Healing of Calvarial Wounds Created by

Er:YAG Laser Irradiation in Comparison

with Conventional Mechanical and

Femtosecond Laser Ablation in presence

or absence of BMPs

by

Martin Cloutier, DMD

A thesis submitted in conformity with the requirements

for the degree of Master of Science

Graduate Department of Dentistry

University of Toronto

© Copyright by Martin Cloutier (2009)

ii

Martin Cloutier

Healing of Calvarial Wounds Created by Er:YAG Laser Irradiation in Comparison

with Conventional Mechanical and Femtosecond Laser Ablation in presence or

absence of BMPs

Degree of Masters of Science

Graduate Department of Dentistry

University of Toronto

2009

Abstract The Er:YAG laser and the USPL are the most promising when considering the previous

study results and their physical characteristics. This investigation compared the healing of

various laser ablation units versus conventional mechanical cutting to explore the future

applications for bone surgery and the effects when combined with rhBMP-7. A full-

thickness circular defect was created on the parietal bones of mice for all the groups.

Hard tissue healing was assessed using a microcomputerized tomography. Wound closure

analyses suggested that the femtosecond laser created wounds displayed slightly healing

delay in closure over the healing period when compared to mechanical instrumentation.

The Er:YAG laser showed a healing rate similar to that of the mechanically ablated

groups. In summary, femtosecond and Er:YAG lasers are two modalities suitable for

bone ablation comparable to mechanical instrumentation.

iii

This work is dedicated to my parents

iv

Acknowledgements

I would like to thank my parents, brothers and sisters. My loving family, who despite the

distance, manage to convey incredible love and support.

A special thank you for Stephanie, whose unconditional love, support, and

encouragement lights up every day of my life.

Many thanks to Dr. Clokie, Dr. Sándor, Dr. Baker and Dr. Miller for your excellent

guidance. Thank you for allowing me to take this project and mold it into my own.

Thank you Bruno for laying the groundwork for this study, and for pointing me in the

right direction with your valuable suggestions and constructive criticism along the way.

Thank you Feryal, Susan and David for all of your help throughout this work. It was

always appreciated beyond words.

v

Table of Contents Abstract............................................................................................................................... ii

Acknowledgements ........................................................................................................... iv

Table of Contents.................................................................................................................v

List of Tables ................................................................................................................... viii

List of Figures.................................................................................................................... ix

Abbreviations.......................................................................................................................x

Symbols ............................................................................................................................ xii

Chapter I: Introduction and Review of the literature...........................................................1

1 Introduction ...................................................................................................................2

2 Review of Literature ......................................................................................................5

2.1 Bone........................................................................................................................5

2.1.1 General Background ........................................................................................5

2.1.2 Bone Healing ...................................................................................................7

2.1.3 Critical Size Defect..........................................................................................8

2.1.4 Animal Model................................................................................................10

2.2 Mechanical Instrumentation .................................................................................12

2.2.1 Ablation and Healing of Calcified Tissues....................................................12

2.3 Laser .....................................................................................................................14

2.3.1 Principle of Laser Ablation............................................................................14

2.3.1.1 Laser Scattering and Absorption ............................................................15

2.3.1.2 Linear Thermomechanical Ablation by Pulsed Irradiation ....................17

2.3.1.3 Kinetics of Phase Transitions .................................................................18

2.3.1.4 Ablation Plume .......................................................................................19

2.3.1.5 Plasma-Mediated Ablation .....................................................................20

2.3.1.6 Control of Precision................................................................................21

2.3.1.7 Thermal Side Effects ..............................................................................22

2.3.2 Ablation of Calcified Tissues ........................................................................22

2.3.2.1 Er:YAG Laser.........................................................................................22

vi

2.3.2.2 Pulsed Femtosecond Laser .....................................................................25

2.3.3 Healing of Calcified Tissues..........................................................................26

2.3.3.1 Er:YAG Laser.........................................................................................26

2.3.3.2 Pulsed Femtosecond Laser .....................................................................29

2.4 Bone Morphogenetic Proteins ..............................................................................30

2.5 Carriers for BMPs.................................................................................................31

Chapter II: Objectives and Hypothesis ..............................................................................33

3 Objectives and Hypothesis ..........................................................................................34

3.1 Objectives .............................................................................................................34

3.2 Hypothesis ............................................................................................................34

Chapter III: Materials and Methods...................................................................................35

4 Materials and Methods ................................................................................................36

4.1 Preparation of the bioimplant ...............................................................................36

4.2 Mechanical Instrumentation .................................................................................37

4.2.1 Carbide Bur....................................................................................................37

4.2.2 Diamond Bur .................................................................................................37

4.3 Laser Systems .......................................................................................................39

4.3.1 Er:YAG Laser................................................................................................39

4.3.2 Femtosecond Laser ........................................................................................40

4.4 Experimental Design ............................................................................................43

4.4.1 Model.............................................................................................................43

4.4.2 Surgery...........................................................................................................43

4.4.3 Animal Sacrifice ............................................................................................45

4.5 Micro CT-Scan 2D and 3D Analysis....................................................................46

4.6 Statistical Analysis................................................................................................48

Chapter IV: Results ...........................................................................................................49

5 Results .........................................................................................................................50

5.1 Initial wound size..................................................................................................50

5.2 Non-BMP group using 2D analysis......................................................................51

5.3 BMP group using 2D analysis ..............................................................................53

5.4 Non-BMP group using 3D bone volume analysis ................................................55

vii

5.5 BMP group using 3D bone volume analysis ........................................................58

5.6 Non-BMP group using 3D mineral content analysis............................................60

5.7 BMP group using 3D mineral content analysis ....................................................62

Chapter III: Discussion ......................................................................................................64

6 Discussion....................................................................................................................65

6.1 Non-BMP group using 2D analysis......................................................................65

6.2 BMP group using 2D analysis ..............................................................................67

6.3 Non-BMP group using 3D analysis......................................................................68

6.4 BMP group using 3D analysis ..............................................................................68

6.5 Future Studies and Development..........................................................................68

Chapter IV: Conclusion .....................................................................................................71

7 Conclusion ...................................................................................................................72

viii

List of Tables

Table 2.1 Studies of Calvarial Critical Size Defect in Mouse Model ...............................11

Table 2.2 Previous studies of Er:YAG on bone ablation and damage ..............................24

Table 2.3 Previous studies of Er:YAG laser on bone healing ...........................................28

Table 4 Previous studies of USPL on bone healing ..........................................................29

Table 5 Er:YAG laser parameters used for this study .......................................................39

Table 6 Femtosecond laser parameters used for this study ...............................................41

Table 7 Distribution of the 4 main groups.........................................................................43

ix

List of Figures Figure 2-1 Absorption curve of hydroxyapatite ................................................................17

Figure 4-1 Preparation of the bioimplant ..........................................................................36

Figure 4-2 Carbide bur (Brasseler, Savannah, GA, #558).................................................37

Figure 4-3 Diamond bur (Brasseler, Savannah, GA, #10839-012) ...................................38

Figure 4-4 Custom-made stabilisation device ...................................................................40

Figure 4-5 Handpiece model #2060 (KaVo, Biberach, Germany)....................................40

Figure 4-6 Surgical procedure ...........................................................................................44

Figure 4-7 2D reconstruction from the MIP feature using the Microview software .........46

Figure 4-8 3D reconstruction used for bone analysis ........................................................47

Figure 5-1 Percent calvarial closure without the use of rhBMP-7 ....................................52

Figure 5-2 Percent calvarial closure with the use of rhBMP-7 .........................................54

Figure 5-3 Bone volume without rhBMP-7.......................................................................56

Figure 5-4 Percent bone volume closure without rhBMP-7..............................................57

Figure 5-5 Bone volume with rhBMP-7............................................................................59

Figure 5-6 Bone mineral content without rhBMP-7..........................................................61

Figure 5-7 Bone mineral content with rhBMP-7...............................................................63

x

Abbreviations

BMC Bone mineral content

BMP Bone morphogenetic protein

cc Cubic centimeter

cm Centimeters

CSD Critical size defect

CO2 Carbon dioxide

CT Computed tomography

CW Continuous wave laser

ECM Extracellular matrix

Er:YAG Erbium Yttrium Aluminum Garnet

Er:YSGG Erbium Yttrium Scandium Gallium Garnet

eV Electron volt

FGF Fibroblast Growth Factor

fs Femtosecond

FTIR Fourier Transform Infrared

FVB Friend Virus B-Type (mouse)

Hb Hemoglobin

H&E Hematoxylin-and-eosin

Hz Hertz

HbO2 Oxyhemoglobin

xi

ICR Institute of Cancer Research

IR Infrared

MIP Maximum intensity projection

mJ Millijoules

MPA Multiphoton absorption

NIH National Institutes of Health

nm Nanometers

ns Nanosecond

PDGF Platelet-Derived Growth Factor

pH Potential of Hydrogen

PMN Polymorphonuclear neutrophils

pO2 Partial pressure of oxygen

ps Picosecond

rhOP-1 Human recombinant osteogenic protein-1

ROI Region of interest

SD Standard deviation

SEM Scanning electron microscopy

SE Standard error

SCID Severe combined immunodeficiency strain

TGF-α/β Transforming growth factor alpha/beta

um Micrometers

USPL Ultrashort pulsed laser

UV Ultraviolet

xii

Symbols

λ Wavelength

δ Optical penetration depth

∈ Absorption coefficient of specific isolated biomolecules

Pulse length

µa Absorption coefficient of a tissue

Sound velocity in water (1.5x103 m/s)

1

Chapter I: Introduction and Review of the literature

2

1 Introduction

Bone cutting and remodelling are required for many procedures in oral and

maxillofacial surgery. Traditionally, varieties of hand instruments, such as low-speed

diamond and carbide drills, or saws, are used to remove, shape or cut bone. Cutting of

bone tissue with mechanical tools can create significant tissue damage, so that efforts are

continuing to improve medical saws and drills, or to substitute them altogether with new

instruments (Gerber et al., 2001; Giraud et al., 1991). There have been reports of Hand-

Arm Vibration Syndrome in many domains including dental technicians and

orthopaedists where the use of drill is widespread (Bernard, 1997). Modern technologies

brought to the surgical field offer alternatives to conventional mechanical

instrumentation. Piezoelectric or ultrasonic devices in the surgical field have been

available for many years but have not yet replaced the conventional mechanical

instrumentation (Gruber et al., 2005). Surgical lasers have been well accepted by some

specialities but their use has been limited primarily to soft tissue ablation in

otorhinolaryngology (Schwab et al., 2004), ophthalmology (Krauss and Puliafito, 1995)

and dermatology (Alster and Lupton, 2001; Kageyama and Tope, 1999).

