dŚĞīĞĐƚŽĨ,ĞĂƚŽŶŽůůĂŐĞŶĂŶĚEĞŽĐŽůůĂŐĞŶĞƐŝƐ
sĂƌŝŽƵƐĞŶĞƌŐǇͲďĂƐĞĚĚĞǀŝĐĞƐĂƌĞƵƐĞĚƚŽƟŐŚƚĞŶͬƚŽŶĞͬůŝŌƚŚĞƐŬŝŶ
Various energy based technologies are available that trigger neocol-‐
ůĂŐĞŶĞƐŝƐĂŶĚŚĞůƉĮƌŵ͕ƟŐŚƚĞŶ͕ƚŽŶĞŽƌůŝŌƚŚĞƐŬŝŶ͕ŶĂŵĞůǇƚŚŽƐĞ
ƵƟůŝǌŝŶŐ ƌĂĚŝŽĨƌĞƋƵĞŶĐǇ͕ ĂŶĚ ĚĞĞƉͲƟƐƐƵĞ ƵůƚƌĂƐŽƵŶĚ ŵŽĚĞƐ ŽĨ energy delivery. ZĂĚŝŽĨƌĞƋƵĞŶĐǇ ĂƉƉƌŽĂĐŚĞƐ ƉƌŽĚƵĐĞ Ă ƌĂŶŐĞ ŽĨ ƚĞŵƉĞƌĂƚƵƌĞƐ͘
dŚĞ dŚĞƌŵĂŽŽů ^ǇƐƚĞŵ ;dŚĞƌŵĂŐĞ͕ /ŶĐ͕͘ ,ĂǇǁĂƌĚ͕ Ϳ (1/
Abraham/169/A)͕ĨŽƌĞǆĂŵƉůĞ͕ƌĞĂĐŚĞƐƚĞŵƉĞƌĂƚƵƌĞƐŽĨĂƉƉƌŽǆŝ-‐
ŵĂƚĞůǇϱϱΣŝŶƚŚĞĚĞƌŵŝƐ(1/Abraham/171/A).1dŚĞĐĐĞŶƚZ&ƐǇƐ-‐
ƚĞŵ;ůŵĂ>ĂƐĞƌƐ͕/ŶĐ͕͘&ƚ͘>ĂƵĚĞƌĚĂůĞ͕&>ͿŐĞŶĞƌĂƚĞƐƚĞŵƉĞƌĂƚƵƌĞƐŽĨ ďĞƚǁĞĞŶϰϬΣĂŶĚϰϰΣ(2/Sadick/183/A).2 dŚĞ hůƚŚĞƌĂΠ ^ǇƐƚĞŵ ;hůƚŚĞƌĂ͕ /ŶĐ͕͘ DĞƐĂ͕ Ϳ ĚĞĞƉͲƟƐƐƵĞ ƵůƚƌĂƐŽƵŶĚŚĞĂƚƐƚŚĞƚĂƌŐĞƚƟƐƐƵĞƚŽхϲϬΣ(3/Laubach/729/A)(4/
White/69/A).ϯ͕ϰdŚŝƐŵŽĚĂůŝƚǇŝƐƚŚĞŽŶůǇĚĞǀŝĐĞƚŚĂƚŚĂƐƌĞĐĞŝǀĞĚ
&ĐůĞĂƌĂŶĐĞĨŽƌĂ͞ůŝŌ͟ŝŶĚŝĐĂƟŽŶ͘
KĨŶŽƚĞ͕ƚŚĞƚĞŵƉĞƌĂƚƵƌĞƚŽǁŚŝĐŚƚŚĞƐĞĚĞǀŝĐĞƐŚĞĂƚƚŚĞƐŬŝŶĐŽƌ-‐
ƌĞůĂƚĞƐǁŝƚŚƚŚĞůĞǀĞůŽĨĐŽůůĂŐĞŶĚĞŶĂƚƵƌĂƟŽŶʹĂŶĚƐƵďƐĞƋƵĞŶƚ
ŶĞŽĐŽůůĂŐĞŶĞƐŝƐʹĂĐŚŝĞǀĞĚ͘
dŚĞƚŚƌĞƐŚŽůĚĨŽƌĐŽůůĂŐĞŶĚĞŶĂƚƵƌĂƟŽŶŝƐĂƉƉƌŽǆŝŵĂƚĞůǇϲϬͲϲϱΣ
ƌƐ ,ĂǇĂƐŚŝ ĂŶĚ ĐŽůůĞĂŐƵĞƐ ĂƐƐĞƐƐĞĚ ƚŚĞ ĞīĞĐƚ ŽĨ Ă ǁŝĚĞ ƌĂŶŐĞ
ŽĨ ƚĞŵƉĞƌĂƚƵƌĞƐ ;ϯϳΣ͕ ϱϱΣ͕ ϲϬΣ͕ ϲϱΣ͕ ϳϬΣ͕ ϳϱΣ͕ ĂŶĚ ϴϬΣͿ
ŽŶ ĐŽůůĂŐĞŶ ĐŽŶƚƌĂĐƟŽŶ ƵƟůŝǌŝŶŐ ƐĂŵƉůĞƐ ĨƌŽŵ ƚŚĞ ŐůĞŶŽŚƵ-‐
ŵĞƌĂů ũŽŝŶƚ ĐĂƉƐƵůĞ (5/Hayashi/107/A).5 ƚ ϲϱΣ͕ ĐŽůůĂŐĞŶ ĐŽŶ-‐
tracted (5/Hayashi/109/A) and architectural changes indica-‐
ƟǀĞ ŽĨ ĚĞŶĂƚƵƌĂƟŽŶ ĐŽƵůĚ ďĞ ŽďƐĞƌǀĞĚ (5/Hayashi/109/B/C). dŚĞƐĞ ĐŚĂŶŐĞƐ ŝŶƚĞŶƐŝĮĞĚ Ăƚ ƐůŝŐŚƚůǇ ŚŝŐŚĞƌ ƚĞŵƉĞƌĂƚƵƌĞƐͶ
ϳϬΣ ĂŶĚ ϴϬΣ (5/Hayashi/109/C)͘ ŵŽŶŐ ƚŚĞ ŚŝŐŚĞƌ ƚĞŵƉĞƌĂ-‐
ƚƵƌĞƐ ƚĞƐƚĞĚ ;ϳϬΣ͕ ϳϱΣ͕ ϴϬΣͿ͕ ŚŝƐƚŽůŽŐŝĐĂů ĂŶĂůǇƐŝƐ ƐŚŽǁĞĚ
ŶŽ ƐŝŐŶŝĮĐĂŶƚ ĚŝīĞƌĞŶĐĞƐ (5/Hayashi/109/B)͕ ƐƵŐŐĞƐƟŶŐ ƚŚĂƚ
ĂĚĚŝƟŽŶĂů ŚĞĂƚ ĚŽĞƐ ŶŽƚ ŚĂǀĞ ĂĚĚŝƟŽŶĂů ĞīĞĐƚƐ ŽŶ ĐŽůůĂŐĞŶ͘
^ŝŵŝůĂƌůǇ͕ ƌƐ sĂŶŐƐŶĞƐƐ ĂŶĚ ĐŽůůĞĂŐƵĞƐ ĂƉƉůŝĞĚ Ă ƌĂŶŐĞ ŽĨ ƚĞŵ-‐
ƉĞƌĂƚƵƌĞƐ ƚŽ ŚƵŵĂŶ ƚĞŶĚŽŶƐ (6/Vangsness/268/A/269/B) and ŽďƐĞƌǀĞĚ ĐŽůůĂŐĞŶ ĐŽŶƚƌĂĐƟŽŶ ĂŶĚ ƐŚŽƌƚĞŶŝŶŐ ũƵƐƚ ďĞůŽǁ ϳϬΣ (6/Vangsness/269/A)ĂŶĚĚĞŶĂƚƵƌĂƟŽŶĂƚŚŝŐŚĞƌƚĞŵƉĞƌĂƚƵƌĞƐ(6/
Vangsness/269/A).ϲ
&ƵƌƚŚĞƌǀĂůŝĚĂƟŶŐƚŚĞƐĞƌĞƐƵůƚƐ͕ƌƐ>ŝŶĂŶĚĐŽůůĞĂŐƵĞƐƵƐĞĚĂƐĞĐ-‐
ŽŶĚͲŚĂƌŵŽŶŝĐŐĞŶĞƌĂƟŽŶŵŝĐƌŽƐĐŽƉĞƚŽĚŝƌĞĐƚůǇŽďƐĞƌǀĞƚŚĞĞīĞĐƚƐ
ŽĨ ŚĞĂƚ ;ďĞƚǁĞĞŶ ϮϱΣ ĂŶĚ ϲϬΣͿ ŽŶ ĐŽůůĂŐĞŶ ĮďĞƌƐ (7/Lin/622/
A/B) ĨƌŽŵƌŽĚĞŶƚƚĂŝůƚĞŶĚŽŶƐ(7/Lin/623/A).ϳ They observed that ĐŽůůĂŐĞŶĮďĞƌƐďĞŐŝŶƚŽĐƵƌǀĞĂƚϱϮΣĂŶĚϱϱΣ(7/Lin/623/B)͕ĂŶĚ
ĐŽůůĂŐĞŶĚĞŶĂƚƵƌĂƟŽŶŽĐĐƵƌƌĞĚĂƚϲϬΣ(7/Lin/624/A).
DŽƌĞƌĞĐĞŶƚƌĞƐĞĂƌĐŚďǇƌƐWĂƵůĂŶĚĐŽůůĞĂŐƵĞƐĂƐƐĞƐƐŝŶŐƚŚĞĞĨ-‐
ĨĞĐƚŽĨŚĞĂƚŽŶĐŽůůĂŐĞŶŝŶƐĂŵƉůĞƐŽĨĂĚŝƉŽƐĞƟƐƐƵĞ;ǁŝƚŚƐĞƉƚĂů
ĂŶĚƌĞƟĐƵůĂƌĐŽŶŶĞĐƟǀĞƟƐƐƵĞͿ͕ĚĞƌŵŝƐ͕ĂŶĚĨĂƐĐŝĂ(8/Paul/88/A/B) ĨƵƌƚŚĞƌ ƐƵƉƉŽƌƚƐ Ă ĐŽůůĂŐĞŶ ĚĞŶĂƚƵƌĂƟŽŶ ƚŚƌĞƐŚŽůĚ ŽĨ ΕϲϬͲϲϱΣ (8/Paul/94/A).ϴ /Ŷ ƚŚŝƐ ƐƚƵĚǇ͕ ƚŚĞ ĐŽůůĂŐĞŶ ĐŽŶƚƌĂĐƟŽŶ ƚŚƌĞƐŚŽůĚ
ĨĞůůďĞƚǁĞĞŶϲϬͲϳϬΣ(8/Paul/94/A)͖ƐƉĞĐŝĮĐĐŽůůĂŐĞŶĐŽŶƚƌĂĐƟŽŶ
ƚĞŵƉĞƌĂƚƵƌĞƐ ǁĞƌĞ ϴϭ͘ϵΣ ĨŽƌ ƚŚĞ ĚĞƌŵŝƐ͕ ϲϭ͘ϱΣ ĨŽƌ ƚŚĞ ĨĂƐĐŝĂ͕
ĂŶĚϲϵ͘ϰΣĨŽƌƚŚĞƐĞƉƚĂͬĂĚŝƉŽƐĞƟƐƐƵĞ(8/Paul/92/A). ŽůůĂŐĞŶĚĞŶĂƚƵƌĂƟŽŶŝƐĨŽůůŽǁĞĚďǇŶĞŽĐŽůůĂŐĞŶĞƐŝƐ
ŽůůĂŐĞŶ ƌĞũƵǀĞŶĂƚĞƐ ŽǀĞƌ ƚŚĞ ŵŽŶƚŚ Žƌ ƐŽ ĂŌĞƌ ƚƌĞĂƚŵĞŶƚ (9/Hayashi/170/A/B).ϵ/ŶĐƌĞĂƐĞĚƐŵĂůůĐŽůůĂŐĞŶĮďĞƌĨŽƌŵĂƟŽŶͶ
ĞǀŝĚĞŶĐĞ ŽĨ ŶĞŽĐŽůůĂŐĞŶĞƐŝƐͶŚĂƐ ďĞĞŶ ŶŽƚĞĚ Ăƚ ϯϬ ĚĂǇƐ ƉŽƐƚ
ŚĞĂƚƚƌĞĂƚŵĞŶƚ(9/Hayashi/170/B)͘ƐĞĐŽŶĚƐƚƵĚǇƚƌĂĐŬŝŶŐƟƐƐƵĞ
ĐŚĂŶŐĞƐĂŌĞƌŚĞĂƟŶŐƚŽƚŚĞĚĞŶĂƚƵƌĂƟŽŶƌĂŶŐĞ(10/Hantash/1/A) ĨŽƵŶĚŶĞŽĐŽůůĂŐĞŶĞƐŝƐ͕ŶĞŽĞůĂƐƚŽŐĞŶĞƐŝƐ͕ĂŶĚĚĞƉŽƐŝƟŽŶŽĨŶĞǁŚǇ-‐
ĂůƵƌŽŶŝĐĂĐŝĚĂƚϭϬǁĞĞŬƐƉŽƐƚƚƌĞĂƚŵĞŶƚ(10/Hantash/3/A/4/A/B).ϭϬ dĞŵƉĞƌĂƚƵƌĞƐďĞůŽǁϲϬΣŚĂǀĞŵŝŶŝŵĂůĞīĞĐƚƐŽŶĐŽůůĂŐĞŶƐƚƌƵĐ-‐
ƚƵƌĞĂŶĚƚŚƵƐĂƌĞƵŶůŝŬĞůǇƚŽŚĂǀĞƐŝŐŶŝĮĐĂŶƚĞīĞĐƚƐŽŶĐŽůůĂŐĞŶĞƐŝƐ
ƌƐ >ŝŶ ĂŶĚ ĐŽůůĞĂŐƵĞƐ ŶŽƚĞ ƚŚĂƚ ǁŚŝůĞ ĐŽůůĂŐĞŶ ĮďĞƌƐ ďĞŐŝŶ ƚŽ
ĐƵƌǀĞ Ăƚ ϱϮΣͲϱϱΣ (7/Lin/623/B)͕ ƐƚƌƵĐƚƵƌĂů ĐŚĂŶŐĞƐ ǁĞƌĞ ŶŽƚ
ƐĞĞŶĂƚůŽǁĞƌƚĞŵƉĞƌĂƚƵƌĞƐ;ϮϱΣĂŶĚϰϬΣͿ(7/Lin/623/B).ϳ ^ŝŵŝůĂƌůǇ͕ƌƐ,ĂǇĂƐŚŝĂŶĚĐŽůůĞĂŐƵĞƐĨŽƵŶĚƚŚĂƚƚĞŵƉĞƌĂƚƵƌĞƐŽĨ
ϯϳΣ͕ϱϱΣ͕ĂŶĚϲϬΣŚĂĚŶŽƐŝŐŶŝĮĐĂŶƚĞīĞĐƚŽŶĐŽůůĂŐĞŶůĞŶŐƚŚ
(5/Hayashi/109/A)ĂŶĚƌĞƐƵůƚĞĚŝŶƐŝŐŶŝĮĐĂŶƚůǇĨĞǁĞƌŚŝƐƚŽůŽŐŝĐĂů
ĐŚĂŶŐĞƐƚŚĂŶĚŝĚŚŝŐŚĞƌƚĞŵƉĞƌĂƚƵƌĞƐ(5/Hayashi/109/B).ϱ
ZĞĨĞƌĞŶĐĞƐ
ϭ͘ ďƌĂŚĂŵDd͕DĂƐŚŬĞǀŝĐŚ'͘DŽŶŽƉŽůĂƌƌĂĚŝŽĨƌĞƋƵĞŶĐǇƐŬŝŶƟŐŚƚĞŶŝŶŐ͘Facial Plast Surg Clin North Am.ϮϬϬϳ͖ϭϱ;ϮͿ͗ϭϲϵͲϭϳϳ͘
Ϯ͘ ^ĂĚŝĐŬ E͘ dŝƐƐƵĞ ƟŐŚƚĞŶŝŶŐ ƚĞĐŚŶŽůŽŐŝĞƐ͗ ĨĂĐƚ Žƌ ĮĐƟŽŶ͘ Aesthet Surg J. ϮϬϬϴ͖Ϯϴ;ϮͿ͗ϭϴϬͲϭϴϴ͘
ϯ͘ >ĂƵďĂĐŚ,:͕DĂŬŝŶ/Z͕ĂƌƚŚĞW'͕^ůĂǇƚŽŶD,͕DĂŶƐƚĞŝŶ͘/ŶƚĞŶƐĞĨŽĐƵƐĞĚƵůƚƌĂ-‐
ƐŽƵŶĚ͗ĞǀĂůƵĂƟŽŶŽĨĂŶĞǁƚƌĞĂƚŵĞŶƚŵŽĚĂůŝƚǇĨŽƌƉƌĞĐŝƐĞŵŝĐƌŽĐŽĂŐƵůĂƟŽŶǁŝƚŚŝŶ
ƚŚĞƐŬŝŶ͘Dermatol Surg.ϮϬϬϴ͖ϯϰ;ϱͿ͗ϳϮϳͲϳϯϰ͘
ϰ͘ tŚŝƚĞtD͕DĂŬŝŶ/Z͕^ůĂǇƚŽŶD,͕ĂƌƚŚĞW'͕'ůŝŬůŝĐŚZ͘^ĞůĞĐƟǀĞƚƌĂŶƐĐƵƚĂŶĞŽƵƐ
ĚĞůŝǀĞƌǇŽĨĞŶĞƌŐǇƚŽƉŽƌĐŝŶĞƐŽŌƟƐƐƵĞƐƵƐŝŶŐ/ŶƚĞŶƐĞhůƚƌĂƐŽƵŶĚ;/h^Ϳ͘Lasers Surg Med.&Ğď͖ϰϬ;ϮͿ͗ϲϳͲϳϱ͘
ϱ͘ ,ĂǇĂƐŚŝ<͕dŚĂďŝƚ'///͕DĂƐƐĂ<>͕ĞƚĂů͘dŚĞĞīĞĐƚŽĨƚŚĞƌŵĂůŚĞĂƟŶŐŽŶƚŚĞůĞŶŐƚŚ
ĂŶĚ ŚŝƐƚŽůŽŐŝĐ ƉƌŽƉĞƌƟĞƐ ŽĨ ƚŚĞ ŐůĞŶŽŚƵŵĞƌĂů ũŽŝŶƚ ĐĂƉƐƵůĞ͘ Am J Sports Med. ϭϵϵϳ͖Ϯϱ;ϭͿ͗ϭϬϳͲϭϭϮ͘
ϲ͘ sĂŶŐƐŶĞƐƐ d :ƌ͕ DŝƚĐŚĞůů t ///͕ EŝŵŶŝ D͕ ƌůŝĐŚ D͕ ^ĂĂĚĂƚ s͕ ^ĐŚŵŽƚǌĞƌ ,͘ Žů-‐
ůĂŐĞŶ ƐŚŽƌƚĞŶŝŶŐ͘ Ŷ ĞǆƉĞƌŝŵĞŶƚĂů ĂƉƉƌŽĂĐŚ ǁŝƚŚ ŚĞĂƚ͘ Clin Orthop Relat Res. ϭϵϵϳ͖;ϯϯϳͿ͗ϮϲϳͲϮϳϭ͘
ϳ͘ >ŝŶ ^:͕ ,ƐŝĂŽ z͕ ^ƵŶ z͕ >Ž t͕ Ğƚ Ăů͘ DŽŶŝƚŽƌŝŶŐ ƚŚĞ ƚŚĞƌŵĂůůǇ ŝŶĚƵĐĞĚ ƐƚƌƵĐƚƵƌĂů
ƚƌĂŶƐŝƟŽŶƐŽĨĐŽůůĂŐĞŶďǇƵƐĞŽĨƐĞĐŽŶĚͲŚĂƌŵŽŶŝĐŐĞŶĞƌĂƟŽŶŵŝĐƌŽƐĐŽƉǇ͘KƉƚ>ĞƩ͘
ϮϬϬϱ͖ϯϬ;ϲͿ͗ϲϮϮͲϲϮϰ͘
ϴ͘ WĂƵůD͕ůƵŐĞƌŵĂŶ'͕<ƌĞŝŶĚĞůD͕DƵůŚŽůůĂŶĚZ^͘dŚƌĞĞͲĚŝŵĞŶƐŝŽŶĂůƌĂĚŝŽĨƌĞƋƵĞŶ-‐
ĐǇƟƐƐƵĞƟŐŚƚĞŶŝŶŐ͗ĂƉƌŽƉŽƐĞĚŵĞĐŚĂŶŝƐŵĂŶĚĂƉƉůŝĐĂƟŽŶƐĨŽƌďŽĚǇĐŽŶƚŽƵƌŝŶŐ͘
ĞƐƚŚĞƟĐWůĂƐƚ^ƵƌŐ͘ϮϬϭϭ͖ϯϱ;ϭͿ͗ϴϳͲϵϱ͘
ϵ͘ ,ĂǇĂƐŚŝ <͕ EŝĞĐŬĂƌǌ :͕ dŚĂďŝƚ ' ///͕ ŽŐĚĂŶƐŬĞ ::͕ ŽŽůĞǇ :͕ DĂƌŬĞů D͘ īĞĐƚ
ŽĨ ŶŽŶĂďůĂƟǀĞ ůĂƐĞƌ ĞŶĞƌŐǇ ŽŶ ƚŚĞ ũŽŝŶƚ ĐĂƉƐƵůĞ͗ ĂŶ ŝŶ ǀŝǀŽ ƌĂďďŝƚ ƐƚƵĚǇ ƵƐŝŶŐ Ă
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169
FACIAL PLASTIC
SURGERY CLINICS
OF NORTH AMERICA
Facial Plast Surg Clin N Am 15 (2007) 169–177
Monopolar Radiofrequency Skin
Tightening
Manoj T. Abraham,
-
-
-
a,b,
MD, FACS
*, Grigoriy Mashkevich,
Overview of the ThermaCool monopolar
capacitive radiofrequency device
Mechanism of action
Clinical experience with monopolar
capacitive radiofrequency treatment
Monopolar capacitive radiofrequency
procedure
Patient selection
Contraindications
Aesthetic improvement in the appearance of facial wrinkles traditionally has been achieved with
various rhytidectomy and ablative resurfacing techniques. The latter group includes dermabrasion,
chemical peels, and laser resurfacing, all of which
diminish facial rhytides via surface re-epithelialization and dermal remodeling.
Although invasive cosmetic approaches have established a proven track record in facial plastic
surgery, extended recuperation and associated complications always have been a consideration. The
development of minimally invasive treatments for
facial rejuvenation has provided an attractive alternative. For instance, the rising popularity of officebased dermal fillers and neurotoxins, such as Botox
is a clear indication of the paradigm shift toward
noninvasive treatments. Implicit in the popularity
-
c
MD
Anesthesia
Technique
Aftercare
Results
Complications
Future directions
Summary
References
of these minimally invasive procedures is the emphasis on minimal recovery and decreased risks, often coupled with the acceptance of less dramatic,
subtle results.
Seen in this light, it is understandable why the
ThermaCool System (Thermage, Inc., Hayward,
California), which uses monopolar capacitive radiofrequency (MRF) energy to tighten skin in a nonablative fashion, has seen growing acceptance over
the past several years. Unlike traditional ablative
methods of skin rejuvenation, the ThermaCool device cools and protects the skin surface while selectively delivering radiofrequency energy to the
deeper dermis. Clinically observed mild to moderate skin tightening and contour enhancement is
thought to arise from thermally induced dermal
collagen contraction and subsequent remodeling.
a
Facial Plastic and Reconstructive Surgery, Department of Otolaryngology—Head and Neck Surgery, New
York Medical College, 40 Sunshine Cottage Road, Valhalla, NY 10595, USA
b
Facial Plastic, Reconstructive & Laser Surgery, PLLC, P.O. Box 2179, Poughkeepsie, NY 12601, USA
c
Department of Otolaryngology—Head and Neck Surgery, New York Eye & Ear Infirmary, 6th Floor, 310 East
14th Street, New York, NY 10003, USA
Video on MRF techniques available on http://www.theclinics.com/.
* Corresponding author. Facial Plastic, Reconstructive & Laser Surgery, PLLC, P.O. Box 2179, Poughkeepsie, NY
12601.
E-mail address: [email protected] (M.T. Abraham).
1064-7406/07/$ – see front matter ª 2007 Elsevier Inc. All rights reserved.
facialplastic.theclinics.com
doi:10.1016/j.fsc.2007.01.005
A
170
Abraham & Mashkevich
Because the epithelium is maintained intact, there
is little recovery, and the potential for complications (eg, infection, pigment changes, and cutaneous scarring) is minimized.
The ThermaCool System was approved by the US
Food and Drug Administration for the treatment
of periorbital rhytides in November 2002. This
approval was followed by authorization for the
treatment of facial rhytides in June 2004 and additional clearance for treatment of all rhytides in
December 2005. The development of noninvasive
treatments to tighten skin remains an area of actively evolving technology. Currently, MRF seems
to have the most published data supporting clinical
efficacy.
Overview of the ThermaCool monopolar
capacitive radiofrequency device
The ThermaCool System is made up of several components that allow delivery of MRF energy to the
skin in a nonablative fashion (Fig. 1).
The MRF generator provides an alternating MRF
signal to the treatment tip. The front panel monitor
displays real-time information pertaining to each
treatment sequence, such as delivered energy levels
and measured tissue impedance. The generator receives and processes feedback data from the tip,
which enables continuous monitoring of skin temperature, contact with tissue, and function of the
cooling system. Energy settings are determined
Fig. 1. The ThermaCool MRF system.
based on anatomic treatment site and are adjusted
easily by using front panel controls.
The cooling module houses and connects a replaceable cryogen coolant canister to the generator.
Cryogen is dispensed to the treatment tip immediately before, during, and after delivery of MRF
energy to cool and protect the skin surface.
The handpiece facilitates delivery of MRF energy
and cryogen coolant from the generator to the treatment tip. It also allows communication between
the generator and the sensors on the treatment
tip. The handpiece is ergonomically designed and
attached to the generator via a flexible cord. The different treatment tips can be fitted onto the
handpiece.
Treatment tips uniformly distribute MRF energy
across treatment areas. These single-use tips are manufactured in a range of sizes (currently 0.25 cm2,
1.0 cm2, 1.5 cm2, and 3.0 cm2) (Fig. 2). Each tip is
designed to provide a specific, uniform depth of
penetration of MRF energy. A critical function of the
treatment tip is to provide contact cooling to the
epidermis during treatment cycles, which occurs
via continuous application of the cryogen spray
onto the inner surface of the tip’s membrane.
Treatment tips also gather real-time data on contact
efficiency with the skin and skin temperature.
Fig. 2. Currently available MRF treatment tips.
Monopolar Radiofrequency Skin Tightening
Mechanism of action
A
The ThermaCool device achieves tissue tightening
by delivering radiofrequency energy to a volume
of tissue situated beneath the treatment tip. The
depth and degree of energy transfer depends on several factors, including the size and configuration of
the treatment tip, selected energy settings, and inherent conductive properties of the tissue.
The MRF electrode is located on the back of
a nonconductive layer inside the treatment tip.
When energy is applied, the nonconductive layer
creates a capacitor with the skin surface and establishes a uniform electric field across the tip’s surface,
which ensures equal transfer of energy at all points
along the area of contact with the skin. During the
treatment cycle, an electromagnetic field is established that alternates polarity at a rate of 6 million
cycles per second, which stimulates movement of
charged particles and creates an electric current
within the treated tissue. The current conducts
most effectively through hydrophilic structures,
such as the dermal collagen framework and underlying fibrous septae, and much less effectively
through subcutaneous fat. Tissue resistance to the
flow of current generates localized heat within the
collagen-based structures [1].
Simultaneously during energy delivery, cryogen
coolant is sprayed onto the inner surface of the contact membrane of the treatment tip, thereby allowing continuous surface cooling of the epidermis
and upper dermal layers. As a result, treatment cycles establish a controlled, volumetric heat gradient
with temperatures above 55! C in the dermis and
20! C to 35! Cin the epidermis (Fig. 3).
The physiologic mechanism underlying skin
tightening and contour enhancement observed
with MRF treatment is essentially understood.
Immediate clinical improvement arises from
thermally induced collagen contraction and
Fig. 3. Thermogram schematic shows the volumetric
heat generated within the upper dermis using a medium-depth MRF treatment tip (darker indicates increased temperature).
denaturation. These processes lead to an inflammatory wound-healing response, which establishes
long-term dermal remodeling and leads to further
tightening of treated areas. Zelickson and colleagues
[2] corroborated this theory by documenting denaturation of collagen fibrils and elevated expression
of type I collagen mRNA in skin samples treated
with MRF. Meshkinpour and colleagues [3] found
increased collagen production (type III > type I) in
biopsies even 12 months after MRF treatment.
Clinical experience with monopolar
capacitive radiofrequency treatment
The clinical efficacy of MRF in the management of
facial rhytides has been well documented in the scientific literature. Several published reports have
described clinical improvements in patients undergoing rejuvenation with MRF. Reports indicate that
mild to moderate skin tightening is achieved, although follow-up is often short, because MRF technology has been available only for the past few
years. Most studies providing data from subjective
patient questionnaires have found good patient satisfaction rates. Several authors have suggested that
proper patient education that focuses on realistic
expectations of radiofrequency treatment significantly improves the subjective experience with
this procedure [1,4,5].
