Anchoring Implants into Bone An in vivo approach

Graduate School for Cellular and Biomedical Sciences

University of Bern

PhD Thesis

Submitted by

Jens David Langhoff

from Germany

Thesis advisor

Prof. Dr. Brigitte von Rechenberg Musculoskeletal Research Unit

Equine Hospital Vetsuisse Faculty of the University of Zrich

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Table of content

Abstract..................................................................................................................................1

I Introduction.....................................................................................................................2

1 Implant function ...............................................................................................................2

1.1 Transient implant use .............................................................................................2

1.2 Permanent implant use ...........................................................................................3

2 Mechanisms and principles of implant anchoring..............................................................4

2.1 Mechanical anchoring and primary stability ...........................................................5

2.2 Screw type implants ...............................................................................................6

2.3 Use of bone cement................................................................................................6

2.4 Implant molding.....................................................................................................7

3 Bone biology ....................................................................................................................8

3.1 Bone healing ........................................................................................................11

3.2 Osseointegration ..................................................................................................12

4 Implant materials ............................................................................................................13

4.1 Steel.....................................................................................................................14

4.2 Titanium ..............................................................................................................14

4.3 Titanium passivation layer ...................................................................................15

4.4 Polylactides..........................................................................................................16

5 Methods of osseointegration analysis ..............................................................................18

5.1 Biomechanical evaluation ....................................................................................18

5.2 Morphological evaluation.....................................................................................20

6 Animal models for osseointegration................................................................................24

6.1 Ethical considerations ..........................................................................................24

6.2 Experimental animals...........................................................................................25

II Rationale of the research projects ..................................................................................29

III Results ..........................................................................................................................34

1 Published paper of Study 1 .............................................................................................34

2 Manuscript of Study 2.....................................................................................................44

IV Discussion and conclusions ...........................................................................................74

1 Improved removabity by surface modification ................................................................74

2 Enhancement of osseointegration....................................................................................75

3 Outlook ..........................................................................................................................77

V Annex ...........................................................................................................................79

1 List of tables ...................................................................................................................79

2 List of figures .................................................................................................................79

3 Bibliography...................................................................................................................79

VI Acknowledgments.........................................................................................................99

VII Curriculum Vitae.........................................................................................................100

Abstract

1

Abstract

The key problem of orthopedic and oral surgery is to reach an appropriate implant stability.

The ideal level can be either moderate or strong anchoring for a specific application.

The goal of the projects was to extend the limits on both sides: Firstly towards moderate

implant anchorage by modification of titanium surfaces and secondly increased anchoring

strength via increased immediate stability, by an innovative implant concept.

The research of project 1 aimed to achieve better removability of implants for transient use.

Modified titanium surfaces were created through different anodization procedures. The new

surfaces were harder compared to standard surfaces. Implants with hard surfaces would be

more resistant to scratches and wear. Differences in surface roughness and chemistry aimed to

induce different bone tissue reaction and subsequent variations in osseointegration to find the

surface with lowest removal force for transient implant use. Screw type implants were used to

test a clinically relevant design in a sheep hind limb model. The performance of the surfaces

were evaluated after 2 months for anchoring strength and bone-surface attachment. The new

surfaces were referenced by a standard titanium surface and a machined stainless steel implant

type. The results showed no significant differences between all titanium surfaces. They

remained with clearly higher bone bonding strength than the stainless steel reference. It can

be concluded that the difference in the surface characteristics were not as pronounced as

necessary to influence the bone response.

In project 2, a novel implant concept was tested to increase the primary stability and

osseointegration of permanent implants. The BoneWelding hybrid concept used implants

consisting of a cylindrical titanium core with resorbable polymer parts attached to the surface.

The implant was inserted with ultrasonic assistance, which molded the polymer into the bone

cavities during insertion. The resolidified polymer provided instant stability and rotation

stability, whereas the titanium core offered large contact area for osseointegration. The hybrid

implant concept aimed to increase the immediate and long term stability of the implants

without loss of intermediate stability. The concept was evaluated in a large animal study,

using a novel sheep pelvis model for implant insertion. The study included several evaluation

points to assess bone healing and implant osseointegration from acute to long term phases.

The implants were evaluated biomechanically or histologically. The hybrid implants always

showed superior performance compared to the controls, especially in the early phases.

Anchoring strength constantly increased with time. The results clearly demonstrated the

BoneWelding hybrid concept as a functional for improved immediate and long term stability.

Introduction

2

I Introduction

1 Implant function

The clinical indication makes a difference to the duration of functionality. Some implants are

intended to fulfill their function for a limited period (transient) while others should work for

the patients lifetime (permanent function). The transient use is typical for osteosynthesis or

reattachment of tendons and ligaments. They should temporarily help the bodies own attempts

of regeneration. Permanent implants are used to replace skeletal structures (e.g. joints or

teeth), when their function could not be restored by other means.

1.1 Transient implant use

Implants, which are used for a limited period, are mainly made to help or enable bone healing

in a physiologic position until the body tissue is ready to do so on its own. Requirements for

the function of an implant for transient use are a biomechanical competence for weeks or

months and basic biocompatibility. The material and surface should allow osseointegration to

provide implant-bone stability. After completed healing, supportive implants should not

remain in the body any longer. If the implant support persists, it will keep the bone from

taking over the full function. This phenomenon is called stress shielding. The reason for this

are the fundamental principles of bone remodeling, which needs mechanical stimulus [1]. It is

discussed as a combination of true mechanical shielding and impairment of vascular supply

[2]. The effect of stress shielding may range from total absence of fracture union to slightly

delayed union or continuous resorption of bone of unloaded bone [3]. The removability is

therefore another essential requirement of transient implants. This results in a conflict of the

ideally unlimited osseointegration strength and the limited material strength. Bone or implant

failure during removal is a major complication, which either leaves a foreign body in the body

or requires large surgical intervention for removal [4,5].

The main anchoring element of transient implant systems is the bone screw that either

connects bone to bone, or bone to another supportive element. These elements of an implant

system are commonly plates for internal fixation or connected bars, rods and/or ring

constructs for internal or external fixation [6].

Some implants have a transient use, but often they are not removed if not hampering the

healing success and removal would be too invasive. This is the case for implants for

reattachment of ruptured or torn out ligaments, tendons or their replacements [7,8]. A

common indication is the replacement of the anterior crucial ligament is a well known

standard procedure in sports medicine. Interference screws are used to clamp the ligament (or

Introduction

3

tendon graft) in a drilled bone canal [9,10]. In other procedures suture anchors are very

common, like the biceps tendon re-fixation [11] or the shoulder rotator cuff repair [12].

Another group of surgical interventions aim to correct deformities in the area of long bones.

Varus rotational osteotomy is a treatment to bring the femur head to a functional position in

patients with hip dysplasia [13]. Also for the realignment of long bones, implants are used to

fix the bony parts at a defined distance or keep the distance during healing as long as the

desired distraction is reached for example the tibial varus correction procedure [14]. In the

craniomaxillofacial field, cranioplasty can offer release of craniosynosthosis in children and

distraction osteotomy in the mandible corrects retrognathia [15].

1.2 Permanent implant use

Permanent implants were designed to replace skeletal elements and to mimic the biological

original as good as possible to restore function and relieve from pain. After completed

osseointegration they remain in the body and become an integral part of the bone structure.

The challenge is to reach this level of perfection of original bones or joints, which represent a

highly optimized state. The self adaptive and regenerative capacity of living tissue could also

not be provided by artificial replacements. Therefore, implants for permanent use require

biomechanical competence for years, while maintaining a high level of biocompatibility. The

target of permanent use is a highly integrated implant. The related challenge is to connect the

artificial replacement to the host bone with a functional interface by implant design and

choice of bone bed quality in the available skeleton location. The optimal interface location is

different for each bone and determines the amount of tissue to be replaced as well as the

complexity of the implant or implant system. Long term osseointegration is threatened by

material failure or late implant loosening [16,17]. Material failure is the consequence of

corrosion and fatigue under the extreme conditions of body fluid and highly dynamic loading.

