Bonding Between Metals and Polymers for Dental Devices

Omar Saleh Alageel

Faculty of Dentistry, McGill University

Montreal, Canada

December, 2013

A thesis submitted to McGill University in partial fulfillment of the requirements of the degree

of M.S.c in Dental Sciences.

© Omar Alageel 2013

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Table of Content

Dedication ...................................................................................................................................... 4

Acknowledgements ........................................................................................................................ 5

List of Figures ................................................................................................................................ 6

List of Tables ................................................................................................................................. 8

Abbreviations ................................................................................................................................. 9

Abstract ........................................................................................................................................ 10

Résumé ......................................................................................................................................... 12

Chapter 1: Introduction ............................................................................................................ 14

Chapter 2: Background and Literature Review ..................................................................... 16

2. Esthetics and Dental Occlusion .................................................................................. 16

2.1. Edentulism ................................................................................................................ 16

2.1.1. Treatment of Edentulous Patients .......................................................................... 17

2.1.1.1. Treatment of Partially Edentulous Patients ......................................................... 18

2.1.1.1.1. Fixed Partial Dentures ...................................................................................... 18

2.1.1.1.2. Dental Implants ................................................................................................ 19

2.1.1.1.3. Removable Partial Dentures ............................................................................ 21

2.1.1.2. Treatment of Completely Edentulous Patients ................................................... 23

2.1.1.2.1. Removable Complete Dentures ....................................................................... 24

2.1.1.2.2. Fixed Complete Dentures ................................................................................ 25

2.2. Malocclusion ............................................................................................................. 26

2.2.1. Treatment of Malocclusion .................................................................................... 26

2.2.1.1. Fixed Appliances ................................................................................................ 27

2.2.1.2. Removable Appliances ....................................................................................... 28

2.3. Materials Used in Dentures and Orthodontic Devices .............................................. 29

2.3.1. Metals ..................................................................................................................... 29

2.3.1.1. Titanium .............................................................................................................. 30

2.3.1.2. Cobalt-Chromium ............................................................................................... 31

2.3.1.3. Stainless Steel ..................................................................................................... 32

2.3.2. Ceramic .................................................................................................................. 32

2.3.3. Polymers and Composites ...................................................................................... 33

2.3.3.1. PMMA ................................................................................................................ 34

2.3.3.2. Bis-GMA ............................................................................................................ 37

2.4. Bonding Systems in Dentures and Orthodontic Appliances ..................................... 39

2.4.1. Mechanical Bonding .............................................................................................. 39

2.4.2. Chemical Bonding ................................................................................................. 40

2.5. Debonding in Dentures and Orthodontics Appliances…........................................... 41

2.5.1. Bonding between Alloys and PMMA .................................................................... 42

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2.5.2. Bonding between Wrought Wire and PMMA ....................................................... 43

2.5.3. Bonding between Brackets and Composite ........................................................... 45

2.6. Aryldiazonium Salts .................................................................................................. 46

2.6.1. Grafting of Diazonium Salts .................................................................................. 47

2.6.2. Diazonium Grafted Layer Properties ..................................................................... 48

2.6.3. Applications of Aryldiazonium Salts ..................................................................... 49

2.6.4. Aryldiazonium Salts as Dental Adhesive .............................................................. 50

2.6.5. Diazonium Grafted Layer Analysis ....................................................................... 50

2.6.5.1 X-ray Photoelectron Spectrometer (XPS) ............................................................ 51

2.6.5.2 Contact Angle Measurement ................................................................................ 54

Chapter 3: Hypothesis and Objective ...................................................................................... 55

3.1. Hypothesis.................................................................................................................. 55

3.2. Thesis Objective......................................................................................................... 55

Chapter 4: List of References ................................................................................................... 56

Chapter 5: Manuscript I: Bonding Metals to Poly-Methyl Methacrylate Using

Aryldiazonium Salts ...................................................................................................... 62

5.1. Abstract ..................................................................................................................... 62

5.2. Introduction................................................................................................................ 63

5.3. Materials and Methods .............................................................................................. 67

5.4. Results and Discussion.............................................................................................. 71

5.5. Conclusion................................................................................................................. 80

5.6. References.................................................................................................................. 81

Chapter 6: Manuscript II: Surface Chemical Treatment of Orthodontic Brackets for

Improved Tooth Adhesion ............................................................................................ 85

6.1. Abstract ..................................................................................................................... 85

6.2. Introduction................................................................................................................ 86

6.3. Materials and Methods .............................................................................................. 89

6.4. Results........................................................................................................................ 94

6.5. Discussion ................................................................................................................. 97

6.6. Conclusion .............................................................................................................. 101

6.7. References ............................................................................................................... 102

Chapter 7: Conclusion.............................................................................................................. 105

Chapter 8: Appendices ............................................................................................................ 106

8.1. Report of Invention ................................................................................................. 107

8.2. Poster I .................................................................................................................... 108

8.3. Poster II ................................................................................................................... 109

8.4. Poster III .................................................................................................................. 110

8.5. Poster IV & V.......................................................................................................... 111

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Dedication

I dedicate my thesis to my parents, brothers, and sisters for their endless encouragement

and support throughout the course of this thesis. Also, I dedicate this thesis to my

wonderful and supportive wife and to my beautiful boy.

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Acknowledgements

First of all I would like to thank Allah Almighty for providing me blessings, help, and courage to

accomplish this thesis and achieve my desired goals.

I would like to express my genuine appreciation to my supervisor Dr. Faleh Tamimi for his

encouragement, supervision, support, and immense knowledge. I am grateful for his generous

guidance from the initial to the final stage of this project. Also, I would thank enormously my

co-supervisor Dr. Marta Cerruti for her suggestions, feedback, and facilities throughout the

completing of my thesis.

In addition, thanks to all my colleagues and friends especially Mohamed-Nur abdalla and Hazem

Eimar for their constant and unconditional help to complete this project. Thanks to Dr. Jean-

Marc Retrouvey, Dr. Rubens Albuquerque and Paige Kozak for their help and guidance.

Finally, I would like to thank King Saud University, Saudi Arabia for the scholarship and the

grant to complete this research.

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List of Figures

Figure 2.1: Replacing missing teeth using fixed partial denture (Bridge).

Figure 2.2: Dental implant components for the single-unit fixed prosthesis.

Figure 2.3: The mandibular removable partially denture (RPD).

Figure 2.4: Complete partial dentures for maxillary and mandibular arches.

Figure 2.5: The implants-supported complete denture (overdentures).

Figure 2.6: Fixed complete dentures that supported on several dental implants.

Figure 2.7: Fixed appliance for malocclusion treatment.

Figure 2.8: Removable appliance for malocclusion treatment (retainers).

Figure 2.9: Scheme of the polymerization reaction of PMMA.

Figure 2.10: Schematic diagram of the chemical reaction for Bis-GMA.

Figure 2.11: Scheme describing grafting of diazonium salts of a substrate.

Figure 2.12: Diagram describing X-ray photoelectron spectroscopy (XPS) components.

Figure 5.1: The custom-made silicone mold used to prepare the PMMA-Ti specimen; and

mechanical test specimen before and after mechanical testing.

Figure 5.2: Scheme depicting reaction sequence performed in first (primer) and second

(adhesive) steps.

Figure 5.3: XPS surveys; elemental compositions; high resolution C 1s spectra on Ti surface for

different groups.

Figure 5.4: Photographs of water droplets placed on different Ti groups.

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Figure 5.5: Tensile strength of the bond between PMMA and treated Ti surfaces.

Figure 5.6: Bond strength of PMMA and stainless steel wires for the control group and

that were treated with diazonium.

Figure 5.7: Drawing shows the acrylic removable partial denture with small volume of

PMMA to support the wire.

Figure 6.1: Photographs showing bracket debonding at the interface between brackets and

composite.

Figure 6.2: Digital photographs illustrating the different types of brackets used in this study; and

schematic drawing of preparation and mechanism of testing the tensile and shear bonding

strength between brackets and Bis-GMA.

Figure 6.3: Schematic diagram of the reactions performed in the first and second solutions of

Bis-GMA/diazonium treatment.

Figure 6.4: XPS general surveys and the elemental compositions for the untreated and treated

brackets.

Figure 6.5: The ultimate tensile force N and bond strength MPa of the different stainless steel

brackets. The ultimate shear force N and bond strength MPa of the different stainless steel

brackets.

Figure 6.6: Photograph of untreated L bracket group and treated S bracket group that were

bonded to the anterior teeth using the adhesive based on Bis-GMA; and drawing shows fixed

orthodontic appliances with different size of brackets.

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List of Tables

Table 2.1: The bond strengths between titanium and PMMA (MPa) by using different bonding

methods.

Table 2.2: The bond strengths (MPa) between PMMA and cobalt –chromium or stainless steel

using different bonding methods.

Table 5.1: The bond strengths between titanium and PMMA (MPa) using different bonding

methods.

Table 5.2: Conditions tested in the second step. The overall solution volume was 12 ml, and was

water-based.

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Abbreviations

AFM ....................................................................................................... Atomic force microscopy

Bis-GMA ................................................................................ Bisphenol A glycidyl methacrylate

CAD/CAM .........................................Computer Aided Design / Computer Aided Manufacturing

Co-Cr ................................................................................................................. Cobalt-Chromium

CP...................................................................................................................... Commercially Pure

DFT ......................................................................................................... Density functional theory

DMLS................................................................................................. Direct Metal Laser Sintering

HCl ..................................................................................................................... Hydrochloric acid

HNO3 .............................................................................................................................. Nitric acid

H3PO2 ......................................................................................................... Hypophosphorous acid

H2SO4 .......................................................................................................................... Sulfuric acid

HEMA ................................................................................................. Hydroxyethyl methacrylate

MMA .............................................................................................................. Methyl methacrylate

µm ................................................................................................................................. Micrometer

MPa .............................................................................................................................. Megapascal

NaNO2 ..................................................................................................................... Sodium nitrite

N .......................................................................................................................................... Newton

PAP................................................................................................................ Polyaminophenylene

pH ....................................................................................................................................... Acidity

PMMA ................................................................................................... Poly-methyl methacrylate

PPD ................................................................................................................. p-phenylenediamine

RPD ...................................................................................................... Removable partial denture

TEGDMA .................................................................................. Triethyleneglycol-dimethacrylate

Tg ........................................................................................................ Glass transition temperature

Ti ....................................................................................................................................... Titanium

ToF-SIMS .......................................................... Time-of-flight secondary ion mass spectrometry

XPS ............................................................................................ X-ray photoelectron spectroscopy

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Abstract

Many dental devices combine acrylic (i.e. poly-methyl methacrylate or bisphenol A-glycidyl

methacrylate) and metallic parts (i.e. titanium or stainless steel) that are bonded together. These

devices often present catastrophic mechanical failures due to weak bonding between their acrylic

and metallic components. These devices include dental prostheses, combining metallic

frameworks (i.e. titanium) and wrought wires with acrylic resin; and orthodontic appliances,

combining acrylic resin with stainless steel wrought wires or composite with stainless steel

brackets. The bonding between metals and polymers in dental devices is usually performed by

the mechanical interlocking, but its bond strength is still too low for dental applications. The

bond strength between them would be high if the chemical bonding, which does not occur

spontaneously, uses in addition to the mechanical interlock. The objective of this study was to

develop a new method of creating a strong chemical bond between alloys and polymers for

dental devices based on diazonium chemistry.

The chemical bond between metals (i.e. titanium or stainless steel) and polymers (i.e. poly-

methyl methacrylate, PMMA or Bisphenol A-glycidyl methacrylate, Bis-GMA) was achieved in

two steps. In the first reaction step (primer), the aryldiazonium salts were chemically reduced to

form aryl radicals which spontaneously got grafted onto the metallic surfaces. The second step of

the reaction (adhesive) was optimized to achieve covalent binding between the grafted layer and

PMMA or Bis-GMA. The chemical composition of the treated surfaces was analyzed with X-ray

photoelectron spectroscopy (XPS), and the bonding strengths between alloys and PMMA or Bis-

GMA were measured.

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XPS characterization and contact angle measurement confirmed the presence of a polymer coat

on the treated metallic surfaces. Whereas, the mechanical test results showed a significant

increase of the tensile bond strength between PMMA and treated titanium or stainless steel wire

by 5.2 and 2.5 folds, respectively, compared to the untreated control group (P<0.05). Moreover,

the bonding strength between metallic brackets and Bis-GMA composite was increased after the

treatment depending on the bracket design by 2 to 3.9 folds compared to untreated brackets.

Diazonium chemistry provides an effective way of achieving a strong chemical bond between

alloys and PMMA or Bis-GMA. The resulting bonding method can be utilized to further improve

the properties of dental devices, reduce debonding of dental prostheses and brackets, provide

more leverage in orthodontic cases with complex mechanics, and allow the use of brackets with

smaller bases.

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Résumé

De nombreux appareils dentaires sont composés d'acrylique (c'est à dire d'un poly -méthacrylate

de méthyle ou de bisphénol A- glycidyle méthacrylate) et de parties métalliques (par exemple en

titane ou en acier inoxydable) qui sont collés ensemble. Ces dispositifs présentent souvent des

défaillances mécaniques catastrophiques en raison de la faiblesse de la liaison entre les

composantes en acrylique et celles en métal. Ces dispositifs comprennent les prothèses dentaires,

alliant des cadres métalliques (c’est à dire de titane) et fils forgé avec de la résine acrylique, et

les appareils orthodontiques, combinant de la résine acrylique avec des fils forgé en acier

inoxydable ou un composite avec des supports en acier inoxydable. La force de liaison entre eux

serait élevée si la liaison chimique, ce qui ne se produit pas spontanément, est utiliser en plus du

verrouillage mécanique.

Dans la première étape de la réaction, les sels d’aryl diazonium sont réduits chimiquement pour

former des radicaux aryles qui sont spontanément greffés sur les surfaces métalliques La

deuxième étape de la réaction a été optimisée pour réaliser la liaison entre la couche greffée et le

PMMA ou le Bis-GMA. La caractérisation XPS et la mesure de l'angle de contact a confirmé la

présence d'une couche de polymère sur les surfaces métalliques traitées. Les résultats des essais

mécaniques ont montré une augmentation significative de la force d'adhérence à la traction entre

le PMMA et le titane traité ou d'un fil en acier inoxydable de 5,2 et 2,5 plis, respectivement, par

rapport au groupe témoin non traité (p < 0,05).

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La chimie de diazonium fournit un moyen efficace d'atteindre une liaison chimique forte entre

les alliages et le PMMA ou le Bis-GMA. Le procédé de collage qui en résulte peut être utilisé

pour améliorer les propriétés des appareils dentaires, réduire le décollement de prothèses

dentaires et des supports, et permettre l'utilisation de supports avec des bases plus petites.

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Chapter 1: Introduction

Dental devices such as dental prostheses and orthodontic appliances are commonly used for

treatment of dental problems such as edentulism and malocclusion [1-6]. Dental prosthesis is an

artificial device used to replace natural teeth for partially or completely edentulous patients.

Dental prostheses such as fixed partial dentures, metal-cast removable partial dentures, and all-

acrylic removable partial dentures are the most common treatment options for partially

edentulous patients while the removable complete dentures are the most common prostheses for

completely edentulous patients [1-5, 7-11]. Furthermore, fixed and removable orthodontic

appliances are the most common treatment methods for malocclusion patients [3, 4].

Dental devices usually are a combination of polymeric (i.e. poly-methyl methacrylate and

bisphenol A-glycidyl methacrylate) and metallic parts (i.e. titanium, cobalt-chromium, and

stainless steel) that are bonded together. The bonding between metals and polymers in dental

devices is usually preformed by mechanical and/or chemical bonds. The metallic framework (i.e.

titanium and cobalt-chromium) in the removable partial dentures (RPD) is usually bonded to

poly-methyl methacrylate (PMMA) by interlocking the PMMA into the irregularities of the

metals that can be prepared by creating small retentions or sandblasting the metals substrates [11,

12]. The metallic wrought wire in the all-acrylic removable partial dentures and removable

orthodontic appliances is usually formed in a zig-zag configuration to provide retention in

PMMA denture base. Orthodontic brackets were developed with a large base designed to

increase the surface area and compensate for the lack of adhesion between brackets and

15

bisphenol A-glycidyl methacrylate (Bis-GMA) composite; however, large brackets have a

negative effect on patient satisfaction and oral health [13-17].

Dental devices often present catastrophic mechanical failures due to lack of bonding between

their acrylic and metallic components leading to prostheses failures and brackets loss [18-21].

The bonding between alloys and polymers in dental devices can be improved using strong

chemical bond (adhesives), which does not occur spontaneously, in addition to the mechanical

interlock. There are several dental adhesives that can be used between metals and polymers for

dental prostheses and orthodontic devices, but the bonding strength reported so far is insufficient

[20, 22-24].