The applications of lasers for hard tissue ablation have been more challenging to

develop than for soft tissue. Only some specific types of lasers are suitable for bone

ablation. For example, the Nd:YAG (neodymium-doped: yttrium, aluminum, and garnet)

and CO2 (carbon dioxide) laser applications are limited due to their thermal side effects,

whereas the Er:YAG (erbium-doped: yttrium, aluminum, and garnet) laser is expected to

show efficiency in medical and dental applications because of its thermomechanical

3

ablation mechanism and the high absorption of its wavelength by water (Hale and

Querry, 1973; Israel et al., 1997; Robertson and Williams, 1971; Zolotarev et al., 1969).

Studies in the dental field have shown the potential applicability of the Er:YAG laser for

hard tissue ablation (Hibst, 1992; Hibst and Keller, 1989; Ishikawa et al., 2004; Kayano

et al., 1991; Keller and Hibst, 1989). In the same way, the Er:YAG laser has been

recognized as a promising tool for bone ablation in orthopaedic and otolaryngology

surgery (Nelson, Yow et al. 1988; Nuss, Fabian et al. 1988; Gonzalez, Van de Merwe et

al. 1990; Li, Reinisch et al. 1992; Buchelt, Kutschera et al. 1994).

Currently, there are few reports regarding the use of Er:YAG lasers in bone

surgery and little is known about the effectiveness of the Er:YAG laser for bone ablation

or the characteristics of the irradiated bone tissue compared to conventional rotary

instrumentation (Sasaki et al., 2002b). There are only a handful of studies that assess the

bone healing following bone ablation (Sasaki et al., 2002b).

The latest developments of laser bone ablation described in the literature show

another class of lasers, the ultrashort pulsed laser (USPL) systems. This new technology

may now offer a revolutionary method to ablate hard tissue. The most recent studies in

this area have shown significant material volume removal while minimizing collateral

damage (Neev et al., 1996a). Similarly to Er:YAG laser, the literature on bone healing

following USPL irradiation is nearly nonexistent and little is known about the laser

interactions with hard tissue.

4

There are many questions that remain to be explored regarding bone ablation with

various laser systems. Currently, mechanical instrumentation remains the gold standard,

but only new studies comparing the emerging laser systems will help to advance

techniques for effective and safe bone ablation in the medical and dental fields. Amongst

all the existing laser systems, the Er:YAG laser and the USPL are the most promising

when considering the previous study results and their physical characteristics.

5

2 Review of Literature

2.1 Bone

2.1.1 General Background

Bone is a connective tissue formed by numerous cells suspended in a matrix of

fibres, ground substance and water. Mineralization of this matrix provides its rigidity.

The relative proportions of each component varies depending on the type and

developmental stage of the bone.

Bone is composed of inorganic and organic matter. Air-dried compact bone is

composed of 71% inorganic matter, 18.6% of collagen, 1.3% of noncollagenous proteins,

and 8.1% of water (Eastoe and Eastoe, 1954). Inorganic matter is referred to as bone

apatite, a compound analogous to hydroxyapatite (Ca10(P04)6(OH)2), with varying ratios

of calcium hydroxide and phosphate (Boskey and Posner, 1984). Hydroxyapatite itself

represents about 80% of the bone apatite (Blair, 1998). Collagen represents

approximately 90% of the organic matrix, while noncollagenous proteins, lipids,

carbohydrates, enzymes, and hormones constitute 10% of this matrix. Collagen is a

structural protein found in a variety of tissues throughout the body and accounts for

approximatively 25% of all proteins in humans (Alberts et al., 1994; Parry and Graig,

1984). Nineteen different types of collagens are now known, and they are distinguished

using roman numerals. Most of them exist as a triple helix made up of three polypeptide

6

chains known as alpha chains with the various types of collagen made up with different

combinations of α -chains (α1 to α3 and other subtypes) (Koolman and Roehm, 2005).

Types I, II, and III represent 90% of all the collagens in the human body. Type I collagen

is a constituent of the bone matrix, dentin, skin and tendons, while type II collagen is

involved in cartilage synthesis. These two types of collagen have the ability to become

mineralized (Reddi et al., 1977). The non-collagenous proteins include phosphoproteins,

glycoproteins, matrix gla proteins (MGP), and proteoglycans, representing slightly less

than 2% of the organic portion of bone. Notably the BMPs are members of the non-

collagenous glycoproteins found in bone and they will be discussed in the following

section (Urist et al., 1979).

Osteoblasts are derived from mesenchymal stem cells. These stem cells

differentiate into osteoprogenitor cells and are responsible for the formation of bone

(Long, 2001). Osteoblasts become encased in lacunae within the newly formed bone and

mature into osteocytes. Osteocytes are responsible for maintaining and preserving

homeostasis as a living tissue (Mulliken et al., 1984). Osteoclasts are formed by the

fusion of hematopoietically derived monocytic cells under the influence of macrophage

colony stimulating factor (m-CSF), and many other factors in an effort to regulate the

bone formation and resorption (Ash et al., 1980; Cappellen et al., 2002; DeLacure, 1994;

Long, 2001; McLean and Urist, 1968; Owen, 1978; Pritchard, 1956).

7

2.1.2 Bone Healing

Bone injury triggers an inflammatory response and activates the complement

cascade. Extravasation of cells is initiated by the damaged blood vessels. Chemotactic

factors attract macrocytes and macrophages. Once activated, the macrophages will

release FGF, stimulating endothelial cells to release plasminogen activator and

procollagenase. PDGF, TGF- and TGF-β are released from the alpha granules of

activated platelets and will stimulate monocytes, macrophages, PMNs and lymphocytes.

The newly formed blood clot will provide growth factors in the injured site to provide an

adequate cell signalling. The hypoxic (pO2 5-10mm Hg) and acidotic (pH 4-6)

environment in injured tissue provides conditions that are needed for proper PMN and

macrophage activity (Marx et al., 1998). The proliferation phase starts on the third day

post-injury and may last up to 40 days following the initial injury. The inner cambium

layer of periosteum contains the osteoprogenitor cells which will proliferate and

differentiate into active osteoblasts. Those osteoblasts will secrete large amounts of

extracellular matrices. The remodeling starts by the fourth week and will continue

indefinitely. This process involves the activation of osteoblasts which vacate an area of

bone and osteoclasts that will resorb these areas. The resulting resorptive pits are known

as Howship's lacunae and will be repopulated by osteoblasts. The osteoid formed by

those osteoblasts is then calcified forming woven bone, which is slowly replaced by

lamellar bone.

8

2.1.3 Critical Size Defect

A critical size defect (CSD) can be defined as an intraosseous defect in a

particular type of bone of a specific species of animal that will not spontaneously heal

during the life span of that animal (Schmitz and Hollinger, 1986). Other authors will

define the CSD as the smallest bone defect which will not heal spontaneously in twelve

weeks (Dalle Lucca and McPherson, 2005; Dalle Lucca et al., 2005). Sanan wrote “when

the defect size is larger than one-tenth of skull, the defect will not heal spontaneously

during the lifetime” (Sanan and Haines, 1997). The concept of a critical size defect is

important to understand because the creation of subcritical defects allow bone formation.

Such bone formation is necessary for this study.

The CSD has been established in many animal models including rat, rabbit, dog

and monkey. The cranium has been the most preferred site although the mandible and

long bones have also been studied. The diameter of a circular CSD in a rat cranium is 8

mm, and 12 mm for a rabbit (Dalle Lucca et al., 2005).

The CSD of mouse cranium has not been clearly established but many studies

have used different sizes for such defects (Table 2.1). Dalle Lucca et al. (Dalle Lucca and

McPherson, 2005; Dalle Lucca et al., 2005) created defects of 3 and 4 mm in diameter in

mice cranium and analysed the percent closure with digital x-ray. The mean percent

closure was 41.07% after 12 weeks for 3 mm defects. Interestingly, they also found a

mean percent closure of 80.01% after one year. This data indicates that a 3 mm cranial

defect may not be a CSD in the mouse. To contrast, the 4 mm cranial defects showed a

9

55.2% closure after 12 weeks and 64.2 % after one year. The 4 mm defects may be the

CSD in a mouse considering that a significant number of the defects remain open at 12

weeks.

Cowan et al. (Cowan et al., 2004) identified 2 mm or larger defects as a critical-

size defects in 2-month-old male FVB mice. These defects showed only minimal (<10%),

unorganized bone formation as seen by X-ray analysis and H&E staining, although

defects of 0.8 mm completely healed within a 12 weeks period.

A significant difference in the ability to heal calvarial defects was seen between 6-

day-old and 60-day-old mice when 3-, 4-, or 5-mm defects were created. The data

suggests that a decreased osteogenic potential of adult (60-day-old) calvarial osteoblasts

may, in part, explain the inability of adult animals to a heal calvarial defect (Aalami et al.,

2004). Adults mice manifested healing only at the periphery of the defects, whereas

juveniles formed bone both as isolated islands in the central areas of the defect and

peripherally (Aalami et al., 2005). Mean percentage of healing in juvenile animals was 59

± 28 percent (SD) in the 3-mm group, 65 ± 29 percent in the 4-mm group, and 44 ± 11

percent in the 5-mm group. Mean percentage of healing in adult animals was less than 5

percent for all defect sizes (Aalami et al., 2004).

A 5-mm-diameter full-thickness circular mouse skull defect created with use of a

dental bur was found to be a critical-sized defect having a maximal healing observed in

any of the control groups of approximately 30% to 40% (Krebsbach et al., 1998).

Cultures of genetically engineered muscle-derived cells producing BMP-2 substantially

10

enhanced the healing of this critical-sized bone defect resulting in a repair greater than

99% of the original defect within 2 weeks (Krebsbach et al., 1998).

In summary, the quality and the quantity of bony repair in a critical size defect is

influenced by the animal species, animal age, anatomic location of the experimental

defects, size of the defect, and intactness of the periosteum (Schmitz and Hollinger,

1986). For our study, we have considered cranial defects of less than 2 mm in diameter to

be non critical.

2.1.4 Animal Model

There are numerous advantages using a mouse model for a bone healing study.

