoscientific Affairs,
Clinique L, Goyang-shi, Gyeonggi-Do, South Korea
Shields  Callahan  Department of Dermatology, Virginia Commonwealth
University School of Medicine, Richmond, VA, USA
Contributors

xvi
Isabella Camacho  School of Medicine, Georgetown University, Washington,
DC, USA
Caroline  V.  Caperton  Miller School of Medicine, University of Miami,
Miami, FL, USA
Jessica Cervantes  Department of Dermatology, University of Miami Miller
School of Medicine, Miami, FL, USA
Rawat Charoensawad  Rawat Clinic, Bangkok, Thailand
Biophile Training Center, Bangkok, Thailand
Sonal  Choudhary  Department of Dermatology and Cutaneous Surgery,
Miller School of Medicine, University of Miami, Miami, FL, USA
Kåre Christiansen  Molholm Research, Molholm Hospital, Vejle, Denmark
Brienne D. Cressey  Department of Dermatology, New York Presbyterian-
Weill Cornell Hospital, New York, NY, USA
Stephen  C.  Davis  Department of Dermatology and Cutaneous Surgery,
University of Miami, Miami, FL, USA
Alpesh  Desai  Heights Dermatology and Aesthetic Center, Houston,
TX, USA
Shraddha Desai  Dermatology Institute, DuPage Medical Group, Naperville,
IL, USA
Tejas Desai  Heights Dermatology and Aesthetic Center, Houston, TX, USA
Roberto Diaz  Department of Neurosurgery, University of Miami, Miami,
FL, USA
Alaleh  Dormishian  Department of Dermatology and Cutaneous Surgery,
University of Miami Miller School of Medicine, Miami, FL, USA
Jeffrey  S.  Dover  Department of Dermatology, Yale University School of
Medicine and Brown University, Chestnut Hill, MA, USA
Meghan  Dubina  Department of Dermatology, Northwestern University,
Chicago, IL, USA
Ariel  E.  Eber  Department of Dermatology and Cutaneous Surgery,
University of Miami, Miller School of Medicine, Miami, FL, USA
Yun Ehrlich  Department of Dermatology and Cutaneous Surgery, University
of Miami, Miami, FL, USA
Ronda  S.  Farah  Department of Dermatology, University of Minnesota,
Minneapolis, MN, USA
Nkanyezi  Ferguson  Department of Dermatology, University of Iowa
Hospital and Clinics, Iowa City, IA, USA
Elizabeth A. M. Frost  Mount Sinai Medical Center, New York, NY, USA
Julie A. Gayle  Department of Anesthesiology, LSUHSC School of Medicine,
New Orleans, LA, USA
Contributors

xvii
Roy G. Geronemus  Laser & Skin Surgery Center of New York, New York,
NY, USA
Department of Dermatology, New York University Medical Center,
New York, NY, USA
Michael C. Giovingo  Department of Ophthalmology, Stroger Hospital of
Cook County, Chicago, IL, USA
Michael  Howard  Gold  Gold Skin Care Center and Tennessee Clinical
Research Center, Nashville, TN, USA
Mercedes E. Gonzalez  Department of Dermatology, University of Miami
Miller School of Medicine, Miami, FL, USA
Salvador  González  Department of Medicine and Medical Specialities,
Alcalá University, Madrid, Spain
Department of Dermatology, Memorial Sloan-Kettering Cancer Center, New
York, NY, USA
Hospital Ramón y Cajal, Alcala University, Madrid, Spain
Emmy M. Graber  Dermatology Institute of Boston, PC, Boston, MA, USA
Michael S. Graves  Southwest Skin Cancer & Vein Clinic, Austin, TX, USA
Ramez  I.  Haddadin  Department of Ophthalmology, Feinberg School of
Medicine, Northwestern University, Chicago, IL, USA
Molly  B.  Hirt  Department of Dermatology, University of Minnesota,
Minneapolis, MN, USA
Vincent  M.  Hsu  Department of Dermatology and Cutaneous Surgery,
University of Miami Miller School of Medicine, Miami, FL, USA
Ran Huo  Broward Dermatology Clinic, Pembroke Pines, FL, USA
Sofia Iglesia  Department of Dermatology and Cutaneous Surgery, University
of Miami Miller School of Medicine, Miami, FL, USA
Viacheslav  Iremashvili  Department of Urology, Jackson Memorial
Hospital, Miami, FL, USA
Michael E. Ivan  Department of Neurological Surgery, University of Miami
Hospital, Jackson Memorial Hospital, and Jackson South Hospital, Miami,
FL, USA
Sanjana Iyengar  Department of Dermatology, Feinberg School of Medicine,
Northwestern University, Chicago, IL, USA
Hana Jeon  Laser & Skin Surgery Center of New York, New York, NY, USA
Jayne Joo  Department of Dermatology, Davis and Sacramento VA Medical
Center, University of California, Sacramento, CA, USA
Laura Jordan  Kansas City University Consortium, Tri-County Dermatology
Residency Program, Cuyahoga Falls, OH, USA
Contributors

xviii
Julie  K.  Karen  Department of Dermatology, CompleteSkinMD, NYU
Langone Medical Center, New York, NY, USA
Alan David Kaye  Department of Anesthesiology, Louisiana Health Sciences
Center, School of Medicine, New Orleans, LA, USA
Ashley Keyes  Department of Dermatology, Weill Cornell Medicine, Lincoln
Medical Center, Bronx, NY, USA
Suzanne  Kilmer  Laser & Skin Surgery Center of Northern California,
Sacramento, CA, USA
Department of Dermatology, Davis Medical Center, University of California,
Sacramento, CA, USA
William T. Kirby  Laseraway Hermosa Beach, Hermosa Beach, CA, USA
Ricardo  J.  Komotar  Department of Neurosurgery, University of Miami,
Miami, FL, USA
Sailesh Konda  Department of Dermatology, University of Florida College
of Medicine, Gainesville, FL, USA
Raymond  J.  Lanzafame  Raymond J.  Lanzafame, MD PLLC, Buffalo,
NY, USA
School of Dental Medicine, State University of New York at Buffalo, Buffalo,
NY, USA
Hans-Joachim 
Laubac  Dermatologische
Abteilung,
Universitäre
Krankenhäuser von Gen, Genf, Schweiz
Ken K. Lee  Dermatologic and Laser Surgery, Oregon Health and Sciences
University, Portland, OR, USA
Angel  Leiva  Department of Dermatology and Cutaneous Surgery, Miller
School of Medicine, University of Miami, Miami, FL, USA
Yakir  S.  Levin  Department of Dermatology, Massachusetts General
Hospital, Boston, MA, USA
Wellman Center for Photomedicine, Massachusetts General Hospital, Boston,
MA, USA
Amanda Abramson Lloyd  Skin and Vein Institute, Encinitas, CA, USA
Maren  C.  Locke  Department of Dermatology, MetroHealth System,
Cleveland, OH, USA
Robert  Magno  Department of Dermatology, University of Miami, Toms
River, NJ, USA
Dieter  Manstein  Wellman Center for Photomedicine, Massachusetts
General Hospital, Harvard Medical School, Boston, MA, USA
Robert Marcovich  Department of Urology, University of Miami, Miami,
FL, USA
Juan-Carlos  Martinez  Dermatologic and Laser Surgery, Oregon Health
and Sciences University, Portland, OR, USA
Contributors

xix
Daniel Michael  Department of Dermatology, Laser and Skin Surgery Center
of Northern California, University of California, Davis, Medical Center,
Sacramento, CA, USA
Summer Moon  Dermatology Specialists of West Florida, FL, USA
Rachael L. Moore  Dermatology Specialists of West Florida, Oregon Health
and Sciences University, Portland, OR, USA
Girish  S.  Munavalli  Dermatology, Laser, and Vein Specialists of the
Carolinas, Charlotte, NC, USA
Timothy  G.  Murray  Department of Ophthalmology/Ocular Oncology,
Retina Vitreous Diseases/Pediatric Ophthalmology, Medical Arts Surgery
Center Baptist, Nicklaus Children’s Hospital, Miami, FL, USA
Vic A. Narurkar  Bay Area Laser Institute, San Francisco, CA, USA
Keyvan Nouri  Dermatology, University of Miami, Miami, FL, USA
Ophthalmology, University of Miami, Miami, FL, USA
Otolaryngology, University of Miami, Miami, FL, USA
Dermatologic Surgery, University of Miami, Miami, FL, USA
Department of Dermatology and Cutaneous Surgery, University of Miami
Medical Group, University of Miami, Miami, FL, USA
Dermatology Services at Sylvester Comprehensive Cancer Center, University
of Miami Hospital and Clinics, Miami, FL, USA
MOHS, Dermatologic and Laser Surgery, University of Miami, Miami, FL,
USA
Surgical Training, University of Miami, Miami, FL, USA
Mahnaz Nouri  Boston Eye Group, Brookline, MA, USA
Peter  O’Kane  Dorset Heart Centre, Royal Bournemouth Hospital,
Bournemouth, Dorset, UK
Carolyn Orgain  Department of Otolaryngology, University of California,
Irvine, Orange, CA, USA
Steven Parker  , Harrogate, North Yorkshire, UK
Krishna B. Patel  Department of Ophthalmology, John H. Stroger Hospital
of Cook County, Lombard, IL, USA
Shalu S. Patel  Department of Dermatology and Cutaneous Surgery, Miller
School of Medicine, University of Miami, Miami, FL, USA
Robert Perez  Department of Dermatology and Cutaneous Surgery, Miller
School of Medicine, University of Miami, Miami, FL, USA
Marina  Perper  Department of Dermatology and Cutaneous Surgery,
University of Miami Hospital, Miller School of Medicine, Miami, FL, USA
Contributors

xx
Roberto Pineda  Department of Ophthalmology, Massachusetts Ear and Eye
Hospital, Boston, MA, USA
Ali Rajabi-Estarabadi  Department of Dermatology and Cutaneous Surgery,
University of Miami Miller School of Medicine, Miami, FL, USA
John  Rawlins  Dorset Heart Centre, Royal Bournemouth Hospital,
Bournemouth, UK
Douglas  J.  Rhee  Department of Ophthalmology and Visual Science,
University Hospitals, Cleveland, OH, USA
Department of Ophthalmology and Visual Sciences, Case Western Reserve
University School of Medicine, Cleveland, OH, USA
Cornelia Selma de Riese  Department of Obstetrics and Gynecology, Texas
Tech University Health Sciences Center, Lubbock, TX, USA
Nicole E. Rogers  Hair Restoration of the South, Metairie, LA, USA
Edward  Victor  Ross  Department of Dermatology, Scripps Clinic, San
Diego, CA, USA
Vanessa  Rothholtz  Pacific Coast Ear Nose and Throat, Beverly Hills,
CA, USA
R.  Rox  Anderson  Department of Dermatology, Massachusetts General
Hospital, Harvard Medical School, Boston, MA, USA
Wellman Center for Photomedicine, Massachusetts General Hospital, Harvard
Medical School, Boston, MA, USA
Ryan  Rubin  Department of Anesthesiology, Louisiana Health Sciences
Center, New Orleans, LA, USA
Fernanda Hidemi Sakamoto  Department of Dermatology, Massachusetts
General Hospital, Harvard Medical School, Boston, MA, USA
Wellman Center for Photomedicine, Massachusetts General Hospital, Harvard
Medical School, Boston, MA, USA
Amy  C.  Schefler  Department of Ophthalmology, Bascom Palmer Eye
Institute, Miami, FL, USA
Gonzalo  Segurado-Miravalles  Department of Dermatology, Ramón y
Cajal Hospital, Madrid, Spain
Pablo  Fernández-Crehuet  Serrano  Department of Dermatology, Alto
Guadalquivir de Andújar, Andújar, Jaén, Spain
Brian  Simmons  Department of Dermatology and Cutaneous Surgery,
University of Miami, Miami, FL, USA
Anita  Singh  Montefiore Medical Center, Albert Einstein College of
Medicine, Bronx, NY, USA
James M. Spencer  Spencer Dermatology and Skin Surgery Center (Spencer
Dermatology, Carillon Outpatient Center), St. Petersburg, FL, USA
Contributors

xxi
Joseph Stuto  Department of Dermatology, Weill Cornell Medical Center,
New York, NY, USA
Amit  Todani  Department of Ophthalmology, Goodman Eye Medical &
Surgical Center, Milford, MA, USA
Allyne  Topaz  Hackensack University Medical Center, Hackensack,
NJ, USA
On Topaz  Asheville VA Medical Center, Asheville, NC, USA
Abel Torres  Department of Dermatology, Loma Linda University School of
Medicine, Loma linda, CA, USA
John  Tsatalis  Department of Dermatology and Cutaneous Surgery,
University of Miami Miller School of Medicine, Miami, FL, USA
Elizabeth Uhlenhake  The Dermatology Group, Mason, OH, USA
Voraphol  Vejjabhinanta  Department of Dermatology and Cutaneous
Surgery, Leonard M. Miller School of Medicine, University of Miami, Miami,
FL, USA
Department of Dermatology, Siriraj Hospital, Mahidol University, Bangkok,
Thailand
Sebastian H. Verne  Department of Dermatology and Cutaneous Surgery,
University of Miami Miller School of Medicine, Miami, FL, USA
Victor M. Villegas  Bascom Palmer Eye Institute, Miami, FL, USA
Rungsima  Wanitphakdeedecha  Department of Dermatology, Faculty of
Medicine Siriraj Hospital, Mahidol University, Bangkok, Thailand
Molly  Wanner  Department of Dermatology, Massachusetts General
Hospital, Harvard Medical School, Boston, MA, USA
Robert  A.  Weiss  Department of Dermatology, Maryland Dermatology
Laser Skin and Vein Institute, University of Maryland School of Medicine,
Hunt Valley, MD, USA
Andrew Whittaker  Department of Cardiology, Dorset Heart Centre, Royal
Bournemouth Hospital, Bournemouth, Dorset, UK
Heather  Woolery-Lloyd  Department of Dermatology and Cutaneous
Surgery, Miller School of Medicine, University of Miami, Miami, FL, USA
Brian J. F. Wong  Department of Otolaryngology, University of California,
Irvine, Orange, CA, USA
Christopher Wong  Miller School of Medicine, University of Miami, Miami,
FL, USA
Charles  C.  Wykoff  Department of Ophthalmology, Bascom Palmer Eye
Institute, Miller School of Medicine, University of Miami, Miami, FL, USA
Roger B. Yandell  Department of Obstetrics and Gynecology, Texas Tech
University Health Sciences Center, Lubbock, TX, USA
Contributors