The first reports of nonablative capacitive MRF
treatments for skin tightening and contour improvement date back to only a few years ago. Fitzpatrick and colleagues [6] published an initial series of
patients treated with the ThermaCool system. In
their multicenter trial of 86 enrolled subjects, a single treatment of periorbital areas resulted in a measurable eyebrow elevation in 62% and clinical
improvement in rhytides in 83%. These outcomes
favorably compared with patient satisfaction rates.
Subsequent reports assessing the upper one third
of the face further substantiated observed tissue
tightening with MRF. Ruiz-Esparza [7] claimed improvement in all nine patients treated for flaccid
lower eyelid skin and commented on ‘‘remarkable’’
patient satisfaction. Abraham and colleagues [8] reported a statistically significant brow height elevation, ranging from 1.6 to 2.4 mm, recorded 12
weeks after the treatment with the ThermaCool
device. Similarly, Bassichis and colleagues [9]
described a significant improvement in brow height
position; however, they raised concerns about the
unpredictable degree of tissue tightening. Studies
focusing on rejuvenation of the middle and lower
facial thirds also have showed favorable aesthetic
results with the ThermaCool device. Fritz and colleagues [5] reported a statistically significant improvement in the appearance of nasolabial folds
171
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Abraham & Mashkevich
Fig. 4. Patient before (left) and after (right) MRF treatment of the face and neck.
in nine patients 4 months after two treatments
spaced month apart. In their study, 75% of patients had a positive experience with MRF treatment
and considered paying for additional sessions. In
another publication, Nahm and colleagues [10]
treated one side of the face with MRF in ten patients;
at 3 months, they documented mean jowl surface
area reduction of 22.6% compared with the nontreated side. Additional reports demonstrating
clinical improvement in wrinkle reduction and
skin tightening in the face and neck have been
published by numerous authors [1,4,11–21].
MRF technology seems safe when used in conjunction with various dermal fillers. Alam and colleagues
[22] reported histologic outcomes after MRF treatment of skin injected with hyaluronic acid (Restylane) and calcium hydroxylapatite (Radiesse).
Punch biopsies taken from volunteers showed
Monopolar Radiofrequency Skin Tightening
Fig. 5. Patient before (left) and after (right) treatment with MRF, Botox, and Restylane to the glabella and Restylane to the nasolabial and melolabial folds.
unaltered histologic appearance of dermal fillers.
These findings agree with results from a similar experiment performed in pigs [23]. In addition to observed
preservation of filler materials, histologic assessment
in these animals showed a statistically significant increase in inflammatory and fibrotic response, which
suggested a synergistic effect that could enhance the
skin tightening achieved with MRF.
Monopolar capacitive radiofrequency
procedure
Patient selection
Appropriate selection of patients is critical for attaining successful outcomes with MRF skin tightening. Ideal candidates are patients starting in their
mid-30s who exhibit early signs of aging and have
173
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Abraham & Mashkevich
Fig. 6. Before (left) and after (right) photos of a patient after MRF treatment and midface lift with
blepharoplasty.
mild to moderate facial and neck rhytides. Patients
who have had prior rhytidectomy who are starting
to develop skin laxity also can benefit. Patients
who have significant structural ptosis represent
poor candidates for the MRF procedure, unless it
is performed in combination with other minimally
invasive lifting techniques. ThermaCool treatment
represents an option for patients who are reluctant
or unable to undergo any surgical intervention, although patients must be cautioned that improvements from MRF treatment are currently not as
dramatic as what can be achieved with surgical rhytidectomy. Unlike ablative skin resurfacing techniques and other laser or light-based treatments
that depend on energy absorption, MRF can be performed safely even in patients who have Fitzpatrick
Monopolar Radiofrequency Skin Tightening
IV, V, and VI sun-reactive skin types because MRF
energy is delivered based on tissue impedance [24].
Contraindications
MRF treatment should be avoided if there is any active skin pathology in the area to be treated. The
depth of penetration of MRF energy is unpredictable in pathologically thinned skin. Previous radiation, autoimmune conditions, smoking, and other
factors that compromise healing also inhibit the desired dermal collagen remodeling. Radiofrequency
energy delivered during treatment may interfere
with the function of any implanted medical devices,
and treatment should not be performed directly
over any metallic implants, tattoos, or plates. In
general, we do not treat patients who are or might
be pregnant.
Anesthesia
Patient discomfort has been cited as a significant issue with MRF treatment [9,14,25]. Experience with
different treatment protocols and analgesia has
considerably improved the tolerability of these
treatments, however. With multiple passes using
lower energy settings, improved and more consistent clinical outcomes have been observed while simultaneously improving patient comfort. Oral
sedatives and narcotic analgesics can minimize discomfort experienced by patients. Topical anesthetic
agents negate the desired cooling sensation at the
skin surface, however, and have not been found to
be effective in allowing increased MRF energy levels
[25]. Injection or tumescent local anesthesia is not
advised because the fluid that is infiltrated into the
tissue alters impedance and impairs delivery of MRF
energy [26]. Providers who are experienced can opt
to treat patients with intravenous sedation or deeper anesthesia, but this removes patient feedback
and is not recommended for the novice user (see
Movie 1).
Technique
Before starting MRF treatment, the skin is cleansed
and patients are told to remove any metallic jewelry.
Patients with a previous history of oral herpes infection are given prophylaxis with an antiviral agent.
A temporary ink grid is placed to ensure uniform
coverage of the treatment area. A grounding pad is
applied to the patient, and the MRF system is calibrated. Non–hair-bearing areas of the face and upper neck that manifest skin laxity are treated.
Contiguous treatment around the mouth is avoided
to prevent circumferential tightening and potential
accentuation of vertical perioral rhytides. Skin
around the eyes is pulled onto the bony orbit—
away from the globe—before treatment. If the upper eyelid is to be treated directly, plastic corneal
shields are placed and only the 0.25-cm2 superficial
eyelid tip is used. Energy settings are decreased appropriately in areas of thinner skin and where facial
fat pads are more superficial (eg, over the temple
and cheeks). Coupling fluid is applied to ensure
uniform transfer of MRF energy from the treatment
tip to the skin.
The multiple pass treatment regimen is used to
maximize skin tightening and contour enhancement. The first set of passes covers the entire surface
area to achieve uniform tightening of the skin. The
next set is performed along superior and lateral vectors to provide lifting of facial structures. The final
set is performed to achieve three-dimensional contouring and tightening. Stacking of treatment
pulses one on top of the other may be performed
at this last stage to enhance inward contraction
(eg, to define the submentum). If the patient is unable to tolerate the treatment, MRF energy levels are
lowered appropriately. Treatment endpoints include erythema of the skin and achieving the desired degree of correction.
Aftercare
Because the epidermal layer is preserved, there is no
need for local skin care after MRF treatment. Patients are instructed to avoid applying ice and using
anti-inflammatory medication to maximize the natural healing response and enhance collagen formation. Some immediate skin tightening is observed,
which is caused by thermally induced contraction
of the collagen scaffold, but tightening continues
over several weeks and months as a result of increased collagen production. Patients who wish to
have further tightening may benefit from additional
treatments [5,16].
Results
Some skin tightening is seen initially at the time of
MRF treatment, but the effect peaks at 2 to 3
months and seems to persist for several years. A degree of tightening is always observed but is usuallymore obvious in patients with thin skin who do not
have significant laxity. Tightening produces contour
improvements over underlying structures. Forehead
and periorbital treatment typically produces 2 to
3 mm of brow elevation. Improvements in skin texture and tone and acne reduction also have been
observed (Fig. 4).
Because the ThermaCool effect is currently limited to skin tightening, combining this capability
with other aesthetic procedures serves to enhance
the final result. Microdermabrasion, superficial
and medium depth peels, intense pulsed light,
and nonablative lasers are useful for treating epithelial surface irregularities and dyschromias not directly addressed by MRF treatment. Tissue fillers
175
176
Abraham & Mashkevich
can fill deeper folds, and neurotoxins can help eliminate dynamic expression lines (Fig. 5). Liposuction
and fat transfer can provide additional tissue sculpting. Minimally invasive and percutaneous suture
techniques can be used concurrently to resuspend
underlying ptotic structures, such as the malar
and jowl fat pads, and the platysma in the neck
(Fig. 6).
Complications
Reported side effects are transient and most commonly include mild tissue edema and erythema,
which resolve in a matter of days. Temporary paresthesia, if it occurs, dissipates over a period of a few
weeks as inflammation around sensory nerves gradually subsides. Similarly, focal inflammation of the
platysma and neck soreness can last for a few weeks.
Occasional burns, blisters, and surface irregularities
have been reported but represent rare side effects of
treatment. Small second-degree burns were more
commonly observed in initial reports with higher
energy settings compared with settings used currently [6]. Skin dimpling as a result of focal collagen
contraction or possibly underlying fat atrophy occurs rarely and has been noted to improve with
time without any further intervention [1,18]. Subcision and autologous fat transfer also have been advocated [27]. According to the manufacturer, of the
more than 280,000 ThermaCool procedures performed worldwide, most (>99.8%) have not
involved any adverse events. Undoubtedly, experience with energy settings and treatment techniques
plays a role in obtaining optimal outcomes and
minimizing the potential for side effects.
Future directions
Developments in the capability of the MRF treatment tips to measure tissue impedance should
lead to more precise and better individualized treatment settings with reduced risks and optimized outcomes [26]. Further control and refinement of
radiofrequency energy delivery to target tissues
has the potential to improve on the clinical outcomes currently reported with this technology. Targeted delivery of radiofrequency energy to deeper
structures (eg, adipose tissue) may ultimately allow
contouring of tissue in ways currently unattainable
without surgery. The next step in the evolution of
MRF technology will most likely involve combination with other nonablative treatment modalities
to provide a synergistic result.
Summary
Monopolar radiofrequency skin tightening represents an exciting frontier in facial cosmetic
rejuvenation. Published reports to date document
the clinical efficacy of this noninvasive technology
and support its further development and use in
the treatment of facial and neck rhytides. MRF facial
skin rejuvenation may be used as a stand-alone modality or in conjunction with other invasive and
noninvasive treatments to maximize aesthetic results. High patient satisfaction rates are predicated
on thorough education, emphasis on realistic expectations, and sound treatment planning. For facial plastic surgeons looking to meet the growing
demand for nonablative skin tightening, MRF is
currently the most established option.
References
[1] Abraham MT, Vic Ross E. Current concepts in
nonablative radiofrequency rejuvenation of the
lower face and neck. Facial Plast Surg 2005;
21(1):65–73.
[2] Zelickson BD, Kist D, Bernstein E, et al. Histological and ultrastructural evaluation of the effects
of a radiofrequency-based nonablative dermal
remodeling device. Arch Dermatol 2004;
140(2):204–9.
[3] Meshkinpour A, Ghasri P, Pope K, et al. Treatment of hypertrophic scars and keloids with a radiofrequency device: a study of collagen effects.
Lasers Surg Med 2005;37:343–9.
[4] Burns AJ, Holden SG. Monopolar radiofrequency
tightening: how we do it in our practice. Lasers
Surg Med 2006;38:575–9.
[5] Fritz M, Counters JT, Zelickson BD. Radiofrequency treatment for middle and lower face laxity. Arch Facial Plast Surg 2004;6:370–3.
[6] Fitzpatrick R, Geronemus R, Goldberg D, et al.
Multicenter study of noninvasive radiofrequency
for periorbital tissue tightening. Lasers Surg Med
2003;33:232–42.
[7] Ruiz-Esparza J. Noninvasive lower eyelid blepharoplasty: a new technique using nonablative
radiofrequency on periorbital skin. Dermatol
Surg 2004;30:125–9.
[8] Abraham M, Chiang S, Keller G, et al. Clinical evaluation of non-ablative radiofrequency facial rejuvenation. J Cosmet Laser Ther 2004;6:136–44.
[9] Bassichis BA, Dayan S, Thomas JR. Use of nonablative radiofrequency device to rejuvenate the
upper one-third of the face. Otolaryngol Head
Neck Surg 2004;130:397–406.
[10] Nahm WK, Su TT, Rotunda AM, et al. Objective
changes in brow position, superior palpebral
crease, peak angle of the eyebrow, and jowl surface area after volumetric radiofrequency treatments to half of the face. Dermatol Surg 2004;
30:922–8.
[11] Alster TS, Tanzi E. Improvement of neck and
cheek laxity with a nonablative radiofrequency
device: a lifting experience. Dermatol Surg 2004;
30:503–7.
Monopolar Radiofrequency Skin Tightening
[12] Ruiz-Esparza J, Gomez JB. The medical facelift:
a noninvasive, nonsurgical approach to tissue
tightening in facial skin using nonablative radiofrequency. Dermatol Surg 2003;29:325–32.
[13] Iyer S, Suthamjariya J, Fitzpatrick RE. Using a radiofrequency energy device to treat the lower
face: a treatment paradigm for a nonsurgical
facelift. Cosmetic Dermatology 2003;16:37–40.
[14] Jacobson LG, Alexiades-Armenakas M, Bernstein L,
et al. Treatment of nasolabial fold and jowls with
a noninvasive radiofrequency device. Arch Dermatol 2003;139:1371–2.
[15] Hsu TS, Kaminer MS. The use of nonablative radiofrequency technology to tighten the lower
face and neck. Semin Cutan Med Surg 2003;22:
115–23.
[16] Koch RJ. Radiofrequency nonablative tissue
tightening. Facial Plast Surg Clin North Am
2004;12:339–46.
[17] Narins DJ, Narins RS. Non-surgical radiofrequency
facelift. J Drugs Dermatol 2003;2:495–500.
[18] Weiss RA, Weiss MA, Munavalli G, et al. Monopolar radiofrequency facial tightening: a retrospective analysis of efficacy and safety in over 600
treatments. J Drugs Dermatol 2006;5:707–12.
[19] Fisher GH, Jacobson LG, Bernstein LJ, et al. Nonablative radiofrequency treatment of facial laxity.
Dermatol Surg 2005;31:1237–41.
[20] Finzi E, Spangler A. Multipass vector (mpave)
technique with nonablative radiofrequency to
[21]
[22]
[23]
[24]
[25]
[26]
[27]
treat facial and neck laxity. Dermatol Surg
2005;31:916–22.
Kushikata N, Negishi K, Tezuka Y, et al. Nonablative skin tightening with radiofrequency in
Asian skin. Lasers Surg Med 2005;36:92–7.
Alam M, Levy R, Pavjani U, et al. Safety of radiofrequency treatment over human skin previously
injected with medium-term injectable soft-tissue
augmentation materials: a controlled pilot trial.
Lasers Surg Med 2006;38(3):205–10.
Shumaker PR, England LJ, Dover JS, et al. Effect
of monopolar radiofrequency treatment over
soft-tissue fillers in an animal model: part 2. Lasers Surg Med 2006;38(3):211–7.
Fitzpatrick TB. The validity and practicality of
sun-reactive skin types I through VI. Arch Dermatol 1988;124:869–73.
Kushikata N, Negishi K, Tezuka Y, et al. Is topical
anesthesia useful in noninvasive skin tightening
using radiofrequency? Dermatol Surg 2005;31:
526–33.
Lack EB, Rachel JD, D’Andrea L, et al. Relationship of energy settings and impedance in different anatomic areas using a radiofrequency
device. Dermatol Surg 2005;31:1668–70.
Narins RS, Tope WD, Pope K, et al. Overtreatment effects associated with a radiofrequency tissue-tightening device: rare, preventable, and
correctable with subcision and autologous fat
transfer. Dermatol Surg 2006;32:115–24.
177
Aesthetic Surgery Journal
http://aes.sagepub.com/
Tissue Tightening Technologies: Fact or Fiction
Neil Sadick
Aesthetic Surgery Journal 2008 28: 180
DOI: 10.1016/j.asj.2007.12.009
The online version of this article can be found at:
http://aes.sagepub.com/content/28/2/180
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Review Article
Tissue Tightening Technologies:
Fact or Fiction
Neil Sadick, MD
Skin laxity is associated with chronological aging and exposure to solar radiation. The authors summarize the
advantages and limitations of current laser, light-, and radiofrequency (RF)-based technologies purported to
treat skin laxity by effecting heat-induced collagen contraction and subsequent remodeling during the months
after treatment. Although penetration of laser or broadband light to the deep dermal layers is limited because
of scattering of the light by epidermal melanin, a new device in which broadband infrared light is minimally
scattered may overcome these limitations. RF energy offers a treatment alternative that has not only been
proven to promote collagen contraction and remodeling but also is not scattered by epidermal constituents.
Recently launched devices that use combinations of optical and RF energy achieve clinical benefits at lower
and therefore safer levels of energy, with only mild pain and few adverse effects. A combined infrared-RF
device takes maximum advantage of both optical and RF technologies to achieve the desired clinical effect.
The electrooptical synergy systems have proven to be safe, effective, reliable, and user-friendly. Other more
advanced powerful technologies may also be effective in this setting. (Aesthetic Surg J 2008;28:180–188.)
T
issue tightening refers to the correction of skin laxity. Suitable patients for nonsurgical skin tightening are those who do not want surgery or are poor
candidates for rhytidectomy.1 In addition, some patients
who have undergone a face lift procedure have found
that postoperative nonsurgical skin tightening enhances
their results.
MECHANISM OF COLLAGEN SHRINKAGE
Collagen is a polymer that exists as a triple helix with
chains held together by hydrogen bonds. These molecules are aggregated and organized as fibrils with tensile
properties attributable to intermolecular cross-links.2
When collagen is denatured by heat, the intramolecular
hydrogen bonds rupture and the triple helices “unwind
to produce a gel of random-coil molecules.”3 Tissue tension in human skin increases because, although the
fibers become shorter,4,5 the heat-stable cross-links
between molecules are maintained, thus increasing the
rubber-elastic properties of the collagen polymer.4 The
heat-modified tissues then undergo remodeling associated with fibroplasia and new collagen deposition.2,3
When denaturation is complete, further increases in
temperature result in additional fiber shortening, probably because of peptide bond hydrolysis.4,5 The mechanism of collagen shrinkage has been described in detail.3
Dr. Sadick is Clinical Professor of Dermatology at Weill Cornell
Medical College, New York, NY.
180 • Volume 28 • Number 2 • March/April 2008
The temperature at which collagen shrinkage occurs is
often quoted as 65° C.2,3 However, collagen denaturation
is described by the Arrhenius equation given by k !
Ae"Ea/RT, in which k is the rate constant, A represents the
frequency of collisions between reacting molecules, Ea is
the energy of activation, R is the gas constant, and T is
the absolute temperature.6 According to this equation,
shrinkage of collagen depends on time, as well as temperature, and collagen contraction occurs at a variety of timetemperature combinations rather than at a specific
temperature.4 That said, it has been suggested that for
millisecond exposures, collagen shrinkage will occur only
at temperatures exceeding 85° C, whereas for exposures
of several seconds, shrinkage will occur at 60° to 65° C.4
TECHNOLOGIES
Treatment options for nonsurgical skin tightening are
based on heat-induced damage to tissue by light,
radiofrequency, or both types of energy. Recent nonsurgical approaches to skin tightening have been reviewed.7
Laser and Light-Based Devices
The design of laser and light devices is based on the
principle of selective photothermolysis, which states that
the laser wavelength must be more strongly absorbed by
the target tissue than surrounding tissues, the amount of
energy (fluence) must exceed the therapeutic threshold
of the target, and energy must be delivered within the
thermal relaxation time of the target tissue.8,9 The therapeutic threshold is the minimum amount of energy to
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Aesthetic Surgery Journal
Table. Investigations of the Titan Device for skin tightening
Reference
(no.)
No. of Treatment
Patients areas (no.)
200617
No. of treatments/
Treatment interval (no.)
Results (no.)
Adverse events (no.)
Excellent (13)
Moderate (3)
Minimal (6)
No change (3)
Pain during treatment at
higher fluences, edema*
25
Eyebrow only (3)
1 session (20)
lower face only (1) 2 sessions (4)
cheek and neck (21) 3 sessions (1)
Taub et al.21
42
Face, lower face,
neck areas
2 sessions/2.5 to 7 weeks (41) None (4)
3 sessions/ 4 weeks (1)
Mild (15)
Moderate (14)
Marked (8)
Outstanding (1)
Mild transient discomfort
during treatment,
edema and erythema
after treatment
Chua et al.22
21
Face and neck
3 sessions/ 4 weeks
Minimal to no pain,
occasional superficial
blistering
Goldberg et al.23
12
Lower neck and face 2 sessions/ 1 month
Ruiz-Esparza
43% good
improvement,
38% moderate
improvement,
19% mild improvement
at 6 months
Obvious clinical improvement Mild transient erythema
in 11 of 12 subjects;
dramatic changes for patients
whose laxity draped separately
from deeper soft tissue
*First 5 patients treated with topical anesthetic.
achieve the therapeutic goal, and the thermal relaxation
time is the time for the target structure to lose 50% of
the delivered energy.9
Ablative treatments of facial skin with CO2 or Er:YAG
laser devices have been shown to cause collagen contraction and remodeling associated with tightening skin
and reducing wrinkles. Although the efficacies of these
modalities are striking, erythema, pigmentary changes,
infection, dermatitis, scarring, and long recovery times
are common.10 The risk of adverse effects depends on
the experience of the treating physician and have been
reduced somewhat by improved laser design.11
Nonablative laser and broadband light devices have
been developed to reduce both recovery time and the risk
of adverse effects associated with ablative treatments.
Beams of these devices inflict thermal damage to the lower layers of the dermis and stimulate collagen production
but do not injure the epidermis.12,13 Wavelengths ranging
from 532 to 1540 nm and intense pulsed light (IPL) have
been used with varying levels of success.12
Although collagen remodeling has been histologically
proven to occur after laser or light-based treatments,12,14
correlation of remodeling with clinical improvement has
been variable.12,15 Full-face treatments with either IPL or
a combination of 532-nm and 1064-nm laser devices
appear to stimulate overall collagen remodeling and provide higher patient satisfaction.12 Stimulation of new
collagen production by laser treatments has been attributed to the release of inflammatory mediators from vascular epithelial cells.16 When stimulated by treatment
with a combination of 532-nm and 1064-nm laser
Tissue Tightening Technologies: Fact or Fiction
devices, collagen remodeling continues for 6 to 12
months and increases with the number and intensity of
treatments, as well as elapsed time.12 Nonablative laser
treatments are suitable for patients desiring short recovery times and minimal adverse effects and willing to
accept mild improvement at considerable expense.15
The Titan system (Cutera, Inc., Brisbane CA) uses
broad-spectrum (1100–1800 nm) infrared (IR) light in
multisecond cycles to heat water in the dermis. With this
band, bulk heating of the dermis is maximized at 1 to 3
mm, and absorption by melanin and hemoglobin is low.
Heating also occurs at 5 mm.17 The epidermis is protected by a cooling head.18 Collagen fibril denaturation after
4 passes has been shown by electron microscopy.19
Multiple passes at low energy levels are associated with
bulk dermal heating and collagen contraction.20 Patient
discomfort is minimal, and anesthesia is not required.18
Studies evaluating the efficacy and safety of the Titan
device are summarized in the Table.
Overall the effectiveness of the Titan system for skin
laxity is apparent for all skin types and ages. Low fluences may be used, thus negating the need for anesthetics. One to 2 sessions 1 month apart seems to be
effective for most patients. Although an immediate tightening effect is achieved, the full effect often does not
appear for 6 months or longer, which makes management of patient expectations crucial.
Radiofrequency Devices
An alternative to nonablative light-based systems are
devices that use radiofrequency (RF) energy to heat, lift,
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Volume 28 • Number 2 • March/April 2008 • 181
and tighten dermal tissue.13,24 RF energy is actually alternating current that flows from the tip of an electrode to
tissue with which it is in contact. In direct current, the
flow of electrons is in one direction, whereas in alternating current the direction of flow cycles back and forth at a
certain frequency. In medicine, the frequency of the alternating current in RF devices is typically 0.3 to 10 MHz.25
In biological tissue, the thermal effects of RF devices
depend on the electrical characteristics of the tissue. In
RF devices, alternating current flows from an electrode
to tissue with which it is in contact. When the current
enters tissue, ions in the tissue try to follow the highfrequency changes in the direction of the current. The
resulting ionic agitation constitutes opposition to the
flow of alternating current (impedance) and results in
the production of heat in the tissue.26,27 The amount of
heat produced depends on the impedance, the square of
the intensity of current, and length of time the skin is
exposed to RF energy.27
RF devices designed to remodel cutaneous tissue have
been reviewed13 and will be summarized here. Unlike
laser and light energy, RF current is not scattered by tissue or absorbed by epidermal melanin. Patients of all
skin types can therefore be treated, and considerable
heat can be generated in the dermal layers to stimulate
collagen contraction and neocollagenesis.13 RF devices
are also considerably less expensive than laser devices.26
In dermatology, RF devices are either monopolar
(unipolar), bipolar, or both. In monopolar systems, current flows from an active electrode in contact with the
treated area to a large grounding electrode positioned on
the skin far from the active electrode.13,25 The current
flows from the electrode contact point, through the
entire body (the path of least resistance), and to the
grounding pad. Most of the heat is produced in the tissue just beneath the electrode, and little heat is generated at the grounding pad13,26 because the energy
diminishes with distance from the active electrode.
Monopolar RF energy is the first nonsurgical procedure
developed for tightening facial skin.28
The advantage of monopolar RF devices is that the
current penetrates to the deeper layers of skin (eg, 5 mm
depth for a 10-mm electrode).25,27 Unfortunately the high
levels of energy and the deep penetration also cause
pain during treatment, necessitating the use of anesthesia.27 The ThermaCool TC (Thermage, Inc., Hayward,
CA) is a nonablative monopolar RF device.
Bipolar RF devices use 2 electrodes positioned at fixed
distances from each other. Both electrodes are in contact
with the area to be treated. In these devices, alternating
current enters the skin from the active electrode and
passes only through the tissue between the 2 electrodes.
Because the current does not pass through the entire
body, grounding pads are not needed.13 In these devices,
the penetration depth of the electrical current is approximately half the distance between the electrodes.25 The
advantage of bipolar systems is that the distribution of
current inside the tissue can be controlled. The Aurora
182 • Volume 28 • Number 2 • March/April 2008
(Syneron Medical Ltd., Yokneam, Israel) is a combined
bipolar RF and optical energy device.
Monopolar radiofrequency. The efficacy and safety of
the ThermaCool device for skin tightening have been
evaluated in a variety of studies.1,6,10, 29-34A histologic
study suggests that collagen fibrils contract immediately
after treatment, and neocollagenesis is induced as part
of a wound-healing response.35 In this device, a 6-MHz
monopolar current signal is produced in a disposable
capacitive membrane tip that treats a 1.0- to 1.5-cm2
area to depths of 3 to 6 mm. A cryogen gas spray device
cools the skin surface before, during, and after the delivery of RF current. The balance between the superficial
cooling and the deep tissue heating creates a reverse
thermal gradient in which the dermis receives the most
intense heat without affecting the skin surface.13
The initial protocol for treating patients with the
ThermaCool device was a single pass at the highest
energy patients could tolerate. With this technique,
patient feedback was negative, and clinical results were
not optimal.36 In a study comparing a new technique of
multiple passes at lower energies with the traditional
single-pass high energy technique, Kist et al37 showed
that the multiple pass-lower fluence technique resulted
in fewer side effects, less pain during treatment, and
more consistent clinical improvement. Bogle et al38
reported similar findings by use of a similar technique to
tighten the skin of the lower face. The use of a 3-cm2 tip
and a generous amount of coupling fluid has also been
reported.38 A new 0.25-cm2 treatment tip has been evaluated for the treatment of human eyelids.39,40
A retrospective study of more than 600 patients treated with the ThermaCool showed that the most frequent
adverse effects are temporary erythema and edema.41
Reliance on patient feedback to adjust treatment settings, with a treatment grid used to prevent overlaps and
delaying for an appropriate time between passes has
been recommended to minimize adverse effects.42
Combined bipolar radiofrequency and optical energy. A
technology that combines RF and optical energies, called
electrooptical synergy (ELOS), is designed to overcome the
limitations of light-based systems alone.13,25 The ELOS
technology is used in the Aurora DS, Polaris WR (Syneron
Medical Ltd.), and the Galaxy (Syneron Medical Ltd.).
The Aurora DS system delivers pulses of IPL
(400–980, 580–980, and 680–980 nm) and bipolar RF
energies simultaneously, but the RF pulse has a longer
duration than the IPL pulse to allow the IPL component
to preheat the dermal target. Preheating increases the
temperature of the target over that of the surrounding
tissue, thus reducing its impedance and preferentially
attracting the RF current.13,25,43 (The higher the temperature of a target, the lower its impedance, and the greater
its attraction to an electrical current.43) In an evaluation
of the Aurora system for skin rejuvenation,44 58% of
patients were satisfied with improvement in skin laxity
after 1 to 2 treatments. In a later study of 108 patients,
Sadick et al45 achieved 62.9% overall improvement in
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Aesthetic Surgery Journal
A
skin laxity, 75.3% overall skin improvement, and an
8.3% minor complication rate.