Septic or aseptic loosening affects the biological counterpart of the implant in several possibly

pathways, all resulting in loss of bone substance replaced by non-functional soft tissue.

Teeth replacing dental implants are relatively simple as a single intraoseous piece, with an

extra-osseous abutment to connect the crown (Fig. 1). They already reach 10-year survival

rates of 95% [18]. A large field of implant anchoring is the total joint replacement (Fig. 1).

The hip prosthesis, as an example, replaces the acetabulum of the pelvis and the femoral head.

The implant system can easily be composed of 10 different pieces. The connections between

them inflict mechanical weak points, production of wear debris and subsequently higher rates

of complications. The consistent higher mechanical load placed on these systems certainly

Introduction

4

contributes to failure rates as well, but demonstrated very low revision rates during the first 10

years after operation [19].

Fig. 1: Classical examples of total joint replacement: knee, hip and dental prosthesis.

Further examples for successful permanent implants under highly loaded conditions are

devices used in spine surgery. They aim to stabilize or replace single vertebral bodies,

intervertebral discs or even provide instrumentation for a large segment of the spinal column.

The replacement of a vertebral body is necessary in cases of bone cancer or collapsed

vertebrae in consequence of metabolic diseases with impairment of the bone structure [20,21].

In cases of degenerated intervertebral discs and related persistent back pain, two main

strategies are followed. Either the disc shall be replaced by bone and two segments are

instrumented to fuse [22,23], or newly developed artificial discs are implanted [24,25].

Furthermore, deformities of the vertebral column can be repositioned and supported the

physiological configuration. A surgical intervention with subsequent implant support offers

the possibility of axis correction in severe scoliosis, as an ultimate treatment, if conservative

approaches are not efficient anymore [26,27]. There are numerous other treatments in

orthopedic and cranio-maxillo-facial surgery which would not be possible without permanent

implants.

2 Mechanisms and principles of implant anchoring

Historically, the methods for anchoring implants to bone originate from metal and wood

engineering. The early approaches aimed to adapt the geometries of the devices and its

material properties to bone. The focus was placed on providing sufficient stability for bone

fragments to provide the general condition for functional bone healing (see chapter 3). The

materials were expected to be bio-inert and to meet basic biocompatibility requirements in the

Stem

Femoral component

Plastic surface

Total knee replacement

Cup Crown

Abutment

Osseal portion

Gingival portion

Total hip replacement Dental implant

Ball

Introduction

5

body, but not necessarily more. However, implant loosening and material incompatibility

issues promoted the search for better materials and to study the interaction of the implant and

host tissue more in depth, leading to the questions of surface topography and chemistry.

2.1 Mechanical anchoring and primary stability

Mechanical stability (or primary stability) of conventional implants is reached by tensioning

the device within the bone bed [28]. A pure form fit would not provide sufficient shear force

to prevent the implant to migrate out of the implant cavity (backing out). This applies for

screw type implants and hammered wires, nails or shafts. The counterpart of the implant is the

bone bed. Its dimension, density and structure are the pre-determinants for the applicable

tensioning forces. In the majority of procedures, the bone is predrilled and pre-tapped to

create a defined cavity dimension, but leaves a random contact pattern to the free trabecular

ends in cancellous bone. For the following comparison to alternative methods it is important

to mention that the tensioning is applied mostly on a rather small contact area between

implant and bone due to differences in the cavity and implant geometry.

The key problem of the instant stability created by tensioning is that tension disappears

quickly. The reason is the material based relaxation of bone [29]: it reacts viscoelastic. This

loss of stability cannot be compensated by higher initial tensioning forces without the risk of

destroying the integrity of the material bone [30]. Cortical bone tends to form cracks, which

might lead to fissures and ultimately to fracture of the entire bone. Breakage of the trabecula

in cancellous bone does not lead to fracture of the entire bone, but reduces the initial

anchoring strength of the implant. The biology of bone tissue also limits the applied forces

with the appearance of pressure necrosis - leading to bone resorption and following non-

integration [31].

The initial stability and the critical load highly depends on the bone quality. The basic bone

qualities are either the cortical bone or the highly porous cancellous bone, but in many

skeletal sites there is a seamless transition [32]. The relative bone volume (total volume / bone

matrix volume) gives a rough indication, but moreover the trabecular architecture, collagen

fiber orientation and degree of calcification are important predictors for bone stability. There

is a large range within those parameters depending also on the individual [33].

However, the mechanism of bone healing is similar in all bone locations. This is based on the

mechanisms of the constantly ongoing process of reconstruction, which aims to adapt and

optimizes the tissue to the current loading situation.

Introduction

6

2.2 Screw type implants

Screw type implants are used as bone or interference screws, suture anchors, or modified

adental implants [9,34,35]. Their working principal is a combination of macroscopical

interlocking and a geometry, which applies the tensioning force in a very controlled fashion

during insertion. The applied tensioning force can be determined by the difference between

thread geometry and tapped bone cavity [36]. Many modern implants combine screw and

taper by providing a tapping tip (self-tapping implants)[37] or some are even self drilling

[38]. Screw geometries range from cylindrical to conical and are single or double threaded,

differently threaded at the proximal and the distal part (Osstem GS System). Numerous

designs were developed to reach an optimal primary stability while providing a maximal bone

response.

2.3 Use of bone cement

Bone cements were developed to support implants in the implant cavity by filling gaps

between bone and implant or within spaces of the adjacent bone. Two different indications for

use of cement can be distinguished. Either bone cavity preparation is hard to match to the

implant geometry [39] or bone bed stability is not sufficient for implant anchoring [40].

In the first case, a lack of implant and bone cavity congruency may lead to insufficient

primary stability. If complex implant geometry is required to reach the required mechanical

strength, e.g. of the femoral shaft of a hip prosthesis, it is not yet possible to machine the bone

bed with the same precision as the medical device. This bears the risk of micromotion with

subsequent implant loosening [41]. Therefore cements were developed to fill this gap as an

adaptive interface material. This became a standard procedure for artificial joint prosthesis

fixation and is also applied for pedicle screws.

The second use of bone cements is to reinforce the bone bed stability in cases of low

anchoring potential of the bone bed by filling the bone marrow cavities near the implant.

Techniques have been developed to either inject cement into the implant cavity, or to apply

the cement via cannulated screws [42-44]. This type of cement use increased anchoring

strength significantly [45,46]. The latter concept is already in use for pedicle screws in spine

surgery [44] and femoral head fracture fixation [47].

Bone cements components are mixed shortly before implant insertion to a dough. It stays

viscous for the period of application and sets thereafter to a solid material. They are applied

by hand or specially designed cement guns onto the implant and pressed into the cavity [48].

The most common bone cements are based on non-resorbable methylmethacrylate (MMA)

[49,50]. A powder of polymethylmethacrylate (PMMA) and the liquid monomer polymerizes

Introduction

7

at room temperature with addition of benzoyl peroxide and tertiary aromatic amines as an

coinitiator {Serbetci, 2003 #662}.

The resorbable calcium hydroxy apatite cements were developed for bone defect filling, but

were also evaluated to increase implant stability [51]. Their mechanical strength is lower than

the PMMA cements [52], which limits their application in highly loaded situations.

Disadvantages of cements are chemical impact [53] or a temperature increase during

polymerization [54] with reduction of their overall biocompatibility [55].

2.4 Implant molding

A novel approach to insert implants is to mold the surface to the structure of the bone cavity

using modern polymers. The BoneWelding technology employs ultrasonic energy to

facilitate implant molding with the insertion process [56]. The ultrasonic vibrations cause

liquefaction of the polymer pin surface in contact with the bone during insertion (Fig. 2). The

liquid phase infiltrates the cancellous cavities of the adjacent bone and adheres to the

structures. Re-solidification initiates as soon as the supply of ultrasonic energy is stopped.