1
© Springer International Publishing AG, part of Springer Nature 2018
K. Nouri (ed.), Lasers in Dermatology and Medicine,
Laser-Tissue Interactions
Amanda Abramson Lloyd, Michael S. Graves,
and Edward Victor Ross
Abstract
The best gauge of laser interactions is the tis-
sue response, and experiment is the most real-
istic manner to address medical treatment
challenges. However, theoretical models are
helpful in planning treatment approaches and
laser parameters. In this chapter we discuss
basics of lasers, their non laser counterparts,
and laser-tissue interactions.
Many physicians choose laser settings out
of habit (or reading it off of a label attached to
the side of the machine—a “cheat” sheet with
skin-type specific parameters), using tissue
endpoints to confirm the appropriateness of
the parameters. For example, when treating a
tattoo with a Q-switched laser, the operator
looks for immediate frosty whitening. Like
driving a car (where the operator may have no
idea about nature of the drive train compo-
nents), successful laser operation does not
demand a complete understanding of the
machine or the details of the light-tissue inter-
action. However, a comprehension of first
principles allows for a logical analysis of final
clinical outcomes—furthermore, more cre-
ative uses of equipment should follow. For
example, with an education in laser tissue
interactions (LTIs) and tissue cooling, one can
deploy the alexandrite (long pulse) laser either
as a hair removal device, vascular laser, or to
remove lentigines.
The reader should note that although the
title of this chapter is “Laser Tissue
Interactions”, the introduction of many new
and diverse technologies make the term some-
what obsolete. We will continue to use the
term, but a more appropriate term is “energy–
tissue interactions.” As both radiofrequency
and ultrasound are increasingly applied in
medicine. We will use both terms interchange-
ably in the remainder of the text.
Keywords
Laser · Radiofrequency · Ultrasound · Skin ·
Lentigines · Hair removal · Vascular ·
Dermatology
A. A. Lloyd
Skin and Vein Institute,
Encinitas, CA, USA
M. S. Graves
Southwest Skin Cancer and Vein Clinic,
Austin, TX, USA
E. V. Ross (*)
Laser and Cosmetic Dermatology, Scripps Clinic,
San Diego, CA, USA
e-mail: ross.edward@scrippshealth.org
1

2
Introduction
1. Light represents one portion of a broader elec-
tromagnetic spectrum.
2. Light-tissue interactions involve the complex
topics of tissue optics, absorption, heat gen-
eration, and heat diffusion
3. Lasers are a special type of light with the char-
acteristics of monochromaticity, directional-
ity, and coherence.
4. Coagulation/denaturation is time and temper-
ature dependent
5. Proper selection of light parameters is based
on the color, size, and geometry of the target
6. Wound healing is the final but not least impor-
tant part of the laser tissue sequence (the
epilogue)
7. Laser-tissue interactions are fluid—the opera-
tor should closely examine the skin surface
during all aspects of the procedure
8. Pulse duration and light doses are often as
important as wavelength in predicting tissue
responses to laser to irradiation
The best gauge of laser interactions is the tis-
sue response, and experiment is the most realistic
manner to address medical treatment challenges.
However, theoretical models are helpful in plan-
ning treatment approaches and laser parameters.
In this chapter we discuss basics of lasers, their
non-laser counterparts, and laser-tissue interac-
tions [1].
Many physicians choose laser settings out of
habit (or reading it off of a label attached to the
side of the machine—a “cheat” sheet with skin-­
type specific parameters), using tissue endpoints
to confirm the appropriateness of the parameters.
For example, when treating a tattoo with a
Q-switched laser, the operator looks for immedi-
ate frosty whitening. Like driving a car (where
the operator may have no idea about nature of the
drive train components), successful laser opera-
tion does not demand a complete understanding
of the machine or the details of the light-tissue
interaction. However, a comprehension of first
principles allows for a logical analysis of final
clinical outcomes—furthermore, more creative
uses of equipment should follow. For example,
with an education in laser tissue interactions
(LTIs) and tissue cooling, one can deploy the
alexandrite (long pulse) laser either as a hair
removal device, vascular laser, or to remove len-
tigines [2].
The reader should note that although the title
of this chapter is “Laser Tissue Interactions”, the
introduction of many new and diverse technolo-
gies make the term somewhat obsolete. We will
continue to use the term, but a more appropriate
term is “energy–tissue interactions.” As both
radiofrequency and ultrasound are increasingly
applied in medicine. We will use both terms
interchangeably in the remainder of the text.
Light
Light represents one portion of a much broader
electromagnetic spectrum. Light can be divided
into the UV (200–400 nm), VIS (400–700 nm),
NIR “I” (755–810 nm), NIR “II” (940–1064 nm),
MIR (1.3–3 μm), and Far IR (3 μm and beyond).
On a macroscopic level, light is adequately char-
acterized as waves. The amplitude of the wave is
perpendicular to the propagation direction. Light
waves behave according to our “eyeball” obser-
vations in day-to-day life. For example, we are
familiar with refraction and reflection. The sur-
face of a pond is a partial mirror (reflection); a
fish seen in the pond is actually deeper than it
appears (refraction) [3]. Normally, the percent-
age of incident light reflected from the skin sur-
face is determined by the index of refraction
difference between the skin surface (stratum cor-
neum n = 1.55) and air (n = 1) [4]. This regular
reflectance is about 4–7% for light incident at
right angles to the skin [3, 5]. The angle between
the light beam and the skin surface determines
the % of reflected light. More light is reflected at
“grazing” angles of incidence. It follows that, to
minimize surface losses, in most laser applica-
tions, one should deliver light approximately per-
pendicular to the skin [3, 6]. One can deliberately
angle the beam, on the other hand, to decrease
penetration depth and also attenuate the surface
fluence by “spreading” the beam. One can reduce
interface losses by applying an alcohol solution
A. A. Lloyd et al.

3
(n = 1.4), water (n = 1.33), or a sapphire crystal
(n = 1.55 μm). This allows for optical coupling
(vide infra). On the other hand, the surface of dry
skin reflects more light because of multiple skin–
air interfaces (hence the white appearance of a
psoriasis plaque).
Light penetrates into the epidermis according
to wavelength dependent absorption and scatter-
ing (vide infra) [1, 6–8]. Because of scattering,
much incident light is remitted (remittance refers
to the total light returned to the environment due
to multiple scattering in the epidermis and der-
mis, as well as the regular reflection from the sur-
face). In laser surgery, light reflected from the
surface is typically “wasted”. This “lost” energy
varies from 15% to as much as 70% depending
on wavelength and skin type. For example, for
1064 nm, 60% of an incident laser beam may be
remitted. One can easily verify this by holding a
finger just adjacent to the beam near the skin sur-
face. Warmth can be felt from the remitted por-
tion of the beam.
To describe laser tissue interactions at the
molecular/microscopic level, light is considered
as a stream of “particles” called photons, where
the photon energy depends on the wavelength of
light.

E
hc
photon =
/ l
(1.1)
Where h is Plank’s constant (6.6 × 10−34 J -s),
and c is the speed of light (3 × 1010 cm/s) [9].
Types of Light Devices
In principle, many non-laser devices could be
used for heating skin [9]. Most properties of laser
light (i.e., coherence) are unimportant insofar as
the way light interacts with tissue in therapeutic
applications. And although collimation (lack of
divergence) of the incident beam might increase
the % of transmitted light with laser versus IPL,
the increasing use of filtered flash lamps in der-
matology suggests that losses from IPL beam
divergence are not critical. In lieu of lasers, some
thermal sources can be used in skin surgery (i.e.,
nitrogen plasma device) for resurfacing (Portrait,
Rhytec, MA). The critical features of any device
are controlling the device–tissue interaction time
to allow for precise heating (vide infra).
Lasers are useful because they allow for pre-
cise control of where and how much one heats
[10]. There are four properties that are common
to all laser types (1) Beam directionality (colli-
mation), (2) Monochromaticity, (3) Spatial and
temporal coherence of the beam, and (4) High
intensity of the beam [11]. The intensity, direc-
tionality, and monochromaticity of laser light
allow the beam to be expanded, or focused quite
easily. With non-laser sources like flashlamps
directed toward the skin surface, the light inten-
sity at the skin surface cannot exceed the bright-
ness of the source lamp. With many lasers, a lamp
similar to the intense pulsed light (IPL) flashlamp
pumps the laser cavity [12]. The amplification of
light within the laser cavity sets laser light apart
from other sources.
For most visible light applications, laser rep-
resents a conversion from lamplight to an ampli-
fied monochromatic form [13]. The high power
possible with lasers (especially peak power) is
achieved through resonance in the laser cavity.
For many dermatology applications requiring ms
or longer pulses delivered to large skin areas,
IPLs are either adequate or preferable to lasers.
The scientific principle on which lasers are based
is stimulated emission. With spontaneous emis-
sion, electrons transition to the lower level in a
random process. With stimulated emission, the
emission occurs only in the presence of photons
of a certain energy. The critical point is maintain-
ing a condition where the population of photons
in a higher state is larger than that in the lower
state. To create this population inversion, a pump-
ing energy must be directed either with electric-
ity, light, or chemical energy.
All lasers contain four main components, the
lasing medium, the excitation source, feedback
apparatus, and an output coupler. With respect to
lasing media, there are diode lasers, solid-state
lasers, dye, and gas lasers. Most solid state and
dye lasers use optical exciters (lamps), whereas
gas and diode lasers use electrical excitation (i.e.,
CO2 and RF). The feedback mechanism consists
of mirrors where one mirror reflects 100% and
the other transmits a small fraction of light [14].
1  Laser-Tissue Interactions

4
An example of a solid-state laser is the alexan-
drite laser. A solid-state laser consists of a rod
that is pumped by a flashlamp. The lamp pumps
the rod for stimulated emission. The rod and
lamp assembly must also be designed for ade-
quate cooling. Lasers typically are finicky
because all of the components are driven near
their damage thresholds (like redlining your car
all the time). As an example of this concept, con-
sider the pulsed dye laser (PDL). As the dye
degrades, the lamps must work harder to generate
higher pulse energies from the dye. Also, mirrors
become contaminated over time such that the
lamps must work harder and harder. These
demands stress the power supply. Thus, eventu-
ally, the dye kit, the power supply, lamps, and dye
are all working at their maximal output. Often
people speak of a tunable dye laser. In fact many
dye lasers are tunable; the manufacturers have
simply chosen one wavelength. An example of a
tunable laser was the Sclero-plus pulsed dye laser
(tunable from 585 to 600 nm in 5 nm increments)
from Candela (Candela, Wayland, MA).
Laser systems differ with regard to duration
and power of the emitted laser radiation. In con-
tinuous wave lasers (CW mode) with power out-
puts of up to 103 W, the lasing medium is excited
continuously. With pulsed lasers, excitation is
effected in a single pulse or in on-line pulses
(free-running mode). Peak powers of 105 W can
be developed for a duration of 100  μs–10  ms.
Storing the excitation energy and releasing it sud-
denly (Q-switch mode or mode-locking) leads to
a peak power increase of up to 1010–1012 W, and a
pulse duration of 10 ps–100 ns [13].
Light emitting diodes (LEDs) are becoming
commonplace in dermatology (Fig. 1.1). Primarily
used as a PDT light source, they are also used in
biostimulation. LEDs are similar to semiconduc-
tor (aka diode) lasers in that they use electrical
current placed between two types of semiconduc-
tors. However, they lack an amplification process
(no mirrors). LEDs do not produce coherent
beams but can produce monochromatic light.
Semiconductor (diode) lasers contain an LED as
the active gain medium. A current passes through
a sandwich of two layers consisting of compounds
(called p type and n type). Below threshold, there
is no oscillation and the semiconductor LASER
acts like an LED. This emission is very similar to
the visible emission of light emitting diodes. If
one adds mirrors it operates as a tiny laser instead
of an LED. The overall efficiency of semiconduc-
tor lasers is quite high, approximately 30% and
among the highest available for any laser types.
Most semiconductor (diode) lasers are operated in
CW mode but can be pulsed. New visible
light semiconductor lasers are available, and also
laser diode arrays are available where scientists
have created large numbers of semiconductor
lasers on one substrate. Some diode lasers are
housed separate from the handpiece and delivered
by fiber optics. Others are configured with the
laser diodes in the handpiece as arrays. Modern
diode lasers achieve higher powers than in the
past, but their peak powers are still lag behind
most pulsed solidstate lasers [14].
Excimer lasers emit UV light and are used for
photomodulation of the immune system. They
have also been used in surgery. The possible muta-
genicity of these lasers has not been well studied.
Materials such as the KTP crystal can be used to
generate harmonics with lasers. The KTP crystal is
used to convert 1064 nm radiation to 532 green
light. Also quality (Q) switching is used for gener-
ating short pulses. Much of the electrical energy
used to create laser emissions is wasted as heat,
which is why water is used for cooling most lasers.
Air cooling is used for some high-powered flash
lamps and many diode lasers. In the future, free
electron lasers might be useful but presently they
are too cumbersome and only generate small
amounts of energy per unit wavelength.
Fig. 1.1  A red LED (OmniLux, Phototherapeutics, Inc.)
A. A. Lloyd et al.

5
Intense pulsed light devices are becoming
increasingly comparable to lasers that emit ms
domain pulses [15]. Absorption spectra of skin
chromophores show multiple peaks (HgB) or can
be broad (melanin) [16], and therefore a broad-
band light source is a logical alternative to lasers.
Proper filtration of a xenon lamp tailors the out-
put spectrum for a particular application. Some
concessions are made with direct use of lamp-
light. For example, rapid beam divergence
obliges that the lamp source be near the skin sur-
face. This subsequent requirement makes for a
typically heavier handpiece compared with most
lasers (Fig. 1.2) (the exception being some diode
arrays where the light source is also housed in the
handpiece-(i.e., Light Sheer, Lumenis, CA)).
Also IPL cannot be adapted to fibers for subsur-
face
delivery.
High
energy
short
pulses
(Q-switched ns pulses) are not possible with
flashlamps. They can, however, be used to pump
a laser, and some modern IPLs feature a laser
attachment where the flashlamp and laser rod are
in the handpiece. In general, the size, weight, and
cost of both laser and flashlamp technology are
steadily decreasing.
Light Device Terminology
Basic parameters for light sources are power,
time, and spot size for continuous wave lasers,
and for pulsed sources, the energy per pulse,
pulse duration, spot size, fluence, repetition rate,
and the total number of pulses [17]. Energy is
measured in joules (J). The amount of energy
delivered per unit area is the fluence, sometimes
called the dose or radiant exposure, given in
J/cm2. The rate of energy delivery is called power,
measured in watts (W). One watt is one joule per
second (W = J/s). The power delivered per unit
area is called the irradiance or power density,
usually given in W/cm2. Laser exposure duration
(called pulse width for pulsed lasers) is the time
over which energy is delivered. Fluence is equal
to the irradiance times the exposure duration
[10]. Power density is a critical parameter, for it
a
b
Fig. 1.2  IPL and green light laser—note smaller size of laser handpiece
1  Laser-Tissue Interactions

6
often determines the action mechanism in cuta-
neous applications. For example, a very low irra-
diance emission (typical range of 2–10 mW/cm2)
does not heat tissue and is associated with diag-
nostic applications, photochemical processes,
and biostimulation. On the other extreme, a very
short ns pulse can generate high peak power den-
sities associated with shock waves and even
plasma formation [18]. Plasma is a “spark” due
to ionization of matter.
Another factor is the laser exposure spot size
(which can greatly affect the beam strength inside
the skin). Other characteristics of importance are
whether the incident light is convergent, diver-
gent, or diffuse, and the uniformity of irradiance
over the exposure area (spatial beam profile). The
pulse profile, that is, the character of the pulse
shapes in time (instantaneous power versus time)
also affects the tissue response [19].
Many lasers in dermatology are pulsed, and
the user interface shows pulse duration, fluence,
spot size and fluence. Some multi-wavelength
lasers also allow for wavelength selection. Some
older lasers, for example a popular CO2 laser,
showed only the pulse energy on the instrument
panel, or in continuous wave (CW) mode, the
number of watts. In these cases one uses the
exposure area and exposure time to calculate the
total light dose (fluence).