This research provides a new way of creating a strong chemical bond between alloys (i.e.

titanium and stainless steel) and polymers (i.e. PMMA and Bis-GMA) for dental devices based

on diazonium chemistry. This new adhesive can be used on titanium or stainless steel surfaces to

increase their bonding strength to PMMA or Bis-GMA composite. Increased bond strength

between alloys and polymers through diazonium treatment would improve the properties of

dental devices, reduce debonding between alloys and polymers in dental prostheses and

orthodontic brackets, provide more leverage in cases with complex mechanics, and allow the use

of brackets with smaller bases resulting in fewer complications associated with esthetics and oral

hygiene.

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Chapter 2: Background and Literature Review

2. Esthetics and Dental Occlusion

Esthetics and well aligned teeth contribute to a healthy masticatory function, pleasant smile, and

adequate phonetics [2]. Loss of teeth (edentulism) and malocclusion can have a negative impact

on patients’ self esteem and masticatory function [6]. These negative impacts can be improved

with dental prostheses which replace missing teeth or orthodontic devices that correct tooth

malocclusion [1-6].

2.1. Edentulism

Edentulism refers to patients missing some or all their natural teeth. A person is completely

(fully) edentulous when missing all the teeth, or partially edentulous when missing some but not

all the teeth. Although the rates of edentulism vary in the world, the number of people who are

completely or partially edentulous is large [25]. It is estimated that 15 % of the global population

is completely edentulous [7]. In particular, 12% of the population in the United States is

completely edentulous which equal to 36,000,000 people [26]. Moreover, 71.5% of USA

population between age 65 and 75 years old is partially edentulous [27]. In Canada, the

population of the completely edentulous patients in 2010 was 6.4% of the whole country and

21.7 % among adults between 60 and 79 years old [25]. The rate of complete and partial

edentulism in the population increase among elderly people, and it will increase in the future due

to aging of the population [27]. Currently, the percentage of elderly people in the United States is

13%, but it is expected to double by the year 2030 [27].

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The incidence of tooth loss in some developed countries is 0.1 to 0.3 tooth for every person per

year [1]. Tooth loss is associated with age, socioeconomic status and lifestyle. The number of

missing teeth is high among elderly people (age 65 years and older) while poor oral hygiene,

tobacco and alcohol consumptions are the most common risk factors for missing teeth [1, 27].

Education, access to dental care, and insurance coverage are the other common factors

contributing loss of teeth [25].

Edentulism has serious of consequences on patents’ general and oral health as well as on quality

of life [25]. Tooth loss contributes to anatomical changes of the mouth and face [25]. For

instance, the bone and residual ridge are expected to shrink after tooth loss affecting patients’

esthetic and challenging future treatments [27]. Having unacceptable aesthetics is the most

common concern for edentulous patients especially when the missed teeth are in the visible

anterior region. Furthermore, patients’ phonetic usually change after tooth loss because the

contacts between the maxillary and mandibular teeth are changed. Losing the ability of chewing

food effectively is the major functional issue; In fact, chewing food effectively involves

subdivision of food by the occlusal force of the teeth, and bringing food into the occlusal surface

of the teeth by the oral tissues including the tongue and the cheek [25, 27].

2.1.1. Treatment of Edentulous Patients

Edentulous patients’ treatment depends on many factors, such as number of missing teeth, oral

tissue structure, and patient’s preferences. The goal of the treatment is to restore function,

esthetics and phonetics using a prosthesis that cooperates with the existing natural teeth and

tissues [11]. There are many treatment options available to replace missing teeth [1-5, 8-11]. In

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this introduction, the treatment options will be discussed for the partially edentulous patients first

and then for the completely edentulous patients.

2.1.1.1. Treatment of Partially Edentulous Patients

The common treatment options for the partially edentulous patients are fixed partial dentures

supported by natural teeth or dental implants, or removable partial dentures retained by natural

teeth or dental implants [1, 5, 8, 10].

2.1.1.1.1. Fixed Partial Dentures

One or more missing teeth can be replaced with fixed partial dentures such as bridges. Bridges

are prostheses that replace missing teeth by anchoring on the teeth or implants adjacent to the

missing teeth (Figure 2.1) [1]. Tooth supported bridges (conventional bridges) are the most

common treatment because of their cost and time needed for completion. Bridges can also

replace two or three adjacent missing teeth according to the edentulous span (the length of the

arch where teeth are missing), occlusal stress, and health of the remaining natural teeth [4, 28].

The principle of bridge prosthesis is preparing the natural teeth adjacent to the edentulous span

as abutments to support the artificial teeth. The bridge (Figure 2.1) usually consists of at least

two retainers which are copings supported on natural tooth abutments, and one or more units

connected to the retainers, called pontics, that replace the missing teeth [1, 4, 28].

Bridges are usually fabricated with metals (full metal), ceramics (full ceramic) or a combination

of both (metal-ceramic restorations) [28]. Full metal bridges (i.e. gold or cobalt-chromium) have

the best mechanical performance but their esthetics is unacceptable especially when the missing

teeth are in the anterior region [1, 4]. The metal-ceramic bridge is the common type, and it

19

consists of a metal framework and a ceramic veneer that is built up over the metal [1, 4]. The

function and esthetics provided by metal-ceramic and full ceramic bridges are excellent and

comparable to that of natural teeth. However, there is a major disadvantage for using bridges that

is the need for trimming the adjacent natural teeth to allow attachment of the bridge [4, 28].

Figure 2.1: Replacing missing teeth using a fixed partial denture (Bridge). A: the prepared tooth for anchoring the bridge

(abutments), B: retainers, C: pontic.

2.1.1.1.2. Dental Implants

A dental implant is a biocompatible device placed in the jawbone to provide support and

retention for the artificial dental teeth and prostheses [29]. Dental implants are not new; in

ancient history Egyptians shaped seashells and inserted them into the jaw [11]. However, the

modern dental implant concept begun in 1952 when Dr. Per-Ingvar Branemark accidentally

discovered that living bone interacts and binds to titanium and this property became to be known

as osseointegration [4, 11, 30]. The first titanium dental implant used to replace a missing tooth

was done in 1965, and since then the application and market for dental implants has been

continuously growing [4, 30]. Dental implants have been subject to significant improvements in

materials and design that raised the success rates for dental implants and make them a very

popular choice for replacing missing teeth [4, 11, 31]. It is estimated that more than 450,000

dental implants are being placed every year [30].

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Osseointegration is the most important factor for the success of dental implants treatment. There

are many factors affecting osseointegration that should be considered before treatment such as

age of the patient, bone quality, smoking, and alcohol consumptions [10]. Dental implants

(Figure 2.2) can be used to support or retain different prosthesis types such as crowns, bridges,

removable partial dentures, and complete dentures [1, 4].

Treatments of edentulous patients with implant-retained dental prostheses eliminate many of the

disadvantages associated with traditional dental prostheses [1, 4, 11]. The traditional removable

partial dentures or complete parietal dentures rely on oral tissues to hold the denture resulting in

discomfort, while implant supported or retained dentures have superior function and comfort for

patients than traditional removable partial dentures. Moreover, using implants to support or

retain dentures are more esthetic than using unesthetic metallic clasps in the traditional

removable partial dentures which located in the labial or lingual sides of teeth [7, 8]. Replacing

one or more teeth using bridge treatment require to trim down the adjacent natural teeth next to

the edentulous span that are not preferred for many patients and it can be avoided by using dental

implants [1].

The most metals used in dental implants are the pure titanium (grade I to IV) and titanium alloys,

such as Ti-6Al-4V [4]. The main components of the dental implants are the implant cylinder

(root), which is inserted inside bone, and the abutment, which retains or supports the dental

prosthesis (Figure 2.2) [4, 10]. The abutment is usually screw-fastened onto the implant’s root.

Dental implants vary according to the design and size of the cylinder and implant-abutment

connection [10]. Implants roots are usually cylindrical in shape with lengths between 6 to 20 mm

and diameters between 3 to 6 mm [10]. The surface of the implants root is treated via many

21

techniques to increase the bonding between implants and bone (osseointegration) including

sandblasting, etching, and spray coating [4, 10].

Figure 2.2: Dental implant components for the single-unit fixed prosthesis. A: implant cylinder (root); B: screw-fastened

implant abutment and; C: artificial crown fixed on the abutment of the dental implant.

2.1.1.1.3. Removable Partial Dentures

A removable partial denture (RPD) is a prosthesis that replaces one or more missing teeth, and is

supported and retained by the remaining natural teeth, tissue, and/or implants. Removable partial

dentures provide high function and pleasing esthetics and they are designed to be removed and

reinserted by the patient [4]. Removable partial dentures are suitable for partially edentulous

patients who are not able to have fixed prostheses because of their health conditions or the length

and location of the edentulous span [7].

The design of removable partial dentures depends on the number and location of missing teeth

and on the health of oral tissues and natural teeth that will support and retain the prosthesis [9].

Partially edentulous arches can be classified into many different classifications according to the

number and location of the missing teeth. The classification named Kennedy is the most

accepted because it is simple and easy to apply; it divides the partial edentulous arches into four

groups [7-9, 32]. The most common groups of Kennedy classification are class I and class II.

Class I defines bilateral edentulous areas located posterior to the remaining teeth while class II

defines unilateral edentulous areas located posterior to the remaining teeth [32]. Class III of

22

Kennedy classification defines tooth bounded unilateral edentulous areas while class IV indicates

a single edentulous area located anterior to the remaining natural teeth and crossing the midline

[32]. Removable partial dentures design depends on the dental arch classification; for example,

the RPD’s design in Kennedy class I and II rely both on teeth and oral tissues for support and

retention of the prosthesis while in class III and IV support and retention is provided solely by

the remaining teeth [7].

The typical removable partial denture, cast-metal RPD, consists of a metal framework, artificial

teeth, and an acrylic denture base. The metal framework consists of four parts (Figure 2.3): major

connectors, minor connectors, direct retainers, and indirect retainers [4, 5, 7, 33]. A major

connector is used to connect all the main parts of the prosthesis and helps distribute the occlusal

force into selected teeth and tissues while a minor connector is used to connect the major

connector or the denture base to other components such as clasps, rests, direct and indirect

retainers [5]. A direct retainer is a component that engage a tooth to provide retention and resist

movements away from the oral tissues and natural teeth, and it can be intracoronal or

extracoronal [33, 34]. Clasps are the most common extracoronal retainers and they usually

consist of a lingual arm, a buccal arm, and a rest, and they are used in different designs including

circumferential clasps, bar clasps, ring clasps, and roach clasps [5, 33]. The rest provides vertical

support and it can be located on the occlusal, lingual, or incisal tooth surface. Indirect retainers

assist the direct retainers to prevent rotation or displacement of distal extensions of the denture

and it is usually composed of a rest [35]. The removable partial denture (RPD) metal framework

is connected to the acrylic-resin denture base (poly-methyl methacrylate; PMMA) and to the

acrylic teeth [36]. The main materials used in the metal framework of cast-metal removable

partial dentures are cobalt-chromium, gold, or titanium.

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Figure 2.3: (Left) the mandibular partially edentulous arch before treatment. (Middle): the mandibular framework for

the removable partial denture and its components of A: major connector; B: minor connector; C: rests; D: direct

retainer; E: indirect retainer. (Right) the removable partial denture (RPD) on the edentulous arch.

An alternative version of the RPDs is the acrylic RPD. The acrylic RPDs, known as temporary

RPDs, are made of an all-acrylic base (poly-methyl methacrylate; PMMA), acrylic teeth, and

wrought wire clasps for retention. The acrylic RPDs are easy to fabricate and less expensive than

the cast-metal RPDs and its esthetic is acceptable [4]. However, the acrylic RPDs are consider

temporary prostheses and not recommended for long-term prosthesis because of their poor

mechanical properties such as strength [4].

Dental implants can be used to provide support and retention for removable partial dentures [4,

8]. Implant-retained removable partial dentures are similar to conventional removable partial

dentures supported by natural teeth and tissue, although they also gain additional support and

retention from the dental implants. Implant-retained RPD provide better function, esthetics and

comfort than conventional removable partial dentures [8].

2.1.1.2. Treatment of Completely Edentulous Patients

The common available treatments for complete edentulism are removable complete dentures and

fixed complete dentures.

24

2.1.1.2.1. Removable Complete Dentures

Complete dentures have been considered the standard treatment option for complete edentulism.

Complete dentures are prostheses that replace all missing teeth in the maxillary or mandibular

arch (Figure 2.4). Complete dentures are supported and retained by the oral tissues and mucous

membranes when the anatomy and the functional tonicity of the patients’ mouth are adequate [7,

26]. Removable complete dentures generally consist of a denture base, made of acrylic (poly-

methyl methacrylate; PMMA) or metal, and artificial teeth, made of acrylic or composite resin

[4, 7]. Removable complete dentures provide acceptable esthetics at a reasonable cost compared

to implant-supported prosthesis; however, several issues are associated with complete dentures

such as low denture stability especially in the lower arch [4, 7, 26]. Moreover, complete dentures

need to be changed or refitted every few years due to the shrinkage and changes in bone and

supporting tissues [4].

Figure 2.4: Complete partial dentures for maxillary and mandibular arches.

Overdentures are complete dentures retained by dental implants. This type of dentures solves

main problem associated with complete dentures that is the lack of stability in the lower jaw in

patients with severe alveolar ridge atrophy [4]. There are two major types of overdentures;

25

removable overdentures and fixed overdentures. Removable overdentures are similar to

conventional complete dentures, but they are retained by two to four dental implants using clip-

bar, spheres, or magnetic attachments (Figure 2.5) [4]. The removable overdentures provide high

stability and retention in the mouth and are less expensive than the alternative implant-supported

prosthesis. However, this type of prostheses must be removed daily for cleaning [4, 7, 26].

Figure 2.5: The implants-supported complete denture (overdentures); two-implants support the mandibular removable

overdentures.

2.1.1.2.2. Fixed Complete Dentures

Fixed overdentures are complete dentures retained and supported directly by dental implants.

Fixed complete dentures (Figure 2.6) consist of a denture-base, made of acrylic or metal, and

teeth, made of acrylic or porcelain, fixed on dental implants [4]. This prosthesis is the best

available treatment for edentulous patients because it provides superior esthetics, comfort, and

function. However, fixed overdentures have several disadvantages such lengthy cost, difficult

fabrication process [4]. Nevertheless, fixed complete dentures are considered the most

recommended treatment for edentulous patients [4].

26

Figure 2.6: Fixed complete dentures supported directly on several dental implants.

2.2. Malocclusion

Malocclusion can be described as a significant deviation from the ideal occlusion, and it is one of

the most common oral disorders [3, 4]. Many factors are involve in defining occlusion such as

the size of maxillary and mandibular arches, size and number of present teeth, and the activity of

lip and tongue [3, 4]. Malocclusion is classified into four classes according to the relation

between upper and lower teeth [3]. The most common malocclusion type is class I which occurs

when the upper and lower teeth bite is normal but teeth are excessively crowded [3]. Class II

malocclusion (overbite) occurs when the upper teeth overlap the lower teeth and class III

malocclusion (underbite) occurs when the lower teeth overlap the upper teeth [3, 4]. In most

cases, malocclusion patients have unacceptable appearance, masticatory discomfort, and speech

difficulties [3, 4].

2.2.1. Malocclusion Treatments

The treatment of malocclusion relies on relieving the crowding, straightening the teeth, and

closing the open spaces, as well as modifying craniofacial growth [3, 4, 37]. Teeth crowding is

usually treated with tooth extractions and subsequent use of fixed (bracket) or removable

27

orthodontic appliances for tooth alignment and space closure [4, 37]. The cases that require

growth modification of the jaws are treated with fixed or removable functional appliances [3,

37]. Finally, removable appliances (retainers and space maintainers) are used to hold the teeth in

their position after the orthodontic treatment is finished.

2.2.1.1. Fixed Appliances

Fixed orthodontic appliances are appliances that are fitted and fixed on the teeth and cannot be

removed by patients. Fixed appliances move teeth from malaligned positions to the correctly

aligned ones. This treatment option can be used to treat most malocclusions; however, they have

many disadvantages in terms of esthetics, oral hygiene, and cost. Fixed appliances usually

consist of bands, metallic wires, and brackets that are cemented to teeth through an adhesive, as

well as other active components such as springs, elastics, and separators (Figure 2.7) [3, 4, 37].

The wires are made of alloys such as stainless steel, cobalt-chromium, cobalt-chromium-nickel,

and titanium [11, 16]. Brackets are made of different materials such as ceramics or plastics,

although most of them are made of metallic alloys (i.e. stainless steel and cobalt-chromium) [16,

37]. Brackets are designed with bases that provide micromechanical interlocking to improve the

bonding between brackets and composite for better adhesion to teeth [17, 38, 39]. The brackets

are usually bonded bucally or lingually to teeth using adhesives based on methacrylate composite

such as Bis-GMA, TEGDMA, and HEMA [3, 4, 16, 37].

28

Figure 2.7: Fixed appliance for malocclusion treatment that consists of A: brackets and B: metallic wires.