Mice are much cheaper to purchase and house than other larger animals. Refined

analytical tools, such as in vivo imaging techniques, microcomputed tomography and

molecular biology are readily available to evaluate mouse healing. Most importantly, the

ability to create both transgenic and knockout mice will greatly enhance an investigator’s

ability to evaluate the functional role of a specific gene in calvarial regeneration in future

research (Aalami et al., 2004). Also, laser or mechanical bone ablation of large surfaces

can be time consuming especially when a large number of animals are used. Finally, the

mouse calvarial defect model has been used in previous studies in our laboratory to

compare mechanical instrumentation and laser instrumentation (Girard et al., 2007a). For

the aforementioned reasons, we choose the mouse model for our experiments.

11

Authors Animal Model Defect size

(diameter) Method of ablation Method of analysis Results

(Seo et al., 2008) NIH-bg-nu-xid (NIH-3) immunocompromised mice

2.7 mm Trephine bur Histological analysis: cross sections

No mineralized tissue at 8 weeks in the control group with hydroxyapatite/ tricalcium phosphate carrier transplant

(Dalle Lucca and McPherson, 2005)

Mice 3 mm Water-cooled trephine Digital radiograph density 41% closure at 12 weeks 80% closure at 1 year

(Dalle Lucca et al., 2005)

Mice 4 mm Water-cooled trephine Digital radiograph density 55% closure at 12 weeks 64% closure at 1 year

(Cowan et al., 2004)

2-month-old male FVB mice ≥ 2 mm (2-3-4-5) 0.8 mm

Trephine/bur Radiographic and histology with H&E staining

< 10% closure at 12 weeks 100% closure at 12 weeks

(Aalami et al., 2004)

6-day-old CD1 mice 60-day-old CD1 mice

3 mm 4 mm 5 mm 3 mm 4 mm 5 mm

Dremel hand drill with brass bits coated with 225/390 diamond or diamond coated trephine bits with irrigation

Radiological and histological analysis with H&E staining

59% closure at 8 weeks 65% closure at 8 weeks 44% closure at 8 weeks <5% closure at 8 weeks <5% closure at 8 weeks <5% closure at 8 weeks

(Lee et al., 2001) Female SCID mice 5 mm Dental bur Histology with Von Kossa staining

Maximal closure of 30 to 40% at 4 weeks

(Krebsbach et al., 1998)

NIH-bg-nu-xid (NIH-3) immunocompromised mice 6-10-week-old females

5 mm Trephine bur attached to an electric Dremel handpiece

Histology with H&E staining Maximal closure of 11% at 12 weeks

(Girard et al., 2007a)

ICR mice 1.2 mm Dental burs and femtosecond laser Microtomographic analysis About 50% closure at 3 weeks and 60% closure at 12 weeks for the mechanical ablation groups

Table 2.1 Studies of Calvarial Critical Size Defect in Mouse Model

12

2.2 Mechanical Instrumentation

2.2.1 Ablation and Healing of Calcified Tissues

In many studies, osteotomies performed by the low speed burs did not produce

any altered tissue (carbonization, dehydratation, necrotic bone); they only produced bone

fragments detached during the bone cutting process. The absence of thermal damage

could be explained by the low speed (1,500 rpm) and by constant irrigation (de Mello et

al., 2008). Lewandrowski showed that the extent of thermally damaged nonvital bone,

measured histomorphometrically by the distance from the osteotomy site to the presence

of vital osteocytes, ranged from 25 to 100 µm in mechanically cut bone specimens. All

those osteotomies were created with a 1 mm diameter cross-cut fissure burs using

copious saline irrigation (Lewandrowski et al., 1996).

Some studies have described a thin dark layer of tissue below another layer of

residue (smear layer) in osteotomies created by low speed burs, probably caused by the

bur rotating faster at 10,000 rpm (Sasaki et al., 2002a; Sasaki et al., 2002b)

(Pourzarandian et al., 2004). Another study using bovine bones and 9.1 mm diameter

burs at 30,000 rpm showed that thermal damage can penetrate up to 1.9 mm (Shin and

Yoon, 2006). Consequently, osteocytes in this area of thermal damage died either by

apoptosis or necrosis. As osteocytes are believed to have a pro-osteogenic influence on

skeletal healing (Li et al., 2004), the broader area of thermal damage will negatively

affect bone regeneration.

13

Thus, the significant difference in bone healing described in the literature may be

proportional to the small area of osteocyte death. That area is greatly influenced by the

technique used for bone ablation including the use of irrigation, speed, shape and

sharpness of the drill, pressure applied by the handpiece and quality of the bone.

14

2.3 Laser

A LASER (an acronym for Light Amplification by Stimulated Emission of

Radiation) is any device generating a high-intensity parallel beam of monochromatic

(single wavelength) electromagnetic radiation. Maiman created the first operational laser

in 1960, a ruby laser emitting a brilliant red beam of light (Maiman, 1960). This was

followed within 3 years by the development of the argon, carbon dioxide (CO2), and

neodymium:yttrium-aluminum-garnet (Nd:YAG) lasers (Geusic et al., 1964). In 1975,

Zharikov introduced the Erbium:YAG (Er:YAG) laser (Zharikov et al., 1975). The active

medium of this laser is a solid crystal of yttrium-aluminum-garnet that is doped with

erbium. Each type of laser interacts differently with biologic materials although some

basic principles apply to all of them. The following sections explain some of those

principles that will ultimately influence bone healing following laser ablation.

2.3.1 Principle of Laser Ablation

Laser ablation has been described by Vogel as “any process of tissue incision or

removal, regardless of the photophysical or photomechanical processes involved” (Vogel

and Venugopalan, 2003). This definition does not include carbonization or

dehydratation/diffuse mass transfer, which result from extended exposure times or low

peak power. Vogel also describes the pulsed ablation as a result of pulse duration of less

than 1 ms (Vogel and Venugopalan, 2003). Current medical lasers use either continuous

wave or long pulse (> 0.1 nanosecond) delivery. Since the ablation process using longer

pulse lasers is dependent on tissue optical and thermal properties (e.g. absorption

15

coefficient, thermal diffusivity) the selection of both wavelength and pulse duration are

critical issues for these types of lasers.

Laser ablation requires the fracture of chemical bonds. The breakage of those

bonds leads to removal of molecules, molecular fragments, molecular clusters or

formation of voids within the material. Vaporization, molecular fragmentation, and void

formation are cause by photothermal, photomechanical, or photochemical mechanisms

(Vogel and Venugopalan, 2003). The spatial distribution of energy density resulting

from pulsed laser irradiation of tissue generates significant thermal and mechanical

transients. These thermomechanical transients are responsible for all laser ablation

processes that are not photochemically mediated and cause a linear thermomechanical

response of the tissue.

2.3.1.1 Laser Scattering and Absorption

Optical absorption and scattering influence the spatial distribution of volumetric

energy density deposited by laser radiation in tissue described earlier. If the scattering is

minimal or absent, the optical penetration depth (δ) of the incident radiation is defined by

the reciprocal of the absorption coefficient and determines the characteristic depth to

which the tissue is heated. However, if optical scattering is significant, δ is smaller

(Jacques, 1993). Optical scattering arises from spatial variations in the refractive index

within the tissue. The composition, size, and morphology of both cellular and

extracellular tissue components affect the refractive index of tissue (Dunn and Richards-

Kortum, 1996; Mourant et al., 2002; Saidi et al., 1995; Schmitt and Kumar, 1996, 1998;

Wang, 2000) and then ultimately the ablation process by a laser.

16

Optical absorption can be quantified using the equation I(x) = I0 exp (-ax) where

I0 is the incident intensity, I(x) is the intensity reaching depth x from the surface, and a is

the absorption coefficient, with units of reciprocal centimeters. The light (at least 63% of

it) is absorbed within a depth of 1/a from the surface. Then the energy can be transported

to greater depth by thermal conduction. The absorption coefficient µ a expresses the

optical absorption properties of tissue, while ϵ is used in reference to the optical

absorption properties of specific isolated biomolecules. Other than the water (~3 µm),

there are two other main absorption peaks that arise from the mineral component (~10

µm) and the organic portion of the bone (~6-7 µm). There is currently no laser available

that has a wavelength that can target the 3 major absorption peaks simultaneously in a

single photon regime.

17

Figure 2-1 Absorption curve of hydroxyapatite

This diagram represents the absorption curve of hydroxyapatite interacting with Er,Cr:YSGG, Er: YAG and

CO2 laser wavelengths. Peaks at 2.7 µm, 2.9 µm, 7,0 µm and 9,6 µm correspond to OH- group, free water,

(CO3)2- group and phosphate group in the molecule, respectively. The water absorption curve is shown by

the dotted line (Adapted from Prof. J. Featherstone, UCSF, California, USA)

2.3.1.2 Linear Thermomechanical Ablation by Pulsed Irradiation

THERMAL DIFFUSION

In the absence of photochemical or phase transition processes, the energy

absorbed by tissue in response to pulsed laser irradiation is completely converted to a

temperature rise. In 1983, a concept was introduced that described a process whereby

spatially confined microsurgical effects can be achieved by the use of laser exposures that

are shorter than the characteristic thermal diffusion time of the heated volume (Anderson

and Parrish, 1983).

18

THERMOELASTIC STRESS

Quick heating of tissue by pulsed laser irradiation also generates thermoelastic

stresses as the heated tissue volume reconfigures to its new equilibrium state. The

magnitude and temporal structure of thermoelastic stresses are influenced by multiple

factors including the laser pulse duration, the longitudinal speed of sound in the medium,

the depth of the heated volume and finally an intrinsic thermophysical property known as

the Grüneisen coefficient (Bushnell and McCloskey, 1968; Carome et al., 1964; Dingus

and Scammon, 1991; Gusev and Karabutov, 1993; Paltauf and Schmidt-Kloider, 1996;

Sisgrit, 1986). The “stress confinement” occurs when the ratio of the laser pulse duration

to the stress propagation time across heated volume is less or equal to one. In that case,

the internal stress do not propagate outside the heated volume during the laser irradiation

(Vogel and Venugopalan, 2003).

2.3.1.3 Kinetics of Phase Transitions

If the rate of volumetric energy deposition in the tissue created by the laser

radiation is higher than the rate of energy consumed by vaporization and normal boiling,

water will enter into a metastable superheated state. If the temperature is lower than the

spinodal point, the liquid can remain metastable. If the temperature is higher than the

spinodal point, the liquid undergoes “spinodal decomposition” and bubble nucleation.

Those processes create a significant pressure rise that can fracture the tissue. This

collective phase transition is referred to as “phase explosion” (Vogel and Venugopalan,

2003).