Fluence
Power
time
area
=
´

(1.2)
With the exception of PDT sources and CW
CO2 lasers, most aesthetic lasers create pulsed
light. In many CW applications (i.e., wart treat-
ment with a CO2 laser), the fluence is not of great
importance in characterizing the overall tissue
effect. A more important parameter is power den-
sity (where higher power densities achieve abla-
tion and lower power densities cause charring),
and the physician stops the procedure when an
appropriate endpoint is reached. On the other
hand, in PDT applications with CW light where
the clinical endpoint might be delayed, the total
fluence and power density are important predic-
tors of the tissue response.
In CW mode, CO2 lasers are used with a
focusing (noncollimated) handpiece such that the
physician can control spot size and tissue effects
simply by moving the handpiece tip toward or
away from the skin. The subsequent rapid
changes in power density offer “on the fly” flexi-
bility and control.
A thorough knowledge of a specific laser’s
operation and quirks is imperative for optimal and
“safe” lasering. Vendors are creating lasers that
are more intuitive to operate. Increasingly, manu-
facturers have added touch screen interfaces with
application-driven menus and skin-­type specific
preset parameters. Some devices permit patient
laser parameters to be stored for future reference.
Most lasers are designed such that the handpiece
and instrument panels are electronically inter-
faced. It follows that the laser control module
“knows” what spot size is being used. Typically
this “handshake” occurs when one inserts the
handpiece into the calibration port, or through a
control cable from the handpiece to the laser
“main frame”. With others, one selects the spot-
size on the display, and the laser calculates the flu-
ence accordingly. For example, one of our erbium
YAG lasers possesses interchangeable lenses for
1, 3, 5, and 7 mm spots. However, there is no feed-
back from the handpiece to the laser control
board. The user “tells” the laser which lens cell is
inserted, and the laser calculates the fluence based
on the selected spot and selected pulse energy. In
this case, if one changes the spot size (for exam-
ple, by exchanging the 7 mm for the 3 mm lens
cell), the laser still “thinks” the 7 mm spot is being
used, and the actual surface fluence is now ~5×
the fluence on the panel. The resulting impact on
the skin surface (the wound depth and diameter)
should alert the enlightened user to reassess his
parameter selection.
Most lasers calibrate through a system where
the end of the handpiece is placed in a portal on
the base unit (Fig. 1.3). This configuration allows
for interrogation of the entire system, from the
“pumping” lamps to the fiber/articulated arm to
the handpiece optics. For example, if a fiber is
damaged, the laser will fail calibration, and an
error message appears. Other systems measure
the output within the distal end of the handpiece
using a small calibration module that “picks off”
a portion of the beam.
A. A. Lloyd et al.

7
There are some simple ways to interrogate for
system integrity. One can examine the aiming
beam as it illuminates a piece of white paper,
checking that the beam edges are sharp—this
suggests that the treatment beam is also sharp and
the profile is according to the manufacturer’s
specifications. Also, burn paper can be used—
here the laser is used with a low energy and the
spot is checked for uniformity from beam edge to
edge. By checking the impact pattern, one can
uncover damaged mirrors in the knuckle of the
articulated arm, or a damaged focusing lens that
renders the laser unstable or unsafe. Likewise, for
scanners, one can ensure that skin coverage will
be uniform.
1. LEDs are becoming commonplace in biomed-
ical applications
2. Solid state lasers generally achieve the largest
peak powers among laser types
3. The laser operator should know every nook
and cranny of a laser’s features to optimize
patient outcomes and safety
4. Power density determines the mechanism for
many LTIs
Beam Profiles: Top Hat Versus
Gaussian
Laser beam profiles vary based on intercavity
design, lasing medium, and the delivery system.
A common profile is Gaussian or bell-shaped.
For many lasers, this profile represents the fun-
damental optimized “mode” of the laser. This
shape is usually observed when the beam has
been delivered through an articulated arm. For
some wavelengths, this is an effective way to
deliver energy (CO2 and erbium). The disadvan-
tage of the rigid arm is limited flexibility, the
typically short arm length, the possibility of mis-
alignment from even minor impact, and a ten-
dency for non-­uniform heating across the spot
[20]. For example, in treating a lentigo with a
Q-switched alexandrite laser equipped with a
rigid articulated arm, one may observe complete
ablation of the epidermis at the center of the
“spot”, but only whitening at the periphery. On
the other hand, sometimes a bell-shaped profile
is desirable, for example, when applying a small
spot FIR beam with a scanner. In this scenario,
the wings of the beam allows for some overlap
without delivering “too much” energy at points
of overlap.
The Gaussian profile can be modified outside
the cavity, which is desirable in many applica-
tions. With a fiber equipped delivery system, the
beam is mixed within the fiber and can be shaped
to be more flat-topped. The lentigo then is more
likely to be uniformly heated (so long as the
lesion itself if uniformly colored!). Although
fiber delivery systems are usually preferred by
physicians, some laser beams are difficult to
deliver through a fiber. Examples include far IR
wavelengths and very short pulses (i.e., few ns
with typical Q switched Nd YAG lasers whose
high peak power exceeds the damage threshold
of most fibers).
a
b
Calibration port
Handpiece
tip
Fig. 1.3  Figures show handpiece before and during insertion into calibration port of a Q switched alexandrite laser
1  Laser-Tissue Interactions

8
Pulse Profiles: Square Versus Spiky
The pulse profile is the temporal shape of the
laser pulse (Fig. 1.4) [21]. In many pulsed laser
applications, the “macro pulse” is comprised of
several shorter micropulses [22]. Depending on
the application, the temporal pulse profile may
impact the tissue effect. For example, simply by
increasing the pulse number from four to six
pulselets, the purpura threshold is increased with
the PDL.  Also, highly energetic spikes tend to
increase the epidermal to dermal damage ratio in
applications such as laser hair reduction. This is
especially true with green-yellow light in vascu-
lar applications.
Summary of Wavelength Ranges
In this section we examine wavelength ranges
that are useful for cutaneous surgery.
1. UV laser and light sources have been used pri-
marily for treatment of inflammatory skin dis-
eases and/or vitiligo, as well as striae. The
presumed action is immunomodulatory. The
XeCl excimer laser emits at 308 nm, near the
peak action spectrum for psoriasis. Other UV
non-laser sources have also been used for
hypopigmentation, striae, and various inflam-
matory diseases [22, 23]. Excimer lasers at
193 nm have been used for skin and corneal
ablation.
2. Violet IPL emissions, low power 410  nm
LED, and fluorescent lamps are used either
alone or with ALA. Alone, the devices take
advantage of endogenous porphyrins and kill
P. acnes [24]. After application, of ALA, this
wavelength range is highly effective in creat-
ing singlet O2 after absorption by PpIX. Uses
include treatment of actinic keratoses, actinic
cheilitis, and basal cell carcinomas [25].
3. Visible light (green yellow) - VIS (GY). These
wavelengths are highly absorbed by HgB and
melanin and are especially useful in treating
epidermal pigmented lesions and superficial
vessels [26–28]. Their relatively poor penetra-
tion in skin (and the even poorer penetration
in blood—see Table  1.1) make them poor
choices for treatment of deeper pigmented
lesions or deeper larger vessels. Their shallow
penetration depths preclude their use in per-
manent hair reduction (with the possible
exception of very large spots (i.e., IPL) that
enhance light depth). The effective portions of
many IPL spectra include the GY range.
By the proper manipulation a laser delivery
device, one can optimize parameters for selec-
tive heating of pigmented versus vascular
lesions. For example, by applying a compres-
sion handpiece without cooling with 595 nm,
blood is depleted as a target and pigment is
Palomar smooth pulse technology
Competing system
Power, au; ∆ Tepi, °C
Power, au; ∆ Tepi, °C
30
25
20
15
10
5
0
30
25
20
15
10
5
0
–5
Time, ms
Time, ms
0
5
10
15
20
25
–5
0
5
10
15
20
25
Skin safety level
Skin safety level
Skin temperature
Skin temperature
Fig. 1.4  Figure shows spiky versus smooth pulse and effect on epidermal temperature
A. A. Lloyd et al.

9
preferentially heated [30]. Also, by (see lime-
light desert mode—Fig. 1.5), one can increase
or decrease the sapphire window temperature
to enhance epidermal versus vascular heating.
By reducing the pulsewidth into the nanosec-
ond range, melanosomes are preferentially
heated over vessels. For example, extremely
short Q-switched 532  nm pulses will cause
fine vessels to rupture, but inadequate heat dif-
fusion to the vessel wall precludes long term
vessel destruction. On the other hand, melano-
somes are sufficiently heated for single-­
session lentigo destruction. By choosing
specific wavelengths with respect to HgB and
melanin, one can achieve some degree of
selective melanin or HgB heating. For exam-
ple, if one wanted to avoid HgB in heating a
lentigo, 694  nm (ruby) represents a better
choice than 532 or 595 nm. This choice might
decrease inflammation by unintended heating
of normal vessels in the dermis.
There are absorption peaks for PpIX in the
green–yellow range, making these wave-
lengths useful for PDT (i.e., sodium lamp,
IPL, frequency doubled Nd YAG, or PDL)
[31, 32]. On the other hand, all visible light
can be used for PDT, as the Soret band and
smaller “Q-bands” can all create singlet O2 on
irradiation of PpIX. Therefore the 532, 595,
and IPL devices, when used adjunctively with
ALA, can all augment the cosmetic result
through both photothermal and photochemi-
cal effects.
4. Red and Near IR (I) (630, 694, 755, 810 nm).
Deeply penetrating red light (630 nm) contin-
uous wave devices are efficient activators of
PpIX after topical application of ALA.  The
694 nm (ruby) laser is optimized for pigment
reduction and hair reduction in lighter skin
types. The 810 nm diode and 755 nm alexan-
drite laser, depending on spot size, cooling,
pulse duration, and fluence can be configured
to optimize outcomes for hair reduction, len-
tigines, or blood vessels [33]. They are posi-
tioned in the absorption spectrum for blood
and melanin between the GY wavelengths and
Table 1.1  Absorption coefficients (cm−1) for various chromophores
Wavelength (nm)
410
532
595
694
755
810
940
1064
Oxy HgB (40% Hct)
1990
187
35
1.2
2.3
3.6
5.2
2.2
Deoxy HgB
1296
138
96
6.6
5.2
2.7
3.0
0.6
Melanina
140
56
38
23
17
13
7
5.7
Water
6.7 × 10−5
0.00044
0.0017
0.005
0.03
0.02
0.27
0.15
Bloodless dermis
10
3
2
1.2
0.8
0.6
0.5
0.4
OPD in skin (μm)
100
350
550
750
1000
1200
1500
1700
a(net epidermis) for moderately pigmented adult—10% mel. volume fraction in epidermis [7, 29]
a
b
In “sun” mode T
is increased
Fig. 1.5  Figure shows user controllable temperature change with an IPL. By increasing the handpiece tip temperature,
pigmented lesion heating is favored over vascular heating
1  Laser-Tissue Interactions

10
1064 nm. They will penetrate deeply enough
in blood to coagulate vessels up to 2 mm [34–
36]; also, they are reasonably tolerant of epi-
dermal pigment in hair reduction (with surface
cooling) so long as very dark skin is not
treated [37]. By decreasing the pulsewidth
into the nanosecond range, the alexandrite
laser is a first line treatment for many tattoo
colors.
5. Near IR (II). 940 and Nd YAG (1064  nm).
These two wavelengths have been used for a
broad range of vessel sizes on the leg and face
[38]. They occupy a unique place in the
absorption spectrum of our “big 3” chromo-
phores, that is blood, melanin, and water.
Because of the depth of penetration (on the
order of mm), they are especially useful for
hair reduction and coagulation of deeper
blood vessels. By varying fluence and spot
size, reticular ectatic veins, as well as those
associated with nodular port wine stains or
hemangiomas, can be safely targeted. On the
other hand, they are not well suited for
epidermal-­pigmented lesions. Also, although
water absorption is poor, it exceeds that of the
VIS and near VIS wavelengths. The result is
that 940 and 1064 nm achieve large volume
temperature elevation in the skin, and with
repeated laser impacts, because of the slow
cooling of this volume (large τ–vide infra),
catastrophic pan-cutaneous thermal damage is
possible. This wavelength (1064  nm) repre-
sents the extreme example of a “what you
don’t see can hurt you laser” [18] (Fig. 1.6).
The Q-switched YAG laser plays an expand-
ing role in the treatment of tattoos, nevus of
Ota, and even melasma.
6. MIR-lasers and deeply penetrating halogen
lamps. These lasers and lamps heat tissue
water. With macrowounding (>1  mm) spot
sizes, depending on where we want to heat, we
can “choreograph” our laser and/or cooling
settings to maximize the skin temperature in
certain skin layers. In general, with more
deeply penetrating wavelengths, larger vol-
umes are heated. On the other hand, achieving
temperature elevations in the volume will
require higher fluences than with highly
absorbing wavelengths. Without surface cool-
ing, unless very small fluences are applied, a
top to bottom thermal injury occurs. The
absorption coefficients for the 1320  nm,
1450 nm, and 1540 nm systems are ~3 cm−1,
20 cm−1, and 8 cm−1 respectively [39], while
the effective scattering coefficients are about
14, 12, and 11 cm−1. The corresponding pene-
tration depths are ~1500, 300, and 700 μm. It
follows that for equal surface cooling and
equal fluences, the most superficial heating
will occur with the 1450 nm laser, followed by
the 1540 and 1320 nm lasers. Newer deeply
penetrating lamps have been introduced (Titan,
Cutera, Brisbane, CA). They emit light over a
1–2 μm wavelength band with relatively low
power densities and long exposures (several
seconds). In a typical scenario, the irradiation
begins after a roughly 2 s period of cooling. At
this point, a band of tissue from roughly 700–
1.5 μm deep in the skin is heated. By varying
the fluence, this relatively large volume can be
heated to different peak temperatures. As part
Fig. 1.6  Note scar after 1064 nm irradiation of a nodule
within a port wine stain
A. A. Lloyd et al.