2.2.1.2. Removable Appliances

Removable orthodontic appliances can be removed by patients for cleaning or on the social

sensitive occasions. They are less expensive and their fabrication consumes less chair-side time,

but they are not recommended for treatment of complex cases [4]. Removable orthodontic

appliances have different designs and functions and they usually consist of an acrylic baseplate

as well as active and retention components [3, 37]. Removable appliances are commonly used to

maintain tooth position after treatment (i.e. retainers) (Figure 2.8), as well as for moving, tipping,

and titling teeth using active components such as springs, screws, or bows [4, 37, 40] Retention

components, such as clasps, and most active components are fabricated from metallic wires (i.e.

stainless steel) bending to the desire shape and embedded to the acrylic baseplate [4, 40]. The

baseplate provides stability and support for the active and retentive components and is usually

made of poly-methyl methacrylate (PMMA) [37, 40].

29

Figure 2.8: Removable appliance (retainers) that consists of A: metallic wires (retention components) and B: acrylic

denture base and used to maintain teeth after treatment.

2.3. Materials Used in Dentures and Orthodontic Devices

The materials used for fabricating dentures and orthodontic appliances can be classified into four

categories: metals, ceramics, polymers, and composites [11, 41]. These categories are different

from each other in terms of their physical and mechanical properties, processing methods, and

cost [11, 41]. Metals are primarily used in dental prostheses and orthodontic devices when

durability and strength are required while ceramics are used when esthetics is important [42].

Polymeric materials are commonly used because they combine excellent esthetic and mechanical

properties at a reasonable cost [42]. Underneath we address in detail each one of these

categories.

2.3.1. Metals

Pure metallic elements alone have inadequate properties for dental applications; thus, dental

alloys that combine various elements suitable to be used in dental prostheses [41]. Generally,

dental alloys should fulfill many different criteria including biocompatibility, corrosion

resistance, strength, hardness, melting temperature, and economic aspects that are useful for

dental applications [11, 42]. Alloys used in dental devices are either laser sinters machines,

30

casted into customized shapes or with wrought wires [11]. Casting alloys used for dental

prostheses are divided into high noble alloys, noble alloys, and base-metals alloy [41, 42]. Alloys

are considered high-noble when more than 60% of their composition is a noble metals such as

gold, platinum, and palladium; noble if the noble metals is 60 to 25%; or base-metal when noble

content is less than 25% [41]. Base-metal alloys are used extensively in all dental prostheses

because of their excellent mechanical properties and low cost. Dental alloys have different

physical, chemical, and biological properties based on their elemental compositions [41, 42].

Underneath we address in detail the main groups of alloys used in dentistry.

2.3.1.1. Titanium

Titanium (Ti) raises great interest in dentistry due to its excellent properties; it is highly

biocompatible which being significantly less expensive than noble metals, such as gold [41-45].

Titanium has excellent mechanical and physical properties, such as high strength and low

density, that helps to withstand the mastication force which make it more comfortable for

patients [42]. Moreover, titanium has low modulus and thermal conductivity with good chemical

stability and corrosion resistance [42-46]. Titanium alloys (i.e. Ti-6Al-4V) are widely used in

dental implants and wrought wires for dental prostheses and orthodontic appliances, and has

recently raised interest as materials for dentures frameworks [41]. However, titanium alloys are

difficult to cast because they require special and expensive furnaces due to their high melting

point; therefore, the use of titanium in casted frameworks for removable and fixed prostheses is

limited [41, 46]. Milling systems, such as Computer Aided Design / Computer Aided

Manufacturing (CAD/CAM), enables the use of titanium in removable and fixed prostheses

frameworks that can be designed through computer software and milled in a machine [41].

Moreover, the Direct Metal Laser Sintering (DMLS) is a new technology that can be used to

31

produce fixed or removable prosthesis frameworks of Ti [41]. DMLS works by applying a high-

power laser to fuse many layers the powdered metal (i.e. Ti and Co-Cr) building up the desired

three-dimensional frameworks [41].

2.3.1.2. Cobalt-Chromium

Removable partials frameworks and metal bases of ceramic-metal restorations such as crowns

and bridges are commonly made of cobalt-chromium alloys (Co-Cr) [41, 46]. Cobalt-chromium

casting alloys used in frameworks of removable partial dentures and in ceramic-metal

restorations may have small differences in their composition used to control specific properties

such as the coefficient of expansion and strength [36, 41]. The percentage of cobalt in the alloy is

usually around 60% and it is responsible for increasing strength, hardness, and elastic modulus

while the chromium content is usually less than 30% and it is responsible for the corrosion

resistance [41, 46]. Beside cobalt and chromium, these alloys include low concentrations of other

elements such as carbon, silicone and molybdenum that help improve their properties (i.e.

hardness and melting point) [41].

Cobalt-chromium alloys are suitable for dental prostheses because of their mechanical properties

and low cost [36, 41, 42]. For instance, mechanical properties of cobalt-chromium alloys such as

tensile strength, yield strength and hardness are excellent. The low density of the cobalt-

chromium which is half of gold density is also an advantage in dental prostheses [41]. Cobalt-

chromium also can be machined with milling systems to fabricate fixed or removable prostheses

through Computer Aided Design / Computer Aided Manufacturing (CAD/CAM) or Direct Metal

Laser Sintering (DMLS) [41, 42].

32

2.3.1.3. Stainless Steel

Steel is an alloy of iron and carbon, and stainless steel is a modification of steel that contains

chromium, manganese and other elements to provide stainlessness [41, 42]. Stainless steel

cannot be cast, and it is frequently used in dentistry in its wrought form or as readymade

products (i.e. orthodontic brackets and bands) that provided by dental suppliers [16, 37].

Wrought wires are used for fabricating orthodontic appliances and acrylic removable prosthesis

(temporary RPD) [41, 42, 46]. Most stainless steel alloys used for dental applications contain

72% iron, 18% chromium, 1% carbon and low concentration of other elements such as nickel,

molybdenum and silicon [41]. The mechanical performance of stainless steel is excellent in

tension, bending, and torsion; however, its high ductility sometimes could be a problem for

stainless steel wrought wires [41, 42].

2.3.2. Ceramic

Dental ceramics (porcelain) are a mixture of three materials: quartz, feldspar, and kaolin, fired at

high temperature [41]. Dental ceramics are classified into two groups according to their

applications: ceramics for the metal-ceramic prostheses (porcelain fused to metal) or for all-

ceramics prostheses [41, 42]. The common examples of metal-ceramic prostheses are crowns and

bridges while the inlays, onlays, veneers, and full ceramics crowns are the common applications

for full ceramic prostheses [41]. Dental ceramics in metal-ceramics restorations are built up on

metal base frameworks, and they are composed of three layers: opaque, dentin, and enamel [41].

The opaque is used to mask the black color of metals, while the dentin is the main bulk of the

restoration. Finally, the enamel is added to add transparency to the artificial teeth.

33

Ceramics are the best available materials used in dental prostheses for matching the esthetics of

human teeth because they can mimic tooth color, shade, and transparency [41, 42]. Dental

ceramics are hygienic, biocompatible, and chemically stable in the oral cavity [42]. However,

ceramics are brittle and weak in tension [42]. The newly developed all-ceramic materials such as

zirconia and alumina have high strength are now widely used as all ceramic prosthesis through

different techniques such as heat-pressing, slip-casting, sintering, and computer aided design/

computer aided manufacturing (CAD/CAM) [41, 42].

2.3.3. Polymers and Composites

Polymers were introduced in dentistry in 1840s when Goodyear discovered vulcanized rubber as

denture-base material for dental prostheses [47]. This polymer was used in denture-base material

for over seventy-five years although it had poor aesthetic and bad taste. In the 1930s, the poly-

methyl methacrylate (PMMA) was introduced in dentistry and became the most frequently used

polymer in dental prostheses [48]. Artificial teeth and denture bases for dental prostheses are the

main area where polymers are used in dentistry [11]. Polymers are also used for different dental

applications such as impression material, impressions trays, fillings, adhesives, and orthodontics

appliances [42].

Dental composites are usually a combination of polymers and ceramics that result a new material

with superior properties [41, 47]. Dental composites are based on methacrylate polymers, such as

bisphenol A glycidyl methacrylate (Bis-GMA), and they are used in as filling materials to restore

damaged teeth and as adhesive for orthodontic brackets [41, 42]. Dental composites are

becoming popular in dentistry because of their esthetics and mechanical properties.

34

2.3.3.1. PMMA

Poly-methyl methacrylate (PMMA) has been used in many important dental applications such as

impression trays, artificial crowns and bridges, and in orthodontic and maxillofacial appliances

[41]. In addition, PMMA is the main material used in dentures, including removable partial

dentures, removable complete dentures, and fixed complete dentures. It is estimated that PMMA

represents 95% of all polymers used in dental prostheses [49].

Poly-methyl methacrylate (PMMA) is basically a chain of repeating units of the monomer

methyl methacrylate (MMA). The polymerization of PMMA is a free-radical addition

polymerization, and the reaction occurs in three stages: initiation, propagation, and termination

[41]. In the initiation stage, the PMMA polymerization starts by adding an initiator (i.e.

peroxide) to the monomer (MMA) that breaks the double bond of the monomer generating free

radicals (Figure 2.9.A). The free-radicals generated in the initiation stages will react rapidly with

other monomers and will continue adding monomers through the propagation stages building up

the molecular weight (Figure 2.9.B) [47]. The termination stage, the final polymerization stages

occurs when the monomer runs out or when another free radical is introduced (Figure 2.9.C)

[41].

PMMA used at dental laboratories is usually supplied in form of a powder and a liquid

component, and it is polymerized by mixing them in a proper ratio. Poly-methyl methacrylate is

classified into three classifications according to the mechanism that initiates the reaction: heat-

cured, chemical (self, auto)-cured, and light-cured polymerizations [50]. For self-cured PMMA,

the polymerization process occurs with a chemical accelerator such as dimethy-p-toludiene to

speed-up the reaction at room temperature [41]. The heat-cured PMMA materials involve using

35

thermo sensitive accelerators that can be activated using a heat source (i.e. hot water bath,

microwave). The light-cured PMMA contains photo sensitive accelerator that are activated upon

exposure to light [41, 47].

Figure 2.9: Scheme of the polymerization reaction of PMMA. A: the initiation stage using initiator; B: PMMA in the

propagation stage; C: the PMMA in the termination stage.

Heat-cured and self-cured PMMA powders are consisted of high molecular weight poly-methyl

methacrylate (PMMA) as main constituent, initiators (i.e. benzoyl peroxide), dyes (i.e. mercuric

sulphide or cadmium sulphide), opacifiers (i.e. zinc oxide or titanium oxide), and plasticizer (i.e.

dibutyl phthalate) [51]. The PMMA liquid consists of methyl methacrylate monomer (MMA),

cross-linking agent (i.e. glycol dimethacrylate), plasticizer (i.e. dibutyl phthalate), and inhibitor

(hydroquinone). The differences between heat-cured and the self-cured PMMA is the chemical

accelerator (i.e. dimethy-p-toludiene) only added into the liquid of the self-cured PMMA [51].

The dyes and opacifiers in dental PMMA are used to provide the required esthetic to mimic the

natural appearance of teeth and oral tissues [41]. The cross-linking agent is added to the liquid to

improve strength, hardness and wear resistance while the plasticiser is added to the PMMA

powder to reduce the rigidity and the glass-transition temperature (Tg) [51]. Finally, adding the

36

inhibitor in the PMMA liquid is to extend the shelf-life of the liquid for long term storage [41,

51].

Unlike heat-cured PMMA, light-cured PMMA is usually supplied as one single component in the

form of a premixed rope or sheet that is activated and polymerized quickly by exposing it to

visible blue light [47]. Recently, the light-cured PMMA has become very popular in many dental

applications because it is easier and faster to process than either heat or self-cured PMMA.

Although the mechanical and physical properties of PMMA are influenced by the concentration

of its components (i.e. monomer and initiator) and curing conditions (i.e. temperature, time, and

cycling process), PMMA properties are very suitable for dental prostheses [52]. The glass

transition temperature (Tg) for the heat-cured (125oC) and self-cured (90

oC) PMMA exceed the

requirements for temperature resistance in the oral cavity [52]. Moreover, the mechanical

properties for PMMA, such as toughness and hardness, are acceptable. The tensile strength (55

and 90 MPa) and flexural strength in the PMMA consider very high compared to different

polymers [52]. The fatigue life and impact strength in the PMMA are the main problems because

they are low and can lead to the prostheses’ failure [52]. Furthermore, the shrinkage percentage

(6.2 %) is another disadvantage of PMMA that can prevent dentures from fitting accurately into

the patient’s mouth [51]. The esthetic properties of PMMA are excellent because PMMA can be

transparent or colored for matching the colors of the teeth and tissues, and can even incorporate

small colored fibers to give a veined appearance [51, 52]. Poly (methyl methacrylate) is a

biocompatible polymer that is non toxic and does not cause irritations to the oral tissue after it is

fully polymerized [50, 51]. However, there is a concern about the biocompatibility of PMMA in

dentures containing small amounts of residual (un-reacted) monomers that cause toxicity,

37

allergy, and irritation to the oral tissues, and they can be transferred to blood through saliva,

affecting organs, such as liver, kidney, and heart [50]. Furthermore, toxicity of self-cured PMMA

is higher than that of heat-cured PMMA because it contains higher amounts of toxic residual

monomer, initiators, and activators [51]. Moreover, dental technicians, who work with PMMA to

fabricate dentures, face more toxicity and allergies because they are exposed to MMA-vapor

while processing it [50]. Generally, the toxicity of PMMA in dental devices is considered very

low and safe when polymerized properly.

The differences between heat-cured and self-cured PMMA is that heat-cured PMMA has higher

molecular weight, strength, fatigue life, and impact resistance than self-cured PMMA [47]. the

porosity, deformation, and distortion in the heat-cured PMMA is lower than self-cured PMMA

especially when it is heated gradually and uniformly during the polymerization process [47].

Light-cured PMMA presents lower shrinkage and faster and easier processing than self-cured

and heat-cured PMMA [47].

2.3.3.2. Bis-GMA

The Bis-GMA (bisphenol A glycidyl methacrylate) is a resin composites based on a methacrylate

that was introduced to dentistry in the 1960s to improve the mechanical and aesthetic properties

of dental polymers. Bis-GMA is used widely in dental prostheses, dental fillings, and adhesives

for orthodontic appliances [41, 48]. Bis-GMA is a combination of one part of bis-phenol and two

parts of glycidyl methacrylate that are polymerized in a free-radical addition reaction (Figure

2.10) [52]. Bis-GMA has high molecular weight and high viscosity because of the hydrogen

bonding. Bis-GMA based dental resins consist of four major components: the organic polymer

matrix, an inorganic filler, a coupling agent, and an initiator-accelerator system [41]. Fillers such

38

as quartz, fused silica, and glasses are the major portion of the composite that can be in macro,

micro, or nano size and helps increase the hardness and reduce thermal expansion and shrinkage

[41, 52]. Coupling agents are added into the Bis-GMA composite to covalently bind the matrix

to the fillers. The light-cured polymerization for the Bis-GMA composite is the preferred

technique while the self-cured polymerization can occurs with peroxide initiators at room

temperature; the dual-cured is a combination of light and chemical activation [41, 52, 53].

Figure 2.10: Schematic diagram of the chemical reaction for Bis-GMA.

The shrinkage percentage in the Bis-GMA (2.7%) is lower than in PMMA (6.2 %) while the

compressive strength for the Bis-GMA (110-160 N/mm2) is higher than PMMA (75 N/mm

2)

[51]. Furthermore, the modulus elasticity of Bis-GMA (11200 N/mm2) is much higher than

PMMA (1800 N/mm2) [51]. Therefore, it is preferred to use the Bis-GMA for dental prosthesis

that withstands a high compression and impact strengths. The polymerization reaction of Bis-

GMA can take up to 24 hours to be full and complete; therefore during this period, this polymer

can present some toxicity due to release of un-reacted reagents [51, 52]. The biocompatibility of

Bis-GMA is better than the biocompatibility of PMMA because it contains a lower amount of

39

residual monomer [51, 52]. Moreover, light-cured Bis-GMA has been found to be less toxic and

irritating to oral tissues than self-cured Bis-GMA [52].

2.4. Bonding Systems in Dentures and Orthodontic Appliances

Dental prostheses and orthodontics appliances usually combine metallic (i.e. Ti, Co-Cr, and

stainless steel) and polymeric (i.e. PMMA and Bis-GMA) parts. Theses metallic and polymeric

parts in dental devices are joint together to prevent mechanical failure at the interfaces between

them and maintain the integrity of the appliance or prosthesis [41, 54]. Weak bonds allow cracks

to form, grow, and split the metal-polymer interfaces causing dental device failure [11]. Strong

bonding between metals and polymers is also important to prevent bacteria colonies to grow at

the interface causing stains and bad smell [41]. It also important in the fixed orthodontic

appliances since strong bond will prevent the debonding between brackets and teeth which cause

bracket loss [3, 4, 13-17, 37, 55]. Bonding at the interface between metals and polymers can be

improved using mechanical or chemical approaches.