Like this, the implant adapts to the cavity instead of preparing the implantation site to the

shape of the implant. In contrast to screw type implants pre-tapping of the drill hole is not

required and saves surgery time [57]. The polymer forms a three dimensional interlocking

within the bone and achieved better stability compared to screws alone [58]. In combination

with plates it provides additional angle stability by fusing with the plate at the end of the

insertion way [59].

Fig. 2: The BoneWelding technology principle of ultrasonically assisted implant insertion.

Introduction

8

3 Bone biology

The primary function of bone is load bearing and distribution, which enables the body for

locomotion. Bones form the skeleton as a structure for the attachment and protection of

important organs. They play an important role for the calcium and phosphate metabolism and

offer space for haemopoesis in the bone marrow. The tissue consists mainly of matrix and

only few cells [60]. Histologically, bone belongs to hard tissues. Bone matrix can be

described a as a composite biomaterial of inorganic matrix (hydroxyapatite and

tricalciumphosphate 50-70%) and organic fiber material (collagen, 20-40%), water (10%) and

lipids (5%). The basic bone qualities are either the compact (cortical) bone or the highly

porous cancellous bone. In the body, also interstages exist as relatively porous cortical bone

(e.g. in the mandible) or very dense cancellous bone (eg. subchondral bone). Cortical bone is

a compact mass of bone matrix which only porosity is a network of narrow nutritive canals.

Cancellous bone is very porous. The trabecular spaces are filled with bone marrow (Fig. 3).

A high range of variability of the bone architecture exists depending also on the age,

individual and location [33,61-63]. Due to the specific mechanical load [64] and thus,

optimized configuration in each body location, material properties vary largely, e.g. the elastic

modulus from 3.81 to 19.6 [65,66].

Fig. 3: Bone macro-structure. Adapted from www.aaos.com

The low cell content of bone can be explained by low metabolic function of the tissue. Only

few cells are needed to renew the matrix and keep its function. Bone tissue needs constant

repair due to the weakening of the bone formation by micro-cracks [67]. Bone repair is driven

Periosteum

Harver s Canal Osteon

Cancellous bone

Cortical bone

Introduction

9

by the process of remodeling, which can start either with the removal of old tissue or

deposition of new bone. The cell species responsible for bone resorption are osteoclasts. They

are multinuclear giant cells, up to 100 m in diameter, which are fused giant cells from the

hemopoetic stem cell line (macrophages). Osteoclasts attach to the bone surface with a limited

part of their surface (Fig. 4) to which acidic vacuoles secrete their content. The drop of pH

and enzymatic activity dissolves the bone matrix [68]. While migrating over the bone surface,

they leave a typical picture of resorption lacunae behind. Osteoclasts, which resorb in cortical

bone represent the proximal end of the remodeling unit (or cutting cone)(Fig. 5). The activity

of osteoclasts is followed by the new bone deposition by osteoblasts.

Fig. 4: Typical picture of a remodeling unit with all relevant bone cell species: The resorption of bone by osteoclasts form a cavity, which is filled with a loose connective tissue and capillaries. Osteoblasts deposit osteoid (pale blueish), which thereafter calcifies (dark blueish). Few osteoblast stay in the newly formed matrix and differentiate to osteocytes.

Osteoblasts are single nuclear cells, which differentiate from mesenchymal osteoprogenitor

cells located in the periosteum and the bone marrow. The are smaller than osteoclast with

about 15-20 m in diameter, but are more numerous. They attach to the bone surface and

secrete a layer of proteoglycan and collagen rich matrix (osteoid), which subsequently

calcifies. Osteoblasts form the sides and the tail of the remodeling unit, replacing the resorbed

bone by newly structured lamellar bone. Some osteoblasts remain in the deposited bone and

differenciate to osteocytes. Osteocytes are thin and oval cells, which are encased in their bone

lacuna, but are highly connected to neighboring cells via cell extensions in tiny canaliculi

[33,69]. Their function is to detect environmental changes (mainly mechanical) and damages

and to signal for remodeling [70].

early Osteocyte

Osteoid

Osteoblast

Osteoclast

Capillary

Introduction

10

Some of the osteoblasts, which remained on the bone surface transform to so called lining

cells after the remodeling process is finished. Lining cells have thin fusiform cell body form.

They form a layer of cells, which also interdigitate with the osteocytes. Their function is seen

in the nutrition of the osteocytes [71] and in signal transduction of load changes [72,73].

Fig. 5: Bone remodeling via the cutting cone. The section at level t1 marks the dense unremodeled bone. During remodeling (t2), bone shows a temporary porosity, which comes back to a dense structure after remodeling (t3). Figure adapted from Perren et al., 2002.

The interaction and coordination of cells in the bone remodeling unit (BRU) works through

local biochemical signaling (Fig. 6). Osteoblasts produce several cytokines that regulate the

osteoclast formation and activity. In a BRU for example, the coordination between bone

resorption and formation relies on this local regulation. Osteoblasts produce the macrophage

colony-stimulating factor (m-CSF). M- CSF, which binds to its receptor on preosteoclastic

cells. This is necessary for osteoclast development [74]. But the most relevant regulatory

system for osteoblasts and osteoclasts is RANKL/OPG. Osteoblasts produce the receptor

activator of nuclear kappa B ligand (RANKL) that resides on their outer cell membrane [75].

RANKL can bind to its receptor (RANK) on osteoclasts precursors and stimulate their

differentiation. So, when an osteoclast precursor encounters an osteoblast, the resulting

interaction between RANK and RANKL stimulates the osteoclast precursor cell to mature

into a fully differentiated, bone resorbing cell [75]. Osteoblastic cells also secrete

osteoprotegerin (OPG), the soluble decoy receptor of RANKL, which can bind to RANK and

disable it and therefore inhibiting osteoclastic differentiation and activation by RANKL (Fig.

6). Therefore does the balance between RANKL and OPG regulate the number and activity of

osteoclasts and thus, the rate at which bone is resorbed [76].

Moreover, bone cells are effectors for sex steroids and other hormones. Estrogen, parathyroid

hormone (PTH) or factors such as 1,25 dihydroxy D (vitaminD) can alter the osteoblasts

RANKL/OPG production ratio and therefore modulate the action of bone cells [77]. For

example, hormonal control allows calcium to be released from bone when needed in the

organism by increasing the bone resorption rate [78].

t3 t2 t1

t1 t2 t3

Introduction

11

Fig. 6: Cellular signaling pathway of bone remodeling cells (adapted from Sommerfeldt, 2005)

3.1 Bone healing

Bone tissue is in a constant process of reconstruction aiming to adapt and optimize the tissue

to the current loading situation [79,80]. Any change stimulates the turnover rate regardless

whether its origin is stemming from a different locomotive activity, a fracture or an implant

[81]. Any type of fracture or surgical trauma activates the remodeling pathway.

Before the stage of adapting bone to the modified load distribution the organism has to pass

the process of wound healing [82]. Regardless if the surgeon was drilling, sawing or

performing any other manipulation on bone, the body reacts towards the trauma with the

initiation of a cascade of actions similar for every tissue: local inflammation. Inflammation is

the essential necessary starting point for the healing process if it is not caused through toxic or

infectious agents or already being chronic in nature. It increases vascular supply and recruits

immune competent cells as well as precursor cells for tissue repair. Bone is able to heal and

remodel directly if no gaps have to be bridged and the situation is sufficiently stable [83].

Only in this configuration, bone will be remodeled with almost absence of bone resorption. In

the ideal case the newly built osteons are created in connection to the already existing osteons.

For most of the cases, bone heals indirectly. New bone formation proliferates and woven

bone is deposited to bridge gaps and stabilize the bone structure [84]. Remodeling will take

place thereafter, such that calcified new woven bone is resorbed by osteoclasts and new bone

deposition by osteoblasts. In this process, the unstructured woven callus bone is replaced

through a lower volume of mechanically optimized orientated lamellar bone [79].