11
of each iteration, post pulse cooling is impera-
tive because such a large volume of skin is
heated that a “thermal wake” advances toward
the skin surface. If one removes the handpiece
prematurely, heat accumulates near the skin
surface with the risks of pain, dermal thermal
injury, and scarring [40]. The 1320  nm
Nd:YAG has been used in the endovenous
ablation of the deep saphenous venous system
as well as laser liposculpture. Recently the
MIR spectral subset has become the mainstay
for fractional non-ablative technologies.
7. Far infrared systems. The major lasers are the
CO2, erbium YAG, and erbium YSGG
(chromium:yttrium-scandiumgallium-garnet)
lasers. Overall, the ratio of ablation to heating
is much higher with the erbium YAG laser.
However, one can enhance the thermal effects
of the Er YAG laser by extending the pulse or
increasing the repetition rate, and likewise one
can decrease residual thermal damage (RTD)
of the CO2 laser by decreasing pw [41, 42].
Where precision is required in ablation, Er
YAG is preferred. On the other hand, depend-
ing on settings, the CO2 laser enjoys a desir-
able blend of ablation and heating. The
thresholds for ablation for CO2 and erbium
lasers vary inversely with their optical pene-
tration depths in tissue (20  μm and 1  μm
respectively). This assumes thermal confine-
ment. It follows that less surface fluence is
required for ablation with the erbium laser.
The CO2 laser at typical operating “pulsed”
parameters performs self-limited controlled
heating of the skin [43, 44], whereas the
erbium laser operates in an almost purely
ablative regime. The erbium YSGG (2.79 μm)
laser has recently been applied to LSR and its
absorption coefficient makes it a kind of
hybrid between CO2 and erbium YAG insofar
as the ratios of heating to ablation. All three
wavelengths (2.79, 2.94, and 10.6 μm) have
recently been integrated into fractional deliv-
ery systems.
There are four key components in the sequence
of most photothermal laser-tissue interactions
(Sects. 7.1–7.4).
Beam Propagation: How the Laser
Energy Gets to the Target
Skin optical properties determine the penetration,
absorption, and internal dosimetry of laser light.
The laser surgeon can divide the skin into two
components, (1) the epidermis (primarily an
absorber of visible light due to melanin) and (2)
The dermis (which can be envisioned as a carton
of milk with red dots in it). Light tissue interac-
tions can be broken down into A. The transport of
light in tissue, B.  Absorption of light and heat
generation in tissue, C.  Localized temperature
elevation in the target tissue (and denaturation of
proteins), and D. Heat diffusion away from the
target [17, 45].
The optical properties of the skin mimic a tur-
bid medium intermixed with focal discrete visi-
ble and infrared light absorbers (blood, melanin,
bilirubin, and dry collagen) [46]. The thermal or
photochemical effects depend on the local energy
density at the target. Once the light penetrates the
surface, it undergoes a series of absorbing and
scattering events. Photons statistically are either
scattered or absorbed in a wavelength dependent
fashion [1, 47]. Scattering is affected by the shape
or size of the particle and the index of refraction
mismatch between the particle and medium. For
most tissues, for λ > 2.5 μm or < 250 nm, absorp-
tion dominates over scattering. For the remainder
of the EM spectrum, scattering is the primary
attenuator of light in tissue with the exception of
focal discrete absorbers (melanosome, HgB, etc.)
The probabilities of absorption or scattering
(designated μa and μs respectively, Table 1.1) are
determined by experiment. For example, for a μa
of 0.3 cm−1, the mean free path before absorption
is 1/μa or 3.3 cm. Generally, light is attenuated as
it propagates through tissue. In turbid tissue (i.e.,
the dermis, where collagen acts as the major scat-
terer), the fluence attenuation can be described
by:

I z
I ke
o
z
( ) =
-
(
)
/d

(1.3)
where I(z) is the local subsurface fluence at
some depth z, k is a constant that accounts for
backscattered light and δ is the wavelength
1  Laser-Tissue Interactions

12
dependent optical penetration depth of light, or
the depth at which there is attenuation to 37% of
the surface value (37% = l/e, where e = 2.7, the
base of the natural logarithm). This depth is
determined by absorption and scattering coeffi-
cients, as related by the simple equation below
[1, 47]:

d
m
m
m
=
+
-
(
)
(
)
1
3
1
a
a
s
g

(1.4)
Where g is the anisotropy coefficient (a mea-
sure of the “mean” direction of the scattered pho-
tons). g = 0.9 for the skin. As μa and μs increase, δ
decreases accordingly. For example, for hair
removal, based solely on depth of penetration,
longer wavelengths such as 800 and 1064  nm
should be preferable to 694 and 755 nm. In the
visible light range, this is why red light can pen-
etrate one’s hand when shining a flash light on
the surface. Scattering decreases roughly propor-
tional to λ3/2, so that, for example, an 800-nm
photon will on average travel about 1.3 times as
far in tissue as a 700-nm photon without being
scattered. It follows that for “more” scattering
wavelengths, there will be greater accumulation
of photons near the surface. In addition to scatter-
ing, this superficial convergence of photons is
based on index of refraction mismatches between
air and tissue [1]. Accordingly, light must be
deposited more slowly with shorter wavelengths
to avoid overheating the superficial tissue.
There is backscattered light that can yield a
higher fluence beneath the tissue than at the tis-
sue surface [48]. This paradox of tissue optics is
that the internal fluence can actually exceed that
at the surface, as below:

I
I
R
o
=
+
(
)
1
6

(1.5)
Where I is the subsurface energy density, Io is
the surface fluence, R = the surface remittance
(0.3, 0.6, and 0.7 for 585  nm, 694  nm, and
1064 nm respectively). (personal communication
from RR Anderson, 1994)
Since neither macromolecules nor water
strongly absorb in the red light and near IR (600–
1200 nm) this range allows deeper penetration in
most tissues [10]. Various skin pigments can play
optical “tricks” on the cutaneous surgeon. For
example, poikiloderma appears to be a mix of
hyperpigmentation and hypervascularity. In fact,
although there is some melanin influence in the
red-brown appearance, the dyschromia is by far
more a disorder of matted telangiectasia. This is
confirmed by the good response of the condition
to the PDL, even with aggressive surface cooling
that should preclude any impact on superficial
cutaneous hyperpigmentation. Additionally, with
diascopy, “poikilodermatous” skin often appears
no browner than the surrounding apparently nor-
mal skin. The explanation is that deoxy-Hb con-
tributes to a “pigmented skin appearance”. This
finding follows from the absorption spectrum of
deoxy-Hb in the 630–700  nm range, which is
very similar to the absorption spectrum of epider-
mal melanin. The size of the vessels in the super-
ficial venous plexus is such that the transmitted
light through these vessels is approximately 50%
lower than the incident intensity. These vessels
therefore appear dark [49].
In most biological systems, tissue constituents
show broad absorption bands with only a few dis-
tinct absorption peaks. From 200 to 290  nm
(UVC), all biological objects (cells and tissue)
absorb energy very strongly. From 290 to 320
(UVB) nm, only a limited number of biomole-
cules show absorption (aromatic amino acids and
nucleic acids). For UVA 320–400  nm, light is
weakly absorbed by colorless skin parts. From
400 to 1000  nm mainly pigments—bilirubin,
blood, and melanin absorb light. The heterogene-
ity of the skin allows for discrete heating over
this range, and therefore selective photothermol-
ysis (SPT) is exploited in this band. For
>1100 nm, all biomolecules have specific strong
vibrational absorption bands. Tissue water is the
primary determiner of the response to laser in this
wavelength range [9].
The absorption coefficient (μa) is the relative
“probability” per unit path length that a photon at
a particular wavelength will be absorbed. It is
therefore measured in units of 1/distance and is
typically designated μa, given as cm−1. The
absorption coefficient is chromophore and wave-
length dependent. For larger heterogeneous vol-
A. A. Lloyd et al.

13
umes, μa can be weighted according to the
fraction of a specific chromophore. For example,
for a dermis a typical blood fraction (f.blood) is
0.2%, assuming that the blood is uniformly dis-
tributed in the skin [7].
Following the descriptive convention of
describing an equivalent average homogeneous
f.blood, the net absorption of the dermis, μa.derm,
is calculated:
m
m
m
a
a
a
.
.
.
.
.
derm
f blood
blood
f blood
skinbaseline
= (
)(
)
+
-
(
)(
)
1

(1.6)
Scattering is responsible for much of light’s
behavior in skin (beam dispersion, spot size
effects, etc.). The dermis appears white because
of light scatter. The main scattering wavelengths
(relative to absorption) are between 400 and
1200 nm. Absorption occurs where the laser fre-
quency equals the natural frequency of the free
vibrations of the particles (absorption is associ-
ated with resonance) [50]. Scattering occurs at
frequencies not corresponding to those natural
frequencies of particles. Scattering is decreased
as wavelength increases [7].
There are four major chromophores (water,
blood, tattoo ink, and melanin) in cutaneous laser
medicine [50]. Water makes up about 65% of the
dermis and lower epidermis. There is some water
absorption in the UV. Between 400 and 800 nm,
water absorption is quite small (which is consis-
tent with our real world experience that light
propagates quite readily through a glass of water).
Beyond 800 nm, there is a small peak at 980 nm,
followed by larger peaks at 1480 and 10,600 nm.
The water absorption maximum is 2940  nm
(erbium YAG).
Hemoglobin: There is a large HgBO2 (oxyhe-
moglobin) peak at 415 nm, followed by smaller
peaks at 540 and 577 nm. An even smaller peak is
at 940  nm. For deoxyhemoglobin (HgB), the
peaks are at 430 and 555 nm. The discrete peaks
of hemoglobin absorption allow for selective ves-
sel heating. Although the 410 nm peak achieves
the greatest theoretical vascular to pigment dam-
age ratio among the other peaks, scattering is too
strong for violet light to be a viable option for
vascular applications.
Melanin: Most pigmented lesions result from
excessive melanin in the epidermis. By choosing
almost any wavelength (<800 nm), one can pref-
erentially heat epidermal melanin. Shorter wave-
lengths will create very high superficial epidermal
temperatures, whereas longer wavelengths tend
to bypass epidermal melanin (i.e., 1064 nm).
Fat: Fat shows strong absorption at 1200 and
1700 nm [51]. Although the ratios of fat to water
absorption are small, the small differences are
exploited with the proper choice of parameters.
1200 nm might represent the best choice due to
decreased overall water absorption and therefore
increased penetration. Sebum is similar to fat but
also is comprised of wax esters and squalene.
Carbon: Carbon is a product of prolonged
skin heating. Once carbon is formed at the skin
surface, the skin becomes “opaque” to most laser
wavelengths (that is, most energy will be
absorbed very superficially). It follows that the
dynamics of surface heating changes immedi-
ately once carbon is formed. This can be used
creatively as an advantage. For example, one can
convert a deeply penetrating laser to one that
would only affect the surface by using a carbon
dye. This has been accomplished with a laser peel
using a Q Switched Nd YAG laser.
Collagen: Dry collagen has absorption peaks
near 6 and 7 μm. With a free electron laser oper-
ating at these wavelengths, collagen can be
directly heated. Ellis et  al. found that this
approach might allow for less tissue irradiation
and less thermal damage than CO2 laser [52].
Heat Generation
Selective Photothermolysis (SPT)
Non-bulk skin heating is based on selective
absorption by discrete chromophores of rela-
tively low concentration (i.e., melanin, hemoglo-
bin). Dr. Leon Goldman showed that color
contrast allowed for selective damage of dermal
targets as early as 1963 [53]. However, it was Dr.
RR Anderson who elegantly described the con-
cept of selective photothermolysis [26]. Selective
photothermolysis offered a mathematically rigor-
ous rationale for tissue-selective lasers. As
1  Laser-Tissue Interactions

14
described by Dr. Anderson, extreme localized
heating relies on: (1) a wavelength that reaches
and is preferentially absorbed by the desired tar-
get structures; (2) an exposure duration less than
or equal to the time necessary for cooling of the
target structures; and (3) sufficient energy to
damage the target. The heterogeneity of the skin
allows for selective injury in microscopic targets.
The focal nature of the heating decreases the like-
lihood of catastrophic pancutaneous thermal
damage. For example, one can apply a 4  mm
laser beam and observe only a 1 mm wide tattoo
line “whiten” with Q switched Nd YAG laser
with a larger round spot (Fig. 1.7)—the skin out-
side the tattoo but within the spot will appear nor-
mal. Also, a darker lentigo will become white but
a lighter lentigo will remain unchanged. The pri-
mary areas where SPT is helpful in dermatology
is in the treatment of vascular lesions, tattoos,
and pigmented lesions. However, even in applica-
tions where water is the chromophore, the prin-
ciples of SPT are useful, as one can design precise
heating and ablation protocols based on wave-
length and pulse duration [54].
Thermal Relaxation Time
The thermal relaxation time (τ) is the interval
necessary for a target to cool to a certain percent-
age of its peak temperature [28]. Larger objects
require longer times than smaller volumes to
cool. For example, a tubful of warm bathwater
requires much longer than a thimbleful to cool to
room temperature. With laser irradiation, we
assume instantaneous heating of the target, so
that τ is the time for cooling after the pulse. If the
pulse is too long, the target cools during the
pulse, akin to one pouring water slowly into a
leaky bucket. If the water represents heat, one
observes that the bucket never fills (akin to a tar-
get never becoming very hot). If one wants to
spatially confine heating one chooses a short
pulse less than τ of the chromophore. For the
same volume, a sphere will cool faster than a cyl-
inder, which will cool faster than a slab. When
defining thermal relaxation time, the target size
and geometry are important. Normally, τ is
defined by:

d
k
2 / g

(1.7)
where δ is the optical penetration depth for
homogeneously absorbing layers (such as tissue
water for IR applications), and κ is the thermal
diffusivity (a measure of heat capacity and con-
ductivity—for tissue, κ ~ 1.3 × 10−3 cm2/s). For
discrete absorbers, i.e., the melanosome or a
blood vessel, τ is defined in terms of the particle
size, and δ represents the diameter of the particle.
κ is the thermal diffusivity, a quantity based on
the thermal conductivity and specific heat of the
medium, and “g” is a constant based on the
geometry of the target (slab, cylinder, or sphere)
[26]. See Table 1.2 for sample thermal relaxation
times for common targets in skin.
The often-used term “thermal relaxation time
of the skin” is meaningful only when used for
specific wavelengths (or specific skin structures,
i.e., the epidermis). With a ubiquitous absorber
such as tissue water, τ should be considered
within the context of the wavelength dependent
optical penetration depth (δ) and the laser source,
not the dimensions of the skin constituents. For
Fig. 1.7  Tattoo treated with 4 mm spot Q YAG laser; note
only 1 mm linear tattoo is heated (indicated by immediate
frosty whitening) 
Table 1.2  Thermal relaxation time of some potential
targets
Erythrocyte—2 μs
200 μm hair follicle—40 ms
Melanosome (0.5 μm)—0.25 μs
Nevus cell (10 μm)—0.1 ms
0.1 mm diameter vessel—10 ms
0.4 mm diameter vessel—80 ms
0.8 mm diameter vessel—300 ms
A. A. Lloyd et al.