The Polaris WR system combines 900-nm diode laser
and bipolar RF energies. The diode laser component
treats superficial rhytids, blood vessels, and pigmentation, whereas the RF stimulates collagen production at
deeper levels.13 In their study of 20 patients treated 3
times at 3-week intervals, Doshi and Alster46 showed
modest improvement in wrinkles 6 months after the
final treatment and progressive improvement in skin laxity during this follow-up period in most patients.
Adverse effects were minimal, and 80% of patients experienced mild discomfort during treatment. In a multicenter trial,47 more than half of 23 patients treated similarly
achieved greater than 50% improvement in wrinkle
appearance, and all patients experienced improvement
in skin texture and smoothness.
In 2006, Alexiades-Armenakas48 treated 28 patients 1
to 5 times, first with the Polaris and then the Aurora or
Galaxy system in each session. With a comprehensive
grading scale for various categories of photoaging, she
achieved 22% overall improvement in skin laxity.
Combined unipolar and bipolar radiofrequency. The
Accent (Alma Lasers, Inc., Ft. Lauderdale, FL) RF system
includes both a unipolar handpiece for volumetric heating of the subcutaneous adipose tissue and a bipolar
handpiece for nonvolumetric heating of the dermis.
During treatment the skin is slowly heated to the
patient’s pain threshold (40° to 44° C) and kept within
that temperature range for approximately 2 minutes
before the physician moves to the next treatment area to
repeat the process. The use of this device for the treatment of cellulite and the subcutaneous tissue of the buttocks and thighs has been evaluated.49
A case study of a patient treated for skin laxity, texture, firmness, and volume reduction with the Accent
device and a monopolar RF device (ThermaCool) has
been reported.50 A 60-year-old woman (skin type III)
underwent a series of 6 Accent treatments (at 2-week
intervals) to tighten the skin of her left upper arm. The
unipolar handpiece was used to a maximum skin temperature of 42.5° C. The right upper arm had been treated in a single session with the ThermaCool device
according to the device’s original protocol and using a
3-cm2 treatment tip. Although the patient was pleased
with the results for both upper arms, she believed that
the skin in her Accent-treated area was tighter and
firmer. An additional 3 treatments with the Accent
device were given to the upper left arm with the combination of both unipolar and bipolar RF handpieces. The
upper right arm received 2 similar treatments. The result
was further tightening without adverse effects. Although
multiple treatments with the Accent device were
required to achieve the desired result, treatment times
were relatively short, and a new disposable tip is not
required for each treatment.
Vacuum-assisted radiofrequency. The safety and efficacy of an investigational prototype of a vacuum-assistTissue Tightening Technologies: Fact or Fiction
ed bipolar RF device (Lumenis, Inc., Santa Clara, CA)
has been evaluated for the treatment of wrinkles and
elastosis.51,52 The device uses functional aspiration-controlled electrothermal stimulation technology in which
skin, with the aid of a vacuum system, is folded to a predetermined depth between the bipolar RF electrodes during treatment. The rationale for this approach is that by
restricting the volume of treated tissue to that positioned
between the 2 electrodes, physicians can treat both
superficial and deep layers with less energy than would
be required from the unfolded skin surface. In other
words, the vacuum system reduces the distance between
the source of RF energy and the dermis, theoretically
reducing pain and increasing safety during treatment.
In this study, 46 patients received 8 facial treatments
at 1- to 2-week intervals. Pain levels were low, patients
were satisfied with the treatment, facial wrinkling and
elastosis scores improved significantly 6 months after
the final treatment, and adverse effects were limited primarily to temporary erythema, burns, and blisters. The
authors are currently using 3-dimensional imaging tools,
measuring skin elasticity, and conducting histologic
studies to further investigate the cutaneous effects of
this device. Preliminary results of an ongoing investigation53 of this device for nonfacial skin tightening are
promising as well.
The device evaluated by Gold et al53 was further evaluated as the Aluma by Montesi et al.27 Thirty patients
received 6 to 8 treatments at 2-week intervals. Biopsy
samples were taken from 15 of these patients before the
first treatment and 3 months after the final treatment.
Among the treated imperfections (periorbital wrinkles,
glabellar wrinkles, slack cheeks, striae distensae, and
acne scars), the most improvement (clinical, histologic,
and immunohistochemical) occurred in the abdominal
striae distensae. In most cases, side effects were limited
to transient rashes and ecchymosis. Improvements
appeared to continue for at least several months after
the final treatment.
Combinations With Other Technologies
Combination infrared and bipolar radiofrequency. A
device that combines 700- to 2000-nm infrared (IR) and
bipolar RF energies (ReFirme ST Applicator, Syneron
Medical Ltd.) has been evaluated for the treatment of
facial laxity in Asian patients.54 In this prospective study,
19 patients (skin types III to V) with skin laxity and periorbital rhytids were given 3 full-face nonablative treatments at 3-week intervals with the combined IR-RF
device. Clinical end points were skin tightening and edema and anesthesia was not used. Standardized photographs were obtained with the Canfield Visia CR system
(Canfield Scientific, Inc., Fairfield, NJ) before treatment
and serially for 3 months after the final treatment.
All subjects completed the study. At 3 months, the
authors observed mild improvement in skin laxity in the
mid and lower face. Statistically significant improvement
was found (by blinded assessors) in the cheek, jowl, and
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Volume 28 • Number 2 • March/April 2008 • 183
nasolabial folds. Patients reported high overall satisfaction, with 89.5% achieving moderate to significant
improvement in skin laxity of the cheek, jowl, periorbital
area, and upper neck. Subjective improvement in skin
laxity was noticed in all patients after the initial treatment session. Most patients experienced mild pain, and
only 3 patients reported moderate pain. Temporary erythema occurred after all treatments, and 3 patients had
edema that disappeared within 24 hours. Superficial
crusting was the primary side effect.
The results of this study suggest that the combined
IR and RF energy device achieves clinical improvement
in skin laxity and rhytids at 10 J/cm2 optical fluence,
which is lower than the 32- to 40-J/cm2 fluences used
by Doshi and Alster46 and the 30- to 50-J/cm2 fluences
used by Sadick and Trelles47 with the Polaris WR system. Yu et al54 speculate that the IR energy, because it
is absorbed by water and possibly collagen, has a more
direct dermal heating effect than the 900-nm diode
energy, which is absorbed by hemoglobin, thus heating
the dermis only indirectly.
In a 2-center study of subjects with skin types I to III,
31 subjects received 2 to 5 treatments with the combined
IR-RF device at 3- to 4-week intervals without anesthesia.
Blinded evaluators compared pretreatment and posttreatment photographs to assess wrinkle clearance rates, and
patients graded satisfaction with the treatment. The overall median clearance rate for wrinkles was 50%, and the
median patient satisfaction rate was 7 on a 10-point scale
in which 1 to 2 is not satisfied and 9 to 10 is exceptionally satisfied. Adverse effects were limited to mild temporary erythema and edema.55 Clinical examples of patients
treated with the Refirme are shown in Figures 1 to 4.
Combination fractional infrared and radiofrequency.
The Matrix IR Fractional Treatment Applicator (Syneron
A
Medical Ltd.) that combines fractional 915-nm diode
laser energy with RF has become available for wrinkle
reduction. Designed for deep dermal heating, the new
ELOS device creates microthermal thermal bands while
leaving surrounding tissues undamaged to promote rapid
healing and minimize downtime. Improvement in wrinkles is noticeable after 2 to 3 sessions.55 In the author’s
experience, the ELOS systems have proven to be safe,
effective, reliable, and user friendly.
The LuxIR and LuxDeepIR Fractional Infrared
Handpieces (Palomar Medical Technologies, Inc.,
Burlington, MA) are designed to deliver IR radiation 1.5
to 3 mm into the dermis and 1.0 to 4.0 mm into the dermis and fat layer, respectively, without damaging the epidermis and upper dermis. The LuxDeep IR Handpiece
offers a longer pulse duration and more powerful cooling.
The resulting soft-tissue coagulation can lead to collagen
remodeling and tighter skin, according to the manufacturer. Each beam of the array of small beams creates a
“lattice of hyperthermic islets” surrounded by undamaged tissue, a pattern considered to expedite healing and
collagen remodeling. Skin is cooled before, during, and
after each pulse to minimize patient discomfort.56
The Affirm (Cynosure, Inc., Westford, MA) is a 1440nm Nd:YAG laser device with combined apex pulse
(CAP) technology for the treatment of photodamaged
skin. The CAP technology produces a pattern of coagulated tissue columns surrounded by unaffected tissue.
The coagulated columns are created by high-fluence
“apexes,” whereas the surrounding uncoagulated
columns are produced by lower background fluences.
The entire treated area is heated, but the coagulated
columns are heated more than the uncoagulated tissue.57
In solar elastosis, bundles of loosely packed collagen58
are present in the dermis at depths of 100 to 400 #m.59,60
B
Figure 1. A, Pretreatment view. B, Posttreatment view 1 month after treatment with a device combining IR and bipolar RF energies.
184 • Volume 28 • Number 2 • March/April 2008
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Aesthetic Surgery Journal
hours, thus minimizing the risk of complications.
Clinical trial results show that efficacy appears to be
comparable to that of an ablative CO2 laser device, and
patients may not require preoperative or postoperative
pain medication. A major advantage is that clinical benefits appear after a single treatment. Preliminary results
of 12-month clinical trials have been presented.63
The ActiveFX Fractional CO2 laser device with
UltraPulse Encore (Lumenis, Inc.) provides noticeable
clinical benefit after a single treatment, with less downtime than traditional CO2 laser devices.64 The novel
device ablates a fraction of the surface, which permits
bridges of undamaged tissue to promote rapid reepithelialization. Clinical trials have shown the efficacy of the
ActiveFX in the treatment of dyschromia, rhytids, and
skin laxity, and that treatment parameters can be adjusted to minimize erythema and edema.
FRACTIONAL INFRARED DEVICES
Figure 2. Split screen view demonstrates patient before and immediately after 1 treatment combining IR and bipolar RF energies.
CAP energy can penetrate up to 400 µm, where most sun
damage occurs.4 Healing after treatment with the Affirm
is rapid, and collagen remodeling is stimulated in the
uncoagulated tissue, as well as in the coagulated
columns, a benefit supported by histologic studies.57,61
FRACTIONAL ABLATIVE DEVICES
The new Lux2940 (Palomar Medical Technologies, Inc.)
is a microfractional Er:YAG laser device designed to ablatively reduce wrinkles, improve skin texture, and reduce
hyperpigmentation with 3 to 4 days downtime, much
less than traditional ablative laser devices.62 Histologic
analyses show that the microcolumns close within 12
A
A fractional IR device (LuxIR Fractional Infrared
Handpiece, Palomar Medical Technologies, Inc.) uses
850- to 1350-nm pulses to cause soft tissue coagulation
and collagen remodeling in deep (1.5–3 mm) dermal layers. The IR light is delivered in an array of small regularly
spaced beams that produce hyperthermic islets surrounded by undamaged tissue.65 A similar device
(LuxDeepIR Fractional Infrared Handpiece, Palomar
Medical Technologies, Inc.) delivers IR light to 4 mm
while cooling the epidermis.66
New IR Devices
A new broadband (800–1400 nm) light source (SkinTyte,
Sciton, Inc., Palo Alto, CA) selectively coagulates soft tissue and targets dermal collagen while cooling the epidermis before, during, and after treatment. The device
B
Figure 3. A, Pretreatment view. B, Posttreatment view 1 month after 1 treatment combining IR and bipolar RF energies.
Tissue Tightening Technologies: Fact or Fiction
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Volume 28 • Number 2 • March/April 2008 • 185
A
B
Figure 4. A, Pretreatment view. B, Posttreatment view 1 month after 5 treatments combining IR and bipolar RF energies.
can be used to treat large surface areas such as the legs,
arms, and abdomen.67
The ST module of the Harmony platform (Alma
Lasers, Inc.) uses near-IR (780–1000 nm) light to shrink
collagen and induce collagen remodeling in the papillary and upper reticular dermis, resulting in skin tightening. The near-IR light heats subdermal connective
tissue and proteins and is absorbed minimally by epidermal water, thus eliminating the need for aggressive
epidermal cooling. The hand piece is stationary while
held against skin during treatment.68 Noticeable tightening requires 5 or 6 treatment sessions.69
OTHER NEW DEVICES
Additional recently launched devices for skin tightening
include the C-Sculpt (DermaMed International, Inc.,
Lenni, PA), which uses a 626-nm LED with cooling and
massage, and the Surgitron Dual Frequency RF (Ellman
International, Inc., Oceanside, NY).70
CONCLUSION
A variety of devices to tighten skin have been either
improved or launched since the first skin-tightening
device became available. Laser and light-based devices,
although effective, have limited efficacy because of scattering of light by epidermal constituents. A combined
IR-RF device used by the author takes maximum advantage of both optical and RF technologies to achieve the
desired clinical effect. Other more advanced powerful
technologies may also be effective in this setting. ◗
DISCLOSURES
The author has no disclosures with respect to the contents of this
article.
186 • Volume 28 • Number 2 • March/April 2008
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Otolaryngol Head Neck Surg 2004;130:397-406.
33. Fritz M, Counters JT, Zelickson BD. Radiofrequency treatment for
middle and lower face laxity. Arch Facial Plast Surg 2004;6:370-373.
34. Finzi E, Spangler A. Multipass vector (mpave) technique with nonablative radiofrequency to treat facial and neck laxity. Dermatol Surg
2005;31:916-922.
35. Zelickson B, Kist D, Bernstein E, Brown DB, Ksenzenko S, Burns J, et
al. Histological and ultrastructural evaluation of the effects of a
radiofrequency based nonablative dermal remodeling device: A pilot
study. Arch Dermatol 2004;140:204-209.
36. Burns AJ, Holden SG. Monopolar radiofrequency tissue tightening—
how we do it in our practice. Lasers Surg Med 2006;38:575-9.
37. Kist D, Burns AJ, Sanner R, Counters J, Zelickson B. Ultrastructural
evaluation of multiple pass low energy versus single pass high energy
radio-frequency treatment. Lasers Surg Med 2006;38:150-154.
38. Bogle M, Ubelhoer N, Weiss RA, Mayoral F, Kaminer MS. Evaluation of
the multiple pass, low fluence algorithm for radiofrequency tightening
of the lower face. Lasers Surg Med 2007;39:210-217.
39. Biesman B, Carruthers J, Baker S, Leal H. Monopolar treatment of
human eyelids: A prospective evaluation. Lasers Surg Med 2006;S18:94.
40. Carruthers J, Carruthers A. Shrinking upper and lower eyelid skin with
a novel radiofrequency tip. Dermatol Surg 2007;33:802-809.
41. Weiss RA, Weiss MA, Munavalli G, Beasely KL. Monopolar radiofrequency facial tightening: a retrospective analysis of efficacy and safety
in over 600 treatments. J Drugs Dermatol 2006;5:707-12.
Tissue Tightening Technologies: Fact or Fiction
42. Narins RS, Tope WD, Pope K, Ross EV. Overtreatment effects associated
with a radiofrequency tissue-tightening device: rare, preventable, and
correctable with subcision and autologous fat transfer. Dermatol Surg
2006;32:115-124.
43. Sadick NS. Electro-optical synergy in aesthetic medicine: novel technology, multiple applications. Cosmetic Dermatology 2005;18:201-206.
44. Bitter P Jr, Mulholland S. Report of a new technique for enhanced noninvasive skin rejuvenation using a dual mode pulsed light and radiofrequency energy source: selective radiothermolysis. J Cosmet Dermatol
2002;1:142-143.
45. Sadick NS. Alexiades-Armenakas M, Bitter P Jr, Hruza G, Mulholland RS.
Enhanced full-face skin rejuvenation using synchronous intense pulsed
optical and conducted bipolar radiofrequency energy (ELOS): introducing
selective radiophotothermolysis. J Drugs Dermatol 2005;4:181-1286.
46. Doshi SN, Alster TS. Combination radiofrequency and diode laser for
treatment of facial rhytides and skin laxity. J Cosmet Laser Ther
2005;7:11-15.
47. Sadick NS, Trelles MA. Nonablative wrinkle treatment of face and neck
using a combined diode laser and radiofrequency technology. Dermatol
Surg 2005;31:1695-1699.
48. Alexiades-Armenakas M. Laser skin tightening: non-surgical alternative
to the face lift. J Drugs Dermatol 2006;5:295-296.
49. Emelia del Pino M, Rosado RH, Azuela A, Graciela Guzmán M,
Argüelles D, Rodríguez C, et al. Effect of controlled volumetric tissue
heating with radiofrequency on cellulite and the subcutaneous tissue of
the buttocks and thighs. J Drugs Dermatol 2006;5:714-22.
50. Mayoral FA. Skin tightening with a combined unipolar and bipolar
radiofrequency device. J Drugs Dermatol 2007;6:212-215.
51. Gold MH, Goldman MP, Carcamo AS, Ehrlich M. Treatment of wrinkles
and skin tightening using bipolar, vacuum-assisted radiofrequency
heating of the dermis. Poster presented at: Annual meeting of the
American Academy of Dermatology; March 3–7, 2006; San Francisco.
52. Gold MH, Goldman MP, Rao J, Carcamo AS, Ehrlich M. Treatment of
wrinkles and elastosis using vacuum-assisted bipolar radiofrequency
heating of the dermis. Dermatol Surg 2007;33:300-309.
53. Gold MH, Biron JS. Vacuum-assisted bipolar radiofrequency therapy for
non-facial skin tightening. Poster presented at: Annual meeting of the
European Academy of Dermatology and Venereology; May 16–20, 2007;
Vienna, Austria.
54. Yu CS, Yeung CK, Shek SY, Tse RK, Kono T, Chan HH. Combined
infrared light and bipolar radiofrequency for skin tightening in Asians.
Lasers Surg Med 2007;39:471-475.
55. Sleightholm R, Bartholomeuz H. Skin tightening and treatment of facial
rhytids with combined infrared light and bipolar radiofrequency technology. Available at http://www.Syneron.com/assets/downloads/pdf/
Refirmewhitepaper/pdf. Last accessed February 21, 2008.
56. Skin Tightening Through Soft Tissue Coagulation. Palomar Medical
Technologies Web site. Available at:
http://www.palomarmedical.com/palomar.aspx?pgID!1049. Accessed
October 9, 2007.
57. Katz, B. Treatment of wrinkles and skin rejuvenation with combined
apex pulse technology. Available at: http://www.cynosure.com/products/
affirm/pdf/2_Katz,%20Bruce.pdf#zoom!67,0,0. Accessed October 9,
2007.
58. Montagna W, Carlisle K. Structural changes in ageing skin. Br J
Dermatol 1990;122 (Suppl 35):61-70.
59. Hardaway CA, Ross EV, Barnette DJ, Paithankar DY. Non-ablative cutaneous remodeling with a 1.45 microm mid-infrared diode laser: phase
I. J Cosmet Laser Ther 2002;4:3-8.
60. Hardaway CA, Ross EV, Paithankar DY. Non-ablative cutaneous
remodeling with a 1.45 microm mid-infrared diode laser: phase II.
J Cosmet Laser Ther 2002;4:9-14.
61. Bene NI, Weiss MA, Beasley KL, Munavalli G, Weiss RA. Comparison
of histological features of 1550 nm fractional resurfacing and microlens
array scattering of 1440 nm. Lasers Surg Med. 2006;38:S18.
62. Wilson F. Palomar breaks new ground in ablative treatments. Aesthetic
Buyers Guide. 2007; September/October: 2–5.
63. Dierickx C. Paper presented at: 26th annual meeting of the American
Society for Laser Medicine and Surgery, held April 11–15, 2007, in
Grapevine, Texas.
Downloaded from aes.sagepub.com by guest on July 20, 2011
Volume 28 • Number 2 • March/April 2008 • 187
64. Goldberg D. Reduced down-time associated with novel fractional ultrapulse CO2 treatment (ActiveFX) as compared to traditional CO2 resurfacing. J Am Acad Dermatol 2007;56(2):AB206.
65. The Palomar LuxIR Fractional Infrared Handpiece. Available at:
http://www.palomarmedical.com/palomar.aspx?pgID!1043. Accessed
November 29, 2007.
66. The Palomar LuxDeepIR Fractional Infrared Laser Handpiece. Available
at: http://www.palomarmedical.com/palomar.aspx?pgID!1011.
Accessed November 29, 2007.
67. Sciton showcases new products and its new ProFractional Technology
in San Francisco [press release]. Palo Alto, CA: Sciton, Inc.; October
13, 2006.
68. Technologies. Available at: http://www.almalasers.com/harmony_
technologies_ST.jsp?nav!1&subnav!0&prodnav!1. Accessed
November 30, 2007.
69. Jesitus J. Broadband skin-tightening device provides combination
treatments. Dermatology Times. March 1, 2006.
70. Aesthetic Buyers Guide, May/June 2007, p. 184.
Accepted for publication December 11, 2007.
Reprint requests: Neil Sadick, MD, Sadick Aesthetic Surgery &
Dermatology, 911 Park Ave, New York, NY 10021.
Copyright © 2008 by The American Society for Aesthetic Plastic Surgery, Inc.
1090-820X/$34.00
doi:10.1016.j.asj.2007.12.009
188 • Volume 28 • Number 2 • March/April 2008
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Aesthetic Surgery Journal
Intense Focused Ultrasound: Evaluation of a New Treatment
Modality for Precise Microcoagulation within the Skin
HANS J. LAUBACH, MD,! INDER R. S. MAKIN, MD, PHD,y PETER G. BARTHE, PHD,y
MICHAEL H. SLAYTON, PHD,y AND DIETER MANSTEIN, MD!
BACKGROUND AND OBJECTIVE Focused ultrasound can produce thermal and/or mechanical effects
deep within tissue. We investigated the capability of intense focused ultrasound to induce precise and
predictable subepidermal thermal damage in human skin.
MATERIALS AND METHODS Postmortem human skin samples were exposed to a range of focused
ultrasound pulses, using a prototype device (Ulthera Inc.) emitting up to 45 W at 7.5 MHz with a nominal
focal distance of 4.2 mm from the transducer membrane. Exposure pulse duration ranged from 50 to
200 ms. Thermal damage was confirmed by light microscopy using a nitroblue tetrazolium chloride
assay, as well as by loss of collagen birefringence in frozen sections. Results were compared with a
computational model of intense ultrasound propagation and heating in tissue.
RESULTS Depth and extent of thermal damage were determined by treatment exposure parameters
(source power, exposure time, and focal depth). It was possible to create individual and highly confined
lesions or thermal damage up to a depth of 4 mm within the dermis. Thermal lesions typically had an
inverted cone shape. A precise pattern of individual lesions was achieved in the deep dermis by applying
the probe sequentially at different exposure locations.
DISCUSSION AND CONCLUSION Intense focused ultrasound can be used as a noninvasive method for spatially confined heating and coagulation within the skin or its underlying structures. These findings have a
significant potential for the development of novel, noninvasive treatment devices in dermatology.
Ulthera Inc. provided the prototype intense ultrasound device for this study. Inder Makin, Peter Barthe, and
Michael Slayton are employees of Ulthera.
L
aser- and light-based devices have been introduced during the past years for noninvasive
heating of the dermis without epidermal damage.1–4
Epidermal protection is achieved by skin surface
cooling during exposure, creating an inverse temperature gradient within the skin. While optical
beams can be superficially focused, photon scattering prevents deep focusing of light within the skin.
High-intensity focused ultrasound (HIFU) has been
investigated as a tool for the treatment of solid benign and malignant tumors for many decades, but is
only now beginning to emerge as a potential noninvasive alternative to conventional therapies.5–20 The
histologic morphology of tissue destruction induced
by focused ultrasound (US) shows coagulative ne-
crosis with precisely defined, sharp margins to
normal tissue.21 The primary physical mechanism
responsible for tissue necrosis with focused US
treatment is heating due to absorption of acoustic
energy, although some concomitant inertial cavitational response of tissue from an intense US field is
probably also present.22–25 The US beam increases
the tissue temperature within a focal volume to the
point at which a wide spectrum of tissue modification can take place. The spectrum of cellular changes
depends on temperature rise and exposure duration
and range from necrosis to more subtle ultrastructural cell damage with modulation of cellular cytokine expression.26 These findings are similar to the
thermally induced changes within the skin after ablative and nonablative laser or light treatments.3,27
!Wellman Center for Photomedicine, Massachusetts General Hospital, Harvard Medical School, Boston, Massachusetts;
y
Ulthera Inc., Mesa, Arizona
& 2008 by the American Society for Dermatologic Surgery, Inc. ! Published by Blackwell Publishing !
ISSN: 1076-0512 ! Dermatol Surg 2008;34:727–734 ! DOI: 10.1111/j.1524-4725.2008.34196.x
727
I N T E N S E F O C U S E D U LT R A S O U N D
The classic HIFU applications described in the scientific literature relate primarily to the delivery of
a high-powered focused US field to ‘‘debulk’’ tissue.
The sources characteristically deposit (focus) acoustic
energy at a location distal to the source plane over a
period of seconds, whereby a region of tissue necrosis
is achieved. This process is repeated over a significant
volume of tissue (typically several cubic centimeters),
to achieve thermal destruction of the entire target
pathology. These HIFU procedures typically take between 30 and 180 minutes to complete6,17,28 depending on the target volume of treatment. In contrast to
the traditional HIFU treatment, the US approach described in this study deposits short pulses of intense
focused ultrasound (IFUS) in the millisecond domain
(50–200 ms). Avoiding cavitational processes, a frequency in the megahertz (MHz) domain is used instead
of the kilohertz (KHz) domain frequencies as commonly utilized in HIFU. The nominal energy level deposited at each site with this approach is also
significantly lower (0.5–10 J) compared to HIFU
(100 J). The goal of this study is to investigate the
ability of this US therapy approach to noninvasively
induce precise thermal damage in human skin.
Materials and Methods
Intense US Prototype Device
Experiments were performed in vitro, on postmortem human skin samples with a custom-developed
IFUS prototype device (Ulthera Inc., Mesa, AZ). An
US probe is connected to a generator system operating in the MHz frequency regime. The US energy is
coupled from the transducer (operating at 7.5 MHz)
to skin tissue by ultrasound coupling gel applied to
the skin surface. The nominal focal depth for
this study was 4.2 mm below the skin surface
(Ulthera Inc.).
Tissue Samples and Tissue Processing
For the in vitro evaluation we used cryopreserved
("801C), full-thickness skin samples of different body
sites and a Fitzpatrick Skin Type II to V. Dermal
thickness of tissue samples ranged from 2 to 5 mm and
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D E R M AT O L O G I C S U R G E RY
total thickness including subcutaneous tissue was up to
20 mm. Tissue samples were defrosted from "801C,
and care was taken that the entire tissue sample was
heated to 351C. After exposure, tissue was again frozen
immediately to "801C and processed by frozen sectioning in the upcoming days. After the exposure, tissue
specimens were processed for frozen sections. Crosssections of 10 mm thickness were collected every
200 mm for further processing and histologic evaluation. This approach allowed for a three-dimensional
understanding of the histology of the treated zone.
Thermal damage patterns in the tissue were assessed for
microscopic evaluations with a nitroblue tetrazolium
chloride (NBTC) assay for cell viability as described by
Neumann and coworkers.29 Furthermore, tissue sections for histologic evaluation were counterstained
with eosin to increase contrast and show possible
collagen denaturation. Cross-polarized light was
also used to confirm collagen denaturation by loss of
birefringence.
US Exposures
Exposures were performed in vitro on postmortem
human skin samples at tissue temperatures of 351C.
The temperature of skin sample was kept constant
by a heating plate on which the sample was placed
and monitored before exposures with a contact
thermometer. The prototype probe was acoustically
coupled to the human skin sample, using US coupling gel. Single exposures were performed at set
power levels and exposure durations. Exposure duration was varied from 50 to 200 ms, output power
was set to a maximum of 45 W, and no active cooling
was used before, during, or after the exposure. Exposure locations were then marked and sectioned for
further histologic evaluation.
Numerical Simulations
The propagation of the focused US beam in skin
tissue was modeled using an approach described
by Hasegawa and coworkers.30 The acoustic field
simulation accounts for the geometric focusing as
well as the attenuation of energy in the epidermis,
dermis, and hypodermis. The thermal gradients
LAUBACH ET AL
resulting from absorption of acoustic energy in tissue
and conversion to heat were calculated using the
Bioheat equation.31 The 4601C contours indicating
complete collagen disruption were chosen to represent the zone of thermal injury.23,31,32
A
Results
B
0.2
Epidermis
Dermis
2.2
Skin depth [mm]
Exposure durations of 150 ms and above resulted in
a palpable and macroscopically visible intracutaneous nodule of approximately 1-mm diameter. The
skin surface was slightly raised over US-induced
dermal nodules but did not show any evident blister.