2.4.1. Mechanical Bonding

The mechanical bond between metals and polymers can be formed by penetration and

interlocking of the polymer into the irregularities of the metal surface [11, 12]. Surface

irregularities also help increase the surface area of metals and consequently the overall bond

strength. Orthodontic brackets are designed with bases have micromechanical interlocking to

improve the bonding between brackets and composite [17, 38, 39]. Accordingly, the most

common methods for creating a mechanical bond between metals and polymers can be done by

creating surface irregularities on metallic substrates using sandblasting (air abraded) or chemical

etching. Sandblasting can be performed by applying a stream of aluminum oxide particles with a

40

size between 50 to 250 µm against the metallic substrates under high pressure for 10 to 60

seconds, this roughens the metallic surfaces and provides mechanical bond to the polymers [54].

This technique also helps to remove all the rust and loose particles from the metal surface after

casting, and it is commonly used on metal frameworks of dental prosthesis such as fixed partial

and removable partial dentures. Another way to create surface irregularities is using the Rocatec

System, a silica-coating to metals at high temperature [54]. Chemical-etching with acids such as

H2SO4, HCl and HNO3 at pH≈1.0 is another way of creating surface irregularities micro to nano-

size (0.5 to 2 µm). Etching is an effective way to increase the mechanical bond of polymers to

metals [56].

2.4.2. Chemical Bonding

Chemical bonding involves the formation of covalent, ionic, or hydrogen bonds on the surfaces

interface. However, chemical bond between alloys and polymers does not occur spontaneously.

Achieving a chemical bond at the interference between alloys and polymers usually requires the

use of an adhesive on the metal substrates [18, 23, 57-70]. Adhesives are materials that are

applied on surfaces to permanently join two or more parts together through a bonding process

[20, 22]. Using adhesives between alloys and polymers for dental prosthesis is not common, but

it has recently raised interest [23].

Dental adhesives are mostly a composed of a hydrophilic monomer carried in solvents that react

violently with an initiator in free radical polymerization [24]. Dental adhesive containing

molecule 4-META (Methacryloxy ethyl trimellitate anhydride) was the first commercial metal-

adhesive launched in the market in 1982 under the name name Super-Bond C&B [23, 24]. Then,

the chemical component MDP (Methacryloyloxydecyl dihydrogen phosphate) was added in 1983

41

with 4-META to enhance the bonding of metals to polymers; however, these primers were only

used with non-noble metals [23]. The primers that contain VBATDT (Vinylbenzyl-n-propyl

amino triazine dithione) knows as V-Primer or Alloy Primer were marketed in 1994 to be used

with noble and non-noble metals [23, 70]. More recently, many different metal-adhesives based

on phosphonic acid monomer or phosphonates such as MHPA (Methacryloxyethexy

phosphonacetate), MEPS (Methacryloxydecly thiophosphate derivative) under commercial

names such as the AZ Primer have become available in the market [23, 24]. MDDT

(Methacryloxydecly dithiooctanoate), commercially known as Metal Link Primer, are suitable to

be used with noble and non-noble metals [24, 70].

Dental silane coupling agents that contain MPS (Methacryloyloxy propyltrimethoxy silane) or

MATP (Methacryloxypropyl-trimethoxysilane) are used in dentistry to enhance the bond

between polymers and metals or ceramics [24, 71]. The silane group provides a covalent bonding

between polymers and silica-based materials (ceramics) or active metallic substrates, but the

bonding to silica-based materials is significantly higher than the non-silica based such as metals

[71, 72]. There are many different commercial dental silane coupling agents used in dentistry

such as RelyX, Bisco Porcelain Primer, Cimara, ESPE Sil, and Pulpdent [71, 72].

2.5. Debonding in Dentures and Orthodontics Appliances

Composite materials that combine polymers with alloys often suffer from mechanical failure at

the interface between them. In fact, dental devices often present catastrophic mechanical failures

due to lack of bonding between their metallic and polymeric components [18, 19]. These devices

include dental prostheses, combining metallic frameworks (i.e. titanium and cobalt-chromium)

and wrought wires with acrylic (PMMA); and orthodontic appliances, combining acrylic

42

(PMMA) with stainless steel wrought wires or Bis-GMA composite with stainless steel brackets.

Chemical bonding between alloys and polymers in dental devices does not occur spontaneously.

Therefore, the bonding between alloys and polymers in dental devices is usually provided by the

micromechanical interlocking or sandblasting which barely creates a mechanical bond on the

metallic surface that can bind to polymer. However, this bond is insufficient to prevent the

debonding at the metal-to-polymer interface [18, 19].

2.5.1. Bonding between Alloys and PMMA

Several bonding methods are currently used to increase the bonding strength between PMMA

and alloys in dental prostheses [18, 57-69]. Still, the bond strength achieved between PMMA and

metals so far is insufficient. The highest tensile strength in the literature for the bond between

PMMA and titanium using a combination of sandblasting and bonding agents phosphonate-based

adhesives (MHPA, MDP and VDT; table 2.1) was only 23.5 MPa [58, 60, 62-64, 66, 69]. That is

much lower than the tensile strength of PMMA that is around 65 MPa [73].

Table 2.1 summarizes the bond strengths obtained with different bonding agents between

titanium and poly-methyl methacrylate (PMMA) for dental prostheses. Most bonding methods

reported in the literature require sandblasting the metallic surface; and all of them are based on

molecules containing either silane or phosphonate [18, 57-69]. Sandblasting increases the surface

area of titanium while silane and phosphonate covalently bind the acrylic to the titanium; thus,

increasing the overall bonding strength [56].

In the literature, the bond strengths between titanium and PMMA have been measured with

different mechanical tests including shear bond, four-point bending, and tensile strength tests

[18, 57-69]. However, the strengths reported for each bonding agent depend on the test used.

43

Higher bonding strengths are reported for the four-point bending and shear bond tests (reported

values range between 25.5 to 42.5 MPa and 7.0 to 46.6 MPa, respectively), while the lowest

values are obtained with the most challenging test, i.e. the tensile strength test (0 to 23.5 MPa).

The latter test is the most accurate technique to measure bond strength because it applies a direct

and uniform force to the surface [74]. On the contrary, the shear bond and four-point bend tests

do not distribute stress uniformly on the testing surfaces [48].

Table 2.2 summarizes the literature of the bonding strength of poly-methyl methacrylate with

either cobalt-chromium or stainless steel alloys. All the metallic samples reported in table 2.2

were sandblasting, and the bonding agents were similar to the agents in table 2.1 that are based

on molecules that contain either silane or phosphonate [62, 68, 75-80]. The higher bond strength

reported for cobalt-chromium was 29.1 MPa in the shear bond tests using the bonding agents

META and MATP while the highest bond strengths for the stainless steel in the shear bond tests

were 51.0 and 50.3 MPa using the bonding agents BPDM and MAC [62, 76-80].

2.5.2. Bonding between Wrought Wire and PMMA

Wrought wires are used in many acrylic devices such as dental prostheses and orthodontic

appliances [16]. These wires usually are made of stainless steel or cobalt-chromium alloys that

lack the ability to bind chemically to acrylic [41]. For these reasons, dental devices that combine

wrought wires with acrylic, such as acrylic removable partial dentures, face technical limitations

when not enough volume of acrylic is available to support the wire. Surprisingly, very little

research has been done in order to improve the adhesion between these wires and acrylic.

Therefore it would be of great interest to develop a bonding agent that could increase the

adhesion between wrought wires and acrylic.

44

Table 2.1: The bond strengths between titanium and PMMA (MPa) by using different bonding methods.

Bonding Agent

(Commercial name)

Surface

Topography

Type of PMMA used

(Commercial name)

Testing

Technique

Bond

Strength

(MPa)

Ref.

None Sandblasted Self-cured with EGDMA and TBB

(Super- Bond C&B)

Shear bond 38.1±2.3 [68]

“ “ Heat-cured Tensile strength 20.0 [69]

“ “ “ “ 16.1±1.6 [58]

“ “ “ “ 3.2±0.4 [62]

“ “ Self -cured with BP (Multi- Bond) Shear bond 13.6±1.6 [68]

“ “ Self-cured “ 9.9 [18]

MHPA (AZ Primer) “ “ “ 46.6 [18]

MDP and VTD (Alloy Primer) “ “ “ 45.7 “

“ “ Self-cured with EGDMA and TBB

(Super- Bond C&B)

“ 39.8±2.0 [68]

“ “ Self -cured with BP (Multi- Bond) “ 22.0±6.6 “

“ “ Heat-cured “ 27.5 ±4.0 [57]

“ “ “ Tensile strength 16 .0±3.6 [62]

MDDT and MHPA (Metal Link

Primer )

“ Self-cured Shear bond 45.4 [18]

“ “ Self-cured with EGDMA and TBB

(Super- Bond C&B)

“ 39.6±2.5 [68]

“ “ Self -cured with BP (Multi- Bond) “ 16.5±2.3 “

MATP (Espe-Sil) Polished Heat-cured “ 0.0 [67]

“ Sandblasted “ “ 5.9±2.1 “

MATP “ Self-cured Tensile strength 14.3 [66]

MATP (Silicoater M D) “ “ Shear bond 21.9± 1.7 [59]

MATP and Silicate Coating (Espe-

Sil; Rocatec System)

“ Heat-cured “ 16.2±2.3 [67]

Silicate Coating (Rocatec System) “ Self-cured “ 38.7 [18]

“ “ Heat-cured “ 23.8±1.7 [58]

META “ Heat-cured (Trevalon) Four-point bend 31.9 ±1.5 [65]

“ “ Heat-cured (Metadent) “ 42.5±2.2 “

“ “ Heat-cured Tensile strength 21.0 [69]

META (Super bond) “ “ Shear bond 19.1 ±8.9 [57]

META (New Metacolor) “ Self-cured “ 21.5± 2.2 [59]

MDP (Estenia Opaque Primer) “ “ “ 42.7 [18]

“ “ Heat-cured “ 7.0 ±3.0 [57]

MDP “ “ Tensile strength 23.5 [69]

“ “ Self-cured with EGDMA and TBB

(Super- Bond C&B)

“ 21.2±4.7 [64]

“ “ “ “ 16.2±5.9 [63]

MDP (Cesead ) “ Self-cured Shear bond 19.0± 2.2 [59]

MEPS (Thermoresin) “ “ “ 14.0± 0.6 “

MPS and n-propylamine Polished Heat-cured Four-point bend 25.5±6.4 [61]

MAC (MR Bond) “ “ Tensile strength 7.4 ±2.1 [62]

DOPA “ “ “ 1.8 [60]

Abbreviations: BP: benzoyl peroxide; DOPA;3,4-dihydroxyL-phenylalanine; MAC:11-metacryloyloxyundecan 1,1-dicarboxylic; MATP:

Methacryloxypropyl-trimethoxysilane; MDDA:10-methacyloyloxydecyl 6,8-dithioctanoate; MDP:10-Methacryloyloxydecyl dihydrogen

phosphate; MDDT:10- methacryloxydecly 6,8-dithiooctanoate; MEPS: methacryloxydecly thiophosphate derivative; META: Methacryloxy ethyl trimellitate anhydride; MHPA;6-Methacryloxyethexy phosphonacetate; MPS:3-Methacryloxypropyl trimethoxysilane; TBB:tribuylborane;

VTD:10-Methacryloyloxydecyl dihydrogen phosphate.

45

Table 2.2: The bond strengths (MPa) between PMMA and cobalt –chromium or stainless steel using different bonding

methods.

Abbreviations: AEAPS: N-(2-aminoethyl)-3-aminopropyltrimethoxysilane; AETA: 4-acryloxydecly trimelliate anhydride; BPDM:adduct 2-hydroxyethyl methacrylate and 3,4,4’,5’-biphenyl tetracarboxylic anhydride; DGEBA: diglycidylether of bisphenol; MAC-:11-

metacryloyloxyundecan 1,1-dicarboxylic; MATP: Methacryloxypropyl-trimethoxysilane; MDDT:10- methacryloxydecly 6,8-dithiooctanoate;

MDP:10-Methacryloyloxydecyl dihydrogen phosphate; MEPS: methacryloxydecly thiophosphate derivative; META: Methacryloxy ethyl trimellitate anhydride; MHPA;6-Methacryloxyethexy phosphonacetate; TEOS:tetraethoxysilane; VTD:10-Methacryloyloxydecyl dihydrogen

phosphate.

2.5.3. Bonding between Brackets and Composite

Orthodontic brackets are made of stainless steel, and are bonded to teeth through a composite

adhesive based on polymers such as Bis-GMA, that adhere very well to teeth but very poorly to

metal resulting in bracket debonding [3, 4, 16, 37]. Brackets debonding or bracket loss are more

common in complex orthodontic treatments that require intensive forces applied on severely

crowded or angulated teeth to move them into their correct positions [81, 82]. Therefore,

orthodontic brackets were developed with large bases designed to increase the surface area and

Type of metal Bonding Agent

(commercial name)

Type of PMMA

used

Testing Technique Bond Strength

(MPa)

Ref.

Cobalt –Chromium None Self-cured Tensile strength 6.8 ±2.3 [80]

“ “ Heat-cured “ 3.4±0.6 [62]

MDP and VTD (Alloy Primer) “ “ 17.1±2.6 “

“ “ Self-cured “ 19.2 ±6.0 [80]

“ META (UBar) “ “ 8.6 ±2.7 “

“ META and MATP “ “ 29.1±6.3 [79]

“ MAC (MR Bond) “ “ 25.6 ±6.4 [80]

“ “ Heat-cured “ 17.8±4.0 [62]

Stainless steel None Self-cured “ 20.0 [78]

“ “ “ Shear bond 21.3±2.7 [77]

“ “ “ “ 19.2±3.6 [76]

“ MDP (Estenia Opaque Primer) “ “ 35.4±2.0 “

“ “ “ “ 49.4±1.3 [77]

“ MDP and VTD (Alloy Primer) “ “ 48.8±1.6 “

“ “ “ “ 34.3±1.9 [76]

“ MDDT and MHPA (Metal Link Primer ) “ “ 34.4±1.7 “

“ “ “ “ 48.8±1.4 [77]

“ “ “ Tensile strength 38.3 [78]

“ META (Meta Fast primer) “ “ 36.0 “

“ META (Super bond) “ Shear bond 49.4±1.5 [77]

“ MEPS (Metal primer PII) “ “ 48.1±1.3 “

“ “ “ Tensile strength 31.1 [78]

“ MAC (MR Bond) “ Shear bond 50.3±2.3 [77]

“ AETA (Acryl Bond) “ “ 49.9±1.5 “

“ BPDM (All-Bond 2) “ “ 51.0±0.8 “

“ DGEBA and AEAPS + MATP and TEOS “ “ 9.7 ± 1.5 [75]

46

compensate for the lack of adhesion [14, 16]. The resulting large brackets are more likely to

affect the quality of life of the patients in terms of food impaction and esthetic concerns, whereas

food impaction around and beneath the brackets lead to increased rates of demineralization and

tooth caries [13-16]. Most of the research conducted to increase the adhesion of brackets has

been limited to modifications of the surface pattern of the bracket base in order to increase

micromechanical interlocking that only achieved minor improvements [14]. Accordingly, a

chemical adhesive system that would increase the bond strength between orthodontic brackets

and dental composite would be of great interest because it could prevent brackets loss and result

in a reduction of the size of brackets that translates into less dental problems associated to them.

2.6. Aryldiazonium Salts

Aryldiazonium salts or diazonium compounds are organic compounds that have the molecular

structure R-N2+

which R is Aryl group contains a phenyl group [83]. Diazonium salts can be

prepared by the reaction knows as diazotization or diazotisation [83, 84]. Diazonium was

described in the middle of the 19th

century, and the name diazonium was from French “diazote”

which means two nitrogen atoms [84]. Then, it has been able to generate azo compounds from

diazonium salts that are important coloring materials, and they were recognized by dyes

industries [83, 84]. In the last two decades, aryldiazonium salts have raised interest for

modifying materials surfaces in numerous applications [83, 85-88]. The interest of using

aryldiazonium salts is because their ease of preparation and reduction, strong covalent bonding

between various materials, large choices of functional groups on the compound, and cost [83, 89-

91].

47

2.6.1. Grafting of Diazonium Salts

There are different methods for grafting diazonium salts on various substrates, and the most

common methods are electro-grafting, ultrasonication grafting, heating grafting, photo grafting,

and redox grafting [83, 84, 86, 92]. Initially, diazonium grafting was performed using

electrochemical reduction that occurs by transfer one electron wave (0.02 to 0.56 V) to

aryldiazonium salts in aprotic or acidic mediums [84, 86, 92, 93]. This reduction leads to reduce

aryldiazonium salts generating free radicals that bind on substrate surfaces [84, 92]. Moreover,

grafting aryldiazonium salts has been achieved under ultrasonication in an acidic solution, being

the ultrasonic frequency of 20 KHz the most efficient one for reducing aryldiazonium salts [84,

86, 94]. Heating or microwave heating an acidic solution containing aryldiazonium salts in

presence of sodium nitrate (NaNO2) at temperature of 60 to 80oC has been proven to graft

aryldiazonium salts on different surfaces [84, 86, 95]. Also, it has been shown that it is possible

to graft aryldiazonium salts on various surfaces using UV (ultraviolet) light or mechanical

assistance [83, 86, 96]. Finally, redox grafting is the more recent and simpler than other grafting

methods. In this method, diazonium salts can be grafted using a reducing agent (i.e. iron powder

or hypophosphorous acid) that activate the aryldiazonium salts and form aryl radicals that

covalently bind to different surfaces (Figure 2.11) [84, 86, 92, 97-99].