15
example, if one uses the 1540 nm laser, the entire
epidermis and large portions of the dermis are
heated, and τ is on the order of seconds, because
δ is several hundred micrometer. So even though
τ of the epidermis is about 10 ms based on its
thickness, a thicker slab of skin is heated at
1540 nm, the epidermis will take several seconds
to cool because there is little temperature gradi-
ent between it and that of the dermis.
For most targets a simple rule can be used: the
thermal relaxation time in seconds is about equal
to the square of the target dimension in millime-
ters. Thus a 0.5 μm melanosome (5 × 10−4 mm)
should cool in about 25  ×  10−8  s, or 250  ns,
whereas a 0.1  mm PWS vessel should cool in
about 10−2 s, or 10 ms. Recall that τ is derived
from a solution of a differential equation and
does not represent an absolute cooling time, but
rather provides approximate pulsewidths for
varying degrees of thermal confinement [13].
Once the local subsurface energy density has
been determined (Eq. 3), heat generation can be
predicted by energy balance (conservation of
energy), pulse duration, thermal relaxation time,
and the wavelength specific absorption for that
target.
The temperature increase of a desired target
can be roughly calculated by knowing the absorp-
tion and scattering coefficients, surface light
dose, size of the target, and the length of the
pulse, as follows:

DT
F
c
z
a
r
r
p
g
=
+
æ
è
çç
ö
ø
÷÷
m
r
t
t
t
/2

(1.8)
where Fz is the local subsurface fluence, ρ is
the density, c is the specific heat “g” is a geomet-
ric factor (“1” for planes, “2” for cylinders, and
“3” for spheres), τp is the laser pulse duration,
and τr is the thermal relaxation time of the target
(time for target to cool to 37% of peak tempera-
ture), defined by Eq. 1.7. Thus one can perform
some quick algebraic calculations to estimate the
peak temperatures of local targets in the skin. The
temperature generally decays as a function of
diameter and time from the target. For ex, in hair
removal the shaft and bulb, heavily invested with
melanin, reach high temperatures, and as the
wave of heat diffuses from this cylinder, the tem-
perature decreases.
Spatially selective temperature elevation is
possible when (1) the absorption coefficient of
the target exceeds that of collateral tissue (selec-
tive photothermolysis), or (2) when the “innocent
bystander” tissues are cooled so their peak tem-
peratures do not exceed some damage threshold
or (3) with very small microwounds (fractional).
Localized heating, for example, in telangiectasia
and lentigines, follows from the concentrations
of blood and melanin there, respectively, such
that μa is focally increased. Verification of the
models can be made by real-time measurements,
thermocouple needles, thermal cameras, etc.
The geometry (and therefore the microscopic
characteristics) of lesions is important—for
example in the treatment for a nevus versus a len-
tigo, the nevus is composed of melanocytes in
aggregates as nodules (collectively the nodules
are often several hundred micrometer in diame-
ter) whereas the lentigo is a mere sheet of mela-
nocytes some 10  μm thick. For example, in
treating nevus with a long pulsed alexandrite
laser with a high fluence, the TRT will approach
a second. From the above equation, it follows that
thermal confinement will be high, and the peak
temperature will rise accordingly. More impor-
tantly, the thick slab of melanocytes will take
long to cool, such that the will be considerable
heat diffusion away from the target. On the other
hand, the lentigo represents a slab only tens of
microns thick; there will be heat diffusion during
the long pulse and rapid cooling after the pulse.
Thus, with ms-domain fluences, the nevus case
might result in scarring, and a lighter lentigo
might not become hot enough for clearance. If
one applies ns pulses to the two lesion types, the
lentigo shows a good response with possibly
complete clearing, whereas the nevus will require
multiple sessions, as each laser application will
result in heat confined to the most superficial part
of the lesion.
Two “offshoots” of SPT are the concepts of
thermal damage time and thermokinetic selectiv-
ity (TKS).
Thermal damage time. In some applications
the immediate absorber and the intended target
1  Laser-Tissue Interactions

16
are not collocated (i.e., hair shaft and hair bulb/
bulge). Thermal damage time is defined as the
pulsewidth that achieves irreversible target dam-
age with sparing of the surrounding tissue. The
thermal damage time represents the interval
when the outermost part of the target reaches the
target damage temperature through heat diffusion
from the heater. In this case the eventual target
and the heater (for example, hair shaft) are differ-
ent and at a considerable distance from each
other [55]. Using this model, the thermal damage
time can be many times longer than the thermal
relaxation time. For example, for laser hair
removal, with a 100 μm shaft and 30 μm follicle,
the TDT can approach several hundred millisec-
onds [55].
Thermokinetic selectivity: Along the same
lines is the concept of thermal kinetic selectivity
(TKS). Using this principle, one selects larger or
smaller targets for heating based on pulse dura-
tion. For example, if one wants to damage larger
targets while sparing relatively smaller ones, the
pulse duration is extended beyond the thermal
relaxation time of the smaller target. In this man-
ner, i.e., a melanosome will be heated to a lower
temperature than the subjacent vessel.
Molecular Basis of LTI
Most devices for cosmetic rejuvenation are based
on photothermal or “electrothermal” mecha-
nisms, that is, the conversion of light or electrical
energy to heat. Two fundamental processes gov-
ern all interactions of light with matter: absorp-
tion and scattering. Absorption and excitation are
necessary for all photobiologic effects and laser-­
tissue interactions. Energy is proportional to fre-
quency and inversely proportional to wavelength.
Thus a 532 nm photon (532 nm is the distance
between two of the transverse waves in a stream
of light) is twice as energetic as a 1064  nm
photon.
Macroscopically, the atomic events in LTIs
are not identifiable, but on the molecular level,
EMR exchanges energy only in discrete quanti-
ties (photons). The molecular basis of LTIs is
based on electronic transitions for the ultraviolet
(UV) and visible (VIS) wavelengths. For exam-
ple, hemoglobin is excited to a higher electronic
state, followed by a complete relaxation into
vibrational
modes
(internal
conversion).
However, NIR wavelengths and beyond are
absorbed via rotational and vibrational excita-
tions in biomolecules (all of which are hydrocar-
bons with the exception of pigments). These
reactions can be considered a two-step process.
In the first the molecule is “pumped” to an excited
state. Then, through a process known as non-­
radiative decay, there are inelastic collisions with
nearby molecules [50]. The temperature rise
results from the transfer of photon energy to
kinetic energy.
For thermal reactions to occur, the energy
must be randomized with a large ensemble of
molecules through statistical processes. With
HgB, the electronic excited state gives way to
vibrational modes. With longer wavelengths, the
quantized energy packets correlate with vibra-
tional transitions from NIR and MIR.
Reaction Types and Effects of Heating
• Photochemical effects (usually 10–1000  s;
10−3–10 W/cm2)
• Photothermal effects (1 ms–100 s; 1–106 W/
cm2)
• Photomechanical and photoionizing effects
(10 ps–100 ns; 108–1012 W/cm2)
Photothermal Effects
Photothermal processes depend on type and
degree of heating, from coagulation to vaporiza-
tion. With a very short pulsewidths (pw), lasers
vaporize targets. For example, in treating blood
vessels, rapid heating results in acute vessel wall
damage and petechial hemorrhage (with Q
switched 532  nm) [56–58]. With intermediate
length pulses (0.1–1.5 ms), one can gently heat
targets without immediate rupture of the vessels.
Still intravascular thrombosis can create purpura
and delayed hemorrhage. With still longer pulses
(6–100 ms), the ratio of contraction to thrombo-
sis increases and side effects are less likely. On
the other hand, too long pulses with very small
targets can create two problems. With highly
absorbing targets, (i.e., tattoo inks)—the heat
A. A. Lloyd et al.

17
generation is so great and long-lived that signifi-
cant diffusion occurs to the surrounding dermis
(Fig. 1.8). On the other hand, using a long pulse
YAG for a nevus of Ota results in an insufficient
temperature rise as the pigmented nevus cells
cool off too fast during the delivery of the pulses
(also melanin absorption is much weaker than
black ink).
A mild–moderate temperature increase results
in denaturation of enzymes and function. If the
heating is very fast, a phase change occurs [50,
59]. Depending on the rate of energy delivery,
photovaporization occurs with or without inertial
confinement (vide infra), where time is short
compared to the time for pressure relaxation.
Here the laser induced pressure causes compres-
sive stresses in tissue. Microcracks in the tissue
are the result of these large stress gradients [13].
Whitening after ns irradiation is thought be gas
vacuoles with scattering that resolve as the as the
“spaces” are refilled with interstitial fluid.
Hemoglobin undergoes a complex set of reac-
tions when heated. Formation of met HgB and
deoxy HgB occur during irradiation on the order
of ms, representing a real-time change in tissue
optics. During the heating phase, hemoglobin
absorption undergoes a bathochromic (red) shift
of the 580 nm absorption peak (the 540 nm peak
does not shift). This has important implications
for designing optimal blood vessel protocols. For
example, it has been suggested that 585  nm
though 600  nm should penetrate deeper than
shorter wavelengths and are therefore preferable
for treatment. In fact, as the blood temperature
rises during laser irradiation, 532 nm light trans-
mission increases and longer wavelength (i.e.
595 nm) transmission decreases [29, 60].
Even after the heating source is removed,
whole blood optical properties change. A macro-
scopic coagulum emerges comprised of dena-
tured HgB, cell membranes of erythrocytes and
plasma proteins. Met HgB formation is also
important, with absorption peaks near 630–
635  nm. With increasing met HgB and deoxy
production, 755–1064  nm lasers penetrate less
and will generate more heat in vessels. This is
one argument for sequential pulses with 1064 nm
lasers. One potential downside of this is that an
abrupt increase in absorption, particularly with
940–1064  nm, causes too much collateral
damage.
Thermal Injury to Cells
There is a range of measurable effects on skin
based on temperature. Below 43  °C, the skin
remains intact, even for very long exposures [45,
61]. The first change is a conformational change
in the molecular structure that occurs at tempera-
tures from 43 to 50  °C.  After several minutes,
there will be tissue necrosis as described by the
Arrhenius equation (an equation that quantita-
tively describes conversion of tissue from a native
to denatured state). Thermal denaturation is a rate
process: temperature increases the rate at which
molecules denature, depending on the specific
molecule. For example, at 45 °C, cultured human
fibroblasts die after about 20 min. However, the
same cells can withstand over 100 °C for 10−3 s.
In general, a temperature of >60° lasting for at
least 6 s leads to irreversible damage.
Coagulation
Temperature, directly related to the average
kinetic energy of molecules, is a critical factor in
tissue coagulation. Denaturation depends on time
and temperature, and at least for exposures >1 s,
conforms to a rate process as described by the
Arrhenius equation. The characteristic behavior
of the Arrhenius-type kinetic damage model is
that, below a threshold temperature, the rate of
Fig. 1.8  Figure shows extensive damage from long
pulsed 1064 nm irradiation of tattoo particle
1  Laser-Tissue Interactions

18
damage accumulation is negligible; and it
increases precipitously when this value is
exceeded. An example of coagulation is the cook-
ing of an egg white. Thermal denaturation is both
temperature and time dependent, yet it usually
shows an all or none like behavior. Most denatur-
ation reactions follow first order rate kinetics. For
a given heating time there is usually a narrow
temperature region above which complete dena-
turation occurs. As a rule, for denaturation of
most proteins, one must increase the temperature
by about 10 °C for every decade of decrease in
the heating time to achieve the same amount of
thermal coagulation [13].
An absolute temperature for coagulation-­
denaturation does not exist. For very short times,
higher temperatures than the oft-quoted “62–
65 °C” should be required. Early signs of micro-
scopic damage are vacuolization, nuclear
hyperchromasia and protein denaturation (recog-
nized as a birefringence loss for collagen).
Moderate temperature-induced damage phenom-
ena in tissue are difficult to assess with conven-
tional light microscopy. In fact, histology
represents and conveys the overall reactions of a
complex system and cannot be related to molecu-
lar species. Specimens obtained 24 h after irradi-
ation tend to be more sensitive than those obtained
immediately after treatment, as often a day is
required to show sign of necrosis; also, an inflam-
matory response might be the most sensitive indi-
cator of injury. Particularly in light of newer large
volume low intensity heating devices for rejuve-
nation, more sensitive tools might be required to
characterize subtle thermal effects. Beckham
et al. [62]. found that over a narrow temperature
range, heat shock protein (HSP) expression cor-
related with laser induced heat stress, and that the
HSP production followed the Arrhenius integral.
Thus HSP expression (in addition to tissue ultra-
structure, i.e., EM) might be an excellent tool to
examine low intensity high volume heat injury.
Vaporization
At a certain threshold power density, coagulation
gives way to photovaporization (ablation). Water
expands as it is converted to steam. Vaporization
is beneficial in that much of the heat is carried
away from the skin. The evaporation of tissue
water acts as a sort of buffer, reducing the peak T
to just over 100 °C. When there is vaporization
there is also increasing pressure as the water tries
to expand in volume. The expansion leads to
localized microexplosions. At the surface, parti-
cles are ejected at supersonic velocities. At tem-
peratures beyond 100  °C (without further
vaporization), carbonization takes place, which is
obvious by blackening of adjacent tissue and the
escape of smoke. Carbon is the ultimate end
product of all living tissues being heated and car-
bon temperatures often reach up to 300 °C. When
treating a wart at low power densities with the
CO2 laser, one can observe almost simultane-
ously incandescence and combustion. In water
free structures, such as char, temperatures can
reach 1000  °C, and incandescence can be
observed with continued irradiation of char at
long pulse cw CO2 lasers. Normally, this should
be avoided, because the depth of tissue injury
will extend well beyond the blackened skin sur-
face [50]. This is particularly true, for example,
when treating a rhinophyma or performing laser
skin resurfacing.
Photomechanical Effects
With very short pulses, there is insufficient time
for pressure relaxation. Mechanical damage is
observed with high-energy, submicrosecond
lasers for tattoo and pigmented lesion removal.
The time threshold for inertial confinement is
predicted by the relation [1]:

d / v
(1.9)
where δ is the target diameter and v is the
velocity of sound in tissue.
Inertially confined ablation occurs when there
is high-pressure at constant volume. In a very
short pulse, the energy is invested so quickly one
that there is no time for the pressure to be relieved.
Under these conditions of inertial confinement,
there’s not enough time for material to move—
this can lead to the generation of tremendous
pressures and relief through shock waves. For
example, one can feel the recoil during laser tat-
too treatment if one touches the skin surface near
the impact site.
A. A. Lloyd et al.