However, bone healing is a complex process, which does not succeed in any case. Incomplete

healing may be observed due to many reasons. Systemic changes may be responsible,

such as in metabolically compromised diabetes or elderly patients, where the general catabolic

active osteoclast

passive osteoclast pre-osteoclast progenitor stem cell

osteoblastic stromal cell

RANK

c-fms

RANKL

OPG OPG

PTH Vit D PGE2 IL-1

m-cfs

RANKL +

_

RANK

Introduction

12

tendency can hamper sufficient bone response [85]. In case of mechanical instability bone

resorb and will be replaced by connective tissue.

3.2 Osseointegration

Osseointegration is differently defined depending on the point of view. From the

biomechanical perspective, an implant is osseointegrated if there is no progressive relative

motion between implant and surrounding bone under functional loading. The biological view

understand the term as the close apposition of new and remodeled bone in congruency with

the implant with no connective tissue interposed to achieve a functional connection. A brief

and sound definition was proposed by Brnemark as the a direct structural and functional

connection between ordered, living bone and the surface of a load-carrying implant [86].

The pathway of osseointegration is derived from the mechanisms of bone healing and is

expected to follow the principle of direct bone healing [87]. The implant insertion means a

trauma or local fracture, to which the body reacts with a local inflammation. Blood supply

increases and immune competent cells are recruited as well as precursor cells for tissue repair

[88]. The bone healing and remodeling pathways start thereafter [89].

The main difference to fracture healing is that the implant cannot be remodeled as old or new

bone. First of all, it is a body and surface, which has to be included in the healing processes

and must not be percepted as a foreign body.

The first interaction of the body with the implant surface takes place on the molecular level

immediately after contact with blood or wound fluid during implantation. Various molecules

bond to the surface due to their polarity [90]. The so formed layer of proteins will be

colonized by cells, which migrated to the wound site [91]. These cells are basically immune

competent cells and vascular pericytes accompanied with the ingrowing vessels [66].

Precursor cells will differentiate according to their biochemical surrounding, tissue integrity

and motion pattern [92]. Mesenchymal precursor cells can differentiate to bone cells or

connective tissue cells. For this step it is well documented that osteoblast differentiation and

bone matrix production highly depend on the surface topography of the implant [93-95]. Cells

need to find surface patterns, which allow for attachment and migration. Furthermore the

surface design aims to stimulate osteoblasts to express their typical genetic profile and to

deposit high quality bone matrix with good contact and attachment to the implant surface

[96,97]. The bone matrix finally forms the interface to the implant, which means that the

consecutive function of the implant surface is, indirectly via the induction of bone matrix, the

micro-structural load distribution to the new and remodeled bone. This secondary or

biological stability is essential for permanent implants. In case of loaded implants over-

Introduction

13

loading and mechanical over-stimulation can lead to bone resorbtion and the interface will be

replaced by connective tissue. Implant loosening and non-integration is the consequence

[31,98].

4 Implant materials

The implant material needs to fulfill biomechanical and biological requirements. This

includes corrosion resistance, elasticity, rigidity, and fatigue behavior. The material properties

highly depend on chemical and structural composition, which can also be influenced or

modified by the production and processing. They determine the advantages and disadvantages

of the material, their limits in implant design and their possible side effects in the body.

Classical implants are made of metal materials (Table 1). In orthopedic surgery they are still

predominant. The major groups are steel, cobalt-chrome alloys, titanium and titanium alloys.

In recent years polymers became more and more popular in medical applications. They

provide a wide range of mechanical properties, depending on their molecular chain length and

cross linking [99]. They also offer various processing options to achieve complex device

geometries. The lower e-module is closer to the properties of bone, which is considered to be

beneficial for load transfer and hence osseointegration. On the other side, polymer materials

are weaker (10 time lower tensile strength), less long term stable and less resistant to constant

load (creeping, increasing brittleness) than metals and therefore limited in their application

[100]. From the vast spectrum of technical polymers, only few fulfill the requirement of

biomaterials for medical applications. The chemical synthesis has to reduce the content of

catalysts, solvents, side products and remaining monomers to a absolute minimum to avoid

toxic effects and foreign body reactions [101]. They also need to be biocompatible, as well as

appropriate for sterilization [102]. Polymers show lower bio-functional performances

compared to metals due to their tendency to be covered by a layer of soft tissue preventing

direct bone-implant contact, which is essential for good anchorage. A very important medical

classification of polymers is the one of biological durability. This property distinguishes them

as biodegradable (resorbable) or non-biodegradable (non-resorbable).

Examples for resorbable polymers are the polylactic acid (PLA), polyhydroxybuthyrate

(PHB), polycaprolactone (PCL), polyetherurethane, polyhydroxyvalerate (PHV) and the

polydioxanone (PDS). They are used for suture material [103], tissue engineering scaffolds

[104] and as screws and plates in the small fragment osteosynthesis [105-107]. Due to the

limited mechanical properties, the application is not advisable for load bearing indications

[99]. Although load bearing vertebral interbody cages and plates of resorbable polymers were

developed, they are accompanied with serious concerns about the risk of failure in cases of

Introduction

14

delayed vertebral fusion [108-110]. The advantage of resorbables is the elimination of the

implant material after a defined period [111]. In the field of osteosynthesis, this avoids a

second operation of implant removal for the patients comfort and cost reduction [112,113].

Module

[N/mm2]

Tensile resistance

[N/mm2]

Elongation

[%]

Steel 210000 800 20

Co-Cr 200000 1200 10

Ti 102000 240-680 15-24

Ti6Al4V 113000 900-1100 10

PLLA 3000 - 5000 60 6

Bone 7000 - 40000 90-120 1

Table 1: Mechanical properties of metallic materials in comparison to bone [114-118].

4.1 Steel

The most common implant steel is an iron based alloy with a low carbon content (0.01 - 0.03

% - by weight) and 17- 19 % chrome and 12 14 % nickel fraction (ISO 5832-1:2008) [115].

It also contains parts of molybdenum, vanadium and copper. The alloy additions enhance

hardness and corrosion resistance. The latter ones have a rather adverse biological effect in

implant materials, as they are known to be toxic (Cu, V, Mo) or causing allergic reactions

(Ni) [119]. Stainless steel has very good mechanical properties. It is favored in highly and

dynamically loaded applications as with bone plates and marrow nails. The material is readily

available and is easy to machine, giving it an additional benefit. However, it suffers

considerably from corrosion problems in biologic environment [120,121].

4.2 Titanium

The biological and mechanical requirements of loaded implants led the developers of dental

and orthopedic implants to use titanium and its medical approved alloys. It supplies a high

level of mechanical strength and a good basic biocompatibility for short and long term

applications. Due to their high rigidity and biocompatibility, titanium alloys are used for

nearly all implant systems from joint prosthesis, especially non-cemented systems, over

interbody devices to bone screws and plates [122]. The low e-module of titanium and its

alloys makes it more similar to bone [123].

Pure titanium (Ti-cp) is available in 4 grades, depending on the O, N and Fe content (ISO

5832-2:2000); grade 1 one has the highest Ti content (> 99.5%) [124]. Little increase of the

O and N content increases the tensile and fatigue strength, but decrease the deformability.

Introduction

15

Pure titanium is very ductile, but not strong enough for endoprosthetic use. The good

compatibility, corrosion resistance and noticeable mechanical properties of grade 4 titanium

made it a suitable material for dental implants [125].

To increase the mechanical properties, while maintaining the excellent biological safety of

pure titanium, several titanium alloys were developed. Their content of oxygen and nitrogen

is very similar to pure titanium, but alloys contain aluminum, vanadium (V), ferrite (Fe) or

niobium (Nb). The massive increase of stability goes along with a decreased ductility and

biocompatibility. A very commonly used alloy is the Ti-Al-V alloy, containing 6% in weight

aluminum and 4% in weight vanadium. The addition of vanadium includes a cytotoxicity

agent into the material, which is normally not part of the oxide layer and is not necessarily

released. However, in case of corrosion, where the oxide layer is destroyed, it is a

disadvantage for its compatibility properties [126]. Ti-Al-Nb alloy may provide better

biocompatibility, but these differences to Ti-Al-V are comparably small [119].