Exposure durations of 125 ms and below could not
be detected by clinical examination (observation and
palpation) of the skin samples. Histologic evaluation
by both NBTC and standard H&E-stained light
microscopy showed that intradermal lesions created
by the single US exposure pulses in this study, although different in size, were typically inverted
cone–shaped. Thermal lesions consisted of a core
defined by an area of thermal cell necrosis and collagen denaturation as determined by the loss of
NBTC staining and loss of birefringence, respectively
(Figures 2A, 3A, and 3B). The lesions typically beginA
in the deep reticular dermis at a depth of approximately 3 to 4 mm (Figure 2A). Serial steps sectioning
as described under Materials and Methods through
the entire skin samples did not detect any epidermal
damage. Increasing the exposure time and energy
delivered caused the thermal lesions to extend from
the deeper reticular dermis toward the papillary
dermis. At exposure durations of 175 ms and above,
the lesions consisted of overt damage of the entire
dermal thickness and overlying epidermis (Figure
1A). A single US exposure of 50 ms produced wellconfined thermal lesions with NBTC staining loss of
approximately 200 # 300 mm at a depth of 2.7 mm
deep within the reticular dermis (Figures 3A and 3B).
Figure 4 shows the result of multiple exposures
within one skin sample of Fitzpatrick Skin Type V.
Lesions can be placed independently from each other
without confluent damage. If the dermal thickness is
less than the focal depth of the intense US device,
thermal lesions are placed within the underlying
4.2
6.2
8.2
10.2
−4
−2
0.0
2
4
Figure 1. (A) NBTC assay with eosin counterstain; # 12.5 magnification; single US exposure, 45 W, 200-ms pulse duration,
7.5 MHz, 4.2-mm focal depth. Complete loss of NBTC staining
and collagen denaturation throughout the entire dermis (black
circle). Please note the artifact due to tissue preparation with
loss of epidermal tissue in necrotic zone due to the thermal
alteration of tissue integrity and resulting friability. (B) Numerical simulation of the thermal response of skin to source conditions corresponding to the experimental results in A. The
zone of thermal coagulation in this simulation is represented
by the 601C temperature contour in the skin tissue.
structures, e.g., subcutaneous adipose tissue. The
theoretical size and location of thermal lesions
predicted from numerical simulations compared well
with the observed experimental zone of thermal
coagulation (see Figures 1–4). The computationally
predicted lesions in these results are nominally
longer axially, compared to the experimentally
observed thermal lesions. This discrepancy is most
likely due to the fact that the simulations do not
account for the change in tissue properties subsequent to the change of temperature.23,34
3 4 : 5 : M AY 2 0 0 8
729
I N T E N S E F O C U S E D U LT R A S O U N D
A
B
0.2
Epidermis
Dermis
Skin depth [mm]
2.2
4.2
6.2
8.2
10.2
−4
−2
0.0
2
4
Figure 2. (A) NBTC assay with eosin counterstain; # 12.5
magnification; single US exposure, 45 W, 75-ms pulse duration, 7.5 MHz, 4.2-mm focal depth. Well-confined zone of thermal damage within the dermis (black circle). (B) Numerical
simulation of the thermal response of skin tissue to source
conditions corresponding to the experimental results in A. The
zone of thermal coagulation in this simulation is represented
by the 601C temperature contour in the skin tissue.
Discussion
Classic HIFU treatment is using the concept of
thermal tissue injury due to the absorption of US
energy.35–37 It has been investigated as a noninvasive
treatment modality tool for benign and malignant
tumors for many decades and has now been applied
as a noninvasive alternative to conventional therapies for nearly a decade.13,24,38,39 Van Leenders and
colleagues,26 for example, have shown that it is
possible to thermally confine an US-induced thermal
treatment zone within the prostate gland. This allowed HIFU to become one of clinical treatment alternatives in the treatment of benign prostatic
730
D E R M AT O L O G I C S U R G E RY
hyperplasia and prostate cancer.5–8,40 While HIFU
can been used to thermally ablate tissue on a macroscopic scale (in the range of several cubic centimeters),25,26,41 we investigated in this study the
potential for focused US as a treatment modality to
induce micro-thermal tissue denaturation within the
human skin. As demonstrated by the numerical
simulation results and confirmed by the characteristic coagulative change shown in the histology from
our study, the US treatment regime (e.g., frequency,
power and exposure duration) caused well-defined
zones of thermal injury within the dermis. Figures 2–
4 show those confined zones of microscopic tissue
ablation induced by focused US using relatively high
acoustic intensity delivered within milliseconds.
Owing to the relatively short exposure duration as
well as the sharp focusing, it is possible to deliver US
energy at significantly lower energies than classic
HIFU to achieve a microscopically small volume of
thermally ablated tissue (o1 mm3). Compared to
classic HIFU, exposure durations used are significantly lower (in the millisecond domain), the total
energy delivered per pulse is considerably smaller
(below 15 J/pulse), and the focal spot within the skin
achieves a zone of thermal tissue effect on the order
of 1 mm3 and smaller. By choosing the appropriate
exposure parameters with the IFUS approach, we
were able to spare the epidermis as well as avoid
damage to the papillary dermis without simultaneous skin cooling, while creating a zone of thermal
coagulation deep within the reticular dermis
(Figures 2 and 3). With increase in the exposure
time, the thermal lesion grows typically in its axial
dimension, progressing proximally toward the skin
surface (Figure 1) while shorter exposure times not
only decrease the lesion size (Figures 2 and 3) but
also minimize the risk of uncontrolled bulk heating
and thermal diffusion into adjacent tissue. Chen and
coworkers42 reported that typically cigar-shaped
lesions are observed in tissue phantoms after HIFU
exposure. These lesions take a tadpolelike appearance once boiling temperatures are reached. We observed similarly shaped lesions within the dermis in
our study (Figure 2). The lesions were typically cigaror inverted cone–shaped and started in the lower
LAUBACH ET AL
A
B
C
0.2
Epidermis
Dermis
Skin depth [mm]
2.2
4.2
6.2
8.2
10.2
−4
−2
0.0
2
4
Figure 3. (A) NBTC assay with eosin counterstain; # 12.5 and # 100 magnification; single US exposure, 45 W, 50-ms pulse
duration, 7.5 MHz, 4.2-mm focal depth. Small and well-confined thermal lesion deep within the deep reticular dermis (black
circle). (B) Same # 100 magnification close up as in A with corresponding cross-polarized image showing complete loss of
birefringence in thermal damage zone (white circle). (C) Numerical simulation of the thermal response of skin tissue to
source conditions corresponding to the experimental results in A. The zone of thermal coagulation in this simulation is
represented by the 601C temperature contour in the skin tissue.
3 4 : 5 : M AY 2 0 0 8
731
I N T E N S E F O C U S E D U LT R A S O U N D
A
B
0.2
Epidermis
Dermis
Skin depth [mm]
2.2
4.2
6.2
8.2
10.2
−4
−2
0.0
2
4
Figure 4. (A) NBTC assay with eosin counterstain, # 12.5
magnification, two separate US pulses deposited 3 mm
apart, 45 W, 125-ms pulse duration, 7.5 MHz, 4.2-mm focal
depth. Two spatially distinct zones of thermal damage within the dermis and the subcutaneous fat (black circles).
Please note that the dermis of this skin sample is thinner
than that in Figures 1–3. Therefore, a focal depth of 4.2 mm
places the thermal damage zones within the subdermal tissue. (B) Numerical simulation of the thermal response of
skin tissue to source conditions corresponding to the experimental results in A. The zone of thermal coagulation in
this simulation is represented by the 601C temperature contour in the skin tissue.
reticular dermis. It is noteworthy that the thermal
lesions on histologic analysis were found slightly
above the geometric focus as predicted by the beam
geometry and the computer simulation. One possible
explanation for this observation can be found in
Bush and colleagues,43 who described that when
tissue heating occurs, the attenuation within that
volume increases, altering the absorbed energy
distribution. The region in which heat is deposited is
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D E R M AT O L O G I C S U R G E RY
expected to alter its absorption properties during the
heating process, thereby shifting the region of
intensity maximum toward the transducer.44,45 Since
the tissue properties change dynamically as US
energy is deposited, future studies should more
extensively investigate multiple lesion formation and
variability of response at fixed source conditions.
Another difference between classic HIFU treatments
and IFUS is that in the clinical setting of HIFU
therapy convective and conductive energy losses play
an important role since exposure durations are in the
order of seconds and longer. Owing to the short
exposure durations (on the order of several milliseconds), the coagulative tissue effect with IFUS is
mostly independent of these losses and is also not
accounted for in our numerical modeling.
Comparing the outline of the lesion determined by
loss of collagen birefringence in comparison with the
loss of NBTC staining behavior has been examined,
and a small size difference could be observed. The
lesion as determined by loss of collagen birefringence
appeared to be consistently smaller than the lesion
determined by loss of NBTC stain. To determine
the exact difference in between these two lesions,
although interesting, is out of the scope of this study.
One inherent advantage of the noninvasive therapy
with US compared to light-based devices is its independence of chromophores for energy absorption. As
demonstrated in Figure 4, even a skin sample with
Fitzpatrick Skin Type V was treated, and welldefined lesions could be created within the deeper
dermis and the subcutaneous tissue without simultaneous skin cooling. No damage is observed in the
upper dermis and the overlying epidermis. The
advantage for the dermatologic use of IFUS is that
the absorption of US energy is independent of the
melanin content of skin. Its absorption is rather
determined by the microscopic and bulk mechanical
properties of tissue.37,46 Therefore, in contrast to
light-based devices, the action of IFUS is independent
of skin color and chromophores. The ‘‘colorblind’’
IFUS treatment approach might be helpful in overcoming some of the difficulties encountered with the
light-based treatment of darker skin types. In addition
LAUBACH ET AL
to its independence of chromophores, IFUS creates a
sharp focus of the US beam several millimeters within
the skin. Hence the power density of the converging US
beam is much lower as it passes through epidermis
than in its focal point. Therefore, only minimal energy
absorption and tissue heating occurs at the epidermal
level insufficient to create significant thermal damage.
This consequently obviates the need for skin cooling
for epidermal protection for any skin type as it is used
with other devices inducing unexpected thermal alterations within the skin.
As demonstrated in Figure 4, several separate lesions
can be placed next to each other within the skin using
IFUS. This allows for the creation of a number of
unique thermal damage patterns. Tissue may be altered
by arrays of microscopically small focal damage from
IFUS rather than ablating an entire macroscopic area
allowing a rapid healing response from tissue immediately adjacent to the thermal lesions, conceptually
similar to laser fractional photothermolysis.47 It remains furthermore to be determined in how far US
‘‘see-and-treat’’ systems, as they are already established
for HIFU therapy,40,48 can also be used for the guidance and monitoring of IFUS in the dermatologic use.
Intense focused ultrasound provides the possibility to
thermally coagulate a target deep within the skin or
below without affecting the intervening tissue. Compared to similar nonablative therapies based on light or
radiofrequency, IFUS has the capability of precisely
controlling the amount and location of thermal injury
at a known depth below the skin surface. IFUS is a new
treatment modality, offering the potential for novel,
noninvasive treatment concepts in dermatology.
Conclusion
Intense focused ultrasound can be used as a noninvasive method for spatially confined heating and coagulation within the skin or its underlying structures.
Acknowledgments The authors thank Qiqi Mu,
MD, for her support with tissue sectioning and Bill
Farinelli for his unceasing technical succor.
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Address correspondence and reprint requests to: HansJoachim Laubach, MD, Massachusetts General Hospital,
Wellman Center for Photomedicine, BAR #305, 50 Blossom Street, Boston, MA 02114, or e-mail: hlaubach@
partners.org
Lasers in Surgery and Medicine 40:67–75 (2008)
Selective Transcutaneous Delivery of Energy to Porcine
Soft Tissues Using Intense Ultrasound (IUS)
W. Matthew White, MD,1 Inder Raj S. Makin, MD, PhD,2 Michael H. Slayton, PhD,2
Peter G. Barthe, PhD,2 and Richard Gliklich, MD1*
1
Division of Facial Plastic and Reconstructive Surgery, Department of Otology and Laryngology,
Massachusetts Eye and Ear Infirmary, Harvard Medical School, Boston, Massachusetts
2
Ulthera, Inc., Mesa, Arizona
Objective: Various energy delivery systems have been
utilized to treat superficial rhytids in the aging face.
The Intense Ultrasound System (IUS) is a novel modality
capable of transcutaneously delivering controlled thermal
energy at various depths while sparing the overlying
tissues. The purpose of this feasibility study was to evaluate
the response of porcine tissues to various IUS energy
source conditions. Further evaluation was performed of the
built-in imaging capabilities of the device.
Materials and Methods: Simulations were performed
on ex vivo porcine tissues to estimate the thermal dose
distribution in tissues after IUS exposures to determine the
unique source settings that would produce thermal injury
zones (TIZs) at given depths. Exposures were performed at
escalating power settings and different exposure times
(in the range of 1–7.6 J) using three IUS handpieces with
unique frequencies and focal depths. Ultrasound imaging
was performed before and after IUS exposures to detect
changes in tissue consistency. Porcine tissues were examined using nitro-blue tetrazolium chloride (NBTC) staining
sensitive for thermal lesions, both grossly and histologically. The dimensions and depth of the TIZs were measured
from digital photographs and compared.
Results: IUS can reliably achieve discrete, TIZ at various
depths within tissue without surface disruption. Changes
in the TIZ dimensions and shape were observed as source
settings were varied. As the source energy was increased,
the thermal lesions became larger by growing proximally
towards the tissue surface. Maximum lesion depth closely
approximated the pre-set focal depth of a given handpiece.
Ultrasound imaging detected well-demarcated TIZ at
depths within the porcine muscle tissue.
Conclusion: This study demonstrates the response of
porcine tissue to various energy dose levels of Intense
Ultrasound. Further study, especially on human facial
tissue, is necessary in order to understand the utility of this
modality in treating the aging face and potentially, other
cosmetic applications. Lasers Surg. Med. 40:67–75,
2008. ! 2008 Wiley-Liss, Inc.
Key words: nonablative devices; porcine tissue; ex vivo;
aging face; collagen; ultrasound; Intense Ultrasound; skin;
SMAS; muscle
! 2008 Wiley-Liss, Inc.
INTRODUCTION
Ultrasound-based imaging systems for clinical diagnosis
have been used for several decades, whereby this energy
modality is considered to be one of the safest and
used routinely for fetal obstetric and general clinical
examinations [1]. However, by using a highly directive
source geometry with the source energy settings increased
significantly, ultrasound energy can be focused spatially in
a tightly confined region (on the order of 1 mm3) to cause
selective tissue thermal coagulation (Fig. 1). This Intense
Ultrasound (IUS) approach enables the creation of well
defined thermal injury zones (TIZs) at depths within soft
tissue while leaving the surrounding regions unaffected.
IUS is similar to fractional laser resurfacing [2] in that
thermal lesions are created, yet IUS is unique in that the
thermal lesions are created below the surface and can
be of variable geometry. The ultrasound waves induce a
vibration in the composite molecules within tissue during
propagation, and the friction developed between intrinsic
molecules is the source of the generated heat. It has been
well established in the literature that Intense Ultrasound
(IUS) fields can be transcutaneously directed into visceral
soft tissue to produce coagulative necrosis resulting
primarily from thermal mechanisms [3,4]. For most of the
work in this area, the effort has been to develop intense
focused ultrasound as a noninvasive surgical tool to treat
human whole organ tumors, such as liver, breast, and
uterus [3,4].
In order to achieve an effective energy delivery for
cosmetic applications, a novel ultrasound therapy device is
described herein to deposit energy localized to the first few
Presented in part at the April 2006 Annual Meeting of the
American Society of Lasers in Surgery and Medicine in Boston,
MA.
Contract grant sponsor: Ulthera, Inc., Mesa, AZ.
The authors have disclosed potential financial conflict of
interests with this study.
*Correspondence to: Richard Gliklich, MD, Division of Facial
Plastic and Reconstructive Surgery, Massachusetts Eye and Ear
Infirmary, 243 Charles Street, Boston, MA 02114.
E-mail: richard _ [email protected]
Accepted 27 December 2007
Published online in Wiley InterScience
(www.interscience.wiley.com).
DOI 10.1002/lsm.20613
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WHITE ET AL.
Fig. 1. Intense Ultrasound beam profile. Visualization of an
ultrasound beam using a Schlieren system enables the
mapping of the power density (Intensity) of the field [22,23].
The Schlieren map of one of the prototype probes used in this
study demonstrates that most of the ultrasound energy
(approximately 95% of total energy), can be focused spatially
into a tightly confined region. As shown in this result, the focal
zone of the ultrasound beam was measured to be 1.8 mm
axially, while the beam was 0.5 mm in the radial direction
(inset: magnified view of focal region). [Figure can be viewed in
color online via www.interscience.wiley.com.]
resurfacing has been proven largely successful for the
nonsurgical treatment of rhytids, the undesirable postoperative intense inflammatory response has caused
the demand for this procedure to drop dramatically. For
this reason, physicians and surgeons alike have tried to
develop various methods (e.g., radiofrequency), termed
‘‘Nonablative Skin Resurfacing,’’ to induce collagen shrinkage and remodeling while preserving the epidermis in an
effort to minimize these post-operative changes. These
modalities have produced variable efficacy at best [7,8,13].
In this initial study, we wanted to determine if the IUS
system could be used to create subsurface, discrete TIZs.
Porcine tissue was used since it is a well-established model
in terms of tissue properties being close to that of human
skin [10,11,14]. Numerical modeling was first performed to
simulate IUS energy–tissue interaction. Homogeneous
porcine muscle tissue was then utilized to examine the
dose–response profile by varying unique source conditions
of the IUS system (e.g., frequency, time, source power).
Lesions were measured, quantified and compared with
the theoretical predictions. Ultrasound imaging was
performed before and after IUS exposure to determine if
TIZ localization was possible. A subset of the energy dose
range from experiments conducted in porcine muscle was
then repeated in porcine skin tissue using the three IUS
handpieces. This work serves as an introductory feasibility
study investigating superficial tissue response to IUS
exposure.
MATERIALS AND METHODS
millimeters of the superficial skin tissue [5,6]. This device is
able to focus energy within tissue to produce a 25 mm line of
discrete TIZs spaced 0.5–5.0 mm apart. Furthermore, both
imaging and selective energy exposure can be accomplished
with the same handpiece. The IUS System can therefore
target and deliver focused energy to a specific soft tissue
region or layer.
Various nonsurgical modalities have been utilized to
treat facial rhytids (peels, microdermabrasion, and lasers)
[7,8]. All of these modalities however have focused on
treating the superficial layers of skin (i.e., epidermis and
dermis) due to their limited penetration depth. The gold
standard for nonsurgical facial rejuvenation has been the
Carbon Dioxide (CO2) Laser. The CO2 laser has been used
extensively for facial resurfacing for the treatment of
rhytids. The mechanism of CO2 laser rejuvenation of the
skin is thought to be: (i) ablating and removing the most
superficial layer of skin (epidermis), and (ii) delivering
energy to the deeper superficial papillary dermis to create a
lesion in the collagen [9–12]. This lesion incites a ‘‘wound
healing’’ response through the liberation of several cytokines which stimulate fibroblasts to synthesize and lay
down new collagen. This collagen remodeling process is a
crucial step in facial skin rejuvenation.
Despite the efficacy of the CO2 laser, treatment results in
complete ablation of the entire epidermis with a wound
that lasts for 7–10 days. Following this, post-treatment
erythema or ‘‘scalded skin’’ appearance that can persist
for months after CO2 laser resurfacing. Although CO2 laser
Fresh, frozen specimens of porcine muscle and skin were
obtained according to the policies of the Massachusetts Eye
and Ear Infirmary Institutional Review Board (IRB).
Specimens were stored in a freezer, and allowed to thaw
to room temperature (258C) prior to experiments.
Intense Ultrasound System
The IUS device is designed to target and deliver focused
ultrasound energy within tissue (Ulthera, Inc., Mesa, AZ).
The IUS handpiece contains a transducer that has two
functioning modes: imaging (which is used to image the
region of interest before the therapeutic ultrasound exposures) and treatment (which is the mode that delivers
a series of higher-energy ultrasound exposures). A series of
selective thermal ablative zones can be produced along a
straight line at a given depth within the tissue (25 mm line of
discrete lesions spaced 0.5–5.0 mm apart). For each series of
exposures, the following source conditions can be varied:
power output (W), exposure time (ms), length of exposure
line (mm), distance between exposure zones (mm), and time
delay after each exposure (ms). Three handpieces were used,
in order of most superficial focus to the deepest: (i) 7.5 MHz
3 mm focal depth, (ii) 7.5 MHz 4.5 mm focal depth, and
(iii) 4.4 MHz 4.5 mm focal depth in tissue.
Numerical Simulations
The acoustic field and the resulting thermal distribution in
tissue resulting from absorption of the focused beam were
simulated numerically for each of the three handpieces.
SELECTIVE TRANSCUTANEOUS DELIVERY OF ENERGY
The thermal effects of a focused field propagating through
skin/superficial tissue have been extensively modeled using
well-established acoustic beam propagation schemes
[13–16] as well as using the bio-heat equation [15–18]. A
multi-layer approach has been applied for numerical
simulations, whereby the significant differences in tissue
attenuation between the epidermis (nominal thickness
0.2 mm), dermis, and subcutaneous tissue have been
accounted for. The 608C temperature contour in tissue is
representative of the region with thermal coagulation within
the source conditions used in this study.
IUS Exposure Procedure
The porcine skin tissue was tattooed (India ink) prior to
IUS exposures to create a grid over the proposed treatment
area (Fig. 8). Porcine muscle specimens were first selected. A
range of source conditions were selected and planned for
each area. Prior to each treatment, ultrasound imaging of
the soft tissue was performed to identify the target tissues.
Ultrasound imaging was performed on each planned treatment area and still images were captured and stored. All IUS
exposures were performed along the parallel axis of the
tattooed grid. Immediately after IUS exposure lines were
delivered, the handpiece was left in place and the axis of the
69
exposure line delivered was marked by the use of Wite-Out1
Correction Fluid (Bic Corporation, Milford, CT; Fig. 8).
After all IUS exposures for a given porcine specimen had
been completed, each treated region was excised. The
tissue bloc was then placed on an acrylic plate and kept in a
!15 degree Celsius freezer for approximately 2 hours.
Using a surgical blade, fine 1 mm thick sections were cut
parallel to the IUS exposure lines. In the case of the porcine
muscle experiments, the zones of thermal coagulation
are reveled as discrete whitened regions. These slices are
photographed and the digital images assessed in terms of
the size and shape of the thermal zone.
For porcine skin experiments, the grossly sectioned thin
strips of skin tissue was placed in NBTC stain overnight for
viability staining. This method has been described in the
literature for identifying thermally affected tissue both in
vivo as well as for ex vivo studies [19–21]. Viable tissue
stains blue and TIZs are demarcated by a pale color or lack of
blue staining. Digital photographs (Nikon D-70, Nikon USA,
Melville, NY) were taken of the gross tissue sections after
they NBTC staining. Depth and dimensions of TIZ were
evaluated from digital photographs of post-NBTC-stained
gross porcine tissue strips using image processing software
(NIH Image J, http://rsbweb.nih.gov/ij/).
A
Fig. 2. IUS simulation. Numerical simulation of the thermal
response of porcine skin tissue to a focused Intense Ultrasound
beam. The epidermis and dermis is modeled as a single layer
of higher attenuation (2.0 dB/MHz/cm), compared to the subdermal tissue (1.5 dB/MHz/cm). Irreversible tissue coagulation
is represented in these numerical simulations as the region of
porcine skin tissue that attains a temperature "608C using one
of the three handpieces. The y-axis represents the depth
(millimeters) from the probe–tissue interface. The x-axis is the
width (millimeters) of the TIZ. A: 7.5 MHz/3.0 mm focus probe,
B: 7.5 MHz/4.5 mm focus probe, and C: 4.4 MHz/4.5 mm
focus probe. [Figure can be viewed in color online via
www.interscience.wiley.com.]
70
WHITE ET AL.
RESULTS
Numerical Simulations
Multiple simulation runs were performed to predict the
zone of thermal coagulation in porcine muscle as well as
skin tissue. Representative results for the three handpiece
configurations used in this study are shown in Figure 2.
These simulations represent propagation through porcine
skin tissue. The simulations predict a spatially confined
zone of thermal treatment within the tissue. The numerically predicted results (Fig. 2) compare favorably with
regions of tissue coagulation demonstrated in gross tissue
specimens (Fig. 9).
Porcine Muscle
Porcine muscle was chosen for the initial experimentation due to its homogeneous composition. In this experiment, an IUS probe with a source frequency of 4.4 MHz and
focal depth of 4.5 mm was chosen. Exposure lines were
delivered to porcine muscle specimens as power levels were
increased. Visual analysis of porcine tissue revealed that
as the source power was increased from 2.3 to 7.6 J, the
TIZ became larger and extended closer to the tissue surface
(Fig. 3). For lower source settings (2.3 J), the thermal
injury region is relatively small and corresponds closely
to the geometric focal zone of the ultrasound field. As
the energy of the IUS source is increased (2.3–7.6 J), the
initial tissue coagulation at the focus presents a region of
significantly higher acoustic attenuative characteristics,
thereby ‘‘screening’’ the thermal injury region from extending post-focally. The TIZ then progressively extends proximally towards the Intense Ultrasound source plane.
Figure 4 shows a series of exposures using the 7.5 and
4.4 MHz (4.5 mm focal depth) handpieces. Lesion locations
and dimensions were measured and were observed to
be dependent on particular source conditions such as
source frequency, focal depth, power, and exposure duration (energy). Imaging scans were made immediately
pre- and post-exposure in each case. The images in each
case, demonstrate a selective hyper-echogenic string of
regions, which correspond reasonably well with the grossly
visualized TIZs.
In order to better characterize the tissue effect, the
proximal and distal range of the TIZ as well as the
corresponding areas are quantified for each exposure
condition. Analysis of these tissue samples shows a dose–
response effect of tissue ablation with source conditions
for the 7.5 MHz (Fig. 5) and 4.4 MHz (Fig. 6) handpieces.
For both the sources, the proximal edge of the TIZ
progressively extends upwards with increasing source
energy, and the lesion becomes larger (greater area).
Porcine Skin
Porcine skin is considered a reasonable model for
investigating the effect for energy for cosmetic applications
[9,10]. Figure 7 shows histologic demonstration of TIZs
from an in vivo experiment utilizing the 7.5 MHz 4.5 mm
focal depth probe. Figure 7A shows an image of a
hematoxylin and eosin stain (H&E) slide, whereas the
Figure 7B shows a gross section of the skin sample stained
with vital stain (NBTC). Both the images are captured
at 10# magnification. The source conditions for the
IUS-treated regions in both samples are the same. Note
that the nominal depth and dimensions of the TIZs
identified in both the separate skin samples were created
using the same source conditions, yet compare favorably
when evaluated by different staining techniques (Fig. 7).
The hematoxylin and eosin staining example shows a
region of selective and distinct thermal coagulative change,
where the collagen fibers are indistinct and fused together,
whereas the loss of vital staining (nitro-blue tetrazolium
chloride—NBTC) within the region of the TIZ indicates
Fig. 3. Dose–response in muscle. Digital photographs of gross tissue sections (approximately
1 mm thick) of porcine muscle reveal profile of changes in geometry of TIZ as the source energy
is increased from 2.3 to 7.6 J. Within the homogenous orange-colored muscle tissue, the white
inverse-pyramidal regions of coagulated tissue are the TIZs resulting from the ultrasound
exposure (4.4 MHz, 4.5 mm focus handpiece). [Figure can be viewed in color online via
www.interscience.wiley.com.]
SELECTIVE TRANSCUTANEOUS DELIVERY OF ENERGY
71
Fig. 4. Image guided therapy—ultrasound imaging of TIZ. Geometry of the TIZs over different
source conditions could be viewed with the integrated ultrasound imaging modality of two IUS
handpieces. Both ultrasound images and digital photographs of gross tissue sections (1 mm
thick) of porcine muscle are shown. [Figure can be viewed in color online via www.interscience.
wiley.com.]
thermal denaturation. No damage to the skin surface was
observed in this in vivo porcine experiment.
In Figure 9, a series of representative results are shown in
gross tissue sections using each of the three handpieces
investigated in this study. The three handpieces help
achieve TIZ depth and shape unique to the probe geometry
(3.0 or 4.5 mm focal depth) as well as source frequency (7.5 or
4.4 MHz). As predicted in the numerical simulations
(Fig. 2), the 7.5 MHz handpiece with 4.5 mm focal depth
results in a TIZ that is nominally shallower with a shorter
axial dimension compared to the TIZ from the 4.4 MHz
(4.5 mm focus) handpiece. This observation is confirmed
by comparing the axial range of TIZs in Figure 9B,C.
Note that in each case (Fig. 9A–C), the epidermal layer is
spared. As described in the Materials and Methods Section,
multiple TIZs were placed along a 25 mm exposure line
at each source condition using a particular handpiece.
Figure 8 shows that the porcine skin surface after placing
two exposure lines each, for the three probes using
the energy settings from Figure 9. No skin surface damage
was observed during exposure with either of the three
handpieces.
DISCUSSION
In this work, we introduce a new approach that has
the potential for use in facial cosmetic procedures. We have
shown that IUS is capable of creating thermal coagulative
zones at depths within porcine soft tissues. Trials with
different handpieces demonstrate that for the same source
geometry (e.g., 7.5 vs. 4.4 MHz probes with 4.5 mm focal
depth), lower frequency exposures tend to produce TIZs
that extend deeper within tissue (compare Figs. 5 and 6).
The results in Figures 5 and 6 respectively, have a low
variability and demonstrate creation of well controlled TIZs
at each source condition.