48

Figure 2.11: Scheme describing grafting of diazonium salts of a substrate: (a) diazotisation of p-phenylenediamine (PPD);

(b) formation of aryl radicals by reducing aryldiazonium using iron powder or hypophosphorous acid; (c) attachment of

aryl radicals to substrates surfaces; (d) growing of diazonium salts as multi-layers.

2.6.2. Diazonium Grafted Layer Properties

One of the most interesting points of diazonium salts is their ability to be reduced and graft

simply and rapidly onto almost any surface, including metals, glass, and carbon. The grafted

layer forms a strong covalent bond to any surfaces that can be used as coupling agents between

different substrates [83, 84, 89, 93, 100-102]. Diazonium structure allows large choices of

functional groups and has two active sites that can be used as a self adhesive layer between

different materials [84, 92]. Moreover, the grafted aryl-layer can be used to improve the physical

and chemical properties of different material surfaces [84, 86, 89, 102].

49

The grafted layer obtained with diazonium salts has been proved to be stable for at least 6

months in harsh conditions including ultrasonication and heating. It was also found that these

grafted layers are partially stable after removing or scratching by electrical or mechanical

methods [84]. The covalent bond formed between diazonium layers and substrates has been

previously observed by XPS (X-ray photoelectron spectroscopy), Raman, AFM (atomic force

microscopy), ToF-SIMS (time-of-flight secondary ion mass spectrometry), and with DFT

(density functional theory) [83, 84, 91, 100, 103].

The grafted layer is usually formed as a multi-layers rather than mono-layer [84]. The multi-

layers occur because the free radicals generated from the reduction can also react with the first

grafted layer and form other layer on it (Figure 2.11.d). The grafted layers grow until electrons

are unable to be transferred through the grafted layers and stop the reaction [84]. Grafted layers

are not homogenous, and their thickness can range from few nanometers to one micron [83, 84].

The thickness depended on the amount of aryldiazonium salts and type of substrates that

aryldiazonium salts graft to it. However, the thickness of grafted layers on metals surfaces is

usually found to be 2-6 nm [83, 84, 87, 104, 105].

2.6.3. Applications of Aryldiazonium Salts

Surface modification with aryldiazonium salts have been used for multiple applications, such as

corrosion resistance, abrasion resistance, and electrical insulation, in many biomedical,

environmental, and industrial applications [84]. There are more than 100 patents on

aryldiazonium salts in these applications, and many of these patents are base on the use of

diazonium salts as coupling agents between different material surfaces [83, 84, 86]. Diazonium

salts were used in many industrial applications, such as inkjet ink and fuel cells [83, 84]. Carbon

50

blocks treated with diazonium result in useful elastomers with high abrasion resistance that can

be used with tires to reduce fuel consumption [84]. Diazonium salts were used in the biomedical

filed for drug delivery and modifying surgical stents [83, 84]. They were used also in biosensors,

functional polymer coatings, protein arrays, molecular electronic junctions [83, 84, 86, 106-108].

2.6.4. Aryldiazonium Salts as Dental Adhesive

Aryldiazonium salts have not been used in dentistry; however, they would be of great interest

due to their ability to act as coupling agent between metals and polymers for dental applications.

Aryl diazonium salts have more attracting properties and features than other available dental

adhesives. First, aryl diazonium salt can react rapidly with surfaces, produce a very active

radical, and give strong adhesion on any surfaces (i.e. noble metals, base metals, ceramics,

polymers, and composites) [83, 89]. Therefore, dental adhesives based on aryldiazonium salts

could works with any type of materials as a multi-purpose adhesive.

The thickness of the diazonium layer is less than 1 µm that is another advantage of this dental

adhesive because it will help to reduce the thickness and weight of dental prostheses [84].

Another advantage of using aryl diazonium as dental adhesive is the cost which would be

reasonable since the raw materials to produce this adhesive are commercially available at a low

cost [84].

2.6.5. Diazonium Grafted Layers Analysis

Several methods can be used to characterize aryl diazonium layer such as X-ray photoelectron

spectroscopy (XPS), Raman spectroscopy, IR-ellipsometry, and contact angle measurements.

However, the best way to assess diazonium grafted layers on metallic surface is by XPS (X-ray

51

photoelectron spectroscopy) because of its unique ability to characterize external thin chemical

coating [83, 84, 91, 100, 103].

2.6.5.1. X-ray Photoelectron Spectrometer (XPS)

X-ray photoelectron spectroscopy (XPS) is the most commonly used method for surface analysis

because it is commercially available and provides clear information on surface chemical

composition [109]. This technique was called ESCA (electron spectroscopy for chemical

analysis) but it changed because there are many surface-electron spectroscopy techniques and the

name for each technique should be precise [109]. XPS is a sensitive technique that provides

quantitative information on the elemental and chemical composition of the outermost 10 nm or

less of n the solid surface [109-112]. XPS can detect all elements in the periodic table except

hydrogen and helium [109, 111].

The surface to be analyzed with XPS is irradiated with X-ray source that penetrates the surface

for a depth ~1 micrometer and emits electrons (photoelectron) which have a binding energy less

than the X-ray energy [109, 110]. The electron analyser determines the binding energy (Eb) of

the photoelectron. The photoelectron kinetic energy (Ek) can be determined through an equation

(Ek=hv-Eb) using the binding energy (Eb) and the energy of the incident radiation (hv) [109-111].

The resulting binding energies are unique from different atoms that show on XPS spectrum as

peaks [109]. The elemental identity, quantity of an element, and chemical state are also

determined.

The main components of XPS (Figure 2.12) are an X-ray source, an ultra-high vacuum (UHV),

and an electron-energy analyzer [109]. Al Kα or Mg Kα X-ray sources are the most common

52

ones because the width and line energy for Al Kα or Mg Kα are suitable for XPS [109, 110].

XPS is a very sensitive technique; therefore, it requires vacuums of the order of 10-8

Pa or lower

to protect samples from surface contamination [109]. The electron-energy analyzer is used to

analyze the energies of electrons ejected from samples which will be quantified in a detector

system (Figure 2.12).

Figure 2.12: Diagram describing X-ray photoelectron spectroscopy (XPS) components.

XPS general survey is a scan that covers the entire binding energy (0 to 1000 eV) of a surface at

low energy resolution that can be used to identify the elements and the quantity of each atoms on

the surface [109, 110]. Moreover, high resolution XPS spectra can be obtained for each element

(i.e. C1s and O 1s) in specific regions to provide specific chemical bonding information [109,

110].

In the first study of this thesis, aryl diazonium layer was analyzed by a monochromatic X-ray

photoelectron spectrometer K Alpha (Thermo Fischer Scientific Inc, East Grinstead, UK). XPS

53

was used for determining the relative quantities and chemical environments of the elements on

the Ti and surfaces. Control and treated Ti samples were carefully cleaned in an ultrasonic bath

with acetone, ethanol, and distilled water for 5 minutes, and they were stored in contamination

free tubes for less than 24 hours before XPS analysis. The setup was equipped with an Al Kα X-

Ray radiation source (1486.6 eV, 0.834 nm), a micro-focused monochromator and an ultrahigh

vacuum chamber (10-9

torr). Survey scans were obtained over the range of 0-1350 eV with pass

energy of 200 eV at a step of 1.0 eV, and high resolution scans were collected with pass energy

of 50 eV at a step of 0.1 eV. Energies were calibrated by setting the binding energy of the carbon

bonded to hydrogen or carbon (C-(H, C)) at 285.0 eV on all samples. Data analysis and peak

fitting were performed using Avantage (5.41v, Thermo Fischer Scientific Inc, East Grinstead,

UK) chemical surface analysis software.

In the second study of the thesis, the aryl diazonium layer was analyzed similarly to the first

study, but it was done with a different XPS model (VG Scientific ESCLAB 3 MIKII). XPS was

used for determining the relative quantities and chemical environments of the elements on the

stainless steel brackets surfaces. The XPS in the second study was equipped with Mg Kα X-Ray

radiation source (1253.6 eV), a micro-focused monochromator and an ultrahigh vacuum chamber

(10-9

torr). Survey scans for all samples were obtained over the range of 0-1350 eV with pass

energy of 100 eV at a step of 1.0 eV. The high resolution scans were collected with pass energy

of 20 eV at a step of 0.05 eV. All spectra were referenced to C-C/C-H of the C 1s spectrum at

285.0 eV.

54

2.6.5.2. Contact Angle Measurements

Contact angle measurement is used to quantify the angle formed by a liquid drop at a solid

surface. Contact angles measurements are useful to understand the wettability and hydrophobicity

of a solid surface because they are sensitive to the first layers of the materials surfaces and the

technique is easy to apply [84]. Therefore, contact angles measurement is a helpful technique to

analyze the grafted diazonium layers on different surfaces.

In this thesis, the contact angle measurements was recorded and analyzed at room temperature on

contact angle meter (OAC 15, Data Physics, Germany). Experimentally, a liquid drop of distilled

water (2ml) was deposited on the Ti surfaces of each samples (control and treated) and the angles

between the drop and surfaces was measured. The static contact angle was automatically

calculated and the side view images were captured. From the contact angle measurements results,

it would be easy to distinguish the properties of first layer of different Ti surfaces that are related

to hydrophobicity of surface [84].

55

Chapter 3: Hypothesis and Objective

3.1. Hypothesis

Aryldiazonium salts could be used to improve the binding between metals and polymers used in

dental applications.

3.2. Thesis Objective

The objective of this study was to develop a new method of creating a strong chemical bond

between alloys and polymers for dental devices based on diazonium chemistry. The objective of

the thesis consists of the two specific objectives addressed in two separate manuscripts that were

divided based on their application in dentistry. Manuscript I assesses the bonding between metals

and PMMA for dental prostheses while Manuscript II investigates the bonding between metals

and Bis-GMA for orthodontic appliances.

56

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62

Chapter 5: Manuscript I

Bonding Metals to Poly-Methyl Methacrylate Using Aryldiazonium Salts

Omar Alageel a, Mohamed-Nur Abdallah

a, Zhong Yuan Luo

b, Marta Cerruti

b, Faleh Tamimi

a

a Faculty of Dentistry, McGill University, Montreal, QC, Canada

b Department of Mining and Materials Engineering, McGill University, Montreal, QC, Canada

5.1 Abstract

Many dental devices, such as partial dentures, combine acrylic and metallic parts that are bonded

together. These devices often present catastrophic mechanical failures due to weak bonding

between their acrylic and metallic components. The bonding between alloys and polymers (e.g.

poly-methyl methacrylate, PMMA) usually is just a mechanical interlock, since they do not

chemically bond spontaneously. The objective of this study was to develop a new method to

make a strong chemical bond between alloys and polymers for dental prostheses based on

diazonium chemistry. The method was based on two steps. In the first step (primer),

aryldiazonium salts were grafted onto the metallic surfaces. The second step (adhesive) was

optimized to achieve covalent binding between the grafted layer and PMMA. The chemical

composition of the treated surfaces was analyzed with X-ray photoelectron spectroscopy (XPS),

and the tensile bonding strength between metals and poly-methyl methacrylate was measured.

XPS and contact angle measurements confirmed the presence of a polymer coating on the treated

metallic surfaces. Mechanical tests showed a significant increase in bond strength between

63

PMMA and treated titanium or stainless steel wire by 5.2 and 2.5 folds, respectively, compared

to the untreated control group (P<0.05). Thus, diazonium chemistry is an effective technique for

achieving a strong chemical bond between alloys and PMMA, which can help improve the

mechanical properties of dental devices.

Keywords

Dental prosthesis; bonding diazonium; poly-methyl methacrylate; titanium; stainless steel

5.2. Introduction

Chemical bonding between alloys and polymers does not occur spontaneously; in fact, composite

materials that combine polymers with alloys often suffer from mechanical failure at the interface

between them. One example of this challenge is dental devices, which often present catastrophic

mechanical failures due to weak bonding between their metallic and polymeric components [1-

3]. These devices include dental prostheses, combining metallic frameworks and wrought wires

with acrylic resin; and orthodontic appliances, combining acrylic resin with stainless steel

wrought wires.

Poly-methyl methacrylate (PMMA) is extensively used in denture materials for dental prostheses

and orthodontic devices because of its biocompatibility, excellent esthetic, and mechanical

properties [4]. Titanium (Ti) is increasingly used in dental implants, implant abutments, and

milled prostheses because of its excellent mechanical properties (i.e. strength to weight ratio) and

biocompatibility [5]. PMMA and Ti in dental prostheses are usually bonded by mechanical

interlocking the PMMA into the irregularities of the Ti surface [2, 6, 7]. Further improvement in

the bonding strength between Ti and PMMA is still needed to prevent debonding, which are

64

otherwise common in clinical practice, and reduce microleaks at the Ti/PMMA interface that

causes accumulation of oral debris and discoloration of denture base materials [1, 2, 8]. Ti-

PMMA bond can be strengthened by adding a chemical link between PMMA and Ti, since Ti

and PMMA do not chemically bind together spontaneously [1, 7, 9-13].

Several methods have been tested to increase the bond strength between polymers and alloys in

dental prostheses [1, 9, 12, 14-24]. Table 5.1 summarizes the strengths obtained by binding

PMMA and Ti with different methods, measured with shear bond, four-point bending, and

tensile strength tests. In general, higher bond strengths are reported for the four-point bending

and the shear bond tests (values ranging between 25.5 to 42.5 MPa and 7.0 to 46.6 MPa,

respectively), while the lowest values are obtained with the tensile strength test (0 to 23.5 MPa).

Indeed, the latter test is the most challenging one; however, it is also the most accurate technique

to measure bond strength because it applies a direct and uniform force to the surface [25]. On the

contrary, the shear bond and four-point bend tests do not distribute stress uniformly on the

surfaces being tested [26].

Most of the methods reported in Table 5.1 require sandblasting the metallic surface, and all of

them use either silane or phosphonate groups to create a chemical bond between the two surfaces

[1, 9, 12, 14-24]. Silanes and phosphonates covalently bind to Ti, while sandblasting increases

the surface area of the exposed Ti, thus increasing the overall bonding strength [27]. The highest

bond strengths reported were achieved using phosphonate-based adhesives (MHPA, MDP and

VDT; see Table 5.1) in combination with sandblasting. Specifically, the highest tensile strength

reported without sandblasting was 7.4 MPa [12], while using a combination of sandblasting and

bonding agents the tensile strength went up to 23.5 MPa [12, 15, 17, 19-21, 24]. These values are

65

still too low for dental applications. Overdentures, for example, have to resist biting forces of up

to 662.2 N, and pressures of up to 51.1MPa [28]. This implies that masticatory forces can exceed

the strength of the Ti-PMMA bond and lead to prosthesis failure. An ideal goal would be to have

a metal/PMMA interface that is at least as strong as PMMA alone, which has a tensile strength of

65 MPa [29].

Another example of metal-acrylic interface found in dental applications is that between wrought

wires and acrylic-based dental devices such as dental prostheses and orthodontic appliances [30].

Wrought wires are usually made of stainless steel or cobalt-chromium alloys, which both lack

the ability to bind chemically to acrylic resins [31]. To overcome this problem, dental devices

combining wrought wires with acrylic such as acrylic removable partial dentures cannot be made

when not enough volume of acrylic is available to support the wire. Surprisingly, hardly anyone

has looked into improving the adhesion between wrought wires and acrylic.

In this paper we will show a technique to improve the binding between PMMA and alloys used

in dental applications based on diazonium chemistry. Aryldiazonium salts have been used to

modify material surfaces for many applications [32, 33]. Diazonium ions can be produced from

aromatic amines and grafted onto almost any surface, including metals, glass, and carbon [34-

37]. Initially, diazonium grafting was performed using electrochemical reduction, but recently

this has been achieved using chemical reducing agents in acidic solutions [38]. The reducing

agents transform the aryldiazonium salts into aryl radicals, which can covalently bind to the

surface of interest [39, 40]. If an extra amino group is present on the aryldiazonium precursor, a

polyaminophenylene (PAP) layer is formed on the metallic surface. The amino groups sticking

out from the PAP layer can be further activated in a second step, and used to bind a second layer

66

onto the original surface [39, 41]. In this work, we optimize such second step to bind PMMA

and metals for dental applications.

Table 5.3: The bond strengths between titanium and PMMA (MPa) using different bonding methods.

Bonding Agent

(Commercial name)

Surface Topography Type of PMMA used

(Commercial name)

Testing

Technique

Bond

Strength

(MPa)

Ref.