19

The phase transition is altered by the presence of extracellular matrix (ECM). The

bubble growth in tissue requires an internal pressure higher than pure liquids. This is

explained by the elastic strain necessary to deform the tissue matrix surrounding the

nucleation center. The pressure buildup occurs until it exceeds the ultimate tensile

strength of the extracellular matrix. This process is called “confined boiling” (Hibst and

Kaufmann, 1991; Majaron et al., 1999; Venugopalan, 1995; Walsh and Deutsch, 1989;

Walsh et al., 1989).

2.3.1.4 Ablation Plume

The ablation plume consists of the material removed at the ablation site by the

phase transitions. This ejected material consists mainly of particulate fragments (Nahen

and Vogel, 2002; Noack et al., 1997; Puliafito et al., 1987; Walsh and Deutsch, 1991).

The particle velocities range from about 500 to about 2000 m/s but can approach 5000

m/s for very high radiant exposures (Cummings and Walsh, 1992; Walsh and Deutsch,

1991). The scattering and absorption of the incident light by the ablation plume may

reduce the amount of energy deposited in the target and limit the ablation efficiency at

high radiant exposure. The ablation plume process is also responsible for acoustic

transients that will evolve into shock waves. The shock wave is usually emitted after the

end of the laser irradiation but high radiant exposures shock wave may begin during the

laser pulse (Cummings and Walsh, 1992). If the volumetric energy density becomes

sufficiently high, plasma formation will create a luminous plume (Cummings and Walsh,

1992). This plasma formation in front of the target may lead to a further decrease of the

optical transmission to the target, (Walsh and Deutsch, 1988; Willmott and Huber, 2000)

thus reducing the ablation efficiency (Green et al., 1990; Walsh and Deutsch, 1988). In

20

all studies, the plasma originated from the tissue surface rather than the ablation plume

itself (Green et al., 1990; Venugopalan et al., 1995; Walsh and Deutsch, 1988). This

occurs because the duration needed for plasma ignition (≤ 100 ns) is shorter than the time

commonly required for material ejection.

2.3.1.5 Plasma-Mediated Ablation

Laser-induced plasma-mediated ablation is achieved when the irradiance

threshold of a specific material is exceeded (Shen, 1984; Vogel, 1997, 2001). The

concept of plasma formation is important to understanding the ablation mechanisms of

high power lasers (Ready, 1971). This plasma formation may be initiated by thermionic

emission of free electrons if the material is strongly heated through linear absorption. In

this instance, the plasma will form at the surface of the target and act as a “shield” for

further energy deposition. The other way to form plasma is by multiphoton ionization or

avalanche ionization in materials that are usually transparent. Plasma formation allows a

highly localized energy deposition, especially in transparent and low-absorbing materials

(Gitomer, 1991; Niemz et al., 1992; Vogel, 2001; Willmott and Huber, 2000). Depending

on the mechanism and the target material, plasma formation has been reported with

radiant exposures as small as 0.25 J/cm2 (Venugopalan, 1995) especially when the target

exhibits a very high linear absorption. In contrast, linear absorption of the material does

not seem to influence the plasma threshold to a great extent when ultrashort pulses of

approximately 100 fs are used (Oraevsky et al., 1996). This is explained by the

multiphoton ionization that occurs early in comparison to the ionization avalanche,

making a significant difference when the pulses are extremely short. To summarize, the

linear absorption does not serve to lower the threshold of the plasma formation via

21

avalanche ionization. However, even for certain plasmas such as those created by

femtosecond lasers, multiphoton ionization only predominates the initial part of the pulse,

while avalanche ionization is the mechanism that will produce the majority of free

electrons.

Plasma plume refers to the plasma formation that can extend into the surrounding

air, which may significantly decrease the amount of light reaching the target (Root,

1989). With femtosecond lasers, this plasma plume does not act as a shield because the

pulse duration is too short to allow the formation of a plasma plume during the pulse

itself (Feit et al., 1998).

For femtosecond laser exposure, the pulse duration is shorter than the electron-

cooling and recombination times, resulting in a minimal transfer of energy through

collision and recombination of electron during the pulse. This leads to an ablation with

minimal plasma energy density, and minimal collateral damage. Interestingly, plasma-

mediated processes have been applied for the removal of hard tissue surfaces in air (Kurtz

et al., 1997; Loesel et al., 1998; Lubatschowski et al., 2002; Neev et al., 1996a; Niemz,

2002; Niemz et al., 1991) achieving ablation with little thermal damage of the residual

tissue (Vogel and Venugopalan, 2003).

2.3.1.6 Control of Precision

High precision of pulsed laser ablation requires a small ablation depth per pulse

along with minimal thermal and mechanical side effects. If linear absorption is utilized,

high precision is achieved by choosing a laser wavelength with a very small optical

22

penetration depth, combined with short pulse duration to provide thermal confinement. If

the nonlinear absorption is utilized (plasma mediated ablation), one should use small

pulse energies to minimize disruptive mechanical effects. The low thresholds are made

possible by using ultrashort pulses focused at large numerical apertures (Vogel and

Venugopalan, 2003).

2.3.1.7 Thermal Side Effects

The Arrhenius equation describes the temperature dependence of the rate of

irreversible heating-damage to many biological materials. The thermal damage of tissue

following laser ablation is described by the Arrhenius equation (Wright, 2003):

k (T) = Ae-Ea/RT

where k is the damage rate, T the absolute temperature, A the frequency factor

(s-1), Ea the activation energy, and R the universal gas constant. For example, normal-

spiking-mode pulses (approximately 200 microseconds) typically leave 10-50 microns of

collagen damage at the smooth wall of the incisions created by a Er:YAG laser. To

contrast, Q-switched pulses (approximately 90 nanoseconds) caused less thermal damage,

at about 5-10 microns of damage (Walsh et al., 1989).

2.3.2 Ablation of Calcified Tissues

2.3.2.1 Er:YAG Laser

The wavelength of the Er:YAG laser lies within the invisible portion of the

spectrum (infrared). This laser has a wavelength of 2940 nm and is highly absorbed in

23

water (Hale and Querry, 1973). From the visible spectral region, the absorption of water

increases by nearly 6 orders of magnitude and reaches a maximum at 2940 nm, where the

absorption coefficient is µa= 12,000 cm-1 (Vogel and Venugopalan, 2003) . There is

minimal tissue degeneration with very thin surface interaction that occurs after Er:YAG

laser irradiation due to its high absorption by water. Also, the temperature rise is minimal

in the presence of water irrigation, thus allowing hard tissue preparations without

carbonization (Aoki et al., 1998). Hibst (Hibst and Keller, 1989) and then Sasaki (Sasaki

et al., 2002a) have explained the theory of “microexplosion” regarding the mechanism for

hard tissue ablation. According to this theory, the energy is selectively absorbed by water

molecules. Internal pressure from the steam created causes explosive destruction of

inorganic substance before the melting point is reached. This theory accounts for the

events in Er:YAG laser ablation that cannot be fully explained by thermal effects alone.

A number of studies have demonstrated that the Er:YAG laser cuts bone precisely, with

minimal thermal damage of 5-30 um (Keller et al., 1991; Nelson et al., 1989; Nelson et

al., 1988; Sasaki et al., 2002a; Walsh and Deutsch, 1989; Walsh et al., 1989).

Histologically, only a thin altered layer is produced by Er:YAG laser on irradiated rat

calvaria in comparison with bone ablated using a CO2 laser that demonstrated extensive

thermal damage (Sasaki et al., 2002a). Sasaki et al. (Sasaki et al., 2002b) used Fourier

Transformated Infrared Spectroscopy (FTIR) to compare bone ablation from Er:YAG

laser, CO2 laser and bur drilling. The chemical composition of the bone surface after

Er:YAG laser ablation compared to bone drilling was found to be similar. However, with

CO2 laser irradiation, toxic substances were detected.

24

Authors Model Laser Pulse

duration (µs) Energy per pulse (mJ/pulse)

Radiant exposure (J/cm2)

Repetition rate (Hz)

Bone damage/altered layer (µm)

(Nuss et al., 1988) In vitro calvarium of guinea pig

Er:YAG 250 N/A 8-102 2 15

(Walsh et al., 1989) In vitro ginea pig bone Er:YAG 0.09 200

N/A N/A

0.5-10 4-80

1 2

5-10 10-50

(Nelson et al., 1989) Rabbit tibiae Er:YAG 200 N/A 140 5 4-10 (Devlin et al., 1994) Sprague-Dawley rats Er:YAG 250 100 N/A 5 A narrow zone of necrotic bone devoid of

osteocytes was seen adjacent to the defects. No carbonized debris was present in the lased bone

(Lewandrowski et al., 1996) Sprague-Dawley rats Er:YAG 1 53 60 2 25-100 (el Montaser et al., 1997) Sprague-Dawley rats Er:YAG N/A 75 38 N/A An amorphous, mineral-rich carbon layer

surrounds the lased bone defect (Sasaki et al., 2002a) 10-week-old male

Wistar rats Er:YAG 200 30-350 N/A 10 13.2 to 30 (mean 21.9)

(Sasaki et al., 2002b) Wistar rats Er:YAG 200 100 N/A 10 Chemical composition of the bone surface after Er:YAG laser ablation was much the same as that following bur drilling

(Pourzarandian et al., 2004) Male Witstar rats Er:YAG 200 100 N/A 10 11.4 to 28 (mean = 20.6) (Wang et al., 2005) New Zealand white

rabbits Er:YSGG 140 - 200 N/A 80 20 80

(Mizutani et al., 2006) Beagel dogs Er:YAG 200 30 to 350 10-100 30 No carbonization (de Mello et al., 2007) Male Witstar rats Er:YAG N/A 500 N/A 10 ~ 24

Table 2.2 Previous studies of Er:YAG on bone ablation and damage

Published studies of Er:YAG/ Er:YSGG on bone ablation and thermal damage including the laser parameters, the models used and the ablation outcomes

(Er :YAG λ = 2.94 µm, Er :YSGG λ = 2.78 µm)

25

2.3.2.2 Pulsed Femtosecond Laser

For a laser with ultrashort pulses (pulses shorter than 10 ps) and a high intensity

of radiation (peak power > 200 megawatts), the physics of the interaction changes. The

direct production of free electrons via multiphoton absorption (MPA) is crucially

important. Even when the total number of electrons produced by MPA is small they

provide the initial seed electrons for avalanche ionization. As a result, the ablation

threshold for USPL is independent of material defects and insensitive to linear absorption

(Neev et al., 1996b). In comparison, the conventional long pulse laser ablation is often

dominated by strong molecular absorption. As a result, the ablation depends on classic

tissue optical properties (Armstrong et al., 2002) allowing efficient ablation process only

with specific tissues. Neev et al. (Neev et al., 1996a) described many major advantages of

the USPL tissue ablation method. For example, the USPL allows efficient ablation due to

the small input of laser energy per ablated volume of tissue and the resulting decrease of

energy density needed to ablate material. At pulse duration of 170 fs, the smallest fluence

with which erosion can be achieved is below 1 J/cm2. With increasing pulse duration the

necessary threshold energy also rises (Neev et al., 1996b). Consequently, the USPL

causes minimal collateral mechanical and thermal damage due to the efficient ablation

and the extremely short deposition time (Schwab et al., 2004). Also, a large fraction of

the deposited thermal and kinetic energy is carried away with the ablated tissue, which

further decreases collateral damage. Girard et al. described a damage zone of 14 ± 5 µm

using freshly excised mice calvarial tissue. To describe that damage zone, staining using

fluorescent markers (calcein-acetoxymethylester and ethidium homodimer) for cellular

membrane integrity was used (Girard et al., 2007b). Femtosecond lasers can also cut at an

26

extremely precise depth because only a small amount of tissue is ablated per pulse and

the number of pulses can be controlled by feedback mechanisms. Finally, the USPL is

associated with a low acoustical noise level (as compared to the acoustical noise

produced by the high-speed dental drill or other laser systems) and minimal operative

pain due to localization of energy deposition and damage.