4.3 Titanium passivation layer

The basic chemical material properties of the implant core are relevant when corrosion in the

body fluids foster ion release to the implant surface. This process alters the material and can

cause toxic tissue reaction as known from aluminum and vanadium. Most of the metal

surfaces form a passive oxide layer, which chemical composition often differs from the core.

The surface of the implant means the outer nano- to micrometers. This small part of the

implant determines the main interaction with the host [93,127].

Titaniums most striking feature is the ability to form a insoluble oxide layer, which is

responsible for the corrosion resistance being highly bio-inert and biocompatible [125].

Already simply roughened titanium surfaces showed very good formation of direct implant-

bone contact [128]. Moreover, titanium offers various surface modifications to influence bone

response. The currently available surface treatments aim to combine optimal bone response

with high chemical stability by modifications of the composition, oxide layer thickness,

surface topography (roughness and porosity) and hardness [93,127]. Sand blasting, etching,

plasma spraying or anodization and electrochemical anodization are the most common

conventional treatments of titanium devices currently available on the market [129].

A common method, originally developed in the aerospace industry for material refinement, is

the electrochemical oxidation also called anodization. Anodization uses an electrochemical

bath, where the material is exposed to an electrical field in an alkaline or acidic electrolyte

environment. Depending on the electrical current, voltage and time settings, the oxide layer

can reach the micrometer scale [94]. A thicker surface layer is generally considered to release

Introduction

16

less ions from the bulk material [130,131] and provide a better protection from scratches and

wear [132]. Also parameters as the electrochemical bath chemistry (alkaline or acidic

electrolyte and other additives) and temperature will influence the chemical composition

[132,133]. The anodized surface is comparably smooth, but very porous.

Examples on the market are the TiMax Surface (DePuy SARL, LeLocle, Switzerland) and the

Ti-Unite surface (Nobel Biocare, Zurich, Switzerland).

Sandblasting and acid etching is frequently used in combination, leading to a spiky surface

with good micro and macro roughness [134]. The chemical composition of the surface

depends on the material used for blasting and the chemicals for etching. It was one of the first

methods, but is still playing an important role on the market [135].

Examples are the SLA surface (Straumann, Basel, Switzerland) and the SPI surface

(Thommen Medical, Waldenburg, Switzerland).

4.4 Polylactides

The poly-lactic acid (PLA) forms a achromatic transparent polymer. It belongs to the group of

thermoplastic polymers and reaches the glass transition temperature already at 55 C.

Several variations of lactide polymers are available, ranging form pure D-isomer to pure L-

PLA. During synthesis, the monomers D, DL or L are polymerized to long chains (Fig. 7).

The chain length determines the molecular weight (MW), which is often given indirectly by

the intrinsic viscosity (iV). The final MW of an implant can be significantly reduced by the

processing and sterilization impact. Common polymers have a ratio of the L and D,L

monomers of 50/50 or 70/30. Medical grade polylactides need to be intensively purified from

monomers and low molecular weight fraction to avoid tissue reactions known from the early

days of PLA use [136].

Fig. 7: Synthesis of a poly [L-DL] lactide. Racemic monomers and the resulting polymer.

The L-lactides tend to form crystalline structures that are more stable have a higher glass

transition temperature (TG) and degrade slower (Table 2). In contrast, the DL-lactides

(PL[DL]LA) are intrinsically amorphous, softer and are resorbed faster. The racemate 50/50

DL (e.g. Resomer R208, Boehringer, Ingelheim, Germany), which was evaluated in study 2,

is a medical grade polymer of this type. Racemic mixtures of both types combine the higher

Introduction

17

mechanical stability (initially and during degradation) and the low crystalinity the in study 2

of this thesis has a ratio of 70% L to 30% DL polylactide (PL[L-DL]LA, e.g. Resomer

LR708), which is the most frequently used PLA on the medical market.

The final material properties are dependant on the manufacturing methods which can reduce

the molecule chain length, as for instance injection molding and gamma-sterilization.

PL[DL]LA

(R208)

PL[L-DL]LA

(LR708)

PLLA

(L210)

Nature amorphous amorphous 5-30% crystalline

Degr. Time 12 - 18 months 15 - 30 months 2 5 years

Mw (g/mol) 250000* 700000* 130000*

Inherent Viscosity (dL/g) 1.5-2.0 5.5-6.5 3.3-4.3

TG (C) 52 54 56

TM (C) amorphous amorphous 179

tensile strength 54-65 55-72 72

E module 2500 2500 5000

Breaking elongation (%) 1.3- 6 2.2-6

Table 2: Chemical and mechanical properties of degradable polymer materials. * MW is depending on the processing of the material. [102,106,137,138]

The primary degradation mechanism of PLAs is the chemical process of hydrolysis [139]. It

is a so called bulk degradation, where the shape of the material is preserved while the material

looses molecular weight. It requires water and the degradation rate depends on the

temperature and pH [140]. The hydrolytic degradation starts after water uptake of the

polymer, which can be inhomogeneous depending on the polymer structure in the implant and

which is faster for machined material compared to injection molded parts [141]. The end

product is the lactate monomer, which diffuses away and is regularly metabolized to carbon

dioxide and water [142]. In vivo degradation can be also accelerated by enzymes [143],

cellular activity [144] and cell-induced local pH drop due to tissue damage [145] or

inflammation [146]. Only if the material is highly fragmented and advanced degraded, the

host tissue can influence the degradation time by uptake of particles and oligomers by

macrophages [147].

Introduction

18

5 Methods of osseointegration analysis

The implant osseointegration is the product of the structural adoption of the host bone and its

bond to the implant. The interface properties are determined by the specific reaction of the

organism and the result of cellular and metabolic activity. In this sense, the anchorage is the

final outcome and the bone material and structure are responsible for these.

5.1 Biomechanical evaluation

The mechanical outcome of osseointegration is the most relevant in clinical implant

applications. Therefore, the measurement of osseointegration is more often understood as a

biomechanical parameter. The implant osseointegration and the resulting bone structure were

optimized by remodeling to the present load profile at the implantation site induced by the

body weight and muscle action. The initial difficulty is to know these forces to create a

relevant test load setup. These forces cannot be measured by conventional test machines intra

vitally, but can only be approximately simulated by finite element analysis. The consequence

is often a simplification of test loading conditions by reducing the setup to one directional

force application like torque, pulling, pushing or bending (Fig. 8) with a static loading

protocol.

The methods for dynamic loading and multi-axial testing are already available and frequently

used for spinal, knee and hip devices [148,149], but rarely for dental implants [150]. Test

protocols of dynamic loading are even less standardized than for static loading which makes it

hard to compare results of different research groups [151].

Measured release forces are highly depending on the choice of the test method and the precise

application of test force, e.g. the alignment of the implant and test machine axis. Therefore,

test method and the resulting strength of implant anchorage have to be critically discussed if

appropriate for the goal of a study and the targeted clinical application [152]. A good example

is push out testing of screw type implants. If the implant is not completely cylindrical, the

interlocking effect of the threads will largely contribute to the anchoring effect. The threads

would perform load transduction deeper into the bone structure and failure would likely occur

in the trabecular structure. This test setup would not be appropriate to test the interface

strength to a modified surface.

Introduction

19

Fig. 8: Basic test setup for biomechanical testing. The force is applied via a clamping or stamping tool to the implant in the bone block (yellow), which is fixed in a holder attached to the force measuring unit. Rotational (a), pull out (b), push out (c) and bending (d) are the most common test approaches. The sample, containing an implant is attached to a holder, which is connected to the load cell.