The tissue response characterized by thermal coagulation change, achieved by IUS exposure is similar to that
from other energy-based devices used in the cosmetic
arena such as lasers, radiofrequency (RF) and combination
laser–RF devices [9,10,19]. However, in contrast to the
other known energy based devices used for cosmetic
applications, the IUS field is sharply focused, thereby
depositing most of the energy in the form of heat around the
focal zone of the beam, leaving the surrounding regions
72
WHITE ET AL.
Fig. 5. Graph of coagulative zone characteristics with the
7.5 MHz, 4.5 mm handpiece. Effect of varying source condition
on the size and depth of the TIZs (N ¼ 39). The red bars
represent the average measurement of the most superficial
portion (labeled ‘‘Top’’), of the TIZs from the surface of the
sample tissue. The blue bar is the average measurement of
the deepest portion of the TIZs from the tissue surface
(labeled ‘‘Bottom’’). The Green bar is the average estimated
area of the TIZs at various source settings. Each error bar
corresponds to one standard deviation for each measurement.
[Figure can be viewed in color online via www.interscience.
wiley.com.]
unaffected [3,4]. In this manner, the overlying epidermal
surface is spared, and the thermal coagulation is achieved
only at depth (on the order of millimeters). The hypothesis
is that this approach of selective thermal injury at depth
will avoid unwanted side effects seen with ablative skin
resurfacing modalities (e.g., skin pigmentary change and
sloughing).
The primary biophysical processes leading to thermal
coagulation during propagation of ultrasound energy as
considered in this investigation are: beam focusing and
acoustic absorption. In the case of a tightly focused beam,
the maximum rate of acoustic energy deposition in the form
of heat is around the focal plane. The remaining prefocal and post-focal areas of the beam in tissue remain
unaffected, since the acoustic power density is insufficient
to achieve thermal tissue coagulation in regions other
than the focal zone (Figs. 3, 7, and 9). Acoustic absorption is
a frequency dependent phenomenon, nominally increasing
linearly with frequency in tissue [14]. Therefore, the
4.4 MHz handpiece (4.5 mm focal depth), tends to achieve
a TIZ deeper compared to the 7.5 MHz (4.5 mm focus)
handpiece (Figs. 2 and 9).
In this study, porcine tissue was chosen as for experiments, since it is an established tissue model for cosmetic
tissue applications [9,10,20]. For example, porcine skin
Fig. 6. Graph of TIZ characteristics with the 4.4 MHz, 4.5 mm
handpiece. Effect of varying source condition on the size and
depth of the TIZs (N ¼ 73). The red and blue bars represent the
proximal (‘‘top’’) and distal (‘‘bottom’’) extents of the TIZ. Green
bar is the average area of the TIZs at a particular source
setting. [Figure can be viewed in color online via www.interscience.wiley.com.]
Fig. 7. Thermal coagulative region with the 7.5 MHz,
4.5 mm focal depth handpiece (3.6 J). Comparison of thermal
coagulative change with H&E (histology) and NBTC staining
of a gross tissue section of porcine skin (both figures in vivo
treatment). Note homogenization of the collagen fibrillar
structure with hematoxylin and eosin staining (A), and
thermal damage in a comparable zone (no blue-dye uptake)
with NBTC staining (B). The inset panel shows a magnified
view of the coagulation zone using H&E staining. [Figure can
be viewed in color online via www.interscience.wiley.com.]
SELECTIVE TRANSCUTANEOUS DELIVERY OF ENERGY
Fig. 8. Skin surface after exposure with 7.5 MHz, (3.0 and
4.5 mm focal depths) and 4.4 MHz, 4.5 mm focal depth
handpieces. Note no damage to skin surface after placement
of multiple TIZs along two exposure lines using each handpiece. The vertical arrows indicate the borders of a single
exposure line. Photographs were taken prior to sectioning and
staining to observe TIZs in skin tissue in Figure 9. [Figure can
be viewed in color online via www.interscience.wiley.com.]
has a similar layered structure as human skin: epidermis,
dermis and glandular components, underlying connective
tissue and muscle. Even though the attenuation coefficient
of porcine muscle is much lower than skin tissue (muscle,
0.75 dB/MHz/cm compared to skin, 2.0 dB/MHz/cm) [14],
the porcine muscle tissue is used as a model to understand
the energy–tissue interaction and reliably demonstrate
selective thermal coagulation using a focused IUS beam.
Using porcine muscle, a wide range of source parameters
could be varied and the tissue effect could be easily
evaluated with gross pathology as regions of whitened
coagulum, even without NBTC staining (Figs. 3 and 4). As
described earlier, increasing the energy levels significantly
resulted in propagation of the TIZ towards the surface
(Figs. 3 and 5). This phenomenon of ‘‘tadpole formation’’ is
well documented in the literature for other soft tissue
ablation with IUS [20]. Therefore, the degree of selectivity
of tissue effect in skin can be controlled by an appropriate
choice of source conditions for a particular handpiece.
In the case of porcine skin tissue, the various anatomical
layers, epidermis, dermis and subcutaneous tissue are
represented. The regions of thermal coagulative necrosis
identified using NBTC staining are representative of
73
the TIZs expected in the human facial skin tissue. The
numerical simulation results for formation of thermal
lesions in a porcine skin tissue, accounting for attenuation
and focusing (Fig. 2) are comparable to the actual porcine
skin experimental results (Fig. 9). A key goal of this concept
study was to understand the characteristics of selective
thermal coagulation using Intense Ultrasound in model
tissue (porcine muscle, and porcine skin—in vitro and in
vivo). Tissue contraction following ultrasound exposure
was not investigated since the tissue was not attached to
the natural anatomical attachment points, and will be the
subject of future studies.
Ultrasound imaging is unique to the IUS device in that
it could potentially provide immediate feedback to the
clinician. In these initial experiments, we demonstrated
that it is possible to detect the thermal coagulative
change that occurs in porcine muscle tissue following IUS
exposure (Fig. 4). However, detecting TIZ in porcine
skin tissue by ultrasound imaging was more challenging.
Imaging the lesion size and location can represent added
safety to the clinician, as it is possible to see immediately
where the energy is being deposited. Further work needs
to be performed in optimizing the ultrasound imaging
component of the system, and in understanding the role of
imaging in developing an IUS based cosmetic procedure.
The mechanism of skin rejuvenation has been well
studied in the gold standard CO2 laser. The CO2 laser has
a dramatic ‘‘skin tightening effect’’ which is routinely
observed by clinicians immediately after the delivery of
laser pulses [8–11]. The mechanism of skin tightening is a
heat induced denaturation of the collagen fibers facilitated
by disruption of collagen cross-linking bonds that results
in an immediate shrinkage. It has been well demonstrated
in the literature that thermal induced shrinkage of collagen
by various devices repeatedly occurs when connective
tissue is heated to 65–758C [6–10]. The IUS approach
offers a potential for a similar thermal tissue effect at depth,
with the exception that the detrimental effects of epidermal
disruption can be avoided. Further investigative work must
be focused on the ultrastructural effect of IUS based
collagen denaturation, the degree of tissue shrinkage
produced by different IUS energy doses, and the safety of
this device in treating human patients.
CONCLUSION
This report demonstrates the creation of discrete TIZs
in porcine muscle and skin specimens using Intense
Ultrasound energy. We have tested the response of
these tissue models over a broad range of energy doses.
We were able to demonstrate a dose range that produces
selective, well-circumscribed TIZs at a desired depth (e.g.,
dermis or subcutaneous tissues), without overlying epidermal disruption. Thermal induced collagen denaturation
is an integral step in skin tightening by various laser
and radiofrequency devices. These energy modalities are
however, either depth-limited or energy density limited
to achieve selective thermal coagulation deep within
the skin tissue. Using ultrasound energy, this work is a
74
WHITE ET AL.
Fig. 9. NBTC stained gross tissue sections of porcine skin. The lesion profile of these three
different IUS handpieces demonstrates variability in thermal injury zones (arrows), as the
dermis and underlying tissue is targeted. Note that the TIZs resulting from the 7.5 MHz,
4.5 mm focus handpiece (10B) extends nominally shallower compared to the TIZs from the
4.4 MHz, 4.5 mm focus handpiece (10C). A: 7.5 MHz, 3.0 mm focus (1 J); B: 7.5 MHz, 4.5 mm
focus (2.2 J); and C: 4.4 MHz, 4.5 mm focus (2.6 J). [Figure can be viewed in color online via
www.interscience.wiley.com.]
demonstration of the ability to create controlled thermal
coagulative zones at various depths (order of millimeters),
within skin tissue, targeting respective anatomical
layers, while sparing the overlying epidermis. It is also
possible with the IUS device to detect TIZs with the built-in
ultrasound imaging component of the device.
The range of selective TIZs demonstrated using Intense
Ultrasound in this study extends from the dermis up to
the level of subcutaneous structures of the facial skin
tissue. The potential for using this device in treating the
aging face is encouraging, and needs to be investigated
further in human tissues [24–26].
ACKNOWLEDGMENTS
Funding for this work was provided in part by the
Ulthera, Inc., Mesa, AZ.
REFERENCES
1. McGahan JP, Goldberg BB, editors. Diagnostic ultrasound:
A logical approach. Wickford: Lippincott-Raven; 1997.
2. Manstein D, Herron GS, Sink RK, Tanner H, Anderson RR.
Fractional photothermolysis: A new concept for cutaneous
remodeling using microscopic patterns of thermal injury.
Lasers Surg Med 2004;34:426–438.
3. Kenndy JE, ter Haar GR, Cranston D. High intensity focused
ultrasound: Surgery of the future? Br J Radiol 2003;76:590–
599.
4. Makin IRS, Mast TDM, Faidi WF, Runk MM, Barthe PG,
Slayton MH. Miniaturized arrays for interstitial ablation and
imaging. Ultrasound Med Biol 2005;31(11):1539–1550.
5. Laubauch H-J, Barthe PG, Makin IRS, Slayton MH,
Manstein D. Confined thermal damage with Intense Ultrasound (IUS). Laser Surg Med 2006;38(18):32.
6. White WM, Makin IRS, Barthe PG, Slayton MH, Gliklich RE.
Selective transcutaneous delivery of energy to facial subdermal tissues using the ultrasound therapy system. Laser
Surg Med 2006;38(18):87.
7. Hruza GJ. Rejuvenating the aging face. Arch Facial Plast
Surg 2004;6:366–369.
8. Kim KH, Geronemus RG. Nonablative laser and light
therapies for skin rejuvenation. Arch Facial Plast Surg
2004;6:398–409.
9. Kirsh KM, Zelickson BD, Zachary CB, Tope WD. Ultrastructure of collagen thermally denatured by microsecond
domain pulsed carbon dioxide laser. Arch Dermatol 1998;134:
1255–1259.
10. Ross EV, Naseef GS, McKinlay JR, Barnette DJ, Skrobal M,
Grevelink J, Anderson RR. Comparison of carbon dioxide
laser, erbium:YAG laser, dermabrasion, and dermatome. A
study of thermal damage, wound contraction, and wound
healing in a live pig model: Implications for skin resurfacing.
J Am Acad Dermatol 2000;42:92–105.
11. Ross EV, Yashar SS, Naseef GS, Barnette DJ, Skrobal M,
Grevelink J, Anderson RR. A pilot study of in vivo immediate
tissue contraction with CO2 skin resurfacing in a live farm
pig. Dermatol Surg 1999;25:851–856.
12. Goco PE, Stucker FJ. Subdermal carbon dioxide laser
cutaneous contraction. Arch Facial Plast Surg 2002;6:37–
40.
13. Zelickson BD, Kist D, Bernstein E, Brown DB, Ksenzenko S,
Burns J, Kilmer S, Mehregan D, Pope K. Histology and
The American Journal of Sports
Medicine
http://ajs.sagepub.com/
The Effect of Thermal Heating on the Length and Histologic Properties of the Glenohumeral Joint
Capsule
Kei Hayashi, George Thabit III, Kathleen L. Massa, John J. Bogdanske, A.J. Cooley, John F. Orwin and Mark D.
Markel
Am J Sports Med 1997 25: 107
DOI: 10.1177/036354659702500121
The online version of this article can be found at:
http://ajs.sagepub.com/content/25/1/107
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On behalf of:
American Orthopaedic Society for Sports Medicine
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The Effect of Thermal Heating on the
Length and Histologic Properties of the
Glenohumeral Joint Capsule
Kei
Hayashi,* DVM, MS, George Thabit Ill,t MD, Kathleen L. Massa,* John J. Bogdanske,*
A. J. Cooley,* DVM, John F. Orwin,$ MD, and Mark D. Markel,*§ DVM, PhD
From the *
Comparative Orthopaedic Research Laboratory, School of Veterinary Medicine,
and &Dagger; Division of Orthopedic Surgery, School of Medicine, University of Wisconsin-Madison,
Madison, Wisconsin, and &dagger; Sports, Orthopedic and Rehabilitation Medicine Associates,
Menlo Park, California
Glenohumeral instability is a common and recurring problem, particularly in the young or athletic patient. 3,6,11,34
Multidirectional and unidirectional glenohumeral instability secondary to ligamentous laxity, capsular redundancy, and excessive joint volume are frequent occurrences
that current operative and nonoperative
treatments do not satisfactorily address in certain
groupS.4,7,24,26,43,45 Nonoperative treatment has an unacceptably high recurrence rate in the young and athletic
individuals Open surgical techniques result in high
morbidity and require prolonged rehabilitation. They also
return only a minority of athletes who use overhead movements back to their preinjury levels of activity. 5,24,32,34,46
Arthroscopic procedures have higher rates of failure than
open surgical techniques, require extreme technical expertise, and may be contraindicated in the case of capsular
redundancy-related shoulder instability.2, 8, 34 Therefore,
there appears to be a need for a simply performed, low
morbidity procedure that eliminates capsular redundancy, diminishes joint volume, and helps stabilize shoulders of patients, allowing them to return to their previous
levels of activity or performance.
A recent pilot study has demonstrated that the nonablative application of holmium:yttrium-aluminum-garnet
(Ho:YAG) laser energy to the joint capsule of patients with
glenohumeral instability shrank the joint capsule, stabilizing the shoulder in the majority of the patients treated.41 In this multi-institutional clinical trial, the Ho:YAG
laser, which has been approved for arthroscopic surgery,
was applied under arthroscopic guidance to patients with
glenohumeral instability but with no capsulolabral detachment or full-thickness rotator cuff tears. Although
this study did not have a comparable nonoperated control
population or an open surgical repair group, the results
indicated that at short-term followup (mean, 6 months)
patients improved dramatically after nonablative reduc-
ABSTRACT
A
The purpose of this
study was to evaluate the
effect of
temperature on shrinkage and the histologic properties
of gienohumeral joint capsular tissue. Six fresh-frozen
cadaveric shoulders were used for this study. Seven
joint capsule specimens were taken from different regions from each glenohumeral joint and assigned to
one of seven treatment groups (37°, 55°, 60°, 65°, 70°,
75°, 80°C) using a randomized block design. Specimens were placed in a tissue bath heated to one of the
designated temperatures for 10 minutes. Specimens
treated with temperatures at or above 65°C experienced significant shrinkage compared with those
treated with a 37°C bath. The posttreatment lengths in
the 70°, 75°, and 80°C groups were significantly less
than the pretreatment lengths. Histologic analysis revealed significant thermal alteration characterized by
hyalinization of collagen in the 65°, 70°, 75°, and 80°C
groups. This study demonstrated that temperatures at
or above 65°C caused significant shrinkage of glenohumeral joint capsular tissue. These results are consistent with histologic findings, which revealed significant thermal changes of collagen in the 65°, 70°, 75°,
and 80°C groups. To verify the validity of laser application for shrinkage of joint capsule, studies designed
to compare these findings with the effects of laser
energy must be performed.
§ Address correspondence and reprint requests to Mark D. Markel, DVM,
PhD, Comparative Orthopaedic Research Laboratory, School of Vetennary
Medicine, 2015 Lmden Dnve West, University of Wisconsin-Madison, Madi-
son, Wisconsin 53706.
No author or related institution has received any fmancial benefit from
research n this study. See &dquo;Acknowledgments&dquo; for funding information
107
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108
tion of redundant glenohumeral joint capsule using the
Ho:YAG laser. The mechanism that causes this effect has
not been identified, and the use of laser energy for the
of glenohumeral
treatment
instability remains
controversial.
The potential applications of lasers in surgery and medicine have been evaluated in a wide variety of specialties
since their initial development. Recent scientific studies
evaluating laser energy for tissue welding and thermokeratoplasty have noted that collagenous tissue will
shrink after application of laser energy at nonablative
levels.18,29,39 We previously reported that Ho:YAG laser
energy at nonablative levels can significantly alter joint
capsular length and its mechanical properties in an in
vitro rabbit study.22 Histologic evaluation of this effect
suggested that the application of nonablative Ho:YAG laser energy to joint capsular tissue caused thermal alterations in collagen and fibroblasts.2° The interactions between laser energy and tissue are based on photothermal,
photochemical, photomechanical, and photoacoustic effects.1O,27,30,33 Thermal shrinkage of collagen is a welldescribed phenomenon.’, 12, 14, 17,19,23 Based on these findings, we hypothesized that the shrinkage of the capsular
tissue induced by laser energy is predominantly caused by
the thermal effect of laser energy and is a function of
temperature. To date, temperature change caused by laser
energy, and the temperature profile of joint capsular tissue, have not been reported. The long-term goal of this
project is to correlate temperature with laser energy’s
effect on joint capsular tissue. The purpose of this portion
of the study was to evaluate the effect of temperature on
shrinkage and histologic properties of glenohumeral joint
capsular tissue using
bath.
a
Figure 1. Schematic drawing of the glenohumeral joint capsular tissue illustrating the seven regions used in this study (1
through 7). B, biceps tendon; G, glenoid; SGHL, superior
glenohumeral ligament; MGHL, middle glenohumeral ligament ; IGHLC, inferior glenohumeral ligament complex; AB,
anterior band; AP, axillary pouch; PB, posterior band; PC,
posterior capsule.
temperature-controlled tissue
MATERIALS AND METHODS
Six fresh-frozen cadaveric shoulders
were
used for this
study (age, 52.3 ± 4.9 years; mean ± SD). This study was
approved by the institutional review board. Overlying
musculature was carefully dissected away from the joint
capsule and the joint capsule was opened by an incision
parallel to the biceps tendon. The entire joint capsule was
then completely detached from the glenoid and humerus.
Seven regions of interest were dissected in a radial manner from the glenoid edge to the humeral edge, yielding
specimens of precisely the same width (10 X 30 mm).
Specimens were collected from the superior, middle, and
inferior (anterior band, axillary pouch, posterior band)
glenohumeral ligament and adjacent capsule, and the posterior capsule (inferior portion, superior portion) (Fig. 1),
according to the description by O’Brien et a1.31 Specimens
were assigned to one of seven treatment groups (37°, 55°,
60°, 65°, 70°, 75°, 80°C) using a randomized block design
(N 6). Each specimen was placed in a custom-made jig
with a pulley system designed to provide a constant load
(0.098 N) on the specimen (Fig. 2). Initial length between
the grips was set at 20 mm. After measurement of pretreatment tissue length in a 37°C tissue bath, specimens
were placed in tissue baths of lactated Ringer’s solution
=
Figure 2. Capsular tissue mounted in a custom-made jig
a pulley system designed to provide a constant load
(0.098 N) on the specimen in a temperature-controlled bath
of lactated Ringer’s solution.
with
heated to one of the designated treatment temperatures
(Dual Water Bath Model 188, Precision Scientific, Chicago, Illinois). Changes in tissue length were recorded at
15-second intervals for 10 minutes. Specimens were then
placed in a 37°C tissue bath to determine posttreatment
length. The treatment time and temperature settings
were determined based on a pilot study.
Immediately after treatment, specimens were processed
for histologic staining with hematoxylin and eosin and
Masson’s trichrome, or fixed in modified Karnovsky’s solution (2% paraformaldehyde and 1.25% glutaraldehyde
Downloaded from ajs.sagepub.com by guest on July 13, 2011
109
in 0.1 M sodium phosphate buffer, pH 7.0) for transmission electron microscopy. A subjective scoring system was
used to evaluate the effect of temperature on the histologic
properties of collagen structure. Slides stained with hematoxylin and eosin were graded in a blinded manner on
a scale from 0 to 3 (0, normal; 1, mild change characterized
by diffuse hyalinization with visible fibrous structure; 2,
moderate change characterized by homogeneous bundles
of fibers; 3, severe thermal alteration characterized by
homogeneous mass of tissue).
A paired t-test was used to compare pre- and posttreatment lengths for each group. A two-way analysis of variance (ANOVA) test was used to evaluate the effect of
treatment and tissue region. A one-way ANOVA test was
used to evaluate the differences among treatment groups.
When the ANOVA test revealed significant differences
among groups, Duncan’s multiple range test was performed to analyze these differences. The Kruskal-Wallis
test was used to compare groups for the subjective scores.
When the Kruskal-Wallis test revealed differences among
the groups, the Wilcoxon rank sum test was performed to
analyze these differences. Differences were considered to
be significant at P <_ 0.05.
RESULTS
A
The results of tissue shrinkage measurements are summarized in Table 1 and Figure 3. Treatments with 37°,
55°, and 60°C tissue baths did not cause significant
changes in tissue length (P > 0.05). Temperature treatments at or above 65°C caused significant shrinkage when
compared with the 37°, 55°, and 60°C groups (P < 0.05).
There was no significant difference in this parameter between the 75° and 80°C treatment groups (P > 0.05).
Posttreatment lengths in the 70°, 75°, and 80°C groups
were significantly less than pretreatment lengths (P <
0.05). Shrinkage of the tissue started immediately after
the onset of temperature treatment and reached its maximum within 3 minutes in the 75°C group and 1.5 minutes
in 80°C group. The tissue shrinkage was accompanied by
swelling in the direction perpendicular to that of the
shrinkage. The type of tissue (i.e., from different regions)
had no significant effect on tissue shrinkage (P > 0.05).
TABLE 1
Results of Joint Capsular Tissue Shrinkage
(mean ± SD)*
Analysis
Means within the column with differing letters were signifiother (P < 0.05).
from pretreatment length (P < 0.05).
$ 100 X (Pretreatment length - Posttreatment length)/Pretreatment length.
*
cantly different from each
t Significantly different
Figure 3. Changes in tissue length of the seven treatment
during testing (mean length for each group).
groups
Histologic analysis revealed significant thermal alter- B
ation characterized by hyalinization of collagen in the 65°,
70°, 75° and 80°C groups (Fig. 4). In the 75° and 80°C
treatment groups, collagen showed moderate-to-severe
changes characterized by homogeneous bundles of collagen fibers (Fig. 4d). Histologic scores for collagen were
significantly higher for the 65°, 70°, 75°, and 80°C groups
than for 37°, 55°, and 60°C groups (P < 0.05) (Fig. 5).
There were no significant differences in this parameter
among the 70°, 75°, and 80°C groups (P > 0.05). Specimens stained with Masson’s trichrome revealed an altered
staining pattern of collagen in the 75° and 80°C treatment
groups, with the homogeneous areas staining red rather
than the normal blue. Transmission electron microscopy
revealed significant ultrastructural alterations in collagenous architecture in the 65°, 70°, 75°, and 80°C treatment groups (Fig. 6). The margins of the collagen fibrils
began to lose their distinct edges, while periodic cross- C
striations were still visible in the 65°C treatment group
(Fig. 6b). In the 70°C treatment group, collagen fibrils
became swollen with the loss of cross-striations (Fig. 6c).
Fibrillar structure was almost lost in the 80°C treatment
group (Fig. 6d).
DISCUSSION
This study demonstrated that temperature treatments at
or above 65°C with a tissue bath of lactated Ringer’s
solution caused significant shrinkage of glenohumeral
joint capsular tissue. Higher temperatures caused an increase in capsular shrinkage; however, there was no significant difference in tissue shrinkage between the 75°
and 80°C groups. These results were consistent with histologic findings, which revealed significant thermal
changes of collagen in the 65°, 70°, 75°, and 80°C groups.
Although the anatomic and functional structure of the
glenohumeral joint capsule is not homogeneous, 31,40 there
was no significant effect of tissue region on tissue
shrinkage.
Heat-induced shrinkage of tissue associated with denaturation of collagen is a well-described phenomenon. The
thermal properties of collagen have been extensively stud-
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110
Figure 5. Subjective histologic
of the
seven
differing letters
(P < 0.05).
Figure 4. Light micrograph of joint capsular tissue from a)
the 37°C treatment group demonstrating normal fibrous
structure of collagen, b) the 65°C treatment group demonstrating fusion of collagen with visible fibrous structure, c) the
70°C treatment group demonstrating fused bundles of colicgen, and d) the 80°C treatment group demonstrating homogeneous collagen (hematoxylin and eosin stain; original
magnification, x 50).
ied in a variety of experimental models since the mechanism was proposed by Flory et al. 14-16 Flory et al.14,16
stated that thermal contraction of collagen is brought
about by a molecular structure transition between the
triple helix and a random coil. It has been shown that
the thermal properties of collagen vary with the age of the
animal and the environmental condition.1,23,42 Rosenbloom et al. 36 found, using a chick tendon model, that
hydroxyproline content determines the denaturation temperature of collagen. More recently, Allain et al.’ described collagen network behavior under the influence of
heat during hydrothermal shrinkage and swelling of rat
skin. The investigators proposed that swelling and shrinkage of collagen fibrils are secondary to unwinding of the
triple helix with maintenance of heat-stable intermolecular crosslinks. Horgan et al. 23 reported a strong correlation between thermal properties of tendon and the concen-
are
for collagen structure
(mean ± SD). Bars with
score
treatment groups
significantly different from
each other
Figure 6. Transmission electron micrograph of joint capsular tissue from a) the 37°C treatment group demonstrating
circular distinct fibrils on cross-section (left) and periodic
cross-striations on longitudinal section (right), b) the 65°C
treatment group demonstrating loss of the fibrils’ distinct
edges with visible cross-striations, c) the 70°C treatment
group demonstrating increases in fibril diameter and loss of
striations, and d) the 80°C treatment group demonstrating
loss of fibrillar structure with amorphous appearance (original
magnification, x24,000; bar 1 p,m).
=
tration of nonreducible crosslinks. To date, the thermal
properties of collagen have been explained mainly in
terms of the crosslinks. A number of different methods
have been used to study this characteristic, including differential scanning calorimetry, ultraviolet difference spectroscopy, isometric tension measurement, and isotonic
contraction measurement. 1, 9, 23, 25, 42
Applications of the thermal shrinkage properties of collagen have been evaluated for ocular surgery. The concept
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111
treatment of keratoconus
from studies that demonstrated corneal stromal collagen shrinks to approximately one-third of its original
length when heated to a temperature of 60°C to 65°C.28,38
Moreira et al. 29 reported that thermokeratoplasty can be
achieved successfully by using the Ho:YAG laser. Thermal
tendinoplasty has been used to shrink extraocular muscle
tendon for the treatment of strabismus. 13 The investigators proposed the advantages of thermal tendinoplasty
over conventional surgical techniques. These studies indicate that long-lasting alterations of collagenous architecture can be achieved by thermal application, with greater
ease and less inflammatory response than conventional
of
thermokeratoplasty for the
arose
surgical techniques.
A recent pilot study has
demonstrated that the application of Ho:YAG laser energy at a nonablative level shrank
the redundant joint capsule of patients with glenohumeral
instability, helping stabilize the shoulder in the majority
of the patients treated.41 The interactions between laser
energy and tissue are based on photothermal, photochemical, photomechanical, and photoacoustic effects. 10, 27,30,33
Our previous studies revealed evident thermal effects of
laser energy on the histologic properties of the joint capsular tissue.’o The histologic properties of collagen altered
by temperature treatments at or above 65°C in this study
were similar to those altered by laser energy. Both nonablative laser energy and thermal heating in a lactated
Ringer’s solution tissue bath at or above 65°C can significantly shrink joint capsular tissue.22 Based on these findings, we hypothesized that the shrinkage of joint capsular
tissue achieved by nonablative laser energy was predominantly caused by the thermal effect of laser energy.
On the other hand, this study demonstrated that ultrastructural changes after tissue bath treatment appear to
be different from those induced by nonablative laser treatment. Transmission electron microscopy revealed loss of
the collagen fibril’s distinct edges and their periodic crossstriations with increased fibril diameters after either tissue bath treatment or nonablative laser application.
These findings have been reported in previous studies
evaluating ultrastructural alterations of collagen fibril
caused by hydrothermal or laser treatment. 25,37,44 Detailed transmission electron microscopic studies indicate
that the ultrastructural changes caused by hydrothermal
heating vary depending on the environmental condition,
degree of crosslinks, and area within the fibri1.25,42 Our
previous laser-ultrastructural study revealed that, although the fibril diameters increased dramatically and
their edges were less distinct, individual circular collagen
fibrils were evident after nonablative laser treatment.21
In the present study, however, transmission electron microscopy revealed that distinct fibrillar structure was lost
at or above 70°C, indicating transition of fibrillar collagen
to an amorphous state. These different findings may be
due to the different mode of heating, tissue differences, or
the effects of laser energy other than the purely thermal
effect. In addition to collagen, other components of the
tissue must be involved in the interaction. Further studies
are needed to clarify the different effects between laser
energy and hydrothermal heating.