USPL systems have great potential for many clinical applications, since

femtosecond ablation occurs with little thermal or mechanical injury (Armstrong et al.,

2002). Armstrong et al. evaluated the ablation of ossicular tissue using a 1053 nm

Ti:Sapphire configured to deliver pulses of 350 femtoseconds. They showed a precise

bone ablation at a rate of 1.26 um/pulse, with virtually no evidence of photomechanical

injury (Armstrong et al., 2002). However, the ablation rate was lower per pulse than some

other conventional lasers. Although, there was subtle evidence of hydroxyapatite melting

at high power, intact Haversian canal systems were identified, and no cracks or fractures

indicating explosive trauma from vaporization of gazes were present.

2.3.3 Healing of Calcified Tissues

2.3.3.1 Er:YAG Laser

Some authors (Lewandrowski et al., 1996; Pourzarandian et al., 2004) have

reported that healing rate following Er:YAG laser irradiation may be equivalent or even

faster than that following bur drilling. In 2005, de Mello et al. demonstrated through

histological analysis that their experimental group (Er:YAG laser) presented a more

advanced bone repair than that observed in the control group (bur drilling) (de Mello et

27

al., 2007). This could be explained by the unusual morphologic features of this tissue,

which may act as a mechanically favourable surface to which the coagulum adheres at the

very beginning of the tissue repair process (Pourzarandian et al., 2004).

Conversely, there have been reports of delayed healing following Er:YAG laser

ablation, likely due to lack of water irrigation as well as higher output energy (el

Montaser et al., 1997; Nelson et al., 1989). Nelson et al. noted histologically that there

was a delay in healing of laser osteotomies groups (Er:YAG) with respect to saw

osteotomies groups (Nelson et al., 1988). Similarly, el Montaser et al. described how

bone filling of the lased defect was retarded by delayed resorption of an amorphous,

mineral-rich carbon layer (el Montaser et al., 1997). Again, saw osteotomies performed in

the rat model resulted in more callus formation than Erb:YAG osteotomies (Buchelt et

al., 1994). Finally, Lewandrowski et al. used histologic evaluation to demonstrate no

significant difference in the amount of newly formed woven bone at the osteotomy site or

screw holes made by the laser or the drill (Lewandrowski et al., 1996). A summary of

those published studies regarding bone healing following Er:YAG laser ablation can be

seen in Table 2.7.

28

Authors Model Laser Pulse duration (µs)

Energy per pulse (mJ/pulse)

Radiant exposure (J/cm2)

Repetition rate (Hz)

Bone healing

(Nelson et al., 1989) Rabbit tibiae Er:YAG 200 N/A 140 5 Delay in the healing process after laser osteotomy (Buchelt et al., 1994) Sprague

Dawley rats Er:YAG 250 300 N/A 2

Biomechanical measurements of bone treated by power saw or Erb:YAG laser osteotomies, respectively, showed no significant statistical difference in the stability of bone between the two groups. However, saw osteotomies resulted in more callus formation than Erb:YAG osteotomies

(Devlin et al., 1994) Sprague-Dawley rats

Er:YAG 250 100 N/A 5 Delayed healing of Er:YAG laser osteotomies compared with conventionally prepared osteotomies

(Lewandrowski et al., 1996)

Sprague-Dawley rats

Er:YAG 1 53 60 2 Same healing as mechanical ablation

(el Montaser et al., 1997)

Sprague-Dawley rats

Er:YAG N/A 75 38 N/A Bone infilling of the lased defect was retarded by delayed resorption of the amorphous, mineral-rich carbon layer.

(Pourzarandian et al., 2004)

Male Witstar rats

Er:YAG 200 100 N/A 10 Initial bone healing following Er:YAG laser irradiation occurred faster than that after mechanical bur. These results demonstrated that new bone formed after Er:YAG laser irradiation was almost 1.65 times greater than that formed by mechanical bur.

(Wang et al., 2005) New Zealand white rabbits

Er:YSGG 140 - 200 N/A 80 20 Minimal delay before the subsequent healing process begins, the overall postoperational healing is favorable

(Mizutani et al., 2006) Beagel dogs Er:YAG 200 30 - 350 10-100 30 Histologically, the amount of newly formed bone was significantly greater in the laser group than in the curette group, although both groups showed similar amounts of cementum formation and connective tissue attachment.

(de Mello et al., 2007) Male Witstar rats

Er:YAG N/A 500 N/A 10 At 7 and 14 days after surgery, the experimental group presented earlier bone repair in comparison to the control group. At 21 days, the histological features of the two groups were very similar

Table 2.3 Previous studies of Er:YAG laser on bone healing

Published studies of Er:YAG/ Er:YSGG on bone healing including the laser parameters, the models used and the ablation outcomes (Er :YAG λ = 2.94 µm,

Er :YSGG λ = 2.78 µm). Overall, the bone healing following Er:YAG laser ablation is comparable to the mechanical ablation.

29

2.3.3.2 Pulsed Femtosecond Laser

The scientific literature reviewing the bone healing after USPL ablation is very

limited. A review of the literature revealed only two articles on the subject. Girard et al.

(Girard et al., 2007a) published their results in 2007 regarding bone healing following

ablation with a 200 femtosecond pulsed laser. The bone volume measurements were

slightly lower for the femtosecond laser group than for the mechanical groups. However,

statistically significant differences were seen only at week 6. Also, no significant

differences in closure were noted for the different methods in the BMP treated groups. A

study published in 2007 used a 1 picosecond pulsed Ti:Sapphire laser (Leucht et al.,

2007). The data demonstrated that corticotomies performed with Ti:Sapphire lasers were

associated with a reduced initial inflammatory response at the injury site, leading to

accelerated osteochondroprogenitor cell migration, attachment, differentiation and

eventually matrix deposition. The model used for that study was murine tibial cortical

defect. Those 2 aforementioned studies are difficult to compare considering that the

animal model, the laser and the type of defects are different.

Authors Year Model Laser Pulse duration (femtosecond)

Energy per pulse (µJ/pulse)

Repetition rate (Hz)

Spot size diameter (µm)

Bone healing

(Girard et al., 2007a)

2007 Mouse calvarium

Ti:Sapphire < 200 100 1000 45 Delayed healing at 6 weeks but no significant difference with the mechanical ablation group at 9 and 12 weeks

(Leucht et al., 2007)

2007 Murine tibia

Ti:Sapphire 1000 650 240

50 Accelerated bone matrix deposition in the laser group

Table 4 Previous studies of USPL on bone healing

Published studies of picosecond and femtosecond laser bone healing including the laser parameters, the

models used and the ablation outcomes.

30

2.4 Bone Morphogenetic Proteins

To date, twenty members of the bone morphogenic protein (BMP) family have

been identified and characterized (Ming Zhao and Mundy, 2004). BMPs are members of

the transforming growth factor β (TGF-β) superfamily, and have been implicated in the

development of various tissues, most notably bone (Dayoub and et, 2003; Linkhart et al.,

1996; Sakou, 1998; Urist, 1965). Physiological roles of BMPs and BMP receptor

signalling in normal bone formation have been investigated. The discovery of the Smad

proteins has allowed for a better understanding of how BMPs function (Celeste et al.,

1990).

BMP-2 and BMP-7 are morphogens known to stimulate new bone formation and

are commercially available as human recombinant proteins (Nussenbaum et al., 2005).

Clinical studies have shown that BMP-2 can be utilized in various therapeutic

interventions such as bone defects, non-union fractures and spinal fusion (Kleeman et al.,

2001). Injection of BMP-2 over the surface of murine calvaria induces periosteal bone

formation without a prior cartilage phase (Chen et al., 1998). A gene therapy study

showed that bone defects are healed by the implantation of a bioresorbable polymer

mixed with bone marrow mesenchymal stem cells to which adenovirus BMP-2 was

transferred (Chang et al., 2003). Another study demonstrated that cells derived from

muscle tissue that were genetically engineered to express BMP-2 elicited the healing of

critical-sized skull defects in mice (Lee et al., 2001).

31

In 1990, human complementary DNA of BMP-7 was isolated and cloned

(Ozkaynak et al., 1990; Wang et al., 1990). This form of the BMP-7 was derived from a

recombinant Chinese hamster ovary cell line and consisted of two 139-amino-acid

monomers connected by a disulfide bond (Implant, 2001). Since then, human

recombinant BMP-7 was successfully used in promoting normal wound healing in animal

experiments (Blockhius et al., 2001; Bostrom and Camacho, 1998; Cook et al., 1994).

Also, the first reported case using a bone morphogenetic protein bioimplant in a human

was by Moghadam et al (Moghadam et al., 2001) in 2001. BMP-7 was used to

reconstruct an osseous defect of the mandible, followed by histological confirmation of

new bone formation. Recently, use of rhBMP-7 has been approved in many countries for

certain applications in humans. The approved formulation relates to powdered bovine

bone-derived type I collagen as a carrier for rhBMP-7 (OP-1 device; Stryker Howmedica,

Mühlheim, Germany). In membranous bone healing, BMP-7 is expressed more strongly

than other the BMPs (Spector et al., 2001), suggesting that rhBMP-7 is an appropriate

candidate for experimental use in calvarial healing. The present study addresses

recombinant human bone morphogenic protein-7 (rhBMP-7), which has been shown to

be effective in promoting bone formation (Ripamonti and Duneas, 1998; Ripamonti et al.,

1996; Sampath et al., 1992; Terheyden et al., 1999).