Creating a rotational failure mode is one of the most common approaches of biomechanical

testing for cylinders and screw type implants. Torque force is applied to the implant head with

a defined speed or force until the bond to the bone breaks. The macroscopical interlocking

effect and the entire integration in the host bone structure on a millimeter scale is less

represented in the test method when the failure occurs in the interface zone. A common

obstacle is the attachment of the implant to the test machine. If the implant head was not

designed for this purpose, as found in nearly all implants for clinical use, a tight connection

for load transduction is very difficult to realized without any slipping and shear forces.

Pushing or pulling an implant out of the bone bed axially is also a very well established

method and is suited for non-rotational and screw type implants. Pull out testing requires also

an implant head suitable to be properly attached to the test machine and to provide sufficient

holding power, whereas the application of force is much more simple for push out testing

since no clamping tool is required. On the other hand does push out testing require more

sample preparation to open up the distal end of the implant (Fig. 8). Preparation of the sample

without affecting the bone-implant interface is hard to ensure and this is often a source of

bias. The dimension of the bone tissue surrounding the implant and the type of sample

fixation is more important to be standardized than in the torque testing method.

All static test methods have in common to record the resulting momentum with a force

measurement unit (load cell) for calculation of standard parameters like the maximum release

force, stiffness and shear force.

Another attempt of mechanical testing is the non-destructive testing of the implant bone

anchoring. Techniques, which employ resonance analysis of the bone-implant interface, need

direct access to the implant head. The resonance frequency analysis (RFA)[153] and the

Periotest (PT)[154] could not create sufficient confidence yet to predict mechanical stability

Introduction

20

but they can indicate changes by repeated measurements [155] and allow samples to be

evaluated with other methods, like histology.

The simulation of the implant stability in the bone is the topic of finite element modeling,

based on the structural information of micro CT scans [156]. Complex models were created to

predict the mechanical situation extremely detailed [157]. The experience points out the

limitation that the results are just as good as the model (the underlying algorithm) and that the

higher the complexity the higher the risk not to recognize or find errors.

5.1.1 Methods used in the studies

The test method of the initial study was using a well established procedure in the Robert

Mathys Foundation Laboratory (Bettlach, Switzerland), conducted with Gianni Bigolin and

supervised by Dr. Lukas Eschbach. The sample was fixed in the holder with a series of

screws, while attached to the torque machine with a stiff tool. The stiff tool was replaced by a

flexible tool with cardan joint. Lateral forces during sample fixation could not be excluded

due to the tolerance of the implant-tool connection. The cardan joint probably reduced further

bias. The implants in study 1 provided a flat implant head with a inner hexagonal coupling. In

some cases, the implant head was already overgrown by calcified bone tissue and needed a

careful removal of the overlying tissue to enable coupling to the test machine.

The mentioned experimental uncertainties could be excluded in study 2 by the method

established in the MEM-center (Maurice Mller Institute, Berne). These measurements were

conducted with Philippe Gdet and supervised by PD Stephen Ferguson. This method

demanded no change of torque tools. The implant head was clamped to the test machine and

fixed exactly aligned to the torque axis by an on-site embedding with a quickly solidifying

liquid (Woods metal)[158]. The implant head design in study 2 was improved for mechanical

testing. It was shaped to be clamped from outside without any tolerance and a healing cap

prevented bone tissue overgrowth.

The measurements of both methods were successful in all cases, but with a higher degree of

reliability in study 2.

5.2 Morphological evaluation

Analysis of bone morphology, the so called histomorphometry, aims to describe and measure

the structures of bone, which are responsible for implant anchoring stability. Several

parameters of the bone architecture are defined for 2- dimensional evaluation (classical

histology) and 3-dimensional evaluations (confocal microscopy or CT scan) [159].

Quantification of bone structures, as done by the bone-implant-contact (BIC) and the bone

Introduction

21

area measurement, enabled to assess the bone response and to compare values of different

implant types and between time points after implantation.

Morphological evaluations are also performed to understand the biologic response of bone

tissue on a cellular level, e.g. by counting osteoblasts or osteoclasts adjacent to the implant.

The cellular evaluations can only be done by the histological method. The strength of the

histochemically stained section is also to allow clear discrimination between old and new

bone and to make microstructures visible.

The limitation of histology to a two-dimensional analysis, which is not necessarily

representing the complete situation, made the three-dimensional micro computer tomography

( CT) very attractive for morphological analysis [160]. The high resolution of modern

tomographs made it possible to visualize trabecular architecture in detail, which was not the

case with clinical CT [161,162]. For more resolution, but very limited sample dimensions, the

synchrotron beam computer tomography became available for regular projects (e.g. the Swiss

light Source at the Paul Scherrer Institute, Villigen, Switzerland)[163,164]. The imaging of

the CT scan bases on the x-ray transmissibility of the bone. This is determined by the degree

of mineralization, which is slightly different in old an new bone. But the contrast is very poor

for reliable detection. The major limitation of CT scans of samples on a centimeter scale is

the resolution to detect the bone-implant interface. In combination with x-ray scattering and

adsorption of metal or ceramic implant materials, it is almost impossible to acquire reliable

images from the implant surface. This a very crucial limitation, because it makes a difference

of anchoring strength, whether the implant is just surrounded by bone, or bone is bond to it.

AREA & VOLUME

The bone area, or volume (bone mass) measurement aims to quantify changes in the bone bed

after implant insertion by calculating the percentage of bone in a defined area or volume.

Comparison of bone mass is feasible between different observation periods, or between

different locations within an implant site of a single time point. For these evaluations, cortical

and cancellous bone needs to be evaluated separately due to the opposing effect on initial

bone mass changes of their healing pathways due to the following reasons:

a) Cortical bone mass initially decreases as a normal step of the remodeling process. This

temporary phenomena must not be interpreted as pathological bone resorption. In the later

phases, cortical bone is remodeled to the original bone density, but possibly resulting in

thicker or thinner layer of cortical bone layer.

b) In contrast, cancellous bone mass is increasing from the very beginning. New bone

formation is build up on top of the existing trabeculae displacing the bone marrow. The

volume of new bone decreases in the late phase of remodeling when the high volume of

Introduction

22

unstructured bone is replaced by lower volume of structured bone.

If performed carefully and interpreted according to the physiological changes with time, bone

mass measurement is a powerful tool to follow the integration process.

BONE IMPLANT CONTACT

Bone implant contact (BIC) measurement aims to quantify the product of implant surface

interaction with the host tissue in the few micrometer thick interface zone. It is generally seen

as a measure for the cellular response toward the implant surface, in means of bone matrix

deposition. The contact is calculated as the ratio of implant surface in direct contact with bone

to the complete implant surface. The assumption behind is that after implant insertion,

osteoblasts or precursor cells migrate towards the implant surface, interact and respond with

differentiation and bone deposition -depending on the surface quality. In the late phase of

bone healing, a 100% ratio is seen as the gold standard. In earlier remodeling phases, bone

can be removed from the implant surface, without being a negative influence by the surface

itself. Care has to be taken in the very early healing phase, especially in cortical bone, when

the old bone has not been remodeled yet. Implant in contact with old bone is a given

precondition by the implantation technique (form fit or press fit) and the present bone

structure. Bone-implant contact is therefore not a sign for good implant performance.

Only few authors presented approaches to normalize the BIC data to an average initial bone

contact [165].

BONE STRUCTURE AND QUALITY

Bone morphology can be judged on a microscopical scale. The basic bone qualities of cortical

and cancellous bone can be distinguished. New bone formation can be primary woven bone,

or already be remodeled to a lamellar structure. The new bone matrix can be non-calcified

(osteoid) or calcified. Depending on the age and species, the lamellar structure forms

plexiform or osteonal bone [166,167].

The trabecular architecture of cancellous bone forms an optimal structure for load

transduction. Depending on the situation there are differences in the relative bone volume,

trabecular thickness, spacing and number. Two different formations are commonly

characterized as rod or plate like structures.