None Sandblasted Self-cured with EGDMA and

TBB (Super- Bond C&B)

Shear bond 38.1±2.3 [23]

“ “ Heat-cured Tensile strength 20.0 [24]

“ “ “ “ 16.1±1.6 [15]

“ “ “ “ 3.2±0.4 [12]

“ “ Self -cured with BP (Multi- Bond) Shear bond 13.6±1.6 [23]

“ “ Self-cured “ 9.9 [1]

MHPA (AZ Primer) “ “ “ 46.6 “

MDP and VTD (Alloy Primer) “ “ “ 45.7 “

“ “ Self-cured with EGDMA and

TBB (Super- Bond C&B)

“ 39.8±2.0 [23]

“ “ Self -cured with BP (Multi- Bond) “ 22.0±6.6 “

“ “ Heat-cured “ 27.5 ±4.0 [14]

“ “ “ Tensile strength 16 .0±3.6 [12]

MDDT and MHPA (Metal Link

Primer )

“ Self-cured Shear bond 45.4 [1]

“ “ Self-cured with EGDMA and

TBB (Super- Bond C&B)

“ 39.6±2.5 [23]

“ “ Self -cured with BP (Multi- Bond) “ 16.5±2.3 “

MATP (Espe-Sil) Polished Heat-cured “ 0.0 [22]

“ Sandblasted “ “ 5.9±2.1 “

MATP “ Self-cured Tensile strength 14.3 [21]

MATP (Silicoater M D) “ “ Shear bond 21.9± 1.7 [16]

MATP and Silicate Coating (Espe-

Sil; Rocatec System)

“ Heat-cured “ 16.2±2.3 [22]

Silicate Coating (Rocatec System) “ Self-cured “ 38.7 [1]

“ “ Heat-cured “ 23.8±1.7 [15]

META “ Heat-cured (Trevalon) Four-point bend 31.9 ±1.5 [9]

“ “ Heat-cured (Metadent) “ 42.5±2.2 “

“ “ Heat-cured Tensile strength 21.0 [24]

META (Super bond) “ “ Shear bond 19.1 ±8.9 [14]

META (New Metacolor) “ Self-cured “ 21.5± 2.2 [16]

MDP (Estenia Opaque Primer) “ “ “ 42.7 [1]

“ “ Heat-cured “ 7.0 ±3.0 [14]

MDP “ “ Tensile strength 23.5 [24]

“ “ Self-cured with EGDMA and

TBB (Super- Bond C&B)

“ 21.2±4.7 [20]

“ “ “ “ 16.2±5.9 [19]

MDP (Cesead ) “ Self-cured Shear bond 19.0± 2.2 [16]

MEPS (Thermoresin) “ “ “ 14.0± 0.6 “

MPS and n-propylamine Polished Heat-cured Four-point bend 25.5±6.4 [18]

MAC (MR Bond) “ “ Tensile strength 7.4 ±2.1 [12]

DOPA “ “ “ 1.8 [17]

Abbreviations: BP: benzoyl peroxide; DOPA;3,4-dihydroxyL-phenylalanine; MAC:11-metacryloyloxyundecan 1,1-dicarboxylic; MATP: Methacryloxypropyl-trimethoxysilane; MDDA:10-methacyloyloxydecyl 6,8-dithioctanoate; MDP:10-Methacryloyloxydecyl dihydrogen

phosphate; MDDT:10- methacryloxydecly 6,8-dithiooctanoate; MEPS: methacryloxydecly thiophosphate derivative; META: Methacryloxy ethyl

trimellitate anhydride; MHPA;6-Methacryloxyethexy phosphonacetate; MPS:3-Methacryloxypropyl trimethoxysilane; TBB:tribuylborane;

VTD:10-Methacryloyloxydecyl dihydrogen phosphate.

67

5.3. Materials and Methods

5.3.1. Materials

Poly-methyl methacrylate (PMMA) and methyl methacrylate (MMA) were obtained from Great

Lakes Orthodontics (Tonawanda, NY), and were used without any further purification. The rest

of the reagents were obtained from Sigma Aldrich (St. Louis, MO). P-phenylenediamine (PPD),

sodium nitrate (NaNO2), sodium dodecyl sulfate (SDS), benzoyl peroxide (BP), and iron powder

(Fe) were used as received. Concentrated hydrochloric acid (HCl) was diluted in distilled water

(DW) to a concentration of 0.5M.

The metallic samples used in the experiments were either orthodontic wrought wires (stainless

steel) or polished rectangular bars (Ti). The wrought wires (Tur-Chrome S.S, Rocky Mountain

Orthodontic, Denver, CO) had a diameter of 0.6 mm and were cut into 200.0 mm long sections.

The Ti samples (Ti alloy grade 2, McMaster-Carr, Cleveland, OH) were obtained as rectangular

bars (6.4, 12.7 and 305.0 mm) and cut into smaller sections (12.7, 6.4 and 6.4 mm) using an

abrasive cutter (Delta AbrasiMet, Buchler, Whitby, ON).

5.3.2 Preparation of the metallic samples

The Ti samples were polished using a six step polishing method to obtain a flat surface. First,

they were polished by means of a water-cooled trimmer and 240-to-600 grit silicon carbide

papers (Paper-c wt, AA Abrasives, Philadelphia, PA). Then, they were further polished on a

polishing wheel (LapoPol-5, Struers, Rodovre, Denmark) using two types of polishing cloths;

rough-to-intermediate polishing cloth (15-0.02μm; TexMet C) and final polishing cloth (1-

0.02μm; ChemoMet), with Colloidal Silica Suspension (≤ 0.06μm; MasterMet; Buchler,

68

Whitby, ON). The orthodontic wrought wires did not undergo any specific preparation prior to

surface treatment besides being cleaned. All metallic samples were cleaned in an ultrasonic bath

(FS20D Ultrasonic, Fisher Scientific, Montreal, Canada) with DW, ethanol, and acetone for 5

minutes in each solution at 37 oC.

5.3.3. Surface treatment of the metallic samples

The surface treatment was performed in a two steps protocol based on p-phenylenediamine

diazotization (primer and adhesive). Both steps were carried out in acidic DW solution at pH≤2,

since diazonium cations are stable at pH ≤2.5, at room temperature in a simple glass beaker [39,

41]. The first step (primer) was conducted as follows: PPD (0.054g; 0.05M) and NaNO2 (0.034g;

0.05M) were dissolved in a glass beaker containing 10 ml of 0.5 M HCl. After ultrasonicating

the solution for 5 minutes, all metallic samples except control group were immersed in the

solution and Fe powder (0.250g) was added as a reducing agent. The samples were left to react

for 15 minutes before ultrasonicating them in DW and acetone for 5 minutes. This first step leads

to spontaneous grafting of a polyaminophenylene (PAP) layer on the metallic samples (i.e.

titanium and stainless steel wrought wire). These samples are referred to as metal-PAP from here

onwards.

Different approaches were investigated in the second (adhesive) step in order to optimize the

adhesion of MMA to metal-PAP samples. These approaches can be summarized in four groups

(Table 5.2). All groups share the following process: NaNO2 (0.034g; 0.05M) was dissolved in

10ml of 0.5 M HCl. Then, the metal-PAP samples were introduced in the solution before adding

Fe powder (0.250g). In the first group, only the monomer (MMA) was added to the solution. In

groups 4, 5 and 6, a surfactant (SDS, 0.026 g) was added along with MMA to help solubilize the

69

hydrophobic monomer [42, 43]. The reaction was allowed to continue for 15 minutes in the

ultrasonic bath and for another 30 minutes on the bench top; during this period the monomer

polymerized and formed a thin layer of PMMA on the metallic surface. In groups 5 and 6, an

initiator (benzoyl peroxide, BP) was added after the fifteen minute sonication stage to accelerate

the polymerization reaction on the metallic surface. Finally, the samples were thoroughly rinsed

with acetone, and then ultrasonicated in DW and acetone for 5 minutes in order to discard any

ungrafted matter.

Table 5.4: Conditions tested in the second step. The overall solution volume was 12 ml, and was water-based.

Groups Metal Solution Abbreviation HCl 0.5 M; NaNO2 0.05M;

Fe 0.25g; MMA 2.0 ml

(ml)

SDS

(M)

[BP]

(mg/ml)

1 Ti None Control 0 0 0 2 Ti None D 0 0 0 3 Ti MMA Emulsion without SDS D+M 12 0 0

4 Ti MMA Emulsion with SDS D+M+E 12 9. 10 -3 0

5 Ti MMA Emulsion with SDS and initiator D+M+E+I 12 9. 10 -3 8-48

6 SS MMA Emulsion with SDS and initiator D+M+E+I 12 9. 10 -3 40

D: diazonium grafting (step 1); M: monomer (MMA); E: emulsifier (SDS); I: initiator (BP); Ti: titanium; SS: stainless steel.

5.3.4. Spectroscopic analysis

A monochromatic X-ray photoelectron spectrometer K Alpha (Thermo Fischer Scientific Inc,

East Grinstead, UK) was used for determining the relative quantities and chemical environments

of the elements on the Ti surfaces. The setup was equipped with an Al Kα X-Ray radiation

source (1486.6 eV, 0.834 nm), a micro-focused monochromator and an ultrahigh vacuum

chamber (10-9

torr). For all the groups (control; D; D+M; D+M+E; D+M+E+I), survey scans

were obtained over the range of 0-1350 eV with a pass energy of 200 eV at a step of 1.0 eV,

while high resolution scans were collected with a pass energy of 50 eV at a step of 0.1 eV.

Energies were calibrated by setting the binding energy of the carbon bonded to hydrogen or

carbon (C-(H, C)) at 285.0 eV on all samples. Data analysis and peak fitting were performed

70

using Avantage (5.41v, Thermo Fischer Scientific Inc, East Grinstead, UK) chemical surface

analysis software.

5.3.5. Contact angle measurement

Hydrophobicity of the Ti surfaces in all groups was evaluated by the contact angle measurement

that was recorded and analyzed at room temperature on contact angle meter (OAC 15, Data

Physics, Germany). The static contact angle was automatically calculated by measuring the angle

produced by a drop of DW (2ml) placed on the Ti surface of each samples, and the side view

images were captured.

5.3.6. Mechanical tests

Tensile test was used to measure the bond strength between PMMA and the metallic surface. To

test the Ti-PMAA bond strength, a custom-made mold was fabricated from a silicone (Exaktosil

N 21, Bredent, Germany) (Figure 5.1.a), and a piece of Ti was fixed in the middle of the mold.

Then, a mix of PMMA powder and MMA liquid monomer (Biocryl Resin Acrylic, Great Leakes,

NY) with ratio of 2:1 was poured to fill the sides of the mold. The PMMA was left to set for

three hours at room temperature and humidity. This procedure generated a final specimen that

was 130 mm long, 13 mm wide and 3 mm thick with two grips of bulk PMMA polymerized at

the sides of the Ti samples (Figure 5.1.b). After complete setting of the acrylic resin, the tensile

bond strength between PMMA and Ti was measured using a universal testing machine (H25K-S,

Tinius Olsen Testing Machine Co., Inc Willow Grove, PA) set up at a constant speed of 10

mm/min. The tensile force was applied to the specimen until fracture occurred at the PMMA-Ti

71

interface, and the strength of the bond between Ti and PMMA was calculated in megapascals

(MPa).

The specimens used to test the bond between wrought wire and PMMA were prepared using a

different custom-made silicone mold. The resulting specimens consisted of 20 mm of wrought

wire embedded vertically into a plate of PMMA (3 mm thick; 20 mm wide; 130 mm long)

(Figure 5.1.c). The acrylic plate and the wire were secured into a universal testing machine (as

described earlier) in order to measure the tensile strength of the bond between the wire and the

PMMA.

Figure 5.1: (a) the custom-made silicone mold used to prepare the PMMA-Ti specimen; (b) the Ti-PMMA specimen

before and after mechanical testing; (c) the stainless steel wrought wire-PMMA specimen before and after mechanical

testing.

5.3.7. Statistical analysis

Statistical analysis on all XPS, contact angle, and mechanical test results was performed using

Origin 8.0 (Origin lab, Northampton, MA). All the data were analyzed using nonparametric tests,

Kruskal-Wallis test and the significance level was set at p<0.05.

72

5.4. Results and discussion

The diazonium chemistry method that we used to bind PMMA and dental alloys consisted of two

steps. In the first step, PPD was first transformed into an amino diazonium cation by adding one

equivalent of NaNO2 (Figure 5.2.a). The diazonium cation was then reduced with Fe to achieve

an aminophenyl radical (Figure 5.2.b). This radical spontaneously grafted onto the titanium

surface (Figure 5.2.c), and kept reacting with itself forming multilayers (PAP, Figure 5.1.d) [39].

The second step (adhesive) of the reaction was optimized in order to achieve covalent binding

between metals and PMMA for applications in dental prostheses. The amino groups of the PAP

layer were reduced to radicals using again NaNO2 and Fe in an acidic environment, as in the first

step (Figure 5.2.e). The radicals reacted with MMA, and formed a thin PMMA layer on top of

the metal-PAP (Figure 5.2.f; SDS was used in some samples to solubilize MMA). In some of the

samples, PMMA chain length was increased by adding BP, which helped the formation of more

MMA radicals (Figure 5.2.g).

73

Figure 5.2: Schematic showing the reaction sequence performed in first (primer) and second (adhesive) steps. (a)

diazotisation of p-phenylenediamine (PPD) in an acidic solution; (b) formation of aryl radicals by reducing aryldiazonium

ions using iron powder; (c) attachment of aryl radicals to the metallic surfaces; (d) growing of polyaminophenylene layer

(PAP) multilayer; (e) reduction of amino groups of the PAP layer to radicals; (f) binding of MMA to the activated PAP

layer; (g) increasing PMMA length with benzoyl peroxide. PAP layer in the scheme showed as single layer for simplifying

but it could be single, multi or mixed layers.

74

XPS results (Figures 5.3.a and 5.3.b) showed that the Ti surface for the control (untreated)

samples was covered by TiO2 and carbon; the carbon has to be related to the unavoidable

contamination upon exposition to air prior to XPS analysis [44]. The bond strength between the

untreated polished Ti and PMMA was very low (1.54±1.02 MPa), indicating that the mechanical

and chemical bonds between Ti and PMMA were minimal (Figure 5.5.a). It is hard to find a

comparison between this value and what reported in the literature (Table 5.1), since all the data

reported for PMMA-Ti bonds on untreated samples refer to sandblasted Ti surfaces, which

provides an increased mechanical bond. After sandblasting, the reported tensile strengths vary

between 3.2 and 20.0 MPa [1, 9, 12, 14-24, 27]. In fact it is extremely difficult to test the bond

between a polished Ti surface and PMMA, since the samples tend to fail very quickly [25].

XPS confirmed the grafting of a PAP layer on the Ti surface after the first step of the diazonium

treatment (Scheme 5.2.d): indeed, the samples from group D showed 4.3% N on their surface,

and an increased C content from 17.9% (control samples) to 64.2% (Figure 5.3.b). Ti content

was decreased from 20.5% in control samples to 6.2% in group D confirming the presence of the

PAP layer covering the Ti surface. The high resolution C 1s spectra (Figure 5.3.c) showed that

the components relative to hydrocarbon (C-C/C-H, centered at 285.0 eV), C-O and C-N groups

(centered at 286.4 eV), and carboxyl groups (O-C=O, centered at 288.8 eV) changed from 61.2,

26.7, and 12.1% respectively in control samples to 79.6, 18.3, and 2.1% respectively in group D.

The drastic decrease of the carboxyl groups is especially indicative of the formation of the PAP

layer, since no carboxyls should be present in this layer.

Contact angle measures gave more evidence of the grafting of the PAP layer on Ti surfaces

(Figure 5.4). A DW contact angle of 53.5±14.8° was measured for polished titanium before

75

treatment; the contact angle changed to 84.7±3.5° for the samples of group D, thus confirming

the formation of the organic, hydrophobic PAP layer on these samples. The presence of the PAP

layer increased the bond strength with PMMA, from 1.54 (control) to 2.33 MPa (Figure 5.5.a).

This increase might be due to some entanglement between the PMMA chains and the PAP

layers. These results indicate grafting of the PAP on the metallic samples was achieved

successfully; however, the mechanical performance of this coating was limited.

Figure 5.3: (a) XPS general surveys and (b) elemental compositions of Ti surface for different groups. (c) High

resolution C 1s spectra on Ti surfaces for different groups; and (d) peaks concentrations from high resolution C 1s

spectra. See Table 5.2 for what the groups are. * indicates significant difference between the different groups (p< 0.05).

76

Figure 5.4: Photographs of water droplets placed on different Ti groups; the contact angle for control group was

53.5±14.8°; D= 84.7±3.5°; D+M= 84.1±1.9°; D+M+E= 83.1±0.5°; D+M+E+I= 82.4±0.9°. See Table 5.2 for what the

groups are * indicates significant difference between the different groups (p< 0.05).

Figure 5.5: (a) Tensile strength of the bond between PMMA and treated Ti surfaces. See Table 5.2 for what the groups

are. (b) Tensile strength of the bond between PMMA and Ti as a function of BP concentration in the aqueous phase of

MMA emulsion. * indicates significant difference between the different groups (p< 0.05).