2.5 Carriers for BMPs

A carrier is necessary for the clinical application of BMPs, since the highly

purified BMPs and recombinant BMPs are water-soluble. Without carriers, they diffuse

away quickly in vivo, which prevents the relatively high local BMP concentration

required for bone induction (Reddi, 1995; Wonzey and Rosen, 1998). Also, to improve

32

clinical handling of the BMPs, inert carriers such as glycerol, dry powder microparticles

of poly (D,L-lactide-co-glycolide) (PLGA) (Kenley et al., 1994), hyaluronic acid or

reverse phase poloxamers (Clokie and Urist, 2000), like poloxamer 407 (polyethylene-

polypropylene glycol), are often utilized. Poloxamer 407 (Pluronic F127, BASF,

Parsippany, NJ) exhibits a unique reverse thermal behaviour when used in a

concentration greater than 20% solution. It will increase in viscosity as the temperature

rises above 4ºC and will form a mouldable gel once present above room temperature.

The addition of Pluronic F127 to relatively rigid biomaterial granules allows the

bioimplant to be shaped to fit the defect and prevent migration of biomaterials.

Poloxamer 407 has been studied extensively for over the past two decades. No

toxic reactions have been noted in both acute and chronic exposure; therefore, it is listed

as an inactive ingredient by the FDA. It was decided that the use of poloxamer 407 as a

control group was not required for this study since previous experiments in our laboratory

have demonstrated that, used alone, it does not induce bone formation (Clokie et al.,

2002). Also, poloxamer 407 as the excipient seems superior when compared to various

methods of delivery to optimize the action of the BMPs. A study in 2000 showed that

greater than 90% of the implants were composed of new bone, 28 days after implantation

when the poloxamer was used as the BMP delivery vehicle(Clokie and Urist, 2000). The

other excipients used were human tendon collagen, human demineralized bone matrix

and hydroxyapatite.

33

Chapter II: Objectives and Hypothesis

34

3 Objectives and Hypothesis

3.1 Objectives

The bone healing potential of defects ablated with conventional Er:YAG laser,

femtosecond laser and low-speed rotary instrumentation was evaluated in mouse model.

As the literature has not reported a direct comparison of primary bone healing associated

with femtosecond laser and Er:YAG laser ablation, it was the primary goal of this study

to address this issue.

The secondary objective of this investigation was to determine if the addition of

rhBMP-7 to the bone defects would alter the healing process. This would allow for the

removal of variability in triggering healing that is currently dependent on the laser cutting

profile with respect to cell membranes or the type of biologic damage. We also sought to

confirm that both femtosecond and Er:YAG laser ablation did not interfere with BMP

bone inductive capabilities as compared to conventional mechanical techniques.

3.2 Hypothesis

It is hypothesized that femtosecond laser, Er:YAG laser and mechanical ablation

groups will demonstrate a similar rate of bone healing, and that the addition of rhBMP-7

will lead to a more rapid rate of healing in all the groups.

35

Chapter III: Materials and Methods

36

4 Materials and Methods

4.1 Preparation of the bioimplant

Poloxamer 407 (BASF Corp., Parsippany, N.J.) was prepared by using the cold

method (Schmolka, 1972). Twenty-five grams of Pluronic F-127 powder was slowly

added over a period of 5 minutes into a beaker containing 75 cc of sterile water. This

procedure was performed by using continuous agitation overnight at 4°C to ensure

complete dissolution.

A recombinant BMP-7 (OP-1, Stryker, Kalamazoo, MI) was used for these

experiments. The commercial form of rhBMP-7 is supplied in a glass vial containing one

gram of the device as a sterile dry powder (3.5 µg rhBMP-7/1 mg bovine bone type I

collagen).

Figure 4-1 Preparation of the bioimplant

A) and B) OP-1 by Stryker (rhBMP-7) C) Poloxamer 407 used to deliver the BMP (liquid state).

The OP-1 device was combined with a 30% Pluronic F-127 carrier to achieve a

concentration of 30 µg/ml of rhBMP-7. A micropipette was used to depose 50 µl of the

A

C

B

37

bioimplant per defect being the equivalent of 1.5 µg of rhBMP-7 per site (Schmoekel et

al., 2004).

4.2 Mechanical Instrumentation

4.2.1 Carbide Bur

Calvarial defects in the first group of mice were created using a carbide bur

(Brasseler, Savannah, GA, #558) in a hand-held slow speed electrical handpiece (5,000

rpm) with copious saline irrigation. The bur diameter was 1.2 mm.

Figure 4-2 Carbide bur (Brasseler, Savannah, GA, #558)

4.2.2 Diamond Bur

A second group of mice had their calvarial defects prepared using an end cutting

diamond bur (Brasseler, Savannah, GA, #10839-012) similarly to the carbide group. The

diameter of the bur was also 1.2 mm.

38

Figure 4-3 Diamond bur (Brasseler, Savannah, GA, #10839-012)

Because both carbide and diamond burs are considered the gold standards for hard

tissue ablation in most of the surgical fields, those two modalities were used as control

groups in this experimental design.

39

4.3 Laser Systems

4.3.1 Er:YAG Laser

The Key Laser III (KaVo, Biberach, Germany) (Class 4) is a Er:YAG laser with a

wavelength of 2,94 µm (near infrared spectrum). At low fluence, desiccation can prevent

efficient ablation of bone, as a result of which the irradiation of the calvarium was

performed in non contact mode under saline solution irrigation. Irrigation also served to

decrease the thermal adverse effect of the laser ablation. A custom device was used to

maintain the handpiece at a specific distance from the calvarium coinciding with the focal

length (see Figure 4-4). The handpiece model used was the #2060 (KaVo, Biberach,

Germany) (see Figure 4-5). It is a non contact handpiece with fine spray cooling

integrated. A 655 nm pilot laser beam was used to aim precisely at the specimen. The

repetition rate was set at 2 Hz to avoid overheating of the specimen. The diameter at the

focal point was 1.2 mm with a focal length of 28 mm. Finally, the pulse duration was 200

µs and the output pulse energy set at 400 mJ. The ablation rate was estimated at ~ 25

um/pulse.

Laser Pulse duration (microsecond)

Energy per pulse (mJ/pulse)

Repetition rate (Hz) Spot size diameter (mm)

Er:YAG 200 400 2 1.2 Table 5 Er:YAG laser parameters used for this study

40

Figure 4-4 Custom-made stabilisation device

Used to maintain the Er :YAG laser handpiece at a specific focal length from the calvaria

Figure 4-5 Handpiece model #2060 (KaVo, Biberach, Germany)

4.3.2 Femtosecond Laser

For this investigation, the laser source used was a commercially available

Ti:Sapphire chirped-pulse amplification system, based on a frequency-doubled Erbium

fiber oscillator at 775 nm, operating at a pulse repetition rate of 1 kHz. The low energy

pulses were amplified in a regenerative Ti:Sapphire pumped by a frequency doubled Q-

switched Nd:Yag laser to provide amplified pulses duration of < 200 fs as measured by a

2nd-order autocorrelator. The laser intensity was controlled using a polarizer and a half-

wave plate in the beam path. The beam was steered with highly reflective coated

41

dielectric mirrors (CVI Laser LLC, Albuquerque, NM) and passed through a 10 cm focal

length lens, resulting in a focal spot size of approximately 45 µm as measured using the

razor blade edge technique. The pulse energy was measured using a photodiode

calibrated against a standard power meter (Newport Corporation, Stratford, CT). The

average pulse energy was determined by measuring the output power at 1 kHz over an

integration period of several seconds. The femtosecond laser described has a pulse

intensity of 100 µJ per pulse (~8J/cm2). The error on this measurement was estimated to

be at most ± 1µJ per pulse (± 0.08J/cm2). The overall pulse-to-pulse energy variation

measured in this system was approximately 2%.

Laser Pulse duration (femtosecond)

Energy per pulse (µJ/pulse)

Repetition rate (Hz) Spot size diameter (µm)

Ti:Sapphire < 200 100 1000 45 Table 6 Femtosecond laser parameters used for this study

For this group the heads of the animals were stabilized with a custom made

craniotaxic stabilization table equipped with a magnetic mount, allowing accurate

placement of the mouse under the laser field. The table was carefully moved to ensure a

position whereby the flat surface of the parietal bone was orthogonal to the laser beam

and at the focal point of the lens. The circular movement of the mouse under the beam

was achieved and controlled using x-y actuators (Oriel, Stratford, CT). Laser irradiation

was controlled by a fast rise shutter (Vincent Associates, Rochester, NY) synchronized

with the actuators, such that no irradiation would occur unless the actuators were in

motion. The mouse was moved in a circular path of 1.25 mm diameter at a speed of 200

um/second. A single laser pass resulted in a cut of approximately 50 µm in width on the

skull surface, thus requiring removal of small bone cores with microsurgical instruments.

This step was not required with the carbide, diamond and Er:YAG groups. At 1 kHz

42

repetition rate and speed of 200 um/second, every micron on the parietal bone surface

received 225 pulses. This was sufficient to perforate both cortices of the parietal bone,

which is approximately 250 µm thick for a 4-week-old ICR strain mice. The ablation rate

was estimated at ~ 1 um/pulse.

43

4.4 Experimental Design

4.4.1 Model

Four-week-old female Tac:Icr:Ha (ICR) mice (Taconic Farms, Albany, NY) were

purchased from a local provider (Charles River Laboratories, Wilmington, MA). Animals

were housed in a light and temperature controlled environment. 160 mice were used for

this study and they were divided into 4 groups each having 40 animals. Within each

group, the animals were further divided in 4 subgroups each having ten animals.

3 weeks without

BMPs

3 weeks with

BMPs

12 weeks without

BMPs

12 weeks with

BMPs

Diamond bur 10 10 10 10

Carbide bur 10 10 10 10

Er:YAG laser 10 10 10 10

Femtosecond laser 10 10 10 10

Table 7 Distribution of the 4 main groups

4.4.2 Surgery

The animal protocol was approved by the animal ethics committee of the

University of Toronto (protocol # 02-122). The mice were anesthetized by inhalation

with isoflurane 1-3% for induction and 0.75-1.25% for maintenance. The scalp of the

animals were then shaved using an electric clipper. The eyes were protected from dryness

by applying a tear substitute.