CELLULAR COMPOSITION

The assessment of the cellular response is a fundamental tool to interpret the morphological

results. The assessment of bone cells includes the number of osteoclasts and osteoblast per

bone surface or bone area/volume as the cellular source for morphological changes.

If components of an implant system are unknown, the presence and distribution of immune

Introduction

23

competent cell populations (macrophages, lymphocytes, plasma cells, multinuclear foreign

body cells) enable to judge a possible inflammation on the microscopical level. This is the

main step to be done in vivo to assess the biocompatibility. These assessments were

standardized as requirements for biocompatibility testing by regulatory authorities for medical

devices (ISO 10993-6) [168]. Measurements were performed on the base of high resolution

digital images in a reproducible fashion. Only in few cases, clarification was necessary with

microscopical review of the original slide.

5.2.1 Methods used in the studies

In study 1, the state of the art method of bone-implant contact measurement (BIC) was

employed [169]. The measurement was implemented in a totally manual (non-automatic)

style for each available thread and measured distances were marked and saved as separate

pictures. In the literature, many authors selected the best areas (consecutive filled threads)

[170-172] or used semiautomatic processing [173-175] which incorporates systematic bias.

The method used in study 1 provided a maximum accuracy and reliability.

The implant geometry in study 2 was not predefined as the thread geometry of the test

implants in study 1. The extremely randomly curved surface of the polymer part was not

suitable to be processed with a reasonable effort by the method of study 1. As an appropriate

solution, the implant surface was divided into sectors which were small enough to be

measured semi-quantitatively with repeatable accuracy.

The non-rotational implant geometry of the hybrid implants in study 2 raised the question if

and how the penetration of the polymer part into the bone cavities induced changes in the

bone structure. To asses these changes, histomorphometrical evaluations were extended by

bone area measurements. Bone area measurement is commonly performed in most of the

histological studies. Unfortunately, many authors showed little attention to the application of

this method and missed to include important factors into the evaluation design, like local

changes within the bone bed and the choice of measurement areas [176].

To avoid systematical bias of this kind, the changes in bone bed were measured in an implant-

adjacent sector, which was compared to a control sector in the periphery. Sectors within this

area were iteratively defined after thorough microscopical review of all samples and the

control measurement area was defined as large as possible. Precondition of the comparative

area measurement is an accurate operation planning, sample taking and sample processing,

which provides and preserves sufficient area of non-affected tissue around the implant to act

as internal reference. To avoid any false-detection of non-bone structures, the first step of

image preparation was to manually label the complete bone structure of each section and store

Introduction

24

it in a separate image layer. This step enables to double check the detection threshold anytime

thereafter.

6 Animal models for osseointegration

Animal models for in-vivo applications and validations of implants are used for clarification

of complex interactions between implant and host bone. They present the opportunity to

simulate conditions as close to the human application as possible in the last step of implant

development [177,178]. Only advanced prototypes or finalized designs for preclinical testing

(for regulative marked clearance) reach this stage. The following chapter shall outline the

decision making process for the use of an appropriate animal model.

6.1 Ethical considerations

Prior to animal experiments, new materials have to demonstrate biological safety by chemical

analysis, cell culture (in-vitro), cadaver testing (ex-vivo) or computer models (finite element

load distribution). Those methods have made great advances with the deeper understanding of

biochemistry, cell biology and the possibilities of computer simulations [179,180] and thus,

numerous animal experiments became avoidable in recent years. The current standard of in

vitro bone testing is a regular cell culture in Petri dishes with addition of vitamin D3 on the

culture medium [181]. Osteoblast culture in the micro mass fashion [182] fosters the cellular

interaction and differentiation compared to the monolayer culture, where cells are seeded as

singled cells and interact far later after proliferation. Micromass culture is achieved by

seeding in droplets of 30 l, which let the cells attach in a dense formation from the very

beginning. If this technology is used, it is possible to achieve better differentiation of

osteoblasts on test surfaces and to study their expression pattern towards different materials

[183,184]. Limitations of in vitro testing are given by the very special conditions bone tissue

requires for growing and remodeling. The physiological environment is a three-dimensional

tissue, which offers specific nutrition, oxygen supply, cell signaling, surfaces and loading

conditions. Until now, it was not possible to mimic them satisfactorily in vitro. Cell cultures

cannot even simulate the organ itself, not to mention the interaction with the complete

organism. If the question of an experiment is, for example, the response of the immune

system, the elimination pathway of debris or the adaptation of bone to physiological loading,

animal experiments are still essential.

The Swiss animal protection law follows the international guidelines [185] and defines clear

guidelines and preconditions for animal experiments and the conducting facilities. A detailed

classification system (document 1.04 and 1.05) helps to judge the degree of severity. The

Introduction

25

most important part of the ethical considerations is the balance of interest between the benefit

for the human patient and the suffering of animals [186]. Generally, the number of animals

used in an experiment is limited to the statistical minimum.

6.2 Experimental animals

In-vivo assessment of bone repair is performed on various species. One can distinguish

between large and small animals, domestic and tamed, higher mammals and lower mammals

(Table 3,4). Small animals, like mice, rats and rabbits, offer repeatable conditions due to

known genetic background and standardized husbandry. Their higher metabolism allows

shorter observation periods for the same healing progress than in larger animals and in

human. This makes mice, rats and rabbits ideal for basic scientific evaluations of bone

metabolism and cellular interactions [187]. These animals are also frequently used for

screening tests for new materials and implant surface modifications [188]. Their considerably

small dimensions limit the possibility to test implants in their final design and indication for

human clinical use [189]. Furthermore, their rather simple bone macrostructure and never

closing growth plates of long bones makes implant testing results difficult to transfer to the

human species [190].

Species Dimension,

femur length

Metabolic rate

relative to human

Bone

structure Husbandry Handling

Mouse 2 cm very high simple simple easy

Rat 3 cm very high simple simple easy

Rabbit 8 cm very high simple simple easy*

Dog 12- 20 cm high complex demanding demanding+

Pig 15 - 25 cm comparable complex demanding demanding+

Sheep 23 cm comparable complex simple demanding

Primates 10 30 cm comparable complex very demanding

very demanding+

* hind leg fracture can occur when pedaling uncontrolled during manipulations + potentially aggressive behavior

Table 3: General properties of animals used for bone healing experiments

The criterion of appropriate size is given for most of the larger animals, like dogs, sheep, goat

and pigs. All of them provide sufficient dimensions to evaluate one or more implants per

animal. The choice of a large animal model also includes matching as many parameters of the

target species and choosing the most relevant of the specific bone properties.

The biochemical bone composition is relatively similar in all species [191], but the

macrostructure (bone architecture) varies on a larger scale [33]. Differences in the

Introduction

26

microstructure were found, such as the complexity of the canal system in cortical bone, the

fiber orientation and the presence of cement lines in-between osteons. These structural

features result in interspecies differences of mechanical properties [192]. As mentioned for

the small animals, also large animals differ in their metabolism, which influences the bone

healing rate. The overall reaction pathways of the organism result in different remodeling

pathways. One of the most important differences is found between young and adolescent

animals, respectively skeletally mature and immature animals [190]. The healing capacity and

the tolerance to destructive influences is much higher in young animals. Implant

osseointegration success cannot be claimed for all patient groups if tested in immature bone.

Canine Sheep/ Pig Rabbit

Macrostructure ++ +++ ++ +

Microstructure ++ + ++ +

Bone Composition +++ ++ +++ ++

Bone Remodeling ++ ++ +++ +

+ least similar, ++ moderately similar, +++ most similar.

Table 4: Summary of key attributes in terms of similarity between animal and human bone. Reproduced from Pearce, et al 2007.

Besides the biological differences, the tolerance of the experimental animals towards

experimental treatment must be considered. Since handling and manipulations are a daily

task, (especially perioperatively for medication and surveillance) it is of great importance to

be performed with a minimum of stress. Else, the results of the experiment can be severely

biased. Also the animals typical social behavior can be a limitation and requires thorough

knowledge. It is not clear for everybody that a ruminant might be unsuitable to test dental

implants for human use, since apart from the structure of teeth chewing frequency and

dynamics are completely different. There is no physician- patient interaction, which might

support the healing process or prevents from excessive loading, as reported in a pig study of

dental implants [193].