The second step of the treatment (adhesive) was designed to change the amino ends of the metal-

PAP layer (-C6H4-NH2) into diazonium radicals (-C6H4-N2•) and then grow a few layers of

PMMA on it. As a first attempt, together with the reactants used to achieve the reduction of the

NH2 group into the diazonium radical, we added the MMA monomer alone (group D+M). This

did not lead to the formation of PMMA on top of the PAP layer. In fact, XPS and contact angle

results for the group D+M are quite similar to those of group D (Figures 5.3 and 5.4), and the

mechanical tests showed almost identical bond strengths for the groups D+M and D (Figure

77

5.5.a). The reason for the failure to grow PMMA in this condition is that MMA is a hydrophobic

monomer, hardly soluble in the aqueous solution used to modify the Ti surface.

XPS and mechanical tests indicate that the polymerization of PMMA in group D+M+E, which

includes the addition of SDS to solubilize MMA in the second step, was better than in group

D+M [43, 45]. We added SDS with a concentration of 9x10-3

M; this concentration was above

SDS critical micelle concentration (8.2x10-3

M [42]), so that SDS could disperse MMA in the

aqueous reaction solution by forming micelles around MMA droplets [43, 46]. The droplet size

of such micelles was reported to be in the range of 30 to 100 nm [47].

XPS confirmed the presence of PMMA on the grafted layer in group D+M+E by showing an

increase in oxygen up to 31.7% as well as a decrease in the concentration of Ti down to 3.9%,

indicating that PMMA was polymerized on the metal-PAP layer covering the Ti surface (Figure

5.3.b). Despite the slight decrease in overall C, the high resolution C 1s spectra for group

D+M+E showed an increase in O-C=O group concentration up to 9.1% compared to the previous

groups at concentration of 3.0%, thus confirming PMMA polymerization [48-50] (Figures 5.3.c

and 5.3.d). The contact angle in group D+M+E was 83.1±0.5° confirming the presence of

hydrophobic layer on these samples (Figure 5.4). The formation of a PMMA adhesive layer

increased the tensile bond strength between PMMA and the treated Ti in group D+M+E up to

3.4±1.2 MPa (P<0.05) (Figure 5.5.a). Most likely this was due to the entanglement achieved

between the PMMA chains in solution and those grown on the Ti-PAP surface thanks to the

better solubilization of the MMA monomer.

To further increase the bond strength between PMMA and Ti, we added BP to help

polymerization (group D+M+E+I). Increasing the length of the PMMA chain that grows on the

78

metal-PAP layer was critical to increase the strength of the bond between Ti and PMMA [51].

PMMA polymerizes by free radical addition polymerization, which requires the presence of an

initiator such as BP to start. BP is a relatively unstable compound, which forms radicals simply

upon heating or irradiation. BP radicals react with MMA and create MMA radicals, which then

propagate and grow longer and longer PMMA chains [31]. BP is the most commonly used

initiator for PMMA polymerization [52-54]. Thus, we added BP to the MMA emulsion in

different concentrations, and the bonding strength between Ti and PMMA increased remarkably.

In group D+M+E+I, XPS showed the N concentration was negligible after PMMA

polymerization with BP, which indicates that the PAP layer was covered by the thick layer of

PMMA formed (Figure 5.3.b). The formation of a thick PMMA layer is confirmed by the

increase in O content up to 42.6% on the surface of this sample and the increase in the O-C=O

component up to 17.6% in the high resolution C1s spectrum (Figure 5.3) [48-50]. Contact angle

measures in group D+M+E+I confirmed the formation of the hydrophobic PMMA layer on these

samples, and the contact angle was significantly difference (82.4±0.9°) compared to control

group (53.5±14.8°) (Figure 5.4). The addition of BP leads to the strongest tensile bond between

PMMA and Ti (8.14±1.10 MPa). This bond strength was significantly higher than that achieved

in any other group (P<0.05) (Figure 5.5). The resulting high bond strength indicates that

formation of a thicker PMMA layer containing long PMMA chains that can entangle very

strongly with the PMMA chains that are polymerized in the bulk PMMA casted on the sample.

The highest bond strength for titanium-PMMA (8.14±1.10 MPa) was obtained at BP

concentration of 40 mg/ml (Figure 5.5.b). At concentrations higher than 40 mg/ml, BP caused a

slight decrease on the bond strength between PMMA and titanium; this might be because BP

attacked the PAP layer grafter on the Ti surface.

79

A similar two step method including all the improvements for the second step described for the

Ti-PMMA bond was applied to bind PMMA and stainless steel wrought wire. The bonding

achieved in this case (4.34±0.68 MPa) was significantly higher than if it was left untreated

(1.71±0.23 MPa) (Figure 5.6). Acrylic removable partial dentures and orthodontic removable

appliances that combine stainless steel wrought wires and PMMA are usually cannot be made

when not enough volume of PMMA is available to support the wire [30, 31]. By increasing bond

strength between stainless steel wrought wire and PMMA through this treatment, more leverage

is possible for fabricating acrylic-based dental devices when not enough volume of acrylic is

available to support the wire (Figure 5.7).

Figure 5.6: Bond strength of PMMA and stainless steel wires for the control group and stainless steel wrought wire that

were treated with diazonium in MMA emulsion using the surfactant SDS and the initiator. * indicates significant

difference between the different groups (p< 0.05).

80

Figure 5.7: Schematic showing the acrylic removable partial denture with a small volume of PMMA to support the wire;

the short wire treated according to the proposed optimized two-step diazonium method should be stable enough in

PMMA to allow prosthesis retention.

5.5. Conclusion

The treatment of metallic surfaces (titanium and stainless steel) with diazonium ions in a two-

step procedure where the second step includes an emulsion containing monomer (MMA), an

emulsifier (SDS), and an initiator (BP) increase the bond strength of PMMA to Ti, and PMMA

to stainless steel wrought wire by 5.2 and 2.5 folds respectively compared to untreated control

groups. The bond strength achieved between polished Ti and PMMA is higher than that achieved

with any other method reported in the literature. The increased bond strength achieved between

Ti and PMMA is likely to be able to prevent dental prostheses failure. Also, it might help reduce

microleaks at the Ti/PMMA interface, thus preventing accumulation of oral debris and

discoloration of denture base materials and improving esthetics and oral hygiene [1, 2, 8]. This

bonding technique provide more leverage for fabricating acrylic-based dental devices when not

enough volume of acrylic is available to support the wire Although there are several commercial

methods to bind metals to PMMA for dental prostheses, the method proposed in this paper

significantly improved the bond strength between PMMA and polished Ti compared to all other

available methods. Further improvements may be obtained by combining this technique with

mechanical interlocking.

81

Acknowledgments

The authors would like to acknowledge King Saud University in Riyadh, Saudi Arabia; Natural

Sciences and Engineering Research Council (NSERC) of Canada–Discovery grant (F.T. and

M.C.); Canada Research Chair Foundation (M.C); and the Fondation de l’Ordre des dentists du

Québec (FODQ), Le Réseau de recherche en santé Buccodentaire et osseuse (RSBO) for their

financial support. Thanks to Enrique Lopez Cabarcos and Xuan Tuan Le for their technical

support.

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85

Chapter 6: Manuscript II

Surface Chemical Treatment of Orthodontic Brackets for Improved Tooth Adhesion

Omar Alageel a, Paige Kozak

a, Mohamed-Nur Abdallah

a, Jean-Marc Retrouvey

a, Marta

Cerrutib, Faleh Tamimi

a

a Faculty of Dentistry, McGill University, Montreal, QC, Canada

b Department of Mining and Materials Engineering, McGill University, Montreal, QC, Canada

6.1. Abstract

Adhesives for orthodontic brackets, such as Bis-GMA (Bisphenol A-glycidyl methacrylate)

composite resin, adhere strongly to teeth but very poorly to metallic brackets resulting in

frequent bracket deboning. To compensate for the lack of adhesion, metal orthodontic brackets

were developed with large bases that have an unfavorable impact on esthetics and hygiene.

Objective: The purpose of this study was to develop a new surface treatment, based on

diazonium chemistry that facilitates chemical bonding between metallic brackets and Bis-GMA

composite. Methods: Three models of stainless steel brackets were first coated with diazonium

ions to allow covalent binding on their surfaces. The brackets were then immersed in an

emulsion of Bis-GMA monomer and an initiator to build up Bis-GMA polymer chains on the

surface of the diazonium layer. The chemical composition of the treated metal brackets was

analyzed by X-ray photoelectron spectroscopy (XPS) and the tensile and shear bonding strengths

between Bis-GMA composite and surface-treated metallic orthodontic brackets were measured.

86

Results: XPS result confirmed the presence of a treatment coat on the metallic brackets and the

bond strength between these coated brackets and Bis-GMA was increased after the treatment by

2 to 3.9 folds compared to untreated brackets. Conclusion: The surface treatment method

proposed in this study can be utilized to reduce bracket debonding and to decrease the current

surface area of metal brackets by at least 50%.

Keywords

orthodontic brackets; bonding; diazonium; aryl radical; bisphenol A-glycidyl methacrylate; Bis-

GMA; stainless steel.

6.2. Introduction

Esthetics is one of the main concerns for patient acceptance of any orthodontic appliance [1].

Although metal brackets are considered unesthetic, they have proven to be more reliable and

effective than plastic or ceramic brackets [2, 3]. Metal brackets are usually made of stainless

steel, and are bonded to teeth using adhesive [4]. The most popular bonding system for

orthodontic brackets is a Bis-GMA (Bisphenol A-glycidyl methacrylate) composite [4, 5]. This

composite consists of an organic matrix (Bis-GMA), an inorganic filler, a coupling agent and an

accelerator system that can be polymerized quickly via a free radical polymerization mechanism

using a chemical activator, a photo activator (light), or a dual activator [5, 6]. This bonding

system adheres very well to the enamel of teeth but very poorly to metallic brackets [7-15].

The weak bonding between brackets and composite is the main factor for bracket debonding that

is a common occurrence in daily clinical practice for average rates between 4.7 to 6 percent

(Figure 6.1) [13]. Brackets loss has many negative impacts on the treatment such as extending

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treatment duration and chair time, and costing more money for using new materials [4]. Bracket

loss is more common in complex orthodontic treatment, where the mechanics of severely

crowded or angulated teeth require strong forces to correct the malocclusion [16, 17]. In fixed

appliances therapy, bands are fitted around a tooth, usually on molars, to provide extra strength,

but they have unfavorable impact on patient comfort, esthetics and oral hygiene [4]. However,

using buccal tubes instead of bands would be more sufficient if they bonded well to teeth [4].

To compensate for this lack of adhesion between brackets and composite, orthodontic brackets

were developed with large textured bases designed to increase the surface area and mechanical

retention [7-9]. Larger brackets have a negative impact on patient satisfaction and oral health due

to increased food impaction, rates of demineralization and caries and reduced comfort and

esthetics [4, 7, 8, 14, 15].

Figure 6.1: Photographs showing bracket debonding at the interface between brackets and composite. a: bracket

debonded from a mandibular canine due to its weak bond to composite that remains on teeth ;b: brackets basses showing

little composite on the surface indicating the debonding was as a result of the weak bond of brackets to composite .

Most of the research conducted to increase the adhesion of brackets to composite has been

limited to modifications of the surface pattern of the bracket base in order to increase

micromechanical interlocking, and has only achieved minor improvements [7, 9, 18, 19].

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Accordingly, a chemical adhesive system that could increase the bond strength between

orthodontic brackets and dental composite would be of great interest for clinical orthodontic

practice because it could prevent brackets loss and result in a reduction of the size of brackets

that translates into less dental problems associated to them.

Aryl diazonium salts have been the subject of much recent research, and their ability to modify

many materials surfaces for many applications [20, 21]. The value of aryl diazonium salts lie in

its ease of preparation, rapid reduction, and to the strong covalent bonding [21, 22]. Diazonium

ions can be produced from aromatic amines and grafted onto almost any surface, including

metals, glass, and carbon [22-26]. Initially, diazonium grafting was performed using

electrochemical reduction, but recently this has been achieved using chemical reducing agents in

an acidic solution [27]. The reducing agents activate the aryldiazonium salts and form aryl

radicals that covalently bind to the surface of interest [28, 29]. If an extra amino group is present

on the aryldiazonium precursor, this group can be further activated in a second step, and used to

bind a second layer onto the original surface [28, 30].

The objective of this study was to develop a new surface treatment, based on diazonium

chemistry that facilitates chemical bonding between metallic brackets and Bis-GMA composite.

Improving bond strength between orthodontic brackets and composite resin will reduce bracket

debonding, provide more leverage in cases with complex mechanics, permit the use of buccal

tubes instead of using molars bands, and allow the use of brackets with smaller bases resulting in

fewer complications associated with esthetics and oral hygiene.

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6.3. Materials and Methods

6.3.1 Teeth

Twenty-four extracted human teeth were collected from adult patients with dental conditions that

required tooth extraction. The extraction procedure was performed in the McGill Undergraduate

Dental Clinic after obtaining approval from the McGill University Health Center Ethical

Committee and informed consent of the patients. Upon extraction, teeth were immersed in 10%

formalin solution (BF-FORM, Fisher Scientific, Canada) for 1 week. The teeth were then

cleaned with distilled water in an ultrasonic bath (FS20D Ultrasonic, Fisher Scientific, Canada)

for 60 min at 25oC and polished for 1 min with a low-speed dental handpiece (M5Pa, KAB-

Dental, USA) using SiC cups (Pro- Cup, sdsKerr, Italy) and dental prophylaxis pumice of low

abrasive capability (CPRTM, ICCARE, USA). Next, the teeth were rinsed again in an ultrasonic

bath before storing in labelled Eppendorf tubes with 10% formalin solution for further analysis.

6.3.2 Mold preparation

Teeth were cleaned with water and air-dried. Each tooth was partially embedded in a mold of

plaster of Paris. Teeth were oriented to be perpendicular to the surface of the plaster mold in

order to test shear strength.

6.3.3 Brackets

Three different commercial models of stainless steel orthodontic brackets were used in this

study: buccal tubes baseless brackets for maxillary first molars (n=60; surface area of 0.325 by

0.300 cm; Rocky Mountain Orthodontic, Denver, CO) referred to as group R (Figure 6.2.a);

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maxillary anterior teeth large brackets (n= 64; base dimension of 0.380 by 0.260 cm; 3M Unitek,

Monrovia, CA) referred to as group L (Figure 6.2.b); and mandible lateral teeth small brackets

(n=60; base dimension of 0.275 by 0.255 cm; 3M Unitek, Monrovia, CA) referred to as group S

(Figure 6.2.c). In each group, 40 brackets were used for tensile bond strength test between two

brackets and Bis-GMA, while 20 brackets were used for the shear bond strength test of brackets

bonded to teeth using Bis-GMA adhesive. All brackets were ultrasonicated in distilled water,

ethanol and acetone for 5 minutes prior to surface treatment.

Figure 6.2: Digital photographs illustrating the different types of brackets used in this study: (a) baseless buccal tubes

brackets (referred as group R); (b): 3M Unitek bracket with large basses (referred as group L), (c): 3M Unitek bracket

with small basses (referred as group S). Schematic drawing showing the preparation and mechanism of testing (d) tensile

and (e) shear bonding strength between brackets and Bis-GMA.

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6.3.4 Treatment Materials

P-phenylenediamine (PPD), sodium nitrate (NaNO2), sodium dodecil sulfate (SDS), benzoyl

peroxide (BP), Bisphenol A-glycidyl methacrylate (Bis-GMA), hypophosphorous acid (H3PO2),

and hydrochloric acid (HCl) were purchased from Sigma Aldrich (St. Louis, MO) and were used

without any further purification. Concentrated hydrochloric acid (HCl) was diluted in distilled

water to a concentration of 0.5 M.

6.3.5 Surface Treatment Process

Brackets were treated in two solutions (primer and adhesive). Both solutions were prepared at

room temperature in an acidic deionized water solution at pH≤2, since diazonium cations are

stable at pH ≤2.5 [28, 30]. The first solution (primer) was prepared by dissolving PPD (0.054g;

0.05M) and NaNO2 (0.034g; 0.05M) in a beaker containing 10 ml of 0.5 M HCl (Figure 6.3.a).

Then, hypophosphorous acid (H3PO2; 0.66 ml) was added to the primer solution as a reducing

agent to form the aryl radical (Figure 6.3.b). The stainless steel brackets were then introduced

into the primer solution and left to react for 15 minutes before ultrasonicating them in distilled

water and acetone for 5 minutes. This first treatment solution leads to the formation of a

diazonium layer on the stainless steel brackets (Figure 6.3.c) [28].

The second solution (adhesive) was prepared by dissolving NaNO2 (0.034g; 0.05M) and SDS

(0.026 g, 9x10-3

M) in a beaker containing 10 ml of HCl (0.5 M). SDS was added into the

solution as a surfactant to help solubilize the hydrophobic monomer solution [31, 32]. Then, the

hypophosphorous acid (H3PO2; 0.66 ml) was added to the adhesive solution (Figure 6.3.d). Next,

the brackets treated with primer solution were introduced into the adhesive solution and Bis-

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GMA (2.3g/2.0 ml of ethanol) was added. After 15 minutes of sonication, an initiator (benzoyl

peroxide (BP), 40 mg/ml) was added to accelerate the polymerization reaction on the metallic

surface (Figure 6.3.e). The reaction was allowed to continue for an additional 15 minutes in the

ultrasonic bath and 30 minutes on the bench top. During this time, the monomer polymerized and

formed a thin layer of Bis-GMA on the diazonium-brackets surface (Figure 6.3.f). Finally, the

brackets were ultrasonicated in distilled water and acetone for 5 minutes.