The surgical sites were cleaned with 70% ethanol solution and betadine.

Subsequently, the mouse was moved to a sterile laminar flood hood and, using a

dissecting microscope, a full thickness sagittal incision was made with a number 15C

44

blade. The scalp was dissected to expose the parietal bones and the periosteum was

stripped. The midline skin incision was 8-9 mm in length. A full-thickness circular defect

was created on the parietal bones with minimal penetration of the dura under copious

irrigation with normal saline for three of the groups (carbide, diamond and Er:YAG

laser). The femtosecond group did not require normal saline irrigation during the ablation

process. In each group, 20 animals had rhBMP-7 applied to the defects. The bioimplant

was allowed to solidify prior to repositioning and suturing the skin flaps.

Figure 4-6 Surgical procedure

A) Preoperative: the surgical site was cleaned with 70% ethanol solution and betadine B) Mechanical

instrumentation creating defects in the calvarium C) Two full-thickness circular defects were created in the

parietal bones D) Application of the bioimplant

A

C

D

B

45

The scalp incisions were closed in a single layer using 4-0 Vicryl (EthiconTM,

Johnson & Johnson Co., Sommerville, NJ) interrupted sutures and animals were allowed

to recover under a warm lamp before being returned to the animal facility. The animals

were provided with food and were allowed activity ad libitum. Buprenorphine (0.05-0.1

mg/kg) was administered prophylactically intraoperatively and as needed for 24 hours

following the surgery. No postoperative mortality or morbidity (infection of wound) was

observed in any group during the course of this study.

4.4.3 Animal Sacrifice

Ten animals from each group were sacrificed at 3 and 12 weeks and the skull

specimens were dissected free of soft tissue. The animals were euthanized via exposure to

carbon dioxide. Specimens were fixated in 10% buffered formalin for at least 24 hours at

4°C.

46

4.5 Micro CT-Scan 2D and 3D Analysis

Hard tissue healing was assessed using a microcomputerized tomograph (GE

Healthcare, Amerham, UK) and associated analysis software (Microview 2.029 and

eXplore Reconstruction, GE Healthcare, Amersham, UK). Two different techniques were

used to evaluate bone healing. The first involved a two-dimensional analysis by

reorienting the reconstructed skulls so that the wounded parietal bone surface was parallel

to the z-axis and subsequently was projected to that plane using the maximum intensity

projection (MIP) feature provided with the Microview software. This resulted in an

image equivalent to that obtained from a conventional radiograph (see Figure 4-7).

Figure 4-7 2D reconstruction from the MIP feature using the Microview software

Example of two-dimensional reconstruction of initial wound size and healing calvaria. This allowed

determination of the percent closure.

47

The digitized images were then transferred to SigmaScan software (Systat

Software, Chicago, IL) allowing quantification of the unfilled area, which was then

subtracted from the wound measurement established immediately at time of wounding to

allow determination of a percentage closure measurement.

Three-dimensional bone analysis was obtained by superimposing a region of

interest (ROI) corresponding to the original wound volume and using the bone

measurement feature provided by the Microview software. It was possible to assess the

volume of the newly formed bone as well as the bone mineral content. Bone density was

also calculated from those previous measurements. Control specimens with defect but

without healing were used to calibrate the results.

Figure 4-8 3D reconstruction used for bone analysis

Example of initial wound size (in yellow) superimposed on healing calvaria. This allowed determination of

bone volume measurements within the wounded area and subsequently, the percent closure calculations.

48

4.6 Statistical Analysis

Graphs were generated using SPSS 15.0 for Windows (SPSS Inc., Chicago, IL).

Data were presented as the mean and standard errors of the mean. Statistical differences

between 2 groups were calculated using a Student’s t-test. Statistical significance was

established at p < 0.05.

49

Chapter IV: Results

50

5 Results

5.1 Initial wound size

The measurement of the average defect diameter immediately following

wounding was required in order to ensure minimal differences. Ten calvarial wounds for

each technique were used to calculate the mean diameter via two-dimensional analysis

using the Microview software. The average wound diameter at t=0 was 1.34 ± 0.05 mm,

1.38 ± 0.06 mm, 1.21 ± 0.08 and 1.25 ± 0.01 mm for the carbide group, the diamond, the

Er:YAG laser and the femtosecond (FS) laser group, respectively. These differences were

not statistically significant (p > 0.05).

51

5.2 Non-BMP group using 2D analysis

Over the 12-week healing period, complete closure was not seen in any group.

The graphic representation of these data is found in Figure 5-1. At 3 week, the average

percent closure was 57 ± 5%, 59 ± 5%, 41 ± 4% and 59 ± 3% for carbide, diamond,

femtosecond laser and Er:YAG groups, respectively. The percent of closure for the

femtosecond laser group was significantly less (p < 0.05) than the carbide, diamond and

Er:YAG groups. No statistically significant difference (p > 0.05) was observed between

the other groups.

At week 12, no statistically significant differences between the groups were

identified (p > 0.05). The average percent closure was 69 ± 4%, 72 ± 4%, 62 ± 4% and 70

± 3% for carbide, diamond, femtosecond laser and Er:YAG groups, respectively.

When comparing the wound closure within similar wounding methods between 3

and 12 weeks, a statistical difference was shown for all groups with the exception of the

carbide group (p = 0.66).

52

Figure 5-1 Percent calvarial closure without the use of rhBMP-7

Bar graph representing the mean comparing the two-dimensional percent calvarial closure in mice at week

3 and 12 using four wounding methods without the use of rhBMP-7. The FS group showed a decrease of

the percent closure of the defects in comparison to the other groups at 3 weeks. The error bar represents the

standard error of the mean. The * represents a statistically significant difference (p < 0.05).

53

5.3 BMP group using 2D analysis

The addition of rhBMP-7 to the surgical site resulted in complete closure of the

defects as early as 3 weeks. The graphic representation of these data is found in Figure

5-2. Overall, the average percent closure and standard error of the mean at 3 weeks was

92 ± 3%, 88 ± 3%, 77 ± 7% and 84 ± 5% for carbide, diamond, femtosecond laser and

Er:YAG groups, respectively. None of these differences were statistically significant (p >

0.05) although, when the carbide and FS group were compared, statistical significance

was appreciated (p = 0.053).

At 12 weeks, no statistically significant differences (p > 0.05) were identified

between the groups evaluated (Figure 5-2).

When the same ablation method were compared between the 3rd and 12th weeks,

only the Er:YAG group showed a statistically significant increase (p < 0.05) in percent

calvarial closure.

54

Figure 5-2 Percent calvarial closure with the use of rhBMP-7

Bar graph representing the mean comparing the two-dimensional percent calvarial closure in mice over 12

weeks using four wounding methods with the use of rhBMP-7. The percent closure variable shows no

statistical difference between all the groups at 3 and 12 weeks indicating that wound closure was similar for

all the groups irrespective of the ablative technique and the time point. The error bars represent the standard

errors of the means.

55

5.4 Non-BMP group using 3D bone volume analysis

The graphical representation of these data is found in Figure 5-3. The numbers in

parentheses represent the percent wound closure. Bone volumes measured at 3 weeks of

healing were 0.12 ± 0.02 mm3 (57%), 0.15 ± 0.02 mm3 (59%), 0.07 ± 0.01 mm3 (41%)

and 0.14 ± 0.01 mm3 (59%) for carbide, diamond, femtosecond and Er:YAG

respectively. A statistically significant difference (p < 0.05) for this 3 week cohort was

identified when comparing the femtosecond group and all the 3 other groups.

At week 12, the bone volumes found were 0.22 ± 0.02 mm3 (69%), 0.25 ± 0.02

mm3 (72%), 0.21 ± 0.02 mm3 (62%) and 0.24 ± 0.02 mm3 (70%) for carbide, diamond,

femtosecond and Er:YAG respectively. There were no statistically significant differences

between any of the groups.

When comparing the bone volume formation within similar wounding methods

between 3 and 12 weeks, statistically significant increase were shown for all groups in

this cohort.

56

Figure 5-3 Bone volume without rhBMP-7

Bar graph representing the mean comparing the bone volume formation over 12 weeks without the use of

rhBMP-7. A statistical difference was found at 3 weeks between the femtosecond group and all the 3 other

groups. There were statistically significant increase in bone volume between 3 and 12 weeks within all the

groups. The error bar represents the standard error of the mean. The * represents a statistically significant

difference (p < 0.05). The dotted line represents the mean bone volume in a non-ablated calvarium.

57

Figure 5-4 Percent bone volume closure without rhBMP-7

Reformatted data of Figure 5-3 using percentage closure as vertical axis

58

5.5 BMP group using 3D bone volume analysis

Figure 5-5 shows the data for the bone volume measurements with the addition of

rhBMP-7. The numbers in parentheses represent the percent wound closure. For the week

3 group, measured bone volumes were 0.25 ± 0.03 mm3 (92%), 0.24 ± 0.02 mm3 (88%),

0.20 ± 0.04 mm3 (77%) and 0.24 ± 0.04 mm3 (84%) for carbide, diamond, FS and

Er:YAG group, respectively. There were no statistical differences between those groups.

At week 12, the measured bone volumes were 0.41 ± 0.03 mm3 (94%), 0.40 ±

0.03 mm3 (92%), 0.38 ± 0.03 mm3 (91%) and 0.43 ± 0.03 mm3 (96%) respectively.

Statistical differences were not found between the groups.

Finally, when comparing the bone volume formation within similar wounding

methods between 3 and 12 weeks, statistically significant increases (p < 0.05) were seen

for all groups describing the bone formation process during healing phase.

59

Figure 5-5 Bone volume with rhBMP-7

Bar graph representing the mean comparing the bone volume formation over 3 and 12 weeks with rhBMP-

7. Statistically significant differences were not found between any of the groups. The error bar represents

the standard error of the mean. The dotted line represents the mean bone volume in a non-ablated

calvarium.

60

5.6 Non-BMP group using 3D mineral content analysis

The Figure 5-6 represents the bone mineral content (BMC) measured at 3 and 12

weeks without the addition of rhBMP-7. The BMC measured at 3 weeks were 0.13 ± 0.02

mg, 0.15 ± 0.02 mg, 0.08 ± 0.01 mg and 0.13 ± 0.01 mg for carbide, diamond, FS and

Er:YAG group, respectively. The FS group was the only one that showed a significant

decrease in BMC formation (p < 0.05) in comparison with all other groups.