This study demonstrated that significant shrinkage of
glenohumeral joint capsular tissue can be achieved by
hydrothermal heating at or above 65°C. However, these
results should be interpreted with caution because mechanical properties after the treatments were not evaluated in this study and mechanical properties of the tissue
could be significantly altered after thermal shrinkage.
Our previous laser study showed that, despite significant
tissue shrinkage, the relaxation properties of joint capsular tissue did not change after application of laser energy,
but the stiffness of the joint capsule decreased at higher
energies.22 More importantly, the biological response to
the thermally altered tissue after the treatment must be
considered. Tissue properties will change with time during the inflammatory, healing, and remodeling processes.
Thermal alteration of tissue might result in detrimental
effects on tissue properties and joint stability, although
altered mechanical and architectural properties may return to normal as healing occurs, while maintaining the
tissue at its shorter length. To verify the validity of laser
application for shrinkage of the joint capsule, in vivo studies designed to evaluate the effect of thermal modulation
of the tissues with time must be performed.
CONCLUSIONS
This study revealed the effects of hydrothermal heating on
the shrinkage and histologic and ultrastructural properties
of the glenohumeral joint capsule, demonstrating that temperatures at or above 65°C caused significant thermal alteration of the tissue. The results of this study suggest that the
glenohumeral joint capsular shrinkage induced by the nonablative application of Ho:YAG laser energy is mainly caused
by the thermal effect of laser energy on collagen. Further
studies designed to compare these findings with the effects of
laser energy, including mechanical and thermometric studies, are required to evaluate the mechanism of laser-induced
shrinkage of the joint capsular tissue.
ACKNOWLEDGMENTS
This work was supported by Coherent, Palo Alto, California ; the Department of the Navy-ONR (N000014-90-C0029) ; NIH LAMP Resource RR 001192; NASA NAG2568 ; the Beckman Laser Institute and Medical Clinic,
Irvine, California; University of California, San Francisco,
School of Medicine, San Francisco, California; and The
National Disease Research Interchange, Philadelphia,
Pennsylvania. The authors thank Renate Bromberg.
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A
622
OPTICS LETTERS / Vol. 30, No. 6 / March 15, 2005
Monitoring the thermally induced structural
transitions of collagen by use of second-harmonic
generation microscopy
Sung-Jan Lin
Institute of Biomedical Engineering, College of Medicine and College of Engineering,
National Taiwan University, Taipei 100, Taiwan,
and Department of Dermatology, National Taiwan University Hospital, Taipei 100, Taiwan
Chih-Yuan Hsiao
Institute of Electro Optics, Department of Electrical Engineering, National Taiwan University, Taipei 100, Taiwan
Yen Sun and Wen Lo
Department of Physics, National Taiwan University, Taipei 106, Taiwan
Wei-Chou Lin
Department of Pathology, National Taiwan University Hospital, Taipei 100, Taiwan
Gwo-Jen Jan
Institute of Electro Optics, Department of Electrical Engineering, National Taiwan University, Taipei 100, Taiwan
Shiou-Hwa Jee
Department of Dermatology, National Taiwan University Hospital, Taipei 100, Taiwan,
and Department of Dermatology, National Taiwan University College of Medicine, Taipei 100, Taiwan
Chen-Yuan Dong
Department of Physics, National Taiwan University, Taipei 106, Taiwan
A
Received September 10, 2004
The thermal disruption of collagen I in rat tail tendon is investigated with second-harmonic generation
(SHG) microscopy. We investigate its effects on SHG images and intensity in the temperature range
25° – 60° C. We find that the SHG signal decreases rapidly starting at 45° C. However, SHG imaging reveals
that breakage of collagen fibers is not evident until 57° C and worsens with increasing temperature. At
57° C, structures of both molten and fibrous collagen exist, and the disruption of collagen appears to be complete at 60° C. Our results suggest that, in addition to intensity measurement, SHG imaging is necessary for
monitoring details of thermally induced changes in collagen structures in biomedical applications. © 2005
Optical Society of America
OCIS codes: 190.4160, 170.3880, 180.0180.
In recent years, multiphoton fluorescence microscopy
has gained significant popularity in bioimaging applications. The nonlinear excitation of fluorescence
photons with ultrafast, near-infrared excitation
sources has important advantages in its ability to acquire enhanced axial depth discrimination images,
reduced overall specimen photodamage, and increased imaging penetration depths.1,2 In addition to
multiphoton fluorescence imaging, nonlinear polarization effects from a special class of biological materials also have biomedical significance. In biological
structures lacking inversion symmetry a nonvanishing second-order susceptibility can contribute to a
second-harmonic generation (SHG) signal given by
2
E jE k .
Pi = !ijk
!1"
A variety of biological materials, such as collagen and
muscle fibers, have been shown to be effective in gen0146-9592/05/060622-3/$15.00
erating second-harmonic signals.3,4 In the case of collagen, SHG imaging is of general interest since collagen is widely found in tissues such as tendon, skin,
and cornea and is a major constituent of the extracellular matrix. A particularly interesting application of
SHG imaging is the monitoring of thermally induced
structural transitions of collagen fibers. A number of B
medical procedures depend on heat-induced changes
in collagen fibers to achieve therapeutic results. In
laser-assisted capsulorrhaphy, laser heating of collagen in the shoulder can result in fiber shrinkage
and enhanced stability of the shoulder joint.5 Another
procedure is conductive keratoplasty, in which
current-induced heat is used to change the cornea
curvature for vision correction.6 Finally, heat from a
laser source can be used to tighten and rejuvenate
skin.7 The thermal effect on collagen has been investigated by measurement of the second-harmonic signal. It was found that the collagen SHG signal de© 2005 Optical Society of America
March 15, 2005 / Vol. 30, No. 6 / OPTICS LETTERS
A
creases at approximately 64° C, presumably because
of a structural transition in the collagen internal
structure.8,9 Laser illumination has also been shown
to induce thermal damage to collagen fibers.10 However, to the best of our knowledge, the correlation between the thermally induced decrease in the SHG
signal and collagen structures is not completely understood. In this work we obtain the SHG images
from rat tail tendon after thermal treatment in the
temperature range between 25° C and 60° C. We correlate the structural changes in collagen to changes
in the SHG signal. An understanding of this relationship will help researchers in developing thermal procedures for biomedical applications.
The second-harmonic imaging system used in this
study is a modified version of a home-built laserscanning microscopic-imaging system based on an
upright microscope (E800, Nikon, Japan) described
previously.11 A diode-pumped (Millennia X, SpectraPhysics, Mountain View, California), Ti:sapphire
(ti-sa; Tsunami, Spectra-Physics) is used as the excitation source. The 780-nm output of the ti-sa laser is
scanned in the focal plane by a galvanometer-driver
x – y mirror scanning system (Model 6220, Cambridge
Technology, Cambridge, Massachusetts). Before entering the upright microscope, the laser is beam expanded to ensure overfilling of the objective’s back
aperture. For high-resolution imaging a highnumerical-aperture, oil-immersion objective (S Fluor
40", N.A. of 1.3, Nikon) was selected for SHG
microscopy. To direct the expanded laser spot
to the sample, a short-pass dichroic mirror
(700DCSPXRUV-3p, Chroma Technology, Brattleboro, Vermont) is used to reflect the incident excitation laser source. To ensure even excitation of collagen fibers at different orientations, a # / 4 wave
plate is used to convert the linearly polarized ti-sa laser beam into one with circular polarization. The average laser power at the sample is 5.1 mW, and the
SHG signal generated at this power is found to be
within the quadratic-dependence region of the SHG
signal of the excitation power. The generated SHG
signal is then collected in a backscattering geometry
in which the dichroic mirror, a short-pass filter
(E680SP, Chroma Technology), and a 390-nm bandpass filter (HQ390/20, Chroma Technology) are used
to isolate the SHG signal. The signal photons are processed by a single-photon-counting photomultiplier
tube (R7400P, Hamamatsu, Japan) and a home-built
discriminator.
In our study, rat tail tendon is sliced into small sections and placed in a phosphate-buffered saline
buffer before being subjected to thermal baths for
10 min. In this manner we can ensure rapid and uniform heating and cooling of the tendon specimens
upon placing them into and removing them from the
thermal bath. The temperature range between 25 ° C
and 60° C is chosen for thermal treatment of the
specimens. At the end of the 10-min heating cycle the
tendon specimen is removed, mounted on a glass
slide, and covered with a No. 1.5 cover glass for viewing. We acquire second-harmonic images of the tendon treated at different temperatures. To gain a thor-
623
ough understanding of the effects of heating on
collagen, a large-area scan of each collagen specimen
composed of a 6 " 6 array of neighboring SHG images
is acquired and assembled. The average SHG signal
per pixel is computed and plotted. To eliminate the
effects of sample scattering or refractive-indexinduced spherical aberration on the measured SHG
signals, we acquire the SHG images only at the surface of the tendon specimen.
Shown in Fig. 1 are the rat tail tendon SHG images
(along with histological images) acquired at six temperatures of 25° C, 40° C, 52° C, 55° C, 57° C, and
60° C. The thermal dependence of the SHG signal
over the entire temperature range is shown in Fig. 2.
A qualitative examination of Figs. 1 and 2 shows several interesting results. First, compared with lowertemperature results, the SHG images at 52° C and
55° C indicate that collagen fibers tend to demonstrate a greater degree of curvature with increasing
temperature. Furthermore, although the SHG sig-
Fig. 1. SHG images !390 nm" of rat tail tendon treated for
10 min at different temperatures. Disruption of collagen
structures are indicated by arrows. Histological images are
shown for comparison.
B
624
OPTICS LETTERS / Vol. 30, No. 6 / March 15, 2005
In conclusion, SHG microscopy of thermally
treated rat tail tendon has shown that, despite a decrease in SHG signal at the earlier temperature of
45° C, breakages to collagen fibers are not evident
until 57° C. At increasing temperatures, SHG images
demonstrate further disruption in the collagen structure. Our results show that SHG imaging is an effective method for fully characterizing the thermal effects of collagen fibers and may be developed into an
effective imaging technique for in vivo biomedical applications.
Fig. 2. Dependence of the SHG signal !390 nm" of rat tail
tendon at different temperatures.
A
nals start to sharply decrease at 45° C, breakages in
collagen fibers are not evident until approximately
57° C (indicated by arrows). This observation, together with the decrease in the SHG signal, indicates
that the structure lacking inversion symmetry responsible for the collagen SHG signal has been disrupted. At 57° C our SHG image shows the coexistence of two types of region with different structural
organization. The SHG image still shows the existence of collagen fibers; however, there are regions
within the collagen fibers in which the SHG signal is
absent. This observation is supported by a histological image in which the fibrous and molten regions are
indicated by elliptical and rectangular regions, respectively. At 60° C, collagen denaturation is more
complete as a further decrease of the SHG signal is
correlated to a further disruption of the fibrous structures. Note that in our approach the large-area SHG
scan is critical in identifying the changes to collagen
fibers. In localized microscopy, features such as fiber
breakage and collagen denaturation may not be easily identified.
Our results suggest that SHG intensity and images
need to be combined to assess the overall changes to
thermally treated collagen specimens. In addition,
since the thermal response of each type of collagen
tissue may be different, therapeutic procedures using
thermal effects should be performed to determine the
threshold damage level to collagen.
This work was supported by the National
Research Programs for Genomic Medicine, Taiwan
(grants NSC 92-2112-M-002-018 and NSC 92-3112-B002-048).
S.-H.
Jee’s
e-mail
address
is
[email protected];
C.-Y.
Dong’s
is
[email protected].
References
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Maloney, J. Davidorf, M. Sabry, and the Conductive
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Aesth Plast Surg (2011) 35:87–95
DOI 10.1007/s00266-010-9564-0
ORIGINAL ARTICLE
Three-Dimensional Radiofrequency Tissue Tightening:
A Proposed Mechanism and Applications for Body
Contouring
Malcolm Paul • G. Blugerman • M. Kreindel
R. S. Mulholland
•
Received: 20 April 2010 / Accepted: 6 July 2010 / Published online: 11 September 2010
! The Author(s) 2010. This article is published with open access at Springerlink.com
Abstract The use of radiofrequency energy to produce
collagen matrix contraction is presented. Controlling the
depth of energy delivery, the power applied, the target skin
temperature, and the duration of application of energy at
various soft tissue levels produces soft tissue contraction,
which is measurable. This technology allows precise soft
tissue modeling at multiple levels to enhance the result
achieved over traditional suction-assisted lipectomy as well
as other forms of energy such as ultrasonic and lasergenerated lipolysis.
Keywords Body contouring ! Liposuction !
Radiofrequency energy ! Soft tissue contraction
Introduction
Radiofrequency (RF) thermal-induced contraction of collagen is well known in medicine and is used in ophthalmology, orthopedic applications, and treatment of varicose
veins. Each type of collagen has an optimal contraction
temperature that does not cause thermal destruction of
connective tissue but induces a restructuring effect in
M. Paul (&)
Department of Surgery, Aesthetic and Plastic Surgery Institute,
University of California, Irvine, CA 92697, USA
e-mail: [email protected]
G. Blugerman
Buenos Aires, Argentina
M. Kreindel
Invasix Corp, Toronto, ON, Canada
R. S. Mulholland
SpaMedica, Toronto, ON, Canada
collagen fibers. The reported range of temperatures causing
collagen shrinkage varies from 60 to 80"C [1–7]. At this
temperature tissue contraction occurs immediately after
tissue reaches the threshold temperature. The shrinkage of
tissue is dramatic and can reach tens of percent of the
heated tissue volume. This type of contraction is well
studied in cornea [1], joints [2], cartilage [4, 7], and vascular tissue [5] but its application for skin, subdermal tissue, and subcutaneous tissue tightening has not been
studied.
Noninvasive RF and lasers have been used for skintightening effects since the mid-1990s [6, 8–12]. Because
of superficial thermal safety concerns, the skin surface
temperature is maintained below 45"C. To increase the
temperature in the deep dermis the skin is heated with RF
or laser energy penetrating into the tissues deeper than
1.5 mm, with simultaneous skin surface cooling. This
sophisticated method of transepidermal, noninvasive RF
thermal delivery provides a variable and controversial
tightening effect, which is not usually apparent, if at all,
until dermal remodeling occurs a few months after the
treatment. Noninvasive tissue tightening treatments have
an inherent safety limitation because energy is delivered
through the skin surface and the threshold epidermal burn
temperature is significantly lower than the optimal temperature for the collagen contraction. Studies indicate that
deeper penetrating energy provides better skin contraction
and RF energy, by penetrating deeper than laser radiation,
is a superior method, not only for treatment of facial rhytides and laxity, but also for body tightening [6, 9, 12]. It is
the physical and biological characteristics of RF that
explain its superior three-dimensional mechanism of skin
tightening.
Recently, the use of thermal-induced tissue tightening
was expanded to minimally invasive treatments [13–16].
123
88
Aesth Plast Surg (2011) 35:87–95
Using laser-assisted liposuction (LAL) or radiofrequencyassisted liposuction (RFAL), physicians have attempted to
achieve reduction of subcutaneous tissue with simultaneous
tissue contraction [13, 16]. DiBernardo [13] reported 17%
skin surface shrinkage measured at 3 months follow-up
after LAL treatment. RFAL technology provides much
higher power and more efficient energy transfer than laser
energy systems and thus allows the treatment of larger
volumes of subcutaneous tissue with optimal thermal profiles, facilitating the significant tightening of the tissue.
Paul and Mulholland [16] introduced a RFAL and soft
tissue contraction technology that showed tremendous
promise for thermal contouring. Invasive thermal treatments are superior because the RF conduit (RFAL emitting
cannula) targets the whole volume of treated tissue with
critical thermal energy, not only the superficial subdermal
layer, and the invasive RF treatments can heat deep adipose
and subcutaneous tissue to much higher temperatures
without compromising skin safety.
When considering skin contraction we have to differentiate two-dimensional horizontal x-axis tightening of the
skin surface from three-dimensional x-y-z tissue tightening
of the subcutaneous tissue, where the skin is also more firmly
connected and adjacent to the deeper anatomical structures.
If two-dimensional contraction is a function of collagen
structure changes in the dermis, the three-dimensional tissue-tightening changes involve contraction of different
types of collagenous tissue. We can separate the following
types of collagen tissue in the subcutaneous space:
•
•
•
•
Dermis: papillary and reticular
Fascia: relatively thick layer of connective tissue
located between muscles and skin
Septal connective tissue: thin layers of connective
tissue separating lobules of fat and connecting dermis
with fascia
Reticular fibers: framework of single collagen fibers
encasing fat cells
One of the main objectives of this study was to evaluate the
possibility of immediate thermal-induced subcutaneous
tissue contraction and to estimate the thermal threshold of
the effect. In this study we compare the threshold temperature and contraction level of different types of ex vivo
collagenous tissue samples and the clinical results based on
RFAL results for body contouring.
Materials and Methods
A
Ex Vivo Experiment Setup
An ex-vivo study was conducted to measure subcutaneous
collagenous tissue contraction with simultaneous monitoring
123
of local tissue temperature to determine the threshold temperature of the collagen shrinkage. Three types of collage- B
nous tissue were studied for thermal-induced contraction:
(1) adipose tissue with septal and reticular connective tissue,
(2) dermis, and (3) fascia.
Samples of ex vivo human tissue were taken from an
abdominoplasty surgery and were tested within 10 min of
excision. Immediate thermal testing was performed to
minimize changes in tissue related to long storage and
temperature variation or change of liquid content, including blood and lymphatic content. The tissue samples were
placed between the two BodyTiteTM (Invasix Ltd., Israel)
RF electrodes, where the small-area, internal RF-active
electrodes (cannula) were placed in contact with the studied tissue and the other large-area electrode was applied to
the opposite side, or epidermal side, of the sample. Large
samples of subcutaneous tissue were used, allowing
observation of any contraction behavior in the tissue’s
native environment in connection with its entire matrix
structure. Two marks were placed 1 cm from the active
internal electrode to visualize tissue displacement. The
experiment design setup is shown in Fig. 1.
RF energy was delivered by the BodyTite device. The
delivered power was 70 W at 1 MHz, and energy was
delivered until evaporation of water from the adipocytes
was observed. Video and thermal cameras (FLIR A-320)
were used to monitor tissue displacement and temperature
change during the treatment. The start of tissue displacement was correlated with tissue temperature to determine
the contraction thermal threshold. Each experiment was
repeated three times for each type of tissue to sample tissue
averages and avoid measurements of random events.
In Vivo Evaluation with Radiofrequency-Assisted
Liposuction (RFAL)
Twenty-four consecutive patients, 22 female and 2 male,
underwent RFAL to the abdomen and hips. The average
age was 39.7 years (range = 19-52 years). The average
preoperative weight was 71 kg. The selected patients were
typical patients indicated for a liposuction procedure. All
patients were healthy anesthetic risks and active with no
significant medical diseases. Fifteen of 24 patients had a
normal BMI (\25), while 9/24 patients were moderately
overweight (BMI = 25–30) and 3 patients were obese
(30 \ BMI \ 32).
RFAL was performed using the BodyTite device. The
BodyTite device deploys a handpiece to deliver radiofrequency energy to the adipose tissue and skin. The internal
cannula is coated with dielectric material and has a conductive tip that emits RF energy into the adipose tissue
toward the skin surface. RF energy flows between the tip
of the internal cannula and external electrode creating a
Aesth Plast Surg (2011) 35:87–95
89
Fig. 1 Ex vivo experimental
setup
Fig. 2 Schematic drawing of
RF handpiece inserted into the
body
localized, confined thermal effect between them. The
internal cannula is inserted into the pretumesced fat to be
contoured and is moved gently back and forth at various
predetermined and controlled depths for uniform heating of
the treated volume. There is also an external electrode that
moves along the surface of the skin in tandem vertical
alignment with the tip of the internal cannula (Fig. 2). The
subcutaneous tissue and skin between the electrodes
experience a significant thermal effect which is maximal
near the tip of the internal cannula and decreases in
intensity toward the skin electrode The operator controls
the depth of the internal cannula within a predetermined
range of 5–50 mm and moves the handpiece back and forth
through the desired fat volume to be contoured. The RF
energy coagulates the adipose, connective, and vascular
tissues in the vicinity of the internal cannula tip and gently
heats the dermis below the external electrode. The internal
electrode also serves as an asynchronous internal suction
cannula, aspirating the coagulated adipose, vascular, and
fibrous tissues.
The RF power, in the range of 40–70 W, was used for
uniform heating throughout a thick subcutaneous flap. The
average total energy of about 72 kJ was delivered to
the abdominal area. The temperature around the tip of the
cannula reached 70–80"C. This internal temperature was
observed using thermography on tissue cross section for
preabdominoplasty patients treated with RFAL when the
Fig. 3 Temperature profile inside adipose tissue during the RFAL
treatment
skin surface temperature reached 38–42"C (see Fig. 3,
cross section of lower abdominal tissue showing the thermal image of the skin surface and tissue incision allowing
visualization of the thermal profile of the internal subcutaneous temperature). The target skin temperature was
monitored and controlled with a thermal sensor built into
the external electrode. The sensor provides continuous realtime epidermal temperature monitoring and feedback loop
control of RF power. The system was set to a target temperature of 38–42"C, which was maintained for 1–3 min.
The strong and sustained tissue heating during the
123
90
Aesth Plast Surg (2011) 35:87–95
Fig. 4 Before and after RFAL
and intraoperative two-point
linear contraction registration
points from pubic RFAL
incision point to the lower
pole of the umbilicus
procedure results in thermal stimulation of the subdermal
layer, the entire matrix of adipose tissue, and the vertical
and oblique fibrous septa, eliciting a powerful threedimensional retraction and contraction of the entire soft
tissue envelope.
The distance between the internal and external electrodes was controlled with an eccentric spring-loaded
mechanism that keeps the external electrode on the surface
of the skin at all times. The device also controls vaporization and prevents carbonization around the tip of the
cannula. When evaporation around the internal cannula
occurs, the tissue impedance rises and exceeds the online
monitored high impedance and the device shuts off the RF
energy.
All patients had their treatment area infiltrated with
tumescent anesthesia prior to the RFAL procedure.
Tumescent anesthesia is critical in the technique as the RF
current travels through tissue most efficiently in a salinated
environment.
The objective of this in vivo portion of the study was
to optimize treatment parameters and correlate treatment
soft tissue contraction results with procedure and patient
variables, including amount of deposed RF energy, body
mass index (BMI) of the patients, and amount of aspirated
fat.
A zone measuring as large as 15 9 10 cm (150 cm2)
may be heated to critical target temperature within 3–8 min
depending on the thickness of the treated fat layer and then
uniform volumetric heating can be safely performed to
reach uniform temperature distribution over the entire
treated volume.
All patients from the study were followed up at 6, 12,
and 24 weeks. To measure linear contraction, the distance
between two fixed points was measured preoperatively and
then at the 24-week postoperative visit. Distances between
incision ports and natural ‘‘fixed’’ anatomical registration
points, such as moles or the umbilicus, were measured
before the treatment, after the treatment, and at 3- and
123
6-month follow-up visits. The linear contraction was
measured as relative change of distance between two points
over the curved surface of the body. Distances were measured using a flexible ruler applied over the skin surface.
For the abdominal area, at least three measurements were
taken between three different points and average linear
contraction was calculated (Fig. 4).
Pre- and postoperative photography, weights, and circumferential reduction data were obtained on all patients.
One RFAL study patient had a biopsy of the thermally
treated skin 12 months after the procedure during which
epidermal skin temperatures of 40"C had been attained and
there was an area contraction of 43% at 6 months.
Results and Discussion
Ex Vivo Tissue Contraction Experiments
The adipose tissue with septal and reticular collagen
behavior is shown in Fig. 5. The experiments showed that
the marker movement (contraction) started within 2 s after
the start of RF energy delivery. Tissue contraction was not
symmetrical as the displacement from one side was 8 mm
and from the other side the average displacement was
3 mm. Adipose fibrous septal tissue coagulation and
vaporization started to be observed at 13 s after the initiation of RF energy. Nonsymmetrical behavior can be
explained by the nonuniform structure of connective tissue
and the nonsymmetrical geometry of the studied tissue
sample. The average marker migration and tissue contraction for the three experiments with adipose tissue was
6.5 mm.
Figure 6 shows thermal images of the same sample
taken before the treatment, at the beginning of tissue displacement, and at the end of the treatment showing the rise
in thermal profile with time and onset of contraction. For
fascial tissue, contraction started when the maximal
Aesth Plast Surg (2011) 35:87–95
91
Fig. 5 Adipose-septal tissue behavior during RF energy delivery at different time points
Fig. 6 Adipose-fibrous septal tissue thermal behavior during RF energy delivery at different time points
adipose tissue temperature near the active internal electrode reached 69.4"C. Adipose fibrous septal tissue coagulation and vaporization started when tissue temperature
reached 90-100"C and is most probably associated with
boiling of adipocyte water content.
Fascia contraction is demonstrated in Fig. 7. The displacement of the markers and tissue contraction in fascia
were significantly less than in adipose tissue. The average
movement was 2.75 mm or approximately 2.5 times less
than the mark migration and tissue contraction observed in
adipose tissue. The marker migration and medial contraction started after 3.5 s and maximal temperature near the
active electrode at this moment was 61.5"C.
Skin behavior is presented in Fig. 8. The migration of
markers and medial displacement and tissue contraction on
the skin were similar to the fascia. The average movement
was 2.0 mm or approximately 3 times less than the marker
migration and contraction observed in adipose tissue. The
medial marker movement started after 2.5 s and the maximal temperature near the active electrode during this
contraction was 81.9"C.
Table 1 summarizes the results on subcutaneous tissue
contraction. From the results one can see that the
strongest contraction response was observed in adipose
tissue containing septal connective tissue and reticular
collagen fibers encasing fat cells. The contraction temperature threshold was the highest for dermis. It is clear
that the immediate contraction of dermal collagen is not
possible to achieve without a skin burn, which happened
when the epidermal temperature exceeded 45"C [13].
Fascia and septa can be heated to these high, optimal
contraction temperatures, but it can be done only in a
minimally invasive transcutaneous manner that deposits
the thermal RF energy directly into the adipose tissue
123
92
Aesth Plast Surg (2011) 35:87–95
Fig. 7 Fascia contraction behavior during RF energy delivery at different time points
Fig. 8 Skin contraction
behavior during RF energy
delivery at different time points
and subdermal space, thus avoiding heating the epidermal surfaces.
The contraction temperatures of collagen in our ex vivo
study were in the same range reported for other collagenous tissues. We observed tissue contraction in the area
with a diameter of 2 cm, which corresponds to a spherical
contraction volume of 4.2 cm3. Knowing the tissue volume
and deposited energy before the start of contraction, we can
estimate the energy density required for each cubic centimeter of treated tissue to reach tissue contraction effects.
We can calculate that for 1 L of adipose tissue up to
48.3 kJ is required to start to see immediate and significant
collagen contraction. These calculations of tissue energy
needed to initiate adipose contraction are consistent with
empirical data obtained with LAL treatment where energy
from 50 up to 100 kJ is recommended for treatment of the
abdominal area.
In vivo clinical monitoring of temperature in the adipose
tissue and on the epidermal surface should allow the physician to predict more accurately the thermal treatment
times and reduce the risk of thermal injuries.
123
A
Table 1 Average displacement and contraction threshold
Dermis Fascia Septa/Adipose
tissue
Average displacement (mm)
2
2.75
6.5
Threshold temperature ("C)
81.9
61.5
69.4
Time before start of contraction (s) 2.0
2.9
2.1
Delivered energy before start of
contraction (J)
203
147
140
In Vivo Clinical RFAL Results
The skin biopsies taken from an RFAL study patient at
12 months show normal dermal architecture with healthy
collagen (Fig. 9) and elastin fibers (Fig. 10) in the deep
reticular dermis and no evidence of scar tissue or abnormal
collagen fibers. All RFAL patients demonstrated some
level of contraction. From 8 to 15% linear tightening was
observed at the end of the surgery on the operating table. It
then increased dramatically during the first week when
most of the swelling was reduced. The linear and area
Aesth Plast Surg (2011) 35:87–95
93
Fig. 11 Correlation between aspirated volume and linear contraction
Fig. 9 Normal skin histology 12 months following optimal RFAL
thermal end point
Fig. 12 Correlation between BMI and linear contraction
Fig. 10 Same RFAL patient with 43% contraction and normal elastic
fiber content
contraction process continued for weeks and maximum
contraction was noted at the last follow-up visit 24 weeks
after the treatment.
Linear contraction observed at 6 months follow-up was
much more significant than reported with any other technology and varied from 12.7 up to 47% depending on
patient and treatment variables. It is important to note that
soft tissue area contraction can be calculated as the square
of the linear contraction and represents much higher
numbers. The measured linear contraction was then correlated with three parameters: (1) aspirated volume that
ranged from 0.5 to 3.4 L, with an average volume of 2.0 L,
(2) BMI that varied from 20.8 to 31.7, with an average
index of 25.7, and (3) deposed RF energy that varied from
60 to 96 kJ per abdominal area, with an average RF energy
of 72 kJ.
For statistical analysis of the correlation between the
measured variables and linear contraction, the Pearson
product moment correlation coefficients were calculated.