For example, prey animals like sheep, tend to run scared, move jerkily and hurt themselves

during the attempt to fly. For this reason, especially gregarious animals should be held in

small groups and stables during the initial healing phase to avoid spontaneous fractures in the

operated limb. On the other side, predators (like dogs) can be seriously aggressive and able to

bite off wound dressing or even implant devices. The bottom line is also the experience of a

research facility with animal care, handling and data interpretation, which makes results

Introduction

27

reliable. The choice of the best animal model is difficult, due to various advantages and

disadvantages of each species [191].

6.2.1 Animals and bone models used in the studies

The experimental animal used in the study was the Swiss alpine sheep. This breed is rather

large with comparably strong bones originally bred for meat producing. This ensured

appropriate dimensions and loading of the bones for implants intended for human use. The

sheep itself is widely accepted as a model for bone healing due to comparable bone structure

and biology [194-196]. The availability of sheep must not be neglected as a reason of choice,

since very few research facilities can afford own breeding of large animals. The animals in the

studies were bought from a farmer, where they were in contact with humans and therefore

comparably docile and tolerated handling very well.

The design of study 1 was based on a widely used sheep tibia bone model for bone healing in

cancellous and cortical bone [56,197,198]. The sheep tibia shaft (diaphysis) offers good

conditions to evaluate implants in cortical bone quality. The access to the medial tibial shaft is

quick and easy without large soft tissue preparation, since the bone lies directly underneath

the skin. This means that there are also no anatomical structures, which are at risk to be

impaired or interfere with the implant heads during the healing period. The dimension (length

of ~25 cm and ~20 mm diameter) are sufficient to place implants in original size. The cortical

thickness is almost constant over a length of 10 cm to assure similar conditions for all of the

bicortically-seated implants. The diameter of the tibia decreases in the insertion area

proximally to distally from 22 to 18 mm. Therefore, three different lengths of implants were

manufactured to have the distal implant end in flush with the periosteum. The cavities were

secondly oriented in a zig-zag pattern to reduce the risk of fracture. Overall, the diaphysis part

of the tibia could be confirmed as an appropriate model for cortical bone with a 100% success

rate and no irregular healing behavior.

The proximal tibia and distal femur were used for evaluations in cancellous bone quality.

The access to the bones is a routine procedure and requires minor preparation of muscle,

tendon and ligaments. The joint capsule may be at risk, especially in the distal femur. The

insertion in direct proximity to these motile structures can injure them and also cause tissue

irritations. Hampered bone healing and false results can be the consequence. Singular cases of

crestal bone resorption might have been attributed to this origin. The distal end of the femur

(complete condyle, ~4 cm) offers a large volume of relatively dense and homogenous

cancellous bone structure, which is comparable to the small epiphyseal part of the proximal

tibia (~1 cm). The metaphyseal cancellous bone of the tibia is very narrow (0.5cm) and shows

Introduction

28

less density. The epiphyseal line is a seam of low porosity and thus more holding power,

which needs to be considered. The small area of accessible cancellous bone limited the

number of implant sites to 60% of the tibial cortex. A risk of fracture is not given in

cancellous bone, but precision of implant placement is necessary due to the changes in bone

density. In the course of the first study, even accurate operation planning could not exclude a

rate of 25% misplaced implants, especially in the tibial cancellous locations.

The lack of homogenous cancellous bone locations demanded for a solution. Additional

demands had to be met, like the sufficient dimension, easy surgical access, low impairment of

local structures and a low impact on the experimental animal. Most of the cancellous

locations are adjacent to joints and provide less space than the epiphyseal parts of the stifle

joint or are not accessible like the hip joint. The vertebral body, with pronounced cancellous

bone content, is also difficult to access and is surrounded by important nerves and large

vessels, which is both risky and time consuming for preparation. The sternal bones, which

might fulfill the above mentioned criteria, are not suitable due to the high pressure during

resting of all large animals.

A solution was finally found in the iliac bone of the pelvic girdle (see chapter 7). The cranial

part consists of abundant cancellous bone and offers space for several implants. The covering

gluteus muscle needs to be prepared and retained to access the iliac bone, but the resulting

access distance from skin to bone still allowed the use of standard surgical instruments. The

caudal part of the iliac column, towards the acetabulum becomes more and more cortical and

also brittle. The cortical zone was excluded for implant insertion after spontaneous fractures.

Also the locations of the iliac crest turned out to have high individual differences regarding

the epiphyseal closure and therefore cannot be recommended for regular bone healing

experiments. After appropriate modifications, the sheep pelvis model was shown to be a valid

model for bone healing and implant osseointegration.

The value of improved animal models for preclinical studies should not be underestimated,

despite the limitations of interspecies differences. In case of implants, which represent solely

a new combination of known materials and techniques, the presence of explicit positive

results can justify to replace large human clinical studies. The reduction of experimental effort

is not only a benefit for the number of potentially involved experimental animals or probands,

but also for the cost, which increases the final prices of the market ready medical device. This

aspect will be of higher relevance in the near future, because of the higher regulatory demands

for new medical devices.

Rationale of the research projects

29

II Rationale of the research projects

The challenge of orthopedic and dental implant development is to match material and design

to the bone tissue properties for achievement of optimal osseointegration. The main difficulty

is that bone is not a static biomaterial, but a living tissue that adapts and reacts to the situation

of an inserted implant. Implant design, insertion concept and the material properties must be

seen in the context of bone healing mechanism (see chapter 3). The intended period of

implant use determines to which extend the bone material and its osseointegration properties

need to be regarded (see chapter 1). Implants for transient and permanent use both have to be

stable in the bone bed, but transient implant stability is no longer needed when bone healing

completed. For most of these implants it is reasonable to be removed not to hamper the healed

bone function and not to risk long term complications. The requirement of osseointegration

implies a conflict of goals between anchoring and removal strength. Implant anchoring

stability at the time of removal would be ideally limited to the strength of the implant

material. An exceeding removal strength would cause implant breakage during the removal

procedure. Removability would not be given when removal means a chiseling out procedure

[4,5]. In this sense can the good osseointegration of standard titanium surfaces become an

obstacle for the transient implants. In contrast do stainless steel implants provide better

removability, but include the risk of insufficient osseointegration.

Fig. 9: Schematic drawing of osseointegration strength of implant materials used for transient and permanent implants. Aim of the research was to optimize the osseointegration strength for the intended use in study 1 and 2.

Another issue of titanium surfaces is that they are damageable. This can either be scratches

during insertion or by tribocorrosion during their functional period [199]. The self-passivation

of a damaged titanium surface will limit the release of ions, but will not reach the properties

of the remaining optimized surface with negative effects on the overall implant performance.

Therefore, resistance to scratches and resistance to tribocorrosion is still yet another issue to

be solved. This was the starting point of study 1 to evaluate surfaces of increased hardness

Anc

horin

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tren

gth

Materials and Implant use

Rationale of the research projects

30

and low release force.

Similar for transient and permanent implant use is that the anchoring effect can only be

maintained with a constant stability within the bone bed during bone healing.

The first critical phase after insertion is the primary stability. Initial stability of the implant

needs to be maintained, where the implant is subjected to loading directly after implantation

until the formation of new bone and bone-implant contact stabilizes the situation [200]. Else,

can lack of stability bear the risk of a cascade of processes, which lead to loss of instant

stability and subsequent loss of implant function. One reason are wear particles which are

produced by the relative motion at the articulation face between implant and bone [201]. Wear

particles support the inflammatory tissue response, manly driven by macrophages [202].

The other factor for loosening is exceeding micromotion, which leads to bone resorption and

a fibrous tissue interface [31,203]. Various combinations of bone bed quality, loading strength

and bone healing capacit