Figure 6.3: Schematic diagram of the reactions performed in the first and second solutions of Bis-GMA/diazonium

treatment; (a): diazotisation of p-phenylenediamine (PPD) in an acidic solution; (b): formation of aryl radicals by

reducing aryldiazonium using H3PO2; (c): attachment of aryl radicals to brackets; (d): reduction of amino groups of the

diazonium layer to radicals; (e): binding of Bis-GMA to the activated diazonium layer; (f): polymerization of Bis-GMA.

*The grafted layer in the scheme showed as single layer for simplifying but it could be single, multi or mixed layers.

6.3.6 X-ray Photoelectron Spectroscopy

A monochromatic X-ray photoelectron spectrometer K Alpha (VG Scientific ESCLAB 3 MIKII)

was used to determine the relative quantities and chemical environments of the constituent

elements on four stainless steel brackets of group L. The setup was equipped with Mg Kα X-Ray

radiation source (1253.6 eV), a micro-focused monochromator and an ultrahigh vacuum chamber

(10-9

torr). For all samples, survey scans were obtained over the range of 0-1350 eV with pass

energy of 100 eV at a step of 1.0 eV, while high resolution scans were collected with pass energy

of 20 eV at a step of 0.05 eV. All spectra were referenced to C-C/C-H of the C 1s spectrum at

285.0 eV.

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6.3.7 Mechanical Tests

The mechanical performance of the bonding between brackets and composite were assessed

using tensile and shear strength tests (Figure 6.2). For the tensile strength test, pairs of identical

brackets were bonded together with a light-cured Bis-GMA composite adhesive (Transbond XT

Light Cure; 3M Unitek, Monrovia, CA) and light cured for 30 seconds. The brackets were pulled

apart using a universal testing machine (Instron, 5569, Grove City, PA) set up at a constant speed

of 5 mm/min (Figure 6.2.d).

In order to evaluate the shear strength of the bond between brackets and Bis-GMA composite on

teeth, the buccal surfaces of teeth were first etched with an etchant gel (EZ Etch; Dentsply, York,

PA) for 30 seconds, rinsed with distilled water and air-dried. Then, the Transbond adhesive

primer (3M Unitek, Monrovia, CA) was applied to the tooth surface. Next, Bis-GMA composite

adhesive (Transbond XT Light Cure; 3M Unitek, Monrovia, CA) was applied to the base of a

single bracket, positioned on the tooth surface before being light cured for 30 seconds. A load

applied parallel to the long axis of the tooth with the universal testing machine, as described

earlier, generated a shear force at the interface between teeth and brackets (Figure 6.2.e).

6.3.8 Statistical Analysis

The average of the X-ray photoelectron spectroscopy and bond strength with associated standard

deviation was calculated for each group. The statistical analysis was performed using Origin 8.0

(Origin lab, Northampton, MA). All data were analyzed using t-test and Kruskal-Wallis test and

the significance level was set at p < 0.05.

94

6.4. Results

6.4.1 X-ray Photoelectron Spectroscopy (XPS)

X-ray photoelectron spectroscopy (XPS) results confirmed the presence of a diazonium/Bis-

GMA layer on treated bracket surfaces. The surface elemental composition for the stainless steel

brackets changed after diazonium/Bis-GMA treatments (Figure 6.4.a). The elements present on

the untreated stainless steel (control) brackets were carbon, oxygen, chromium, and iron;

however, only carbon and oxygen were present after treatment. The carbon content (C 1s; at

285.0 eV) increased from 43.2% in the untreated group to 81.4% after treatment, while oxygen

content (O 1s; at 532.0 eV) decreased from 43.2% in the untreated group to 17.2% after

treatment. Chromium (Cr 2p; at 576.5 eV), and iron (Fe 2p3; at 708.5 eV) were present only on

the untreated brackets in concentrations of 9.6% and 4.4% respectively.

The high resolution carbon (C 1s) spectra for the untreated brackets showed different carbon

peaks from the treated brackets (Figure 4.b). The untreated group included the C-C, C-O/C-O-C,

and O-C=O carbon peaks in concentrations of 31.3, 5.4, and 4.2%, respectively, while in the

treated group the peaks were C=C, C-C, C-O/C-O-C, and O-C=O with concentrations of 37.9,

22.3, 16.8, and 3.3%, respectively.

95

Figure 6.4: (a): XPS general surveys and elemental compositions for the untreated (control) and treated brackets; (b):

high resolution XPS C 1s spectra for the untreated and treated brackets. * indicates significant difference between groups

(p< 0.05).

6.4.2 Mechanical Tests

The mechanical tests showed that the tensile and shear bond strengths between Bis-GMA

composite and brackets treated with diazonium/Bis-GMA were significantly higher than

untreated (control) brackets for all bracket groups (Figures 6.5). Tensile bonding strengths

between Bis-GMA and treated brackets for groups R, L, and S were 4.62±0.87, 4.15±1.00, and

4.49±1.30 MPa, respectively, and they are significantly higher than untreated brackets

(1.41±0.64, 1.57±0.99, and 2.22±1.39 MPa respectively) (Figure 6.5). The shear bonding

strengths of brackets in groups R, L, and S that were treated with diazonium/Bis-GMA and

bonded to teeth with Bis-GMA composite were significantly higher 3.83±1.47, 6.67±2.52 and

6.71±2.51 MPa, respectively, than the untreated brackets 1.94±1.25, 1.84±1.07, and 1.72±1.06

MPa (Figure 6.5).

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From the mechanical tests results, tensile and shear bond strengths for the untreated (control)

brackets to Bis-GMA composite were almost similar and not significantly different between

control groups in R, S, and L. However, the baseless brackets in group R are better than brackets

with bases (groups S and L) in the tensile bond strength after diazonium/Bis-GMA treatment

while brackets with bases in groups S and L are better than the baseless brackets (group R) in the

shear bond strength after diazonium/Bis-GMA treatment.

Figure 6.5: (a):The ultimate tensile force N (Left) and bond strength MPa (Right) of the different stainless steel brackets

R, L, and S that were untreated (control) or treated with diazonium/Bis-GMA treatment and bonded to another bracket

with Bis-GMA composite. (b):The ultimate shear force N (Left) and bond strength MPa (Right) of the different stainless

steel brackets R, L, and S that were untreated (control) or treated with diazonium/Bis-GMA treatment and bonded to

teeth with adhesive based on Bis-GMA. * indicates significant difference between the different groups (p< 0.05).

97

6.5. Discussion

Numerous studies have been done to improve the bonding strength between polymers and metals

[33-36]. However, very little studies have been reported for improving the bonding between

metallic brackets and composite adhesive aside from studies based on micromechanical

interlocking with textured bracket bases [9, 18, 19]. Within the limit of our knowledge, this is the

first study improved the chemical bonding between brackets and adhesive composite.

The chemical bonding between metallic brackets and adhesive composite in this study was

achieved in two solutions. The first solution (primer) was formed a grafted layer of

aryldiazonium salts on the bracket, and the reaction started by diazotisation of p-

phenylenediamine (PPD) and sodium nitrite (NaNO2) in acidic solution (Figure 6.3.a). Then, the

aryldiazonium salt was reduced chemically by adding hypophosphorous acid (H3PO2) as a

reducing agent to form aryl radicals (Figure 6.3.b), that bind to metallic brackets (Figures 6.3.c)

[28]. The second solution (adhesive) was designed to change the amino ends of the grafted layer

(-C6H4-NH2) into diazonium radicals (-C6H4-N2•) using H3PO2 (Figure 6.3.d), and then grow

layers of Bis-GMA on it (Figure 6.3.e). Bis-GMA chains were longer after adding the benzoyl

peroxide (BP) that would achieve high bond strength to adhesive composite (Figures 6.3.f).

The presence of diazonium/Bis-GMA treatment on the stainless steel bracket surfaces was

confirmed by XPS. Chromium (Cr 2p) and iron (Fe 2p3), which are the characteristic

constituents of stainless steel alloys, were only showed in the XPS survey of the untreated

brackets. Carbon (C1s) in the untreated group was probably was probable due to the unavoidable

air contamination upon exposed to air prior to XPS analysis since the high resolution carbon

(C1s) spectra revealed the presence of carbon peaks at C-C, C-O, and O-C=O which is consistent

98

with samples of air alone [37]. As expected, the total amount of carbon (C1s) found on the

treated group was higher than that found in the untreated (control) group while chromium and

iron concentrations became negligible most likely due to the presence of a Bis-GMA layer on

bracket surfaces. Furthermore, the high resolution carbon (C1s) spectra for the treated brackets

presented carbon peaks (C=C, C-C, C-O, and O-C=O) that corresponding to pure Bis-GMA

samples (Figure 6.4) [5].

The bonding between stainless steel brackets and Bis-GMA was assessed with two different

debonding tests. The tensile bond strength test measured the direct bond strength between the

bracket and Bis-GMA composite resin while the shear bond strength test mimicked the

debonding of brackets on teeth [38]. Brackets used in this study were either baseless (group R) or

with bases (groups L and S). Typically, bracket bases are designed with micromechanical

interlocking to increase the surface area of brackets and the bonding between brackets and Bis-

GMA composite [9, 38, 39]. The mechanical tests showed that brackets bases gain minor

improvements for tensile and shear bond strength between brackets and Bis-GMA composite

before the treatment, and they were not statistically significant to baseless brackets group.

However, brackets bases help for increasing the shear bond strengths after using diazonium/Bis-

GMA treatment.

Diazonium/Bis-GMA treatment increased the tensile bond strength between Bis-GMA resin and

brackets by 3.3, 2.6, and 2.0 folds in groups R, L, and S respectively (p< 0.05) while the

treatment increased shear bond strength by 2.0, 3.6 and 3.9 folds respectively (p< 0.05) for

similar groups (Figure 6.5). The increase of tensile bond strength in baseless brackets (group R)

was higher than in brackets with bases (group L and S), while the increase of shear bonding

99

strength for brackets with bases (group L and S) was higher than baseless brackets (group R).

This means that baseless brackets (group R) benefit more from the diazonium/Bis-GMA

treatment in tensile bond strength that assess direct bonding between brackets and composite [9,

38]. Whereas, brackets with bases (groups L and S) benefit from the treatment in shear bond

strength probably because the bases was designed with micromechanical interlock providing

high retention to composite and exhibiting more friction during shear testing [9, 38].

Increased bond strength between brackets and Bis-GMA composite through diazonium/Bis-

GMA treatments can be utilized to reduce the debonding of orthodontic brackets, a common

occurrence in clinical practice that prolongs treatment duration and chair time [13]. Thus, fixed

appliances therapy would be more effective and resistance to failure after using diazonium/Bis-

GMA treatment on brackets than without treatment. By increasing bond strength between

brackets and composite, more leverage is possible for complex orthodontic mechanics that

usually failed because the debonding [16, 17]. Moreover, the bands around teeth in fixed

orthodontic treatments could become obsolete when the strong bonding of buccal tubes to molar

teeth is attained using diazonium/Bis-GMA treatment [4]. Using buccal tubes in the orthodontic

treatment provide better esthetics, comfort, and oral hygiene than using molars [4].

Increased bond strength between brackets and Bis-GMA composite can also be applied to

bracket design by decreasing bracket size from 50 to 74% of its original size. S brackets are the

smallest available metallic brackets produced by in 3M Unitek; and they are designed for the

mandible anterior teeth while L brackets are designed for the maxillary anterior teeth with

corresponding bracket bases 28% larger than those of the S bracket. According to our results, it

would be possible to use S-sized brackets on the maxillary anterior teeth (Figure 6.6.b) instead of

100

using the L bracket (Figure 6.6.a). Moreover, the manufacturers of orthodontic brackets would

be able to design brackets with bases of half their current size, or smaller, according to the

bonding strength results presented in this study (Figure 6.6).

Using small brackets size is beneficial for patients and orthodontists; it can improve the patients’

primary concern of esthetics while undergoing fixed orthodontic treatment [1]. Smaller brackets

are also more comfortable for patients since they cause less oral tissue irritation than larger

brackets. Using small orthodontic brackets provides better oral hygiene than large one because

food impaction around and beneath the brackets will be reduced and cleaned easily preventing

tooth caries [7, 14, 39, 40]. The small amount of residual composite on the teeth after finishing

orthodontic treatments using small brackets could helps orthodontists to reduce the chairside

time when removing the residual composite from the teeth; and it also minimize the negative

impacts of the residual composite on teeth when have smaller amount of demineralization and

less etched area of teeth [7, 14].

101

Figure 6.6: Photograph of untreated L bracket group (a) and treated S bracket group (b) that were bonded to the

anterior teeth using Bis-GMA composite. The bracket in (b) was treated with diazonium/Bis-GMA and its shear bond

strength (47.65±17.85 N) was 364% higher than the shear bonding strength (18.03±10.50 N) of bracket (a) while its size is

28% less and the size of (a). (c) drawing shows fixed orthodontic appliances with the original size of metallic brackets

bonded on teeth, and (d) brackets with bases 50% smaller size than original brackets as a results of using diazonium/Bis-

GMA treatment.

6.6. Conclusion

The treatment of the metal orthodontic brackets with diazonium ions in an emulsion containing

Bis-GMA (Bisphenol A-glycidyl methacrylate) monomer, an emulsifier (SDS), and an initiator

(BP) increased the bond strength between brackets and Bis-GMA composite. The

diazonium/Bis-GMA treatment increased the bonding strength between Bis-GMA and brackets

by 2 to 3.9 folds compared to untreated brackets. Although brackets bases provide minor

improvement of the bond strength between brackets and composite, it helps for increasing the

shear and tensile bond strength after using diazonium/Bis-GMA treatment. This bonding

technique could be used to prevent bracket debonding, treat complex orthodontic cases requiring

increased leverage, and decrease the size of metal brackets by at least 50%.

102

Acknowledgments

The authors would like to acknowledge King Saud University in Saudi Arabia, Riyadh; Natural

Sciences and Engineering Research Council (NSERC) of Canada–Discovery grant; the

Fondation de l’Ordre des dentists du Québec (FODQ), and Le Réseau de recherche en santé

Buccodentaire et osseuse (RSBO) for their financial support.

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titanium. Dent Mater J, 2006. 25(1): p. 120-124.

36. Koizumi, H., et al., Bond strength to primed Ti-6Al-7Nb alloy of two acrylic resin

adhesives. Dent Mater J, 2006. 25(2): p. 286-290.

37. Cai, K., et al., Surface structure and composition of flat titanium thin films as a function

of film thickness and evaporation rate. Appl Surf Sci, 2005. 250(1): p. 252-267.

38. Iizuka, Y., et al., Bond strength of an orthodontic bonding material and adhesion energy

of artificial saliva to an experimental titanium bracket. Orthod Waves, 2011. 70(1): p. 21-

26.

39. Bishara, S.E., et al., Effect of time on the shear bond strength of glass ionomer and

composite orthodontic adhesives. Am J Orthod Dentofacial Orthop, 1999. 116(6): p. 616-

620.

40. Schaneveldt, S. and T.F. Foley, Bond strength comparison of moisture-insensitive

primers. Am J Orthod Dentofacial Orthop, 2002. 122(3): p. 267-273.

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Chapter 7: Conclusion

Metallic surface treatment with diazonium ions can be used as substrate for polymerizing dental

polymers such as PMMA and Bis-GMA. This process results in direct covalent bonding between

dental polymers and metals that can be used to improve mechanical strength of bond between Ti

and PMMA, stainless steel wrought wire and PMMA, and stainless steel brackets and Bis-GMA

by 5.2, 2.5, and 2.0 folds respectively compared to untreated control groups. Increased bond

strength between metals and polymers used in dental devices would help to prevent their

frequent debonding and subsequent failures. This new bonding technique achieved strong

binding between metals and PMMA, and its bond strength on polished Ti to PMMA was the

strongest among the available bonding systems. The treatment can be used to make acrylic

removable dentures and appliances when only small volume of acrylic is available to support the

wire. This bonding technique based on diazonium could be used to prevent bracket debonding,

treat complex orthodontic cases requiring increased leverage, and decrease the size of metal

brackets by at least 50%.

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Chapter 8: Appendices

8.1. Report of Invention

A report of invention was received on November, 27, 2012 for the invention entitled “bonding

between metals and polymers”.

8.2. Poster I

A poster was presented at Dentistry Research Day- McGill University on March 29, 2012.

8.3. Poster II

A poster was presented at Dentistry Research Day- McGill University on March 26, 2013.

8.4. Poster III

A poster was presented at Materials Research Society (MRS) fall-meeting 2013 at Boston, USA

on December 4, 2013.

8.5. Poster IV

A poster was presented at the Oral and Bone Health Research (RSBO) Scientific Day at the

Estrimont in Orford, QC, Canada on January 17th

, 2014.

8.6. Poster IV

A poster was presented at the American Association for Dental Research (AADR/CADR)

Annual Meeting, Charlotte, USA on March 20th, 2014.

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