The closer the coefficient is to 1, the higher the linear
correlation between the measured variable and tissue contraction. Analysis shows no or very weak correlation
between aspirated volume and linear skin contraction. The
Pearson coefficient is about 0.22. Figure 11 shows the
correlation between these values and has a random distribution. The Pearson coefficient for correlation between
contraction and patient BMI is much higher and equal to
0.64. Figure 12 demonstrates a much stronger connection
between these parameters and it is easy to understand that a
patient with a larger volume of adipose tissue would have
more tissue available to undergo contraction.
123
94
Aesth Plast Surg (2011) 35:87–95
During this study we had one case of a seroma that was
treated with closed serial aspiration. Seroma is not a rare
side effect for energy-assisted liposuction, especially for
high-volume treatment and may necessitate a lower
threshold for closed drainage systems in selected patients.
Conclusions
Fig. 13 Correlation between total energy and linear contraction
The highest correlation (0.86) was obtained between
deposed RF energy and skin contraction. Figure 13 shows
measurement results that have an almost linear function
between these two parameters. The more energy deposited,
the more linear contraction that was observed. In spite of
improved contraction obtained at higher energies, the amount
of energy used during treatment can and should be measured
and controlled to avoid side effects such as seroma and skin
burn and still achieve optimal linear and area contraction.
Features of an ideal liposuction procedure would include
reduced ecchymosis, pain, and edema from preaspiration
coagulation of adipose and vascular tissue, followed by less
forceful and traumatic extraction forces, as well as significant soft tissue contraction when host tissue elasticity is
compromised. Thermal-based lipoplasty appears to hold
this potential.
In the present study based on volumetric heating, we
reached an average local linear contraction of 31% that is
statistically significantly higher than that reported with
other energy-emitting liposuction technologies. Overall
area contraction was much higher than linear contraction.
We believe that these in vivo results confirm our proposed
mechanism of RF-based tissue tightening and recruitment
of the vertical and oblique fibrous adipose matrix. Our
biopsy at 7 months suggests that the papillary and reticular
dermis is populated with normal collagen and elastin
that have been stimulated and remodeled by subnecrotic
subdermal RFAL temperatures.
About 30% of patients noted minor weight loss but it is
premature to correlate it with the treatment procedure.
The in vitro experiments produced different degrees of
contraction for septal and dermal tissues which emphasizes
the balance between these processes for optimal aesthetic
results. Lower two-dimensional contraction of the skin and
significant three-dimensional contraction of subdermal
adipose connective tissue may cause wrinkling of the skin
surface in high-volume liposuction patients.
123
We believe the study results confirm the hypothesis of
Kenkel [17], i.e., skin tightening and elasticity changes
following thermal lipoplasty are mostly a result of subdermal tissue contraction but not dermal, which experiences lower heating during the treatment. It is clear that
40–42"C on the skin surface cannot result in an immediate
contraction effect. Deep dermal remodeling may account
for some horizontal contraction over time. It is possible
that the dermal-fat junction experiences higher temperatures, but this process requires future investigation. We
believe that the mechanism of subcutaneous collagen
contraction during RF-assisted liposuction is similar to that
witnessed in other types of collagen in that the contraction
process has thermal contraction thresholds in the range of
60–70"C.
It is likely more accurate to talk about tissue contraction
rather than skin tightening because significant area contraction is a result of the strong contribution of deeper
adipose fascial layers. Further studies with accurate 3D
area measurements will tell us more about the RF-mediated
area contraction in this RFAL technology. This RFAL
thermal process and contraction can be effectively applied
during a liposuction treatment in selected cases, improving
patient satisfaction and extending liposuction procedures to
higher-weight patients and patients with compromised skin
conditions.
Disclosures Dr. Paul serves as consultant to and chairman of the
board of the medical advisory board for Invasix, Ltd., and received
consultation fees and stock options. He also serves as consultant to
and chairman of the scientific advisory board for Angiotech/Surgical
Specialties and receives consultant fees. Dr. Mulholland received
consulting fees and technology from Invasix.
Open Access This article is distributed under the terms of the
Creative Commons Attribution Noncommercial License which permits any noncommercial use, distribution, and reproduction in any
medium, provided the original author(s) and source are credited.
References
1. Asbell PA, Maloney RK, Davidorf J, Hersh P, McDonald M,
Manche E (2001) Conductive keratoplasty for the correction of
hyperopia. Trans Am Ophthalmol Soc 99:79–87
2. Obrzut SL, Hecht P, Hayashi K, Fanton GS, Thabit G III, Markel
MD (1998) The effect of radiofrequency on the length and
A
Lasers in Surgery and Medicine 20:164–171 (1997)
Effect of Nonablative Laser Energy on
the Joint Capsule: An In Vivo Rabbit Study
Using a Holmium:YAG Laser
Kei Hayashi, DVM, MS,1 Janet A. Nieckarz, BS,1 George Thabit III, MD,2
John J. Bogdanske, BA,1 A.J. Cooley, DVM,1 and Mark D. Markel, DVM, PhD1*
1
Comparative Orthopaedic Research Laboratory, School of Veterinary Medicine,
University of Wisconsin, Madison 53706
2
Sports, Orthopedic and Rehabilitation Medicine Associates, Menlo Park,
California 94025
Background and Objective: The nonablative application of holmium:yttrium-aluminum-garnet (Ho:YAG) laser energy to the
joint capsule of patients with glenohumeral instability has been
found to shrink capsular tissue and to help stabilize the joint. The
purpose of this study was to evaluate the effect of nonablative
laser energy on the short-term histological properties of joint
capsular tissue in an in vivo rabbit model.
Study Design/Materials and Methods: Eighteen mature New
Zealand white rabbits were used in this study. One randomly
selected stifle was treated with laser energy, and the contralateral stifle was sham-operated. Animals were euthanized immediately after surgery (day 0), at 7 days postsurgery and 30 days
postsurgery. Specimens were processed for histology and transmission electron microscopy.
Results: Laser-treated samples at day 0 showed diffuse hyalinization of collagen with nuclear karyorrhexis of fibroblasts. Lasertreated tissue at 7 days postsurgery revealed fibroblast proliferation around and into acellular hyalinized regions of collagen. At
30 days postlaser treatment, areas of fused collagen were greatly
reduced as large reactive fibroblasts migrated and secreted matrix.
Conclusion: This study illustrates the short-term in vivo tissue
response to nonablative laser treatment, where acellular hyalinized regions of collagen are infiltrated by fibroblasts that have
used the treated collagen as the framework for migration and
secretion of new collagen matrix in order for tissue repair to
proceed. Lasers Surg. Med. 20:164–171, 1997.
© 1997 Wiley-Liss, Inc.
Key words: collagen; fibroblast; histology; tissue response; transmission electron
microscopy
INTRODUCTION
A recent pilot study has demonstrated that
the nonablative application of the holmium:yttrium-aluminum-garnet (Ho:YAG) laser energy to
the joint capsule of patients with glenohumeral
instability shrank the joint capsule, stabilizing the
shoulder in the majority of the patients treated [1].
Glenohumeral instability secondary to ligamentous laxity, capsular redundancy, and excessive
© 1997 Wiley-Liss, Inc.
Contract grant sponsor: Department of the Navy-ONR; Contract grant number: N000014-90-C-0029; Contract grant
sponsor: NIH Lamp Resource; Contract grant number:
RR001192; Contract grant sponsor: NASA; Contract grant
number: NAG-2568; Contract grant sponsors: Coherent,
Beckman Laser Institute and Medical Clinic and, Oratech.
*Correspondence to: Mark D. Markel, DVM, Comparative Orthopaedic Research Laboratory, Department of Medical Sciences, School of Veterinary Medicine, University of Wisconsin, 2015 Linden Drive West, Madison, WI 53706.
Accepted for publication 21 February 1996.
Nonablative Laser Energy on Joint Capsule
joint volume is a frequent occurrence [2–5] that
current closed, open, and arthroscopic treatments
do not address satisfactorily in certain subgroups
[4, 6–10]. In a multi-institutional clinical trial [1],
nonablative application of the Ho:YAG laser,
which has been approved for arthroscopic surgery,
was applied to patients with glenohumeral instability without capsulolabral detachment or fullthickness rotator cuff tears. Laser energy was applied tangentially with the unit set at 10 watts (1.0
J, 10 pulses/sec) to shrink the capsuloligamentous
tissues of the glenohumeral joint without ablation.
For all patients, regardless of arm dominance, age,
sex, or direction of instability, postsurgical subjective scores were significantly higher than presurgical scores. Although this study suffered from
lack of a comparable nonoperated control population or an operated open surgical repair group,
these results indicate that at this short-term follow-up (mean 6 months), patients in this subgroup
improved dramatically after nonablative reduction of redundant glenohumeral joint capsule using the Ho:YAG laser.
We previously reported that Ho:YAG laser
energy at nonablative levels can significantly alter joint capsular length and its mechanical and
histological properties in an in vitro rabbit study
[11]. Laser treatment significantly shortened the
tissue by 9% (5 watts: 0.5 J/10 Pulses per sec),
26% (10 watts: 1.0 J/10 pulses per sec), and 38%
(15 watts: 1.5 J/10 pulses per sec), respectively.
Histological analysis of the tissue revealed significant thermal alteration of collagen and fibroblasts in the laser treatment groups, with each
subsequently higher laser energy causing significantly greater morphologic change over a larger
area. [12]
These results suggested that the predominant effect of nonablative laser energy on joint
capsular tissue is thermally mediated. Thermal
damage can cause denaturation of collagen and
necrotic changes of fibroblasts. Although a pilot
clinical study suggested the effectiveness of laser
treatment in patients with glenohumeral instability, to date no studies have been performed examining the histological and ultrastructural
properties of joint capsular tissue following the
application of laser energy. The purpose of this
study was to evaluate the effect of tissue response
on the laser-induced alterations of joint capsular
tissue. Specifically, we evaluated the effect of
nonablative Ho:YAG laser energy on the shortterm histological and ultrastructural properties
of joint capsular tissue in an in vivo rabbit model.
165
MATERIALS AND METHODS
Eighteen mature New Zealand white rabbits, ranging in weight from 4.3–6.5 kg (4.7 ±
0.52; mean ± SD), were used in this study. This
study was approved by the Institutional Animal
Use and Care Committee. Rabbits were randomly
assigned to one of three groups (0, 7, and 30 days
postsurgery). The animals were anesthetized with
halothane and oxygen, and both stifles of each
rabbit were aseptically prepared for surgery. The
femoropatellar joint was exposed via a patellar
tenotomy. One randomly selected stifle was
treated with laser energy using a Ho:YAG laser
(VersaPulse, Coherent, Palo Alto, CA) and a 1.7
mm hand piece (InfraTome, Coherent, Palo Alto,
CA), and the contralateral stifle was sham-operated. A custom-designed jig that allowed delivery
of the laser energy in a lactated Ringer’s solution
bath was used. Laser energy (5 watts: 0.5J per
pulse / 10 pulses per second) was applied to the
medial and lateral compartments of the femoropatellar joint capsule in a defocused manner. The
laser handpiece was held ∼1.5 mm from the synovial surface by a custom-designed jig and
moved over the tissue in a paintbrushlike motion.
Following the procedure, the joint capsule, subcutaneous tissue, and skin were closed routinely.
Animals were euthanized at three time intervals: immediately after surgery (day 0), 7 days
postsurgery, and 30 days postsurgery. The medial
and lateral portions of the femoropatellar joint
capsule were harvested immediately after euthanasia. Specimens were processed for histology and
transmission electron microscopy. Tissue samples
for histology were fixed in neutral-buffered 10%
formalin, embedded, sectioned on the plane perpendicular to the synovial surface of the specimen, and processed for histological staining with
hematoxylin-eosin. Tissue samples for transmission electron microscopy were fixed in modified
Karnovsky’s solution (2% paraformaldehyde and
1.25% glutaraldehyde in 0.1 M sodium phosphate
buffer, pH 7.0), stored in 0.1 M sodium phosphate
buffer for 8 hr at 4°C, postfixed for 2 hr in 1%
osmium tetroxide, and stained with 1% uranyl acetate. After sequential dehydration in ethanol
and infiltration in epon-araldite and propylene
oxide, specimens were embedded in 100% eponaraldite and polymerized at 60°C. Thick (1 �m)
and ultrathin (70 nm) sections were cut for light
and electron microscopy, respectively. The ultrathin sections were placed on grids, stained
with lead citrate and viewed using a transmission
electron microscope.
166
Hayashi et al.
RESULTS
Histology
Control tissues obtained from animals euthanized immediately after sham operations
showed no significant histological lesions in the
joint capsule tissue (Fig. 1a). Laser-treated samples at day 0 showed significant histological alterations with diffuse hyalinization and fusion of
collagen fibers along with nuclear karyorrhexis
and nuclear streaming of fibroblasts throughout
the treated regions (Fig. 1b). Control tissues at 7
days postsham operation showed granulation tissue, mild fibrosis, and mixed inflammatory infiltration including lymphocytes, plasma cells, and
heterophils (Fig. 1c). Laser-treated tissue at 7
days postsurgery revealed a similar inflammatory response to control tissues along with fibroblast proliferation around and into multifocal
acellular hyalinized collagen regions (Fig. 1d).
Normal fibrous collagen was present in the regions adjacent to the acellular treated region with
increased numbers of large rounded fibroblasts
(Fig. 1d). Control tissue at 30 days postsham operation showed mature granulation tissue and
regular dense fibrous connective tissue in the normal collagenous joint capsule tissue (Fig. 1e). Laser-treated tissues at 30 days postsurgery showed
fibrosis with cellular and disorganized connective
tissue. Fused collagen regions were greatly reduced by 30 days postlaser treatment as large fibroblasts migrated to the site and secreted new
collagen matrix to replace the hyalinized tissue
(Fig. 1f). For both laser-treated and sham-operated groups at 7 and 30 days postsurgery, there
were variations in the degree of inflammatory reaction within groups, although the responses to
the laser treatments were similar within lasertreated groups.
Electron Microscopy
Transmission electron microscopy revealed
no significant ultrastructural alterations in collagen or fibroblast architecture in control tissues
obtained immediately postsham operations (Figs.
2, 3). The typical appearance of cross-sectional regions at day 0 revealed collagen fibrils of a variety of sizes with distinct margins (Fig. 2a). Longitudinal sections of control tissue collagen fibrils
showed normal periodical cross-striations and
normal quiescent spindle shaped fibroblasts with
large condensed nuclei and sparse cytoplasm with
no ultrastructural evidence of active secretion
(Fig. 2a). Tissue samples obtained immediately
after laser treatment revealed significant alter-
ations in collagenous and fibroblast ultrastructure (Figs. 2b, 3a). Cross-sectional regions showed
increases in fibril diameter with a loss of distinct
fibril margins and longitudinal sections revealed
increased fibril diameter with the loss of cross
striations (Fig. 2b). Fibroblasts in laser treated
areas were pyknotic with evidence of nuclear
karyorrhexis and nuclear streaming resulting
from disruption of the nuclear and cellular membrane (Figs. 2b, 3a).
Control tissues at 7 days postsham operation
showed no ultrastructural alterations in collagen
fibrils; however, some active fibroblasts with increased rough endoplasmic reticulum and secretory vesicles were noted. Tissue samples at 7 days
postlaser treatment indicated significant ultrastructural alterations of collagen and fibroblasts
relating to the tissue response and repair process
(Figs. 2c, 3b,c). Areas directly treated with laser
energy showed loss of cellularity and evidence of
cellular degeneration (Figs. 2c, 3b). In this acellular region, striated collagen fibrils and agglomerates of polymerized microfibrils were observed
(Figs. 2c, 3b). Metabolically active fibroblasts were
noted to be most predominate adjacent to these
treated acellular regions (Fig. 3c). Electron microscopy in this area showed increased active fibroblasts and surrounding small collagen fibrils.
Fibroblasts revealed an increase in nuclear and
cytoplasmic area with elaborate rough endoplasmic reticulum, polyribosomes, mitochondria, and
secretory vesicles.
Control samples at 30 days postsham operations revealed no significant collagen or fibroblast
ultrastructural changes. Samples obtained at 30
days postlaser treatment indicated that lasertreated regions had increased cellularity with enlarged fibroblasts with extensive cytoplasm that
showed an increased number of secretory vesicles
along the plasma membrane and an increased arrangement of rough endoplasmic reticulum, golgi
apparatus, and mitochondria (Figs. 2d, 3d). Electron microscopy of cross-sectional regions at the
treatment interface showed very small collagen
fibrils interspersed with larger diameter fibrils
and increased active fibroblast cellularity (Figs.
2d, 3d). Longitudinal sections of collagen revealed
both large collagen fibrils and finer fibrils with
striations and large active fibroblasts with increased cytoplasmic organelles (Fig. 2d).
DISCUSSION
This study illustrates the histological and ultrastructural alterations of the in vivo tissue re-
Nonablative Laser Energy on Joint Capsule
167
Fig. 1. Light micrograph of (a) control day 0 capsular tissue
demonstrating normal collagen and fibroblasts, (b) lasertreated day 0 capsular tissue demonstrating hyalinization of
collagen and karyorrhexis of fibroblasts, (c) control day 7 tissue demonstrating a mild inflammatory response, (d) lasertreated day 7 tissue demonstrating acellular treated region
and adjacent fibrous collagen and active fibroblasts, (e) control day 30 tissue demonstrating regular fibrous connective
tissue, and (f) laser-treated day 30 tissue demonstrating
greatly reduced hyalinized acellular regions with surrounding fibroblasts and fibrosis (hematoxylin-eosin stain, original
×50).
sponse and collagen repair process following nonablative laser treatment of joint capsular tissue.
The nonablative applications of laser energy have
been evaluated primarily in tissue welding and
thermokeratoplasty [13–16]. Rabau et al. [17]
evaluated the healing process of laser-welded in-
testinal anastomoses in a rat model as compared
with sutured anastomoses. The investigators reported that despite significantly lower DNA and
collagen concentrations at the 4th postoperative
day, collagen concentrations on the 7th and 10th
postoperative days were significantly higher in
168
Hayashi et al.
Fig. 2. Transmission electron microscopy of (a) control day 0
joint capsular tissue demonstrating a normal fibroblast and a
variety of sizes of distinct collagen fibrils (a1: cross section)
with characteristic periodical cross striations (a2: longitudinal section), (b) laser-treated day 0 joint capsular tissue demonstrating pyknotic fibroblasts with loss of membrane integrity and collagen fibrils with increased diameter and loss of
distinct edges (b1) and loss of longitudinal cross-striations
(b2), (c) laser-treated day 7 joint capsular tissue in the treated
area demonstrating loss of cellularity with evidence of cellular degradation (c1) and striated collagen fibrils with microfibrillar structures (c2), and (d) laser-treated day 30 joint
capsular tissue at the treatment interface demonstrating
small collagen fibrils interspersed with large fibrils around
an active fibroblast (d1, d2) (×24,000, bar � 1 �m).
Nonablative Laser Energy on Joint Capsule
169
Fig. 3. Transmission electron microscopy of (a) laser-treated
day 0 joint capsular tissue demonstrating a pyknotic fibroblast and collagen fibrils with loss of striations, (b) lasertreated day 7 joint capsular tissue in the treated area demonstrating striated collagen fibrils and loss of cellularity,
(c) laser-treated day 7 joint capsular tissue in the area adjacent to the treated area demonstrating increased numbers of
active fibroblasts, and (d) laser-treated day 30 joint capsular
tissue demonstrating large and small collagen fibrils around
active fibroblasts (×6,900, bar � 1 �m).
the laser treated group than in a sutured group.
They attributed this result to less inflammatory
reaction following more rapid fibroblast proliferation in the laser treated group. To date, in vivo
studies have not been performed examining the
effect of tissue healing on the histological and ultrastructural properties of joint capsular tissue
following the application of nonablative laser energy.
Histology of laser-treated samples at day 0
revealed significant fusion and hyalinization of
collagen caused by thermal damage of the laser
treatment. At 7 days postlaser treatment, acellular hyalinized regions of collagen were infiltrated
with large rounded fibroblasts. Fibrosis with cellular randomly arranged connective tissue continued to replace the treated tissue at 30 days post-
laser application, which significantly reduced the
area of hyalinized collagen as seen via light microscopy. Both control and laser-treated samples
showed some level of inflammatory infiltration,
fibrosis, and granulation tissue invasion postsurgically, indicating that the sham operation probably resulted in a mild inflammatory response.
Transmission electron microscopy revealed collagen and fibroblast alterations that further support
the histological tissue response. As noted in previous studies [18], the most significant change in
collagen observed at day 0 postlaser treatment was
the disruption of regular fibril organization, which
was demonstrated as an increase in fibril diameter
and the loss of the fibril’s distinct edge on crosssection and the loss of periodical cross-striations
on longitudinal section. Although histologically
170
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Hayashi et al.
the collagen bundles appeared fused with light
microscopy, electron microscopy revealed that individual circular collagen fibrils were still evident.
We hypothesized that these changes in collagen
are mainly due to denaturation of collagen caused
by the thermal effect of laser energy. Heat-induced
shrinkage associated with denaturation of collagen is a well-described phenomenon [13–16, 19–
23]. At shrinkage temperatures, thermal unwinding of the triple helices outweighs the constraints
of natural crosslinks, causing the fibrils to denature and shrink [19]. Fibroblastic morphology was
altered in the laser-treated sites with both pyknotic changes and loss of cytoplasmic and nuclear
membrane integrity evident. These changes were
most likely caused by the thermal and/or mechanical effects of laser energy.
The synthesis, accumulation, and degradation of collagen are dynamic processes that occur
intracellularly and extracellularly during morphogenesis, growth, inflammation, and repair
[24–27]. In this study, evidence of active tissue
healing was observed at 7 days and 30 days postlaser treatment. At the interface of the treated
regions and normal tissue, increased numbers of
actively secreting fibroblasts were present, which
was established by an increase of cytoplasmic organelles, including rough endoplasmic reticulum,
mitochondria, golgi, and secretory vesicles. Fine
collagen fibrils adjacent to laser-altered fibrils
may provide evidence of newly secreted collagen
matrix and tissue repair. It appears that reactive
fibroblasts migrate into the treated regions using
the larger denatured collagen fibrils as a scaffold
in order to initiate collagen repair. Over time in
laser-treated sites, pyknotic nuclei fragmented
and degraded to form acellular regions. No macrophages or phagocytic cells were evident within
the treated region; however, active fibroblasts
were significantly increased adjacent to and infiltrating these areas.
In this study, laser energy (5 watts: 0.5J per
pulse / 10 pulses per sec) was applied to the medial and lateral compartments of the femoropatellar joint capsule in a defocused manner in a paintbrushlike motion using a custom-designed jig
that allowed delivery of the laser energy in a lactated Ringer’s solution bath ∼1.5 mm away from
the synovial surface. The distance from the tip of
the handpiece to the tissue was relatively well
controlled; however, in addition to the distance,
other factors such as angle of the beam, spot size,
and intervening solution temperature and thermal conductivity would also play important roles.
Clinically, it would be much more difficult to deliver laser energy at nonablative levels without
inadvertently overheating some regions and underheating other regions. Although the histological alterations of the tissue by laser application
were similar within laser treatment groups, further improvements in the method of energy delivery may be necessary.
This study illustrates the short-term, in vivo
tissue response to laser-treated joint capsular tissue in the rabbit model. Histological and ultrastructural examination revealed alterations in
both collagen and fibroblast morphology in lasertreated regions as demonstrated by hyalinized
collagen with pyknotic cells and nuclear streaming, and indistinct enlarged collagen fibrils with
loss of cross striations. At 7 days postlaser treatment, large acellular region of hyalinized collagen with surrounding large fibroblasts was the
predominant feature at the treated site. Histology
revealed a significant reduction in the area of hyalinized collagen at 30 days postlaser treatment
with increased fibroblast proliferation and fibrosis. Electron microscopy supported this histological finding in which metabolically active fibroB
blasts with increased cytoplasmic area, including
rough endoplasmic reticulum, golgi apparatus,
mitochondria and secretory vesicles, were evident. Small collagen fibrils were also significantly increased and interspersed with larger collagen fibrils in previously treated areas. This
study demonstrated that active healing is ongoing with a residual population of fibroblasts at the
end of this experimental period. This finding supports the concept that new collagen is actively
synthesized around treated collagen fibrils, although this study did not clarify whether the denatured collagen is entirely degraded or some areas of altered collagen may return to their
original fibril organization and function. Further
long-term in vivo studies are needed to evaluate
the condition of collagen and fibroblasts and the
synthetic activity of fibroblasts on collagen after
thermal treatment by nonablative laser energy.
This work was supported by grants from
Coherent, the Department of the Navy-ONR
(N000014-90-C-0029), NIH LAMP Resource RR
001192, NASA NAG-2568, the Beckman Laser
Institute and Medical Clinic, and Oratech.
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Lasers in Surgery and Medicine 41:1–9 (2009)
Bipolar Fractional Radiofrequency Treatment Induces
Neoelastogenesis and Neocollagenesis
Basil M. Hantash, MD, PhD,1,2,3 Anan Abu Ubeid, BS,3 Hong Chang,
and Bradley Renton, PhD2*
1
Stanford University School of Medicine, Stanford, California
2
Primaeva Medical, Inc., Pleasanton, California
3
Elixir Institute of Regenerative Medicine, San Jose, California
A
Background: We recently introduced RenesisTM, a
novel minimally invasive radiofrequency (RF) device, for
the treatment of human skin. The wound healing response
post-fractional RF (FRFTM) treatment was examined in
human subjects.
Study Design: The FRF system delivered RF energy
directly within the dermis via 5 micro-needle electrode
pairs. Tissue temperature was held at 728C for 4 seconds
using an intelligent feedback system. The wound healing
response was evaluated histologically and by RT-PCR up
to 10 weeks post-RF treatment. Neoelastogenesis and
the role of heat shock proteins (HSPs) were assessed by
immunohistochemistry.
Results: FRF treatment generated a RF thermal zone
(RFTZ) pattern in the reticular dermis that consisted of
zones of denatured collagen separated by zones of spared
dermis. RFTZs were observed through day 28 post-treatment but were replaced by new dermal tissue by 10 weeks.
HSP72 expression rapidly diminished after day 2 while
HSP47 expression increased progressively through
10 weeks. Reticular dermal volume, cellularity, hyaluronic
acid, and elastin content increased. RT-PCR studies
revealed an immediate increase in IL-1b, TNF-a, and
MMP-13 while MMP-1, HSP72, HSP47, and TGF-b levels
increased by 2 days. We also observed a marked induction of
tropoelastin, fibrillin, as well as procollagens 1 and 3 by
28 days post-treatment.
Conclusion: Our study revealed a vigorous wound healing
response is initiated post-treatment, with progressive
increase in inflammatory cell infiltration from day 2
through 10 weeks. An active dermal remodeling process
driven by the collagen chaperone HSP47 led to complete
replacement of RFTZs with new collagen by 10 weeks posttreatment. Furthermore, using both immunohistochemical
and PCR studies, we successfully demonstrated for the
first time evidence of profound neoelastogenesis following
RF treatment of human skin. The combination of neoelastogenesis and neocollagenesis induced by treatment
with the FRF system may provide a reliable treatment
option for skin laxity and/or rhytids. Lasers Surg. Med.
41:1–9, 2009. ! 2008 Wiley-Liss, Inc.
Key words: bipolar; fractional; micro-needle electrodes;
minimally invasive; neocollagenesis; neoelastogenesis;
radiofrequency thermal zones; renesis; wound healing
! 2008 Wiley-Liss, Inc.
PhD,
3
Reza Kafi,
MD,
1
INTRODUCTION
Energy based devices have enjoyed increasing popularity
for the treatment of a variety of skin conditions over the
last several decades. This has been driven in part by
demographic changes resulting in increased demand for
aesthetic related procedures to address the effects of
intrinsic aging, excessive sun exposure, and a myriad of
other factors that contribute to unwanted skin laxity and
an accelerated appearance of rhytids. To reduce the appearance of wrinkles, physicians have turned to a number of
treatment options varying in degree of invasiveness and
side effect profile. These include treatment with topical
retinoids, chemical peels, microdermabrasion, noninvasive
and invasive energy based devices, and finally surgical
reconstruction [1–7].
Although a significant number of clinical studies have
investigated the efficacy of each approach for the treatment
of facial rhytids, very little is known about the molecular
events that lead to improvements in the appearance of
wrinkles. Since ablative resurfacing devices (e.g., carbon
dioxide laser, erbium/yttrium-aluminum-garnet laser)
remain the non-surgical gold standard for facial rejuvenation [8–11], a significant effort to better understand the
wound healing response has recently been undertaken.
A number of authors have speculated that coagulation of
dermal collagen may underlie the observed efficacy following treatment with energy based devices [12,13]. Since
then, work by Fisher’s group has shed further light on the
molecular mechanisms that ensue post-ablative resurfacing, implicating an orchestrated series of dynamic changes
beginning with upregulation of metalloproteinase and
Abbreviations: EVG, elastic-Van Gieson; FRF, fractional radiofrequency; H&E, hematoxylin & eosin; HSP, heat shock protein;
IFS, intelligent feedback system; LDH, lactate dehyrodogenase;
PPH, precision planar heating; RF, radiofrequency; RFTZ, radiofrequency thermal zones; RT-PCR, reverse transcriptase polymerase chain reaction.
Dr. Hantash serves on the scientific advisory board of Primaeva
Medical, Inc., and Dr. Renton an employee of Primaeva Medical, Inc.