19
Photochemical Effects
The absorption of light in tissue does not always
generate heat. A term luminescence describes
cases where absorbed light causes emission of
light of a different color. This occurs after elec-
trons are excited from some lower (ground) state,
to an upper level, excited state. If the process is
fast, fluorescence occurs. If the is an intermediary
reaction and longer decay time, phosphorescence
occurs. Most photochemical reactions occur in
the UV and violet portion of the spectrum as the
electronic transitions demand highly energetic
photons.
Photochemical reactions are governed by spe-
cific reaction pathways, whereas thermal reac-
tions tend to be non specific. Photochemical
reactions include fluorescence, phosphorescence
and photodynamic action. The former two
involve the reemission of light after absorption at
lesser wavelengths. Fluorescence spectra are
increasingly used in diagnostic applications. An
example is PpIX fluorescence to determine the
amount of photosensitizer (PS) in the skin after
application of ALA.  Fluorescence spectra can
also be used to assess collagen content in the
skin. Also, keratin fluoresces such that a milium
will display a bright yellow color during irradia-
tion with a green light laser. We have occasion-
ally used this technique to distinguish a
deep-seated milium from a small pearly
BCC. Endogenous fluorescence applications are
increasing and have included assays of NADH,
collagen, and amino acids [63].
In photodynamic reactions, a photosensitizer
(acting as a catalyst) is excited by a certain wave-
length of light. The PS then undergoes several
sequential decays, forming singlet O2. In the
presence of oxygen, oxygen is transformed from
its triplet state, which is its normal ground state,
to an excited singlet state. The excited singlet
state oxygen reacts with biological molecules
and attacks plasma and intracellular membranes
(type II PDT reactions). The most common pho-
tosensitizer (PS) in dermatology is PpIX. This PS
is formed by skin cells by the pro-drug, ami-
nolevulinic acid (ALA). Most photochemical
reactions proceed more efficiently with lower
power densities, such that for example, the Blu-U
light (DUSA, Vahalla, NY) or Omni Lux
(Phototherapeutics, CA, USA) will outperform a
pulsed source (IPL, KTP, or PDL) for AKs with
one treatment. However, largely because most
practitioners posses at least one pulsed visible
light source, they have been widely used in PDT
and shown to be useful in a range of PDT-­
responsive skin disorders [64, 65]. Because
pulsed light sources in dermatology do not meet
the
theoretical
PDT
saturation
threshold
(4 × 108 W/m2) [66], in theory there should be
significant PDT activity [67].
Most photochemical activity is observed
between 320 and 630 nm. The main absorption
peaks for PpIX are 415, 504, 538, 576, and
630 nm. Beyond 800 nm, photochemistry, even
with exogenous photosensitizers or pro-drugs, is
unlikely. PDT reactions are complex; for exam-
ple, with ALA, optimization require an under-
standing of skin pharmacokinetics, conversion
kinetics of ALA to PpIX, and proper delivery of
light dose (including power density, wavelength,
etc.). A new development is a possible role for
vascular specific therapy with PDT.  Both with
hematoporphyrin and benzoporphyrin deriva-
tives, this approach is being investigated for
refractory deeper vascular lesions [68].
Photodisruption or Photodecomposition
This reaction type is usually observed at 107–
108 W/cm2 in the UV range. Usually there is little
residual thermal damage (RTD).
Rather than heating water directly as their FIR
counterparts, these lasers can exceed the bond
energies of many organic compounds. Among the
available wavelengths (193, 241 and 308  nm)
XeCl (308 nm) is more likely to be more thermal
because of the lowered energy per photon.
Excimer Laser (308 nm)
The mechanism of action for the excimer laser
(XeCl) is thought to be the same as narrow band
UV therapy. A reduction in cellular proliferation
most likely plays a role in epidermal cellular
DNA synthesis and mitosis. In Parrish’s original
study in 1981, he showed that wavelengths
between 300 and 313  nm were most effective
[69]. It appears that the excimer laser works
1  Laser-Tissue Interactions

20
through in an immunomodulatory way much like
standard non coherent UVB “light boxes”. There
may be a thermal component as well at fluences
>800 mJ/cm2.
Biostimulation
Biostimulation belongs to the group of photo-
chemical interactions. Increasingly this field is
validated by the number of well-executed investi-
gations. Typical fluences are in the range of
1–10 J/cm2, and temperature elevation is absent.
Potentially positive reactions include (1) Increase
in phagocytic activity (2) Depression in rate of
bacteria replication (3) Increase in repair of skin,
and (4) Stimulation of wound healing. We should
not debunk these applications—certainly they are
being studied more and we are witnessing a tran-
sition from lab to bedside. One example of a “bio-
stimulation” device is the Gentlewaves LED
Photomodulation unit (Light BioScience, LLC,
Virginia Beach, VA). This device uses 590  nm
light in a high repetition rate and low power den-
sity to increase collagen synthesis and enhance
facial tone. Cell culture work supports this con-
cept [70]. Newer LED devices (800 nm) also sup-
port a role for photorejuvenation and even hair
growth. It is unknown if any special features of
laser light are relevant for biostimulation, although
at least one study supports coherence as a possible
factor in photomodulation. Some investigators
have shown that laser showed the best effect while
the non-coherent LED light showed the poorest.
Coherence does not influence the transmission;
rather, because of interference in the scattered
light field, coherency influences the microscopic
light distribution into tissue.
Plasma Induced Ablation
With very high power densities exceeding 108 W/
cm2, optical breakdown occurs. Plasma removes
skin without evidence of mechanical or thermal
damage when choosing appropriate parameters.
Plasmas are sometimes produced by laser tattoo
removal, where one can observe a spark [13].
There is a new resurfacing system that uses
plasma created by an RF excited N2 gas. Unlike
laser-induced plasma, this flame heats tissue by
direct heat transfer from the plasma edge.
Heat Conduction Away
from the Chromophore
This final “physical” step in LTIs in important in
characterizing collateral damage. Once heat is
generated, heat losses are based on heat conduc-
tion, heat convection, or radiation. Radiation can
be neglected in most types of laser applications.
A good example of heat convection is transfer
from blood flow. Heat conduction is the primary
mechanism by which heat is transferred to unex-
posed tissue structures.
Cooling
Surface cooling enhances efficacy and safety in
skin laser surgery, especially for visible light tech-
nologies, (i.e., green–yellow light sources such
as IPL, KTP laser, and PDL) that are popular in
cutaneous laser surgery. They are also the wave-
length ranges where epidermal damage is most
likely. The epidermis is an innocent bystander in
cutaneous laser applications where the intended
targets, such as hair follicles or blood vessels, are
located in the dermis. Specifically, absorption of
light by epidermal melanin causes skin surface
heating.
The first goal is of surface cooling is preserva-
tion of the epidermis. The second and related
goal of surface cooling permits higher fluences to
the intended target (i.e., the hair bulb and/or
bulge or a subsurface blood vessel). Another ben-
efit of surface cooling is analgesia, as almost all
cooling strategies will provide some pain relief
[71–79].
The timing of the cooling relative to the
laser pulse is important. Cooling can be pre,
during the pulse (parallel), or after the pulse
(post) [79]. Post cooling may prevent retro-
grade heating (i.e., from the vessel back to the
epidermis) from damaging the skin surface. A
cooling protection factor (CPF) has been pro-
posed by Dr. Rox Anderson. The cooling pro-
tection factor is the ratio of fluence, with and
without surface cooling, that spares the epider-
mis. It is defined by:

CPF
T
T
T
T
c
ic
c
i
=
-
-

(1.10)
A. A. Lloyd et al.

21
In the above equation, Tic and Ti are basal layer
temperatures before laser irradiation with and
without cooling, respectively. Tc is the critical
temperature at which thermal injury occurs. The
detailed calculations described later indicate that
if the initial skin temperature is 30 °C, contact
cooling reduces the temperature of the basal layer
to about 20 °C. If Tc is assumed to be 60 °C (it is
actually somewhat higher for the brief laser
exposure times in this analysis), this would give
the CPF as (60–20)/(60–30) or 1.33. Similarly,
cryogen cooling reduces the temperature to about
0 °C, thus giving a CPF value as (60–0)/(60–30)
or 2.0. Finally, there is convective air cooling,
where cold air is commonly used in skin chilling.
The Zimmer (Cyro5, Zimmer Medizin Systeme,
Ulm, Germany) directs −10 °C air at the skin at a
rapid rate (1000 L/min). This system proves for
good bulk cooling but spatial localization of the
cooling is poor. The CPF, depending on the air
temperature and nozzle velocity, is near that of
contact cooling.
Some Interesting Concepts
and Ideas in Laser Tissue
Interactions
With Q switched lasers, one might hear a loud
“pop” accompanied by a spark-like emission at
the skin surface. Normally this is the result of
inorganic compounds (make-up) remaining in
the skin during irradiation. Thorough cleaning of
the skin will remove this distraction. One should
ensure that any dark markers are off the skin,
especially when using long pulsed visible light
technologies. For example, one of our trainees
placed a black marking pen to outline an area for
a test spot for laser hair removal. During irradia-
tion, the beam was absorbed almost completely
by the ink—the hot ink cooled at the expense of
the skin surface, such that a very superficial burn
occurred.
Also, one should consider oral medications
both in the genesis of treatable lesions (i.e.,
­minocycline hyperpigmentation), and in the cau-
sation of pigmentation disturbances (i.e., gold).
In the case of gold, the medication is sequestered
in macrophages of the dermis—the combination
of gold and Q-switched lasers produces a
photothermal-­photochemical conversion such
that the gold darkens to a light blue or grey color
(Fig. 1.9). This reaction is a good teaching tool in
that it points out the role of pulse duration on the
laser tissue interaction [80, 81]. As noted earlier,
some reactions are dependent on power den-
sity—with higher power densities, multi-photon
interactions are possible, that is, the energy is
condensed into such a short duration, that simul-
taneous “arrival” of two photons at the same
locale can result in two-photon absorption. In the
case of gold, the chemical compound structure
can be changed (from crystalline to elemental).
Once the reaction occurs, one can apply longer
pulses to diminish the dyspigmentation (even
with the same wavelength!). This reaction also
underscores the importance of beam scattering,
as the “gold” Q-switched laser reaction extends
beyond the diameter of the beam with each pulse.
Fig. 1.9  Note blue macules
1  Laser-Tissue Interactions

22
Focusing the laser beam: A trick to increase
the dermal to epidermal damage ratio is use of a
convergent lens. This tool increases the local
radiant exposure in the dermis (targeting the hair
bulb, a blood vessel, or dermal water).
Theoretically, one should be able to use smaller
incident fluences, therefore achieving some pro-
tection of the epidermis.
Vacuuming the target in the laser beam: A
company (Aesthera, Livermore, CA) has created
a pneumatic device whereby the skin is vacu-
umed into the light path such that the light pene-
tration in skin is enhanced. In this way more
energetic high frequency photons can be deliv-
ered, for example, to the hair follicle, with rela-
tive epidermal sparing. By applying suction, the
absorption coefficient of the epidermis can be
reduced by up to 25%. The technologies have
also been used for acne and pain reduction.
By proper timing of the suction with respect to
irradiation, selective targeting of various chromo-
phores can be achieved, for example, to increase
the dermal blood fraction in pale PWS (and
increase the blood vessel diameter). The very
small vessels in paler PWS have too small vessel
diameters for thermal confinement—that is, the
vessels cool too quickly to reach a critical tem-
perature. By applying suction, the blood volume
fraction increases, not simply a result of the
mechanical force but a physiologic response as
well [82, 83].
Pixilated Injury (aka fractional photothermol-
ysis): One can use a “pixilated” injury with water
as a chromophore in what is called fractional
photothermolysis. Roughly 100  μm spots have
been used with 250–500 μm spacing [84]. The
tissue can recover from this fractional injury
without the widespread epidermal loss observed
after traditional resurfacing applications. A num-
ber of technologies have been introduced. Despite
a wide range of devices, the pitch, wound diam-
eter, wound depth, and other wound features have
not been optimized. Ideally one would design
devices that maximize downtime while maximiz-
ing cosmetic enhancement. One can consider
ablative and non ablative approaches. Early evi-
dence suggests that there is a difference between
ablative and non ablative wound healing even
with similar microvolumes of injury, that is, even
when the same total volume is observed, wound
healing proceeds differently.
In the most common approach, 75–150  μm
wide microwounds are created in the skin
(Fig.  1.10) with densities ranging from 100 to
1500  microwounds/cm2. By spatially confining
Fig. 1.10  Note damage pattern with 1540 nm microbeam
A. A. Lloyd et al.

23
the micro-lesions, deeper wounds can be created
than with a “slab-like” approach, while still man-
aging a larger measure of safety. There are both
ablative and nonablative approaches. Ablative
devices include the Profractional laser (Sciton),
equipped with a scanned microbeam, the Pixel
erbium YAG laser from Alma (Alma lasers,
Buffalo Grove, IL), and a newly introduced
2940  nm technology from Palomar. Reliant
Technologies manufactures a fractional CO2 laser
system (Re Pair) that creates 125  μm diameter
“ablative” wounds as deep as 1 mm. Early inves-
tigations have shown immediate superficial skin
tightening.
“Macrowound” fractional technologies create
wounds >300 μm in diameter. These include the
KTP laser with a scanner (with approximately
700 μm wounds) as well as the active FX CO2 sys-
tem (Lumenis, Santa Clara, CA), which creates an
array of 1.3  mm wounds and covers approxi-
mately 60% of the surface area per session.
Wound depths range from 80 to 150 μm depend-
ing on pulse energy. Fluences with these
approaches range from 5 to 15 J/cm2. The applied
fluences are another means (besides wound diam-
eter) to differentiate microwound injuries from
macrowound injuries. With ablative micro-
wounds, fluences tends to exceed 30× the ablation
threshold, whereas with traditional resurfacing
laser applications (CO2 and erbium) fluences
range from 0.8 to 10× ablation threshold per pass.
The original non ablative fractional laser was
(Reliant Technologies, Mountain View, CA),
deploying a 1550  nm scanned microbeam that
required a surface blue dye for proper tracking
along the skin. The newer Fraxel  technology
achieves deeper wounds and does not require the
dye. Palomar introduced a fractional 1540-nm sys-
tem. This device uses a “stamping” approach,
where each 10  mm macro-spot is comprised of
100 beamlets. With progressive passes, an increas-
ing skin surface area is covered. Another nonabla-
tive example is a 1440-nm/1320-nm Nd YAG laser
(Affirm, Cynosure, Chelmsford, MA) that delivers
hundreds of beamlets interspersed with a relatively
uniform low-fluence background irradiation.
After high-fluence fractional CO2 and erbium
YAG laser (50–200 J/cm2), one observes microcrat-
ers on routine histology. Particularly for the erbium
YAG laser, there is immediate water loss through
these portals of entry [85], and postoperative dis-
comfort is often severe for an hour after the proce-
dure. Pinpoint bleeding is sometime observed,
particularly with higher-­pulse energies and shorter
pulsed erbium YAG applications.
Optical damping: Replacing air (n = 1.0) with a
higherindex medium at the skin surface such as
glass (n = 1.5) or sapphire (n = 1.7) tends to spare
the epidermis. This effect has nothing to do with
heat transfer, but rather is a consequence of optical
scattering behavior. At wavelengths from about
600–1200 nm, most light in Caucasian epidermis is
back- and multiply-scattered light. By providing a
match to the skin’s refractive index, internal reflec-
tion of the back-scattered light is greatly reduced,
decreasing the natural convergence of photons at
the skin surface. This version of optical epidermal
sparing requires a physically thick external medium
such as a sapphire window or heavy layer of gel.
Compacting the Dermis
One can decrease the depth photons must propa-
gate by applying pressure over the treated area.
This maneuver may, for example, decrease the
relative depth of the bulb and bulge of the hair
follicle up to 30% relative to the skin surface.
Disadvantages include variability in the amount
of pressure, such that adjacent treatment areas are
exposed to different subsurface fluences. Also, it
is unclear if compacting the dermis might alter its
scattering properties. In theory compression
should decrease water content and improve der-
mal transmission [86].
Spot diameter: In general the spot size should
be 3–4× > δ (for wavelengths where scattering
dominates absorption), as larger spots make it
more likely that photons will be scattered back
into the incident collimated beam [13]. Photons
scattered out of the beam are essentially wasted.
Traveling “alone”, they carry insufficient energy
to cause macroscopic thermal responses. The
consequences of spot size can be explained best
on surface to volume arguments. Larger beams
(with the same surface fluence as smaller beams)
1  Laser-Tissue Interactions