At 12 weeks, the bone mineral content increased significantly in all the

groups. The BMC at 12 weeks were 0.21 ± 0.02, 0.23 ± 0.02, 0.20 ± 0.02 and 0.24 ± 0.02

for carbide, diamond, FS and Er:YAG group, respectively. Also, there was no statistically

significant differences (p > 0.05) between the different ablative methods.

Finally, when comparing the bone mineral content within similar wounding

methods between 3 and 12 weeks, a statistically significant (p < 0.05) difference was

shown for all groups describing the bone formation process during healing phase.

61

Figure 5-6 Bone mineral content without rhBMP-7

Bar graph representing the mean comparing the bone mineral content without rhBMP-7. The FS group

showed a decrease of bone mineral formation in comparison to the other groups at 3 weeks but this

difference is not significant at 12 weeks. The error bar represents the standard error of the mean. The *

represents a statistically significant difference (p < 0.05). The dotted line represents the mean bone mineral

content in a non-ablated calvarium.

62

5.7 BMP group using 3D mineral content analysis

There were no statistically significant differences (p > 0.05) between all the

groups at either 3 weeks or 12 weeks. Finally, when comparing the bone mineral content

within similar wounding methods between 3 and 12 weeks, a statistically significant

difference (p < 0.05) was shown for all groups.

63

Figure 5-7 Bone mineral content with rhBMP-7

Bar graph representing the mean comparing the bone mineral content with rhBMP-7. Statistically

significant difference was not found between any of the groups. The error bar represents the standard error

of the mean. The dotted line represents the mean bone mineral content in a non-ablated calvarium.

64

Chapter III: Discussion

65

6 Discussion This study evaluated the efficiency of four different calvarial wounding methods

in mice with or without the addition of rhBMP-7. We sought to compare the healing of

various laser ablation versus conventional mechanical cutting to explore the future

applications for bone surgery and combined application of BMPs.

6.1 Non-BMP group using 2D analysis

After 3 weeks, the results suggested that wounds created by carbide, diamond and

Er:YAG heal at a similar rate, which is faster than the femtosecond laser created wounds.

At week 12, no statistically significant differences (p > 0.05) were observed between all

groups in terms of percent defect closure. While no statistically significant differences

were seen at week 12, one can still observe the trend noted in the previous time points at

3 weeks with the femtosecond laser group having less bone formation. When comparing

the wound closure within similar wounding methods between 3 and 12 weeks, a

statistically significant (p < 0.05) increase was shown for the diamond, ER:YAG and

femtosecond lasers. These findings suggest significant bone formation even after 3 weeks

following the initial ablation.

Healing delays have been observed using several other laser systems. This has

generally been attributed to the formation of a damage zone; a zone of carbonized

material that interfered with healing. This was seen repeatedly with CO2 laser systems

(Allen and Adrian, 1981; James et al., 1986). It is interesting to note that there is a

significant difference in bone healing when Er:YAG lasers are compared with

66

conventional CO2 lasers (Pourzarandian et al., 2004). Studies have been reported on the

healing of bone using Er:YAG lasers (Table 2.3) and most of these published results

corroborate our findings. These studies found little or no difference when comparing

Er:YAG laser and mechanical ablation groups (Buchelt et al., 1994; Lewandrowski et al.,

1996; Wang et al., 2005). However, a few studies report either delayed or improved

healing, which may be explained by different laser settings or the different animal models

that were used (Buchelt et al., 1994; de Mello et al., 2008; Mizutani et al., 2006; Nelson

et al., 1989). The literature available discussing femtosecond laser bone healing was

scarce when compared to the Er:YAG laser studies. Only one published study using

similar settings for the femtosecond laser was found to compare our results (Girard et al.,

2007a). Their results were quite similar to those reported here, as they demonstrated an

initial delay in healing with no statistically significant differences in healing after 12

weeks. Another study published in 2007 used a pulse duration of 1 picosecond (1000

femtoseconds), which is about 5 times longer than the pulse duration used for our study

(Leucht et al., 2007). Interestingly, they found a faster onset of bone healing and matrix

deposition when compared to mechanical ablation. Their results do not support our

findings; however, the results are not directly comparable as they used different laser

settings, and a different animal model. Furthermore, they did not mention the use of any

irrigation during their mechanical osteotomies, which may explain the aforementioned

results.

The exact underlying mechanisms involved in the healing delays seen in the

femtosecond laser groups in this investigation are beyond the scope of this study but

67

general assumptions can be made. The thermal damage in living tissue is actually very

minimal, or even non-existent, so it is unlikely that the previously discussed damage zone

is solely responsible for healing delays (Girard et al., 2007b). The femtosecond laser’s

ability to change the bone composition could be responsible for part of the delay in

healing. Further, since femtosecond ablation is driven by multiphoton excitation and

plasma formation, the high ionizing environment may lead to the formation of free

radicals in the surgical zone, which could potentially lead to a delay in healing. Finally,

the shock waves created by laser irradiation may cause deleterious effects on the healing

process. These avenues remain unexplored as potential explanations for healing delays in

laser bone cutting.

6.2 BMP group using 2D analysis

Wound closure was similar for all groups irrespective of the ablative technique at

3 weeks and 12 weeks. BMP seems to be effective in all groups regardless of the

wounding technique with the femtosecond laser group, demonstrating a slightly more

sluggish healing, which was not significant. That same trend appeared in the femtosecond

laser group at 12 weeks but the difference was not significant. It seems that with the use

of BMPs, the modalities of ablation do not play an important role in the healing rate. This

study supports the simultaneous use of BMP and laser ablation. With the exception of

one study published in 2007 (Girard et al., 2007a), no other data are available that

investigates the effects of laser ablation and concomitant use of BMP on bone healing.

68

6.3 Non-BMP group using 3D analysis

The 3D analysis of the specimens provided a complete measurement of wound

closure, and therefore, provided a more accurate view of the healing. The trend of FS

laser healing delay seen with the 2D analysis, was also reflected in the 3D bone volume

and bone content analysis. This difference was only statistically significant at 3 weeks for

the bone volume and the bone mineral content variables. At 12 weeks, the difference was

not significant and the ablation modalities seem to have played a limited role to influence

the rate of bone healing.

6.4 BMP group using 3D analysis

When rhBMP-7 was added to the bone defects, both the bone volume and content

kept increasing in the wound area over the 12-week period, and this increase in bone

volume was statistically significant between week 3 and 12 (p < 0.05). In other words, the

bone volume continues to increase over the 3-12 week period with the use of BMPs in a

very similar way for all the ablative modalities. A non-significant difference was seen

with the FS laser group. Again, with the exception of one study published in 2007 (Girard

et al., 2007a), no other data are available to assess the combination of laser ablation and

concomitant use of BMP on bone healing.

6.5 Future Studies and Development

Our promising results with the femtosecond laser may open new applications in

the broad clinical field of bone osteotomies. For example, a study in 2007 concluded that

both the ablation quantity and quality of cutting bone tissue with femtosecond laser

69

pulses are quite close to clinical expectations (Liu and Niemz, 2007) required for knee

joint replacement.

Techniques including analysis of the acoustic transients, fluorescence

spectroscopy or plasma spectroscopy are only some of the future developments that will

give lasers a stunning advantage over the mechanical instrumentation. A laser device able

to differentiate between tissue types and thus selectively ablate various tissue would have

tremendous implications, revolutionizing the fields of maxillofacial surgery, orthopaedic

surgery and neurosurgery. As recently as 2003, a study implemented a tissue monitoring

system where the Er:YAG laser would shut itself off after perforating the bone cortex

(Rupprecht et al., 2003). This kind of monitoring would most likely be able to reduce the

intraoperative risk. In that regard, however, the laser cutting systems used in this study

were not equipped with a feedback loop system and thus if the laser perforated the cortex,

some damage may have occurred to the dura. This could have been valuable as the mice

calvarial model is small making the assesment of the dura difficult intraoperatively. Since

the dura is responsible for osteogenic regeneration, damage to this layer may lead to

variability in healing among the laser groups, as well as other cutting techniques.

It is hoped that future laser technology will progress and the mechanical

instrumentation currently being used for bone surgery will be left as part of historical

documentation. Therefore, it is important to better understand the healing process

associated with laser ablation so that future studies will have optimal clinical impact. The

assessment of bone formation, in the presence of growth factors like rhBMP-7 following

70

laser ablation, would be part of this goal to enhance tissue healing. The molecular biology

study of bone defects may elucidate some of the healing mechanism and allow

researchers to tailor a laser with optimal setting for bone healing.

The ultimate goals of laser bone ablation should be the following:

1) Efficient ablation reducing surgical time

2) Eliminating the need for irrigation increasing the visibility in the operative field

3) Minimal injury to the tissue decreasing the layer of damaged tissue

4) Selective ablation of various tissue types

5) Promoting rapid bone formation

6) Compatibility with growth factors (BMPs or others)

7) Reduction of tissue debris during the cutting procedure

71

Chapter IV: Conclusion

72

7 Conclusion This study evaluated the wound healing closure of critical size defect in ICR mice

over a 12-week period in the presence or absence of rhBMP-7. Complementary two- and

three-dimensional analysis techniques were used to compare the healing among the

various groups. Wound closure analyses suggested that the femtosecond laser created

wounds displayed slightly healing delay in closure over the healing period when

compared to mechanical instrumentation. The amount of bone formed in the wound was

practically equivalent in both groups. The Er:YAG laser showed a healing rate similar to

that of the mechanically ablated groups. When BMPs were added to the surgical sites,

wound closure proceeded at a similar rate in all test groups. This study suggested that the

BMPs may be used with success regardless of the ablative modality utilized. While some

of the advantages of using lasers over mechanical ablation have not yet been identified

with the current technology, others are well established. It is worth emphasizing that no

irrigation was used when cutting bone with our femtosecond laser-wounding group as

compared to copious amounts that are required to minimize thermal damage using

mechanical ablation. By not using irrigation, the operator of the femtosecond laser has

the advantage over mechanical instrumentation of increased visibility in the operative

field. Finally, since a laser has no moving parts that cut bone, there is no risk of

entangling the adjacent soft tissue. This would provide an extra measure of safety,

particularly when working in tight spaces, such as would be seen during bone ablation for

mandibular sagittal split osteotomies where the infra-alveolar nerve is adjacent to the

bony cuts. In summary, femtosecond and Er:YAG lasers are two modalities suitable for

bone ablation comparable to mechanical instrumentation.

73

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