Contract grant sponsor: Primaeva Medical, Inc.
*Correspondence to: Bradley Renton, PhD, Primaeva Medical,
Inc., 4160 Hacienda Drive, Suite 100, Pleasanton, CA 94588.
E-mail: [email protected]
Accepted 4 November 2008
Published online in Wiley InterScience
(www.interscience.wiley.com).
DOI 10.1002/lsm.20731
2
HANTASH ET AL.
collagenase activity [14]. These enzymes initiate the
dermal remodeling process, helping remove photoaged
dermal tissue and allowing for deposition of new dermal
tissue. Moreover, Fisher and co-workers have recently
extended their original findings to microdermabrasion and
topical retinoid treatments [15–17].
Unfortunately, interest in ablative resurfacing has since
waned, primarily due to the high incidence of side effects
such as prolonged erythema and edema, hyperpigmentation, permanent hyopigmentation, scarring, and infection
[18–20]. These issues combined with the marginal efficacy
of topical retinoids, chemical peels, and microdermabrasion
has led to a search for a novel approach with an improved
side effect profile. Indeed, this has led to the introduction of
fractional photothermolysis, a novel concept that exploits
the healing power of the healthy tissue reserve surrounding
zones of treatment [21,22]. Recent published studies by
Manstein and Hantash demonstrate the importance of
untreated zones to the dermal remodeling process for both
nonablative and ablative infrared devices [23,24]. Although
Zelickson and co-workers did report on the ultrastructural
changes in collagen following monopolar RF treatment of
bovine tendons and human skin [25,26], there remains a
conspicuous absence of published studies characterizing
the wound healing response of skin at the molecular level
post-RF treatment.
We recently introduced a novel bipolar micro-needle RF
device for the treatment of skin [27]. In that seminal study,
we demonstrated the ability to create dermal lesions known
as radiofrequency thermal zones (RFTZTM). By varying
pulse length, we showed that lesion size was tunable.
Through the use of our proprietary real-time intelligent
feedback system (IFSTM), we described for the first time
fractional sparing of dermal tissue using a RF system and
coined this novel dermatologic treatment approach, fractional radiofrequency (FRFTM). To extend this work, we
examined the wound healing response following FRF
treatment in human subjects using histological, immunohistochemical, and molecular techniques. We found that
FRF treatment induced a dramatic wound healing response characterized by increased expression of heat shock
proteins (HSPs) and inflammatory mediators coupled with
dynamic remodeling of collagen and elastin. This is the first
report characterizing the wound healing response of FRF
treatment. Our findings further support the importance of
fractional sparing in achieving a balance between efficacy
and side effects for the treatment of facial rhytids.
MATERIALS AND METHODS
Study Design
This prospective clinical study was conducted at three
independent sites using a protocol that was approved by an
institutional review board. All subjects were consented
prior to participation in the study. Patient consent for
digital photography was also obtained prior to treatment.
Twenty-two healthy subjects were enrolled in the study.
Inclusion criteria consisted of age 18 years or older and
undergoing elective surgical face lift or abdominoplasty.
Exclusion criteria consisted of history of injection with
silicone, fat, collagen, or a synthetic material placed in the
intended treatment area, bleeding disorder, hypertrophic
scar or keloid formation, isotretinoin treatment in prior
12 months, anaphylaxis, or lidocaine hypersensitivity.
Other exclusion criteria included prior, current, or anticipated treatment with anti-coagulants, thrombolytics,
chemotherapeutic(s), systemic corticosteroids, or anabolic
steroids. Patients with a compromised immune system,
impaired wound healing (e.g., diabetics, smokers), collagen
vascular disease, implantable electronic device, or active
infection were disqualified from participation. Only subjects available for longitudinal follow-up during the entire
study length were enrolled.
Treatment Regimen
The wound healing response following FRF treatment
was studied in human subjects. Patients were treated
immediately, 2 days, 14 days, 28 days, and 10 weeks prior to
their pre-scheduled abdominoplasty to capture the temporal evolution of the in vivo wound healing response.
Preceding each treatment, skin was cleansed using 70%
isopropyl alcohol, followed by wiping with 10% povidoneiodine topical anti-septic. Subjects were then infiltrated
with 1–2% lidocaine with or without 1:100,000 epinephrine. The FRF system was used to deliver bipolar RF energy
to the dermis via 5 micro-needle 30 gauge electrode pairs
6 mm in length, each spaced 1.25 mm apart. The angle of
micro-needle insertion into the skin was 208. During RF
energy application, the dermal tissue temperature within
the RFTZ was maintained at 728C for 4 seconds using IFS.
Superficial cooling to minimize epidermal damage was
achieved using a solid state Peltier device equipped with a
heat sink and fan maintained at 158C. Within 40 minutes
post-abdominoplasty, biopsies were taken for all time
points as well as for baseline control.
Histology Studies
The wound healing response was evaluated histologically to determine the thermal effects of RF treatment on
collagen, elastin, and inflammation. Immediately postexcision, three different 6 mm biopsies were performed for
each time point. The 1st biopsy was fixed in 10% v/v neutral
buffered formalin (VWR International, West Chester, PA)
overnight and then embedded in paraffin. The 2nd biopsy
of the set was embedded in optical cutting temperature
compound (Sakura Finetek, Torrance, CA) for frozen
sectioning. The 3rd biopsy was snap frozen in liquid
nitrogen and stored at !1908C in preparation for further
tissue processing. Ten micrometer paraffin sections
were sliced vertically and stained with hematoyxlin &
eosin (H&E), elastic-Van Gieson (EVG), or anti-human
antibodies to elastin, HSP72, and HSP47. Frozen sections
were sliced similarly but stained with lactate dehyrodogenase (LDH) to assess tissue viability post-RF treatment.
Semi-Quantitative RT-PCR Studies
Semi-quantitative reverse transcriptase polymerase
chain reaction RT-PCR was used to elucidate the molecular
FRACTIONAL RADIOFREQUENCY AND WOUND HEALING
changes involved in dermal remodeling following RF
treatment. Frozen punch biopsies measuring 6 mm in
diameter with an imputed length of 2.84 mm (based on the
mean weight and skin density of 1 g/ml) were thawed at
room temperature, quickly weighed (mean of 80.25 "
0.74 mg), then homogenized in 2 ml Trizol (Invitrogen,
Carlsbad, CA) and mixed with 400 ml chloroform. After
centrifugation, total RNA was extracted from the aqueous
phase using the RNeasy mini kit (Qiagen, Valencia, CA)
following the manufacturer’s protocol. Two milligram of
total RNA was used to synthesize cDNA using TaqMan
Reverse Transcription Reagents (Applied Biosystems,
Foster City, CA) with Oligo dT as primer in a 50 ml reaction.
RT reactions were annealed at 248C for 10 minutes,
followed by first-strand cDNA synthesis at 488C for 1 hour
and heat inactivation at 958C for 5 minutes. The resulting
cDNA was stored at !208C until assayed.
The PCR primer sequences and corresponding amplicon
sizes are shown in Table S1. The PCR primer pair for TGF-b
was designed using Primer3 software (Whitehead Institute, Cambridge, MA) while all remaining primer sequences were obtained from previously published articles. The
semi-quantitative RT-PCR reactions were performed on a
DNA Engine Peltier Thermo Cycler (Bio-Rad, Hercules,
CA). Briefly, DNA polymerase was activated at 948C for
2 minutes, followed by 35 cycles of denaturation at 948C
for 30 seconds, annealing at X8C for 30 seconds (where
X depended upon the sequences of PCR amplicons listed
in Table S1), extension at 728C for 30 seconds, and 1 cycle
of further extension at 728C for 10 minutes. For HSP47,
37 cycles of denaturation were used with each step
lasting 45 seconds instead of 30 seconds. The Taq PCR
master mix kit (Qiagen) was used to carry out all PCR
reactions.
For each PCR reaction, a total of 0.25 ml of cDNA per
20 ml total reaction volume was utilized; this concentration
was found to be within the predetermined linear range of
PCR amplification for all genes tested. Electrophoresis was
performed using 1.5% agarose gel containing 0.5 mg/ml
ethidium bromide and imaged using the FluorChem1 HD2
Imaging System (Alpha Innotech Corporation, San Leandro, CA). The densitometry analysis was performed using
AlphaEase FC software (Alpha Innotech Corporation) with
b-actin as an internal control. The amplicon intensity ratios
were calculated by dividing the intensity value for the gene
of interest by the intensity value for b-actin. The calculated
intensity ratios represent the minimum relative expression
value assuming expression was not confined to the RFTZ
but instead involved the entire 6 mm biopsy specimen for
each sample.
Statistical Analysis
All data represent a minimum of four independent
experiments. The means and standard errors of the
amplicon intensity ratios for each gene of interest were
calculated using Microsoft Excel and statistical significance was determined using a paired analysis of variance.
P values were taken to be statistically significant at
P<0.05.
3
RESULTS
Inflammatory Response
In order to characterize the wound healing response postFRF treatment, human subjects were treated immediately,
2 days, 14 days, 28 days, or 10 weeks prior to abdominoplasty. Skin was harvested, paraffin-embedded, sectioned,
and then stained with H&E. Immediately post-FRF treatment, RFTZs were apparent in the deep reticular dermis
(Fig. 1B). Similar to baseline untreated skin (Fig. 1A),
we found virtually no evidence of inflammatory cells at
this time point (Fig. 1B). Scant focal inflammation
surrounding the RFTZs was evident by day 2 post-FRF
treatment (Fig. 1C). Inflammation increased progressively
through day 14 (Fig. 1D). Infiltration of RFTZs by inflammatory cells was observed at day 28 post-FRF treatment
(Fig. 1E,F). In certain instances, tissue specimens were
frozen sectioned and stained with both H&E and LDH.
Consistent with our above findings, we did not see any
significant evidence of inflammation in untreated tissue
(Fig. 2A,C). At day 28, H&E stained sections revealed
generalization of the inflammatory response to include
areas overlying the RFTZs (Fig. 2B). Diffuse presence
of viable nucleated cells was more easily visualized in
LDH-stained sections (Fig. 2D).
Heat Shock Protein Response
Little is known about the HSP response post-FRF
treatment. We therefore stained sections with an antibody
against the early responding HSP72 and the procollagen
chaperone HSP47. Consistent with previous studies [24],
we observed baseline positive staining with anti-HSP72
antibody in all epithelial cells of the skin including
keratinocytes and sweat glands (Fig. 3A). HSP72 staining
was found in the dermis at 2 days post-FRF treatment (Fig.
3C), but not immediately post-treatment (Fig. 3B). No
residual staining could be detected at the site of thermal
injury or surrounding the RFTZ between days 14 and 28
post-FRF treatment (Fig. 3D,E). Similarly, HSP47 staining
could also be observed as early as 2 days post-FRF treatment (Fig. 4C) but not at baseline (Fig. 4A) nor immediately
post-treatment (Fig. 4B). In contrast to HSP72, however,
we observed progressively increasing HSP47 staining between days 14 and 28 (Fig. 4D,E), that remained persistent
at 10 weeks post-FRF treatment (Fig. 4F). Consistent
with the inflammatory cell response noted above, HSP47
staining extended diffusely throughout the dermis at
day 28 and 10 weeks.
Neocollagenesis and Neoelastogenesis
We next determined whether FRF treatment led to neocollagenesis. A significant inflammatory response remained
evident at 10 weeks post-FRF treatment (Fig. 5B). In
addition, a clearly demarcated zone of increased dermal
thickness over baseline (Fig. 5A) was observed at 10 weeks
post-treatment and could be attributed to neocollagenesis
(Fig. 5B). Examination at higher magnification revealed
evidence of newly deposited hyaluronic acid, as indicated by
the presence of a whispy blue-gray staining substance in
A
4
HANTASH ET AL.
Fig. 2. Frozen section images before and 28 days after
treatment with the Renesis system. Human abdominal skin
was processed similar to Figure 1 and then stained with H&E
or LDH at baseline (A,C, respectively) or 28 days posttreatment (B,D, respectively). An absence of blue staining
nuclei in the dermis was consistent with an absence of
inflammation at baseline. A RFTZ is shown in the reticular
dermis at 28 days post-treatment (B,D) and is indicate by the
white arrows. Similar to paraffin-embedded sections,
frozen sections revealed viable cells infiltrating RFTZs at day
28. A generalized inflammatory response throughout the
dermis was more easily appreciated in LDH-stained sections
(D). All images are shown at 4# the original magnification.
Fig. 1. FRFTM lesions post-treatment with a novel bipolar
radiofrequency device. Human abdominal skin was biopsied at
baseline (A), and immediately (B), 2 days (C), 14 days (D), or
28 days (E,F) following in vivo treatment with the FRF system.
Skin was processed according the Histology Studies Section.
Each RFTZ, indicated by the white arrows, represents an area
of coagulated elastin and collagen and is surrounded by viable
tissue. Inflammatory cells (blue staining nuclei) were detected
at the periphery of each RFTZ as early as 2 days post-treatment
(C). The intensity of this response progressively increased
through day 28 post-treatment (D,E). The RFTZ remained
prominent at day 28 (E), at which time inflammatory cells were
observed infiltrating within each RFTZ (F). All images are
H&E stained and shown at 4# the original magnification
except panel F which is at 10#.
the midst of a high density field of nucleated cells (Fig. 5C).
Three distinct zones were typically observed: an old
collagen zone with scant cellularity and hyaluronic acid, a
mixed transition zone with modest cellularity and hyaluronic acid deposition, and a new collagen zone with a high
cellular density and significant amount of de novo hyaluronic acid deposition (Fig. 5C). In addition, subcutaneous
interstium was also thickened without any evidence of fat
necrosis suggesting that both dermal and subcutaneous
collagen remodeled in response to FRF treatment.
To assess for neoelastogenesis, we performed an EVG
stain at baseline and 10 weeks post-FRF treatment. As
shown in Figure 6D, a significant increase in the amount of
elastin (dark brown to black stain) was evident when
compared to baseline skin that had not undergone FRF
treatment (Fig. 6B). The zone of neoelastogenesis coincided
with the zone of dermal remodeling demarcated by the
presence of hyaluronic acid and increased cellularity
(Fig. 6C). This conclusion was further supported by
immunohistochemical studies using a human anti-elastin
antibody. Positive immunostaining for elastin was detected
at 10 weeks post-FRF treatment (Fig. 7D), but not at
baseline (Fig. 7B). The elastin immunostaining again colocalized to the zone of dermal remodeling post-FRF
treatment (Fig. 7C). As expected, no such zone of dermal
remodeling was detected in any of the baseline untreated
A
B
FRACTIONAL RADIOFREQUENCY AND WOUND HEALING
5
Fig. 3. HSP72 response to FRF treatment. Human abdominal skin was processed similar to
Figure 1 and then stained with anti-human HSP72 antibody. Baseline (A) epidermal cells
stained positively for HSP72 [24]. Increased expression in the dermis was observed at day 2 (C)
but not immediately post-treatment (B), and returned to baseline by day 14 (D), remaining at
that level at day 28 (E) post-FRF treatment. All images are shown at 4# the original
magnification.
To better understand the sequence of molecular events
triggered by FRF treatment, we performed PCR at baseline
and various time points following FRF treatment. The
details for tissue harvesting are indicated in the material
and methods section. Since no previous studies had defined
the dermal remodeling response induced by RF or FRF skin
treatment at the gene expression level, we selected an
identical panel of genes as that previously investigated by
Orringer et al. [14] in their seminal study on the effects of
non-fractional CO2 skin resurfacing. However, to minimize
systematic errors, we utilized b-actin as an internal control
for each sample time point. We therefore present only
relative expression ratios (gene of interest/b-actin) as a
semi-quantitative output from our RT-PCR studies. Table 1
shows the means and standard errors at all time points
sampled for a total of 13 genes, each subclassified into 1 of
4 categories as follows: cytokines (TNF-a, IL-1b, TFG-b),
Fig. 4. HSP47 response to FRF treatment. Human abdominal
skin was processed similar to Figure 1 and then stained with
anti-human HSP47 antibody. At baseline (A) and immediately
(B) post-FRF treatment, there was minimal HSP47 expression
in the dermis. Increased HSP47 expression was first detected
at day 2 (C), but unlike HSP72, remained elevated from day 14
onward (D–F). At day 28 and 10 weeks post-FRF treatment,
HSP47 staining became diffuse throughout the dermis and was
not restricted only to the peri-RFTZ regions. All images are
shown at 4# the original magnification.
skin samples stained with H&E (Figs. 6A and 7A), EVG
(Fig. 6B), or anti-elastin antibody (Fig. 7B).
Molecular Events Underlying the Wound Healing
Response Post-FRF Treatment
6
HANTASH ET AL.
Fig. 5. Long-term dermal remodeling and neocollagenesis
post-FRF treatment. Human abdominal skin was processed
similar to Figure 1 and then stained with H&E. At 10 weeks
post-treatment (B), dermal thickness was increased over
baseline (A). Subcutaneous interstitial collagen was also
thickened with no evidence of fat necrosis. Both observations
can be attributed to dermal remodeling and ongoing neocollagenesis post-FRF treatment. Higher magnification revealed
the presence of a whispy blue-gray staining substance
(indicative of de novo hyaluronic acid deposition) in the midst
of a high density field of nucleated cells (C). Panel C also
illustrates the 3 distinct zones that were typically observed: an
old collagen zone, a mixed transition zone, and a new collagen
zone (see Results Section for details). Panels A&B at 4# and
C at 10# the original magnification.
metalloproteinases (MMP-1, MMP-3, MMP-9, MMP-13),
heat shock proteins (HSP72, HSP47), and extracellular
matrix proteins (fibrillin, tropoelastin, procollagen 1, procollagen 3). For untreated controls, none of the selected
genes was expressed at a level greater than that for b-actin
(all relative ratios are <1), suggesting that dermal
remodeling rates were low under basal conditions.
Immediately post-FRF treatment, IL-1b and TNF-a
expression increased by > 42% although the other cytokine
TGF-b1 remained unchanged. Expression of MMP-13 and
procollagen 1, but not of procollagen 3, increased by 47%
and 55% over baseline, respectively. By day 2, TGF-b1 and
MMP-9 expression were 60% and 70% greater than baseline. We observed a 65% increase in fibrillin and 188%
Fig. 6. Long-term dermal remodeling and neoelastogenesis
post-FRF treatment. Human abdominal skin was processed
similar to Figure 1 and then stained with H&E (A,C) or EVG
(B,D). A significant increase in dermal elastin content was
observed at 10 weeks post-FRF treatment (D) compared to
baseline (B). This increase co-localized to the region of dermal
remodeling (white arrows) as evidenced by the increase in
hyaluronic acid and cellularity observed post-FRF treatment
(C) relative to baseline (A). All images are shown at 10# the
original magnification.
Fig. 7. Immunohistochemical evidence of neoelastogenesis
post-FRF treatment. Human abdominal skin was processed
similar to Figure 1 and then stained with H&E (A,C) or
anti- human elastin antibody (B,D). Panel D shows that
elastin immunostaining (brown) in the reticular dermis
co-localized with the RFTZ shown by the black arrows in
panel C. At baseline (A), there is no evidence of an active
dermal remodeling zone and elastin immunostaining was
negative (B). All images are shown at 10# the original
magnification.
FRACTIONAL RADIOFREQUENCY AND WOUND HEALING
7
TABLE 1. Response to FRF Treatment of Various Wound Healing Genes Involved in Dermal Remodeling
Control
Gene
TNF-a
IL-1b
TGF-b1
MMP-1
MMP-3
MMP-9
MMP-13
HSP72
HSP47
Fibrillin
Tropoelastin
Procollagen 1
Procollagen 3
Immediate
2 days
14 days
28 days
Mean
SEM
Mean
SEM
Mean
SEM
Mean
SEM
Mean
SEM
0.62
0.24
0.41
0.33
0.40
0.68
0.35
0.83
0.42
0.90
0.34
0.78
0.91
0.16
0.04
0.08
0.10
0.06
0.20
0.06
0.27
0.12
0.24
0.01
0.23
0.11
0.89
0.34
0.41
0.32
0.37
0.76
0.52
0.95
0.47
0.99
0.40
1.21
1.00
0.14
0.07
0.11
0.05
0.05
0.12
0.08
0.35
0.09
0.28
0.03
0.27
0.42
0.91
0.36
0.65
0.43
0.37
1.15
0.67
1.37
1.03
1.48
0.98
1.23
0.98
0.17
0.02
0.12
0.09
0.06
0.23
0.13
0.29
0.07
0.08
0.05
0.31
0.19
0.73
0.38
0.69
0.47
0.43
1.46
0.76
1.22
1.51
1.56
1.49
1.48
1.11
0.18
0.08
0.07
0.06
0.08
0.25
0.18
0.23
0.02
0.25
0.03
0.37
0.28
1.01
0.51
0.76
0.59
0.65
2.00
0.99
1.44
1.87
1.76
1.67
2.36
1.94
0.10
0.10
0.10
0.07
0.07
0.17
0.15
0.10
0.07
0.11
0.03
0.31
0.28
Relative expression was calculated as the ratio of the expression level of the gene of interest/expression level of b-actin at each
particular time point. For each gene, the mean " standard error (SEM) for four independent samples is shown as ratio units of
relative expression.
increase in tropoelastin expression by 2 days posttreatment. In addition, both HSP47 and 72 expression
levels were increased, with the former at a striking 146%
and latter at 64% over baseline levels. Procollagen 3
expression did not show a significant increase until 28 days
post-treatment (112% over baseline). This was similar to
MMP-3, which was upregulated by 64% at 28 days posttreatment. Once elevated, cytokine expression remained
fairly stable over the ensuing 4 weeks. On the other hand,
MMPs appeared to increase progressively, although this
was less obvious for MMP-3 since its upregulation seemed
to lag behind that of the family members. There was no
significant difference between expression levels at days 2,
14, or 28 for HSP72. In contrast, HSP47 progressively
increased ending at $4.5-fold over baseline by day 28.
Extracellular matrix protein expression levels also seemed
to increase over time, with tropoelastin and procollagen 1
ending at $5-fold and 3-fold greater than baseline levels.
DISCUSSION
Previous studies have investigated the effects of microneedle based bipolar RF treatment on joint capsular tissue
[28], however, very little is known about its effects on
human skin. We recently introduced a novel fractional RF
device, known as Renesis, utilizing a minimally invasive
bipolar micro-needle delivery system for the treatment of
human skin and demonstrated the ability to produce
controlled zones of collagen coagulation in the reticular
dermis at the histological level [27].
Our seminal study suggested that both elastin and
collagen were being regenerated post-FRF. We also found
complete sparing of vasculature as well as adnexal
structures such as sebaceous glands, hair follicles, and
sweat glands. Furthermore, unlike its monopolar counterpart [29], we observed an absence of necrosis in the adipose
layer even though interstitial collagen thickening was
observed. These findings combined with the relative dearth
of literature examining the precise molecular events
triggered by bipolar RF treatment of human skin led us to
conduct the present study.
The skin wound healing response follows a very well
orchestrated set of events in humans and is separated into 3
distinct phases: inflammation, proliferation, and remodeling (reviewed in Ref. 30). Following injury, chemotactic
factors and vasoactive mediators are released leading to the
recruitment of neutrophils over the first 3 days followed by
monocytes through approximately day 7. By day 14 or so,
fibroblasts begin laying down a provisional matrix mainly
composed of collagen and proteoglycans. It is thought that
upon reaching a critical collagen density, fibroblast
proliferation and collagen synthesis are suppressed. While
this series of events takes place with predictable precision
for open wounds, it remains unclear whether this exact
model can explain the events that follow skin injury with
electromagnetic energy.
In our study, we observed a progressive increase in the
density of the inflammatory infiltrate in the region of each
RFTZ between baseline and 10 weeks post-FRF treatment
(Figs. 1 and 5). This increase mirrored that expected for an
open wound injury [28,30]. By day 28, fibroblasts were seen
within the actual RFTZ, suggesting that active dermal
remodeling had begun (Fig. 1F). Interestingly, the inflammatory response was not limited to the plane of dermal
injury as we found evidence of extension several millimeters above the RFTZ (Figs. 2B,D and 5B). These cells
were viable based on positive LDH staining of nuclei within
and above the RFTZ (Fig. 2D). Our findings are consistent
with those of the first author’s previous studies which
showed a generalization of the wound healing response
following treatment with a fractional CO2 device [24].
8
HANTASH ET AL.
HSP72 is known to respond within hours to days of
thermal injury [31]. HSP72 expression peaked by 2 days
and was not detectable at 14 days and beyond (Fig. 3).
Similarly, expression of HSP47, a wound healing regulator
known to function as a collagen chaperone [31], was first
detected at 2 days post-FRF treatment (Fig. 4). Unlike
HSP72, however, HSP47 expression continued to increase
between day 14 and 10 weeks. These data are consistent
with previous findings for fractional CO2 treatment as well
as the inflammatory response data presented herein.
However, direct evidence for neocollagenesis or neoelastogenesis was not reported in that previous study [24]. We
therefore pursued a series of experiments aimed at directly
addressing this important question.
At 10 weeks post-FRF treatment, histological studies
revealed evidence of new collagen deposition highlighted
by increased cellularity and hyaluronic deposition, both
observable by standard H&E staining (Fig. 5). Elastin
immunostaining and EVG studies revealed the presence
of increased elastin content at 10 weeks post-treatment
compared to baseline (Figs. 6 and 7), suggesting that that
FRF treatment induced neoelastogenesis. In fact, the
elastin immunostaining co-localized to the zone of dermal
remodeling observed by H&E (Fig. 7C,D). Moreover,
our PCR studies revealed a nearly fivefold increase of
tropoelastin over baseline by day 28 post-FRF treatment
(Table 1). To our knowledge, this is the first direct evidence
of significant new elastin production following RF
treatment.
This profound increase in neoelastogenesis and neocollagenesis appears to have been initiated by cytokines such as
TNF-a, IL-1b, and TGF-b, although involvement of other
cytokines cannot be ruled out. Degradation of collagen
requires two types of MMPs, a collagenase and a gelatinase.
Orringer et al. [14] showed that MMP-1, an enzyme
that catalyzes the first step of collagen degradation, was
significantly upregulated in the first week following
CO2 resurfacing. We, however, did not observe a similar
increase in MMP-1 expression as levels only increased to
less than twofold by day 28 post-treatment. This suggests
that FRF treatment induces a unique wound healing
response that does not require significant catabolic activity
during the inflammatory and proliferative phases.
MMP-3 breaks down partially degraded extracellular
matrix proteins including collagen, elastin, and proteoglycans. Orringer reported that MMP-3 expression mirrored
that of MMP-1 temporally [14]. Similar to our findings with
MMP-1, we observed a modest 60% increase of MMP-3
expression by day 28 (Table 1). Taken together with our
MMP-1 data, our studies suggest that pre-existing extracellular matrix proteins such as collagen and elastin were
not significantly degraded post-FRF treatment. In the
context of skin tightening and laxity treatment, this may
allow for increased overall dermal volume since degradation of pre-existing collagen and elastin is kept to
a minimum.
Our PCR studies did reveal a concerted uprgulation
of MMP-13 and the gelatinase MMP-9, both of which
increased progressively from immediately post-through
day 28 post-treatment. Our data differ from those previously reported which showed that peak expression of
MMP-9 and MMP-13 occurred 1 and 2 week(s) post-CO2
treatment, respectively [14]. Our findings for MMP-13 are
consistent with its role in long-term dermal remodeling,
which would require sustained activity of the enzyme
beyond 4 weeks. It is possible that any residual collagen
degradation fragments are processed by the modest and
delayed activity of MMP-3 in conjunction with that of
MMP-9; however, further studies are required to clarify
this point.
Overall, our data suggest that a sufficiently high
concentration of cytokine mediators is released in response
to the initial FRF treatment. This response essentially
establishes a passive diffusion gradient centered on the
RFTZ, but capable of traversing through the fluid extracellular matrix to expand throughout the untreated and
viable dermal tissue. If passive diffusion allowed for
involvement of the entire volume ($80 mm3) of the biopsy
specimen, the values shown in Table 1 represent the
minimum relative expression of each particular gene.
However, if passive diffusion did not contribute to involvement of tissue outside the RFTZ, the values shown in
Table 1 would underestimate the relative expression by
at least threefold. This is based on a maximum volume
estimate for the RFTZs in each 6 mm biopsy equal to $33%
of the total volume. The clinical relevance of our findings
remains unknown but will be clarified by longer term
follow-up of in vivo treatments.
In conclusion, we observed a vigorous in vivo wound
healing response with progressive increase in inflammatory cell infiltration from day 2 through 10 weeks post-FRF
treatment. Although coagulated dermal tissue was present
at day 28 post-treatment, an active dermal remodeling
process driven by the collagen chaperone HSP47 led to
complete replacement of RFTZs with new collagen by
10 weeks post-treatment. Furthermore, the results of our
immunohistochemical and PCR studies provide the first
evidence of neoelastogenesis accompanying upregulation
of procollagen secretion in the setting of FRF treatment.
The combination of neoelastogenesis and neocollagenesis
induced by treatment with the FRF system may provide a
reliable treatment option for skin laxity and/or rhytids.
ACKNOWLEDGMENTS
The authors thank James Newman, MD, Braden Stridde,
MD and R. Laurence Berkowitz, MD for their assistance in
conducting the clinical treatments.
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