24
create deeper subsurface cylinders of injury
because there is less surface versus volume for
photons to escape. Basically, for small beams
(narrow), scattered photons are carried out of the
beam path after only a few scattering events. As a
clinical example of the effect of spot size, we
have found for 3 mm vs. 6 mm spots with the
YAG laser that roughly ½ the fluence is required
with the larger spot for leg vein clearance. For
shallow penetrating lasers such as CO2 and
erbium where the δ ≪ spotsize (all cases except
for fractional devices), the diameter of the beam
does not affect the tissue response. That is why
equivalent results can be obtained for skin resur-
facing using pulsed CO2 lasers versus scanned,
tightly focused cw CO2 lasers [44]. Although
studies suggest that large spots increase the ratio
of dermal to epidermal damage (usually desir-
able, for example, when treating a hair bulb),
there are instances where small spots are desir-
able. For example, when treating a smaller vessel
with an Nd YAG laser, a small spot with higher
fluence will result in a higher percentage of the
energy being invested in vessel heating versus
larger spots. For any turbid medium, even if the
spot is “top hat”, there will be an accumulation of
photons near the center of the beam such that a
greater clinical effect will often be noted at the
center of the spot.
Changing optical properties in real-time:
Chromophore concentrations can change during a
treatment session. One should never consider each
laser tissue interaction as an independent event, but
rather a cumulative process where visual endpoints
are the most important ally for the physician.
Optical properties of the skin are like the weather
[3], and one must accommodate the changes in
real-time. For example, the dermal blood fraction
increases after one pass of the PDL, such that for a
second pass, the skin temperature will increase due
to the higher μa. The phenomenon will, for exam-
ple, lower the purpura threshold on a second pass.
On the other hand, general anesthesia can decrease
the blood flow in PWS and require a higher light
dose. A failure to respond to these real-time
changes accounts for many laser treatment short-
comings. In treating a PWS, tetracaine, for exam-
ple, can increase local blood flow, as can applying
heating pad or simply placing a patient in
Trendelenburg position. One of our patients actu-
ally performs jumping jacks prior to her rosacea
laser therapy to increase the response [87].
Most laser tissue interactions are threshold-­
based, that is, a certain amount of energy must be
invested over a specific time to achieve the
desired efficacy. For example, to lighten a lentigo
on the nose, even ten very–low-fluence passes, so
long as the interval between passes is long enough
to preclude cumulative heating, will not result in
clearance. The analogy is a smallish man trying
to push a car up a hill. Even if the man were to
arrive every day at 6 AM to push the heavy car,
the vehicle will remain stationary. There is no
incremental car movement each day. One “laser”
exception to this analogy is perhaps tissue tight-
ening and protein denaturation over large vol-
umes with complex molecules (i.e., collagen),
where repeated low impact low fluence passes
have been shown to increase the percentage of
denatured collagen fibers recruited in to the tight-
ening process. Part of this phenomenon might be
secondary to differential denaturation tempera-
tures of older versus younger fibers.
When treating vascular lesions, multiple
“low” fluence passes can achieve cumulative
improvement. For example, a second pass even
seconds after an initial pass with the YAG laser or
PDL will achieve additional bluing of an angi-
oma. The dynamics of vascular heating is some-
what different than for water and melanin. In
vascular applications, dynamic changes in blood
properties play a role. Met-Hg is produced by
one pass so that additional passes can result in an
increase in absorption. Also the partial clot
enhances absorption venous red blood.
With pigment lesions, repeated laser pulses
delivered over short periods (0.25–1 s) intervals
results in progressive graying or darkening of the
lesions. On the other hand, repeated passes (after
>1 min) will result in cumulative extent.
Both immediate and delayed pigmented dark-
ening (seconds to minute after irradiation) after
application is most likely due to optical property
changes in melanin as well as erythema deep to
A. A. Lloyd et al.

25
the lesions that might add to the darkening per-
ception of (Fig. 1.11).
Optical clearing with hyperosmolar solutions:
Transparency of the skin is enhanced by topical
application or intradermal injection of solutions
such as glycerin [88]. Water and collagen become
less bound such that the effective scattering coef-
ficient of the dermis is reduced. Already this con-
cept has been applied to increase the visibility of
blood vessels from the surface. Possible applica-
tions include tattoo removal, where particles
often are found several mm deep in tissue. More
recently Perfluorodecalin has been applied topi-
cally to accelerate the clearing of the immediate
tissue whitening response after tattoo removal.
Once the whitening response has diminished
(usually about 5 min after the application), a sec-
ond treatment can be applied in the same session
without the tissue scattering created by the first
pass [89].
“Carbonization” at the surface: Carbon will
cause all wavelengths to increase absorption such
that one can convert a deeply penetrating laser to
a more superficially penetrating laser by having a
fine carbon layer at the surface. For example, one
can “convert” a 694 nm ruby laser into a laser
with CO2 laser like effects by applying a fine
layer of graphite from a copy machine to the skin
surface. In this way the 694 nm laser energy is
confined to the surface by the almost 100%
absorption by carbon. This fine layer of heated
materiel then cools much like a superficial layer
of tissue heated by a CO2 laser alone.
Photon recycling: The remittance of human
skin is wavelength dependent (vide supra). These
reflected photons are scattered into the environ-
ment and “wasted” in surgical laser applications.
One can design a simple hemispherical reflector
to return reflected light to the incident spot on the
skin. In theory the gain in total energy available
to skin is a factor of
1
1-
(
)
R R
S
M
, where RS is the
skin reflectance, and RM is that of a hemispherical
mirror. For example, if RS is 0.7, and RM is 0.9, a
gain of
1
1
0 63
-
(
)
.
, or almost threefold, can be
achieved.
Photothermal responses in individual cells.
Most of our characterization of laser-tissue
responses is based on “macroscopic” responses.
That is, individual cells are rarely examined dur-
ing and after laser irradiation. When focal cell
damage has been examined, the following consid-
erations are made. (1) Heterogeneity of cell struc-
ture can lead to extreme localized light absorption
and temperature elevation different from that of a
homogenous medium. (2) Localized overheating
may cause cell damage, even in the absence of
average thermal effects over larger volumes [90].
After absorption of a laser pulse, non-radia-
tive relaxation of optical energy occurs within
10−11 s. Thus heating at the site of absorption is
instantaneous. On the other hand, heat diffusion
is much slower and characterized by the TRT. In
experiment, not unexpectedly, it was found that
that temperature fields in cells were more uni-
form with longer pulses. It follows that short
Darkening
after IPL
Whitening
after Q alex
Fig. 1.11  Figure shows immediate pigment response
after IPL and Q switched alexandrite laser
1  Laser-Tissue Interactions

26
pulses have smaller thermal fields but higher
localized T elevations. The shorter the laser
pulse, the more the final tissue response will
depend on the properties of the local absorbing
components. One interesting phenomenon is that
on a localized level, an initial thermal field does
not provide the maximum amplitude of the inte-
gral photothermal response inside a cell. Rather,
the T response reaches its maximum as a result
of the multiple secondary thermal fields as they
emerge.
Using a polarizing lamp to enhance illumina-
tion. Laser treatment can be enhanced by using a
polarizing lamp during procedures to treat vascu-
lar and pigmented lesions. This is particularly
helpful, for ex. when treating PWS in kids using
general anesthetic, the lamp is [91] helpful to
delineate the edges of the PWS prior to treat-
ment. Also, the visual enhancement tends to
result in more complete elimination of vessels,
therefore patients are more satisfied.
Selective cell targeting. A process called
selected cell targeting has been examined as a
way to destroy selected cells. This precise energy
deposition is achieved by using laser pulses and
light absorbing immunoconjugates tagged to the
respective cells. The investigators in one study
showed, for example, that lymphocytes could be
selectively damaged by attaching iron oxide mic-
roparticles absorbing 565 nm radiation at those
sites [92]. One can imagine, in the future, using
this type of modality to treat T-cell mediated dis-
eases such as atopic dermatitis or psoriasis. In
this way, one makes the “bad guy” more notice-
able to the laser.
Scatter limited therapy—using small micro-
beams. Reinisch [93] proposed the use of beam
diameter to titrate the depth of penetration, For
example, we have studied a fractional 1064 nm
ms laser (100  μm diameter microbeam and
100  mb/cm2) technology to achieve superficial
vessel heating with relative epidermal sparing
with just such a device to limit penetration into
the dermis. By using the aforementioned spot
size arguments, one can exploit the properties of
small spots to change the way particular wave-
lengths behave in the skin. For example, one can
tailor a 1064 nm laser to heat progressively larger
depths of skin by increasing the spot size.
1. LTIs are usually based on varying degrees of
light absorption by tissue HgB, melanin, and
water.
2. Wavelength ranges should be chosen to
achieve as much specificity as possible in tis-
sue heating
Radiofrequency (RF) Technology
With R  radiofrequency energy, local heat gener-
ation depends on the local electrical resistance
and current density. The distribution of the cur-
rent density is determined by the configuration of
the electrodes with respect to the skin anatomy.
There are multiple types of electrode deploy-
ments [94–97].
Monopolar vs. Bipolar
Radiofrequency energy induces tissue heating in
multiple ways, one by creating bulk heating of
the dermis while sparing the epidermis via cool-
ing mechanisms, and alternatively, micro needles
can be placed in the skin to deliver small coagula-
tive injuries [98, 99].
Since radiofrequency energy stimulates the
generation of new collagen and elastin, most
devices require at least three treatments and
effects are not typically fully seen until 1–3 months
after the last treatment. One benefit to using radio-
frequency devices is that the energy is colorblind
and passes through the melanin and hemoglobin
present in the tissue; therefore, all skin types can
be treated. One should note that there is attenua-
tion of the electrical field as a function of depth
(like light and lasers); however, the specific equa-
tions that guide the specifics of the attenuation are
beyond the scope of this chapter.
There are four broad types of RF interventions
in the skin. One is a monopolar system where
a capacitive coupling device is placed on the
surface and heat is delivered for a few ­seconds.
Simultaneously and just after the “pulse”, the
surface is cooled. The second is a bipolar sys-
tem comprised of metal rails on the skin surface
where the current is alternated in a rapid manner
(300 kHz–1 MHz) between the superficial skin
A. A. Lloyd et al.

27
layers. Generally, in this configuration, the depth
of the heating is roughly ½ the distance between
the electrodes. In the third type of intervention,
an array of needle or pin electrodes is placed in
the skin such that very focal heating zones are
created at predetermined depths. A final type
of RF heating is capacitive far field heating,
where a system of electrodes delivers energy at
27.12 MHz to create apoptosis in fat.
For monopolar radiofrequency, the delivery
electrode is in the hand piece and the return elec-
trode is placed elsewhere on the patient’s body
whereas with bipolar radiofrequency, the deliv-
ery and return electrodes are both incorporated
into the hand piece. Therefore, with monopolar
radiofrequency the area of heating is a column/
cylinder extending from the epidermis towards
the subcutaneous tissue. The electrical energy is
most concentrated near the tip of the delivery
electrode and decreases rapidly with distance,
with the penetration depth about half the size of
the delivery electrode. However, the behavior of
the current as it passes through the body to the
return electrode is somewhat unpredictable.
Monopolar skin rejuvenation systems create
large-volume heating. Electrical energy is distrib-
uted uniformly over the electrode surface through
“capacitive coupling”. This type of coupling
reduces the natural accumulation of electrical
energy at the electrode edge [100]. The first non-­
ablative RF device (Therma Cool TC, Solta
Medical, Hayward, CA) uses cryogen spray cool-
ing (CSC), where the spray is started before the
RF current. Interestingly, if an electric field is
induced perpendicular to the skin-fat interface, a
monopolar device can selectively heat large areas
of fat while sparing the skin and muscle [98].
With bipolar radiofrequency, a U shaped area
of heating is created between the delivery and
return electrodes in the hand piece and thus the
current travels a fixed distance. The depth of the
“U” is limited to one half of the fixed distance
between the electrodes. Therefore, the distribution
of, as well as the location of the radiofrequency
current, is controlled and predictable within the
tissue. In one scenario, cooled bipolar electrodes
are combined with a diode laser, halogen lamp, or
intense pulsed light device. In this configuration,
there is synergy between the two energy sources
[99]. The theory predicts that once the selectively
targeted chromophores are heated by the visible
light, their impedance decreases and the subse-
quent RF energy will preferentially heat those tar-
geted tissues (i.e. hair and vessels). A purported
advantage of the treatment is that lower optical
energies can be used to selectively heat sub-sur-
face targets than if a light source were used alone
(thus enhancing epidermal preservation).
In microneedling fractional radiofrequency
devices, the needles become the delivery and return
electrodes and the electric power and current is
divided among the needles. In this configuration,
the device fractionates the radiofrequency energy
among several delivery and return ­electrodes. With
uninsulated needles as electrodes, the entire needle
conducts heat, and wounds are created along the
entire length of the needle. Insulated needles only
allow the tip of the needle to heat and therefore
the epidermis is preserved. There are also devices
where the needles (or “pins”) are deployed very
superficially (only down to 100–300 μm) where
both fine lines and pigment can be reduced with
multiple treatment sessions.
How RF Creates Dermal Heating
Skin has inherent impedance, which is the resis-
tance to electric current, and when an electric
current meets resistance it generates heat in
accordance with Joule’s law,
Q = I2Rt
where Q is the heat generated in joules, I is the
electric current, R is the resistance and t is the
time of application. Ohm’s law is V = IR where V
is the voltage, I is the electric current and R is the
resistance. Ultimately, it is the local current dis-
tribution, time, and resistance (impedance) that
determine the local heat generation.
One can envision local RF effects as P = GV2
where P is the power loss in the tissue, G is the
local conductivity of the tissue, and V is the voltage
drop across the target. If one examines Fig. 1.12,
the local heat generation can always be determined
if one knows the microenvironment of the electrical
system. Oftentimes, we must simplify the ­analysis
1  Laser-Tissue Interactions

28
by characterizing the interaction in terms of macro
scale electrical equivalent circuits.
For example, if we consider a non-contact
capacitive electrode system, where we have a
“series circuit” comprised of capacitive and
resistive parallel parts (Fig. 1.13), the fat will be
preferentially heated according to Joules law. In
this particular capacitive electrode application
(i.e. Vanquish, BTL Aesthetics, Prague, Czech
Republic), spacing of the plates is important. The
spacing should prevent too close positioning to
the skin surface, or the superficial skin might be
overloaded with current. On the other hand, if
sweating occurs under the spacer, local current
on the skin can produce thermal injury.
If we look at the scenario in Fig.  1.14, one
sees the highest T will be achieved in the sub-
cutaneous fat under the electrodes, where the
tissues are modeled as a “series” circuit. On the
other hand, between the electrodes, the superfi-
cial muscles will be heated more. If one examines
the accompanying equations above Fig. 1.14, one
can see that the power delivery in the fat will be
roughly an order of magnitude larger than that in
the muscle. This observation only holds true for
this particular electrode configuration and can-
not be extrapolated to other dissimilar electrode
deployments.
Blood has the highest electrical conductivity
among most body components. At 1 MHz, blood
conductivity is 0.7 siemens/meter (S/m), where
1 S is equal to the reciprocal of 1 Ω. Wet skin is
at 0.25 S/m and finally dry skin, fat, and bone
have the lowest conductivity at around 0.02–
0.03 S/m [99]. Additionally, tissue conductivity
is significantly correlated with tissue tempera-
ture, where every 1 °C of increase in temperature
lowers skin impedance by 2% [101]. Thus, sur-
face cooling drives the electrical current deeper
into the tissue which allows for selective dermal
heating by the radiofrequency current and pro-
vides epidermal protection. This electrothermal
reaction results in selective heating of the dermis,
and when the temperature reaches 65 °C for 1 s,
tissue is coagulated. It should be noted that for
most RF applications (save the needle configura-
tions), the peak temperatures are much lower
(38–42 °C) but delivered over much longer times.
I
I
~
jB
δ
δ
δ
Ic
G
Id
C
a
b
c
d
J
J
V
V
E
~
Physical
Configuration
Equivalent
Circuit
Where

I = Current

G = Admittance

J = Current density

V = Rms voltage

C = Capacitance

Id = Displacement current

B = Susceptance
Elemental
Volume
Fig. 1.12  Heating
homogeneous tissue
with direct-contact,
conducting plates
A. A. Lloyd et al.

29
Monopolar Devices
For the monopolar devices, the radiofrequency
energy can be delivered by a stamped mode, a
continuously gliding movement, or via a fiber
that is placed internally.
Thermage (Comfort pulsed technology, Solta
Medical Hayward, CA)
Thermage (Comfort pulsed technology, Solta
Medical Hayward, CA) is a monopolar radiofre-
quency device where the radiofrequency is deliv-
ered via a stamped method. For each “stamp”
pulse, the hand piece is held in place and within a
short 3–4 s sequence; the hand piece tip releases a
cooling cryogen spray simultaneously with the
delivery of the radiofrequency energy. The cryogen
cooling spray provides epidermal protection and
the device continuously monitors the epidermal
temperature with each pulse so that it does not
exceed 45  °C.  The peak heating depth is most
likely around 1–4  mm and the intention is (1)
immediate tissue tightening and (2) delayed colla-
gen production. There is minimal downtime from
the procedure as the post treatment redness resolves
in a few hours and bruising and discomfort are very
rare. With the initial protocol of a single pass with
a high fluence, there was intense pain during the
procedure, and on occasion scars and surface irreg-
ularities were seen, therefore the protocol was
revised to multiple passes with lower fluence [102].
The new protocol has no reports of scarring, fat
atrophy, and the treatment is much less painful.
Exilis system (BLT Aesthetics, Prague, Czech
Republic)
Exilis (BLT Aesthetics, Prague, Czech
Republic) is a monopolar radiofrequency device
that delivers its radiofrequency via a continuous
motion hand piece which has a cooled gliding
electrode ball tip with an integrated “energy flow
control system” which senses if contact is lost,
therefore eliminating peaks of energy that could
arc and burn the skin [102]. In combination with
the hand piece, a bipolar grounding pad is used to
allow for real time monitoring of impedance to
eliminate spikes of radiofrequency and increase
patient comfort [102]. As the hand piece is con-
tinuously moved over the patient’s skin, it deliv-
ers its radiofrequency energy; the goal of therapy
is to maintain the skin at 40–43 °C for 10–15 min
I
~
Gf
Yt
Ct
Gm
Ym
Cm
Gf
Yt
Ct
C0
C0
Y0
Y0
d
V
~
V
Fat
Air
Fat
Muscle
E0
E0
Physical Configuration
Equivalent Circuit
Where

Y = The admittanace for Fat,


Muscle, and Air layers

E0 = Electric Field strength

I
= Current

V = Voltage

E = The Conductance

C = Capacitance
Fig. 1.13  Heating multiple layers of tissue with non-­
contacting capacitor plates
SKIN
Due to electrical and thermal properties of
Fat Vs. Muscle, there is a tendency for greater
Fat Vs. Muscle, heating in this electrode
Configuration.
FAT
MUSCLE
Fig. 1.14  Cross-sectional sketch showing fields in lay-
ered tissue exposed to shortwave diathermy capacitor-­
type electrodes
1  Laser-Tissue Interactions

30
for each region treated. It is indicated for skin
tightening and body contouring. Side effects
include redness and edema.
Pelleve (Ellman International, Inc., Oceanside,
NY)
Pelleve (Ellman International, Inc., Oceanside,
NY) is a monopolar radiofrequency device that
delivers its radiofrequency via a continuous
motion hand piece. The epidermis is cooled in
this system via two mechanisms, first with the gel
that is applied to the skin and second by convec-
tion to the surrounding air created by the continu-
ous motion use of the device. During the treatment
an infrared laser thermometer is used to fre-
quently monitor the epidermal temperature and
ensure the epidermal temperature is maintained
between 41 and 43  °C for the duration of the
treatment. It is FDA indicated for treatment of
mild to moderate facial wrinkles. The patient
experiences minimal pain during the procedure
and there is little to no down time. Side effects
are minimal and include erythema and edema.
ThermiTight (ThermiTight; Thermi­Aesthetics,
Southlake, TX)
ThermiTight (ThermiTight; Thermi­Aesthetics,
Southlake, TX) is a minimally invasive monopolar
radiofrequency device that delivers its radiofre-
quency via a 600 μm electrode enclosed in a 1 mm
cannula that is inserted into the subcutaneous
layer in order to completely spare the epidermis
and to more directly heat the dermal hypoder-
mal interface. The target temperature is set to
the optimal temperature for the target; the skin
surface T should not exceed about 42  °C,
whereas the T to shrink fibrous septae is
55–65 °C and to melt fat it is 70 °C. The device
features dual T monitoring, where the internal T
is measured firm the probe tip and the skin T is
monitored from an external thermal camera.
This configuration optimizes safety and effi-
cacy. If the temperature exceeds the set temper-
ature by 7 °C, the system is automatically turned
off [102]. Additionally, a forward looking infra-
red camera (FLIR) (FLIR E40; FLIR Systems,
Inc., Wilsonville, OR) provides infrared video
streaming of the entire epidermis in the treat-
ment field so epidermal temperature is simulta-
neously visually monitored. It has its FDA
indication to treat glabellar lines; however, it is
used off label to treat areas of unwanted local-
ized fat and/or loose skin of the arms, abdomen,
love handles, thighs or knees. Side effects
include erythema and edema which typically
resolve in 24  h. The procedure does require
injections with local anesthetic.
TrueSculpt (Cutera, Inc., Brisbane, CA)
TrueSculpt (Cutera, Inc., Brisbane, CA) is a
monopolar radiofrequency device that creates an
electric
field
perpendicular
to
the
skin-­
subcutaneous interface and results in bulk heat-
ing of the adipose layer. During a treatment, the
area is treated with a stamping delivery technique
and with each individual iteration the fat is heated
to 43–45 °C for 15 min, which results in a delayed
adipocyte cellular death response at about day 9
post treatment [98]. The operational frequency
can be adjusted to match the anatomic site of
treatment, using high frequencies in areas that
have a thin layer of adipose and low frequencies
in areas with thicker adipose layers [98]. The
device has target depths of 7–14 mm and has a
16 cm2 and a 40 cm2 hand piece. It is indicated for
noninvasive body contouring. There is minimal
downtime. Reported side effects include redness,
swelling and mild tenderness in treated areas
which typically resolve in a few hours.
Vanquish (BLT Aesthetics, Prague, Czech
Republic)
Vanquish (BLT Aesthetics, Prague, Czech
Republic) is a non-invasive and non-contact,
hands free body contouring device. It uses radio-
frequency energy to selectively heat the adipose
tissue to the point of adipocyte apoptosis while
sparing the epidermis. The device uses a capaci-
tive coupling far field configuration and automat-
ically adjusts its energy output based on the
resistance of adipose tissue to maintain homoge-
neous heating of the tissue [103, 104]. The adi-
pose tissue is heated to 40–45  °C and the
surrounding tissue only experiences temperatures
of 40–41 °C [103, 104]. There is minimal down-
time from the procedure as the device never con-
tacts the patient. Side effects include transient
erythema, warmth and tenderness which last at
most 1 h post treatment.
Infini (Lutronic, Inc., Fremont, CA)
The Infini (Lutronic, Inc., Fremont, CA) is
an insulated microneedle fractional rejuvenation
A. A. Lloyd et al.

31
device with a 49 needle tip and a 16 needle tip.
The microneedles are made of surgical stainless
steel, are coated with gold for increased conduc-
tivity, and then double coated with an insulating
silicone compound except for the 300 μm clos-
est to the point of the needle. Therefore, only the
tip of each needle is active and there is no elec-
trothermal damage to the epidermis. The needles
are each 200 μm in diameter and a point diameter
of 20 μm. The maximum power for the device is
50 W at level 20 and the exposure times can range
from 10 to 1000 ms. It is the exposure time that
provides the user with the most control over the tis-
sue damage as the exposure time can be decreased
for the higher power levels. The maximum cur-
rent density is near the electrodes. At 1 MHz, the
current flows rapidly to and fro between the alter-
nating rows of electrodes and creates small ~0.5–
0.7 mm diameter subsurface microthermal zones.
In these applications we model the micro-currents
as purely resistive and would find the greatest cur-
rent density at the needle tip; as heating proceeds,
the tissue impedance (Z) will decrease so long as
the tissue is not vaporized neat her needle surface.
As we proceed to the midpoints between the
electrode arrays, the absorbed power density (or
heat generation) drops off to a very small value.
By manipulating the applied voltage to the elec-
trode and “on” times, one can optimize localized
heating around the electrodes. In our application
with microneedles, the time scales are short
enough such that blood flow is irrelevant in the
calculations.
The needles can penetrate from a depth of
0.5–3.5 mm, which allows the user to set treat
multiple layers of the dermis. It is currently indi-
cated for the treatment of wrinkles. There is mini-
mal downtime and moderate intra-treatment
discomfort. Side effects include mild edema can
occur for 12–48 h post procedure and erythema
can last 3–5 days post procedure. There typically
is no persistent erythema. Mild crusting as well
as pin point bleeding can occur and typically
resolves spontaneously in 3–5 days. Mild PIH is
possible at the needle insertion points. This
observation brings up an important point. As
noted earlier, the RF immediate tissue interaction
is colorblind; however, depending on where and
how intense the injury is, the skin response might
not be colorblind. Any superficial needle based
RF systems can create hyperpigmentation (PIH)
like their laser counterparts.
EndyMed Intensif (EndyMed Medical,
Cesarea, Israel)
The EndyMed intensif (EndyMed Medical,
Cesarea, Israel) is a fractionated radiofrequency
device with gold plated, sharp tapered noninsu-
lated needles. The needles are 300 μm in diame-
ter and the tips are tapered. The needles can
achieves dermal heating up to 3.5 mm deep with
minimal epidermal damage using a fractionated
pulse mode which enables uniform distribution
of the energy. There is built in constant energy
delivery circuitry. It is indicated for treatment of
acne scarring, deep wrinkles, atrophic scars and
stretch marks. After treatment, there is minimal
micro-crusting and erythema lasts <24 h. There is
no risk for hypopigmentation, and only a mini-
mal risk for hyperpigmentation.
InMode
Fractora
(InMode
MD
Ltd.
Yokneam, Israel)
The InMode Fractora (InMode MD Ltd.
Yokneam, Israel) is a bipolar fractional radiofre-
quency resurfacing device which uses either a 24
needle, a 60 needle and a 126 high density tip.
The 24 pin tip can be coated to create an insulated
needle option. The needles penetrate up to 3 mm
in depth and when used with ablative settings,
they create a zone of coagulation up to 100 μm
around the ablated area. It is indicated for fine
lines, deep wrinkles, scars and discolored red and
brown skin tone. The downtime involves moder-
ate redness and swelling for the first 3–5 days,
however, the redness can last up to a week.
Ultrasound Tissue Interaction
Ultrasound waves can travel several millimeters
deep through tissue. Their propagation results in
vibration at the molecular level that can lead to
tissue heating in a dose-dependent fashion.
Unlike EMR waves, US is comprised of com-
pressive and refracted longitudinal waves that
behave like a “slinky” toy. Generally, as the fre-
quency and the focusing increase, the injuries
become smaller in volume. Different physical
processes are operative in US interactions, such
1  Laser-Tissue Interactions

32
as cavitation in lower frequency, lower power
density, non focused applications. With very low
power density applications, gentle diffuse waves
can be delivered as a physical therapy tool for
heating muscles after injuries (US diathermy).
By utilizing a specialized hand piece to focus
ultrasound waves tightly at a controlled depth
and using much higher fluences than needed for
imaging, selective zones of thermal injury can be
created. The result is collagen denaturation and
coagulative necrosis [105].
One advantage of US over light is the capacity
to deliver energy deeper in the tissues. For exam-
ple, even with relatively deeply penetrating light
(1064 nm), the surface intensity will diminish by
a factor of over 60% over a depth of 2 mm in the
skin. If follows that in the absence of cooling and
very long exposures (>1 s), very deep tissue heat-
ing with laser is challenging, particularly where
water is the target. The only viable way to create
focal deep heating in the tissue is with the assis-
tance of a cannula or other conduit that allows the
laser beam to bypass the highly scattering dermis
and fat tissues.
Two other focused US technologies are avail-
able in the USA and target fat. One targets fat
about 1–2  cm deep in the tissue and the other
(Ultra shape) about 2–4  cm deep in the tissue.
The volumes of damage (although focused) are
larger than the Ulthera device.
Overall, lower frequencies will result in
deeper penetration and larger volumes of injury.
For example, an US system (ultrashape) uses a
200 KHz transducer and cavitation to target fat.
Another system (Liposonix, Solta Medical)
uses HIFU but at a lower frequency (2 MHz ver-
sus 4–7 MHz for Ulthera) and targets tissue in a
plane down to about 1–2 cm.
If one looks at the total amount of tissue
affected, the lower frequency devices create
larger wounds deeper in the tissue, such that after
an Ulthera procedure only about a tsp of tissue is
damaged, whereas the for the deeper heating fat
destruction US devices, a few hundred milliliter
(size of soda can) are damaged.
In the case of HIFU (specifically Ulthera),
studies of this technology on both porcine soft
tissue and cadaveric human skin have revealed
predictable, reproducible and dose-dependent
creation of thermal injury zones at the level of
the subcutaneous fat down to the superficial
musculoaponeurtoic system [106]. Increased
energy levels results in thicker thermal injury
zones. The epidermis is spared from injury by
focusing the beam several millimeters below the
surface of the skin, although with increased
injury the thermal injury zones extend more
superficially [106]. A limited degree of direct
collagen contraction has been observed.
Ulthera (Ulthera, Inc, Mesa, AZ) is currently
the only microfocused ultrasound device on the
market. It is FDA approved for lifting of the eye-
brow, neck and submental areas as well as for
improvement of décolletage lines and wrinkles.
The device allows one to with visualize the skin
with the same ultrasound probe as the treatment
transducer. There is no downtime with treatment
and it is safe for use in all skin types. Potential
side effects tend to be transient and can include
erythema, edema, temporary pain, bruising,
numbness and scarring. It is contraindicated in
patients with open wounds, severe acne or active
implants in the treated area.
Summary
An understanding of the scientific principles in
laser applications empowers the physician to
optimize the use of this very expensive equip-
ment. At every patient encounter, the physician
should craft an approach based on the logical
sequence of laser tissue interactions outlined in
this chapter. By appropriate choreographing of
cooling and heating, the physician can be con-
fidant in predicting the immediate tissue
response. However, all approaches should be
undertaken within the context of an under-
standing of wound healing. The physical
aspects of the interaction are typically more
predictable than the subsequent healing
response. If in doubt, test spots are always an
option and should be considered for the anx-
ious patient (or physician).
A. A. Lloyd et al.