fabrication of 3d printed titanium scaffolds and ... - UPCommons

MASTER´S THESIS

Master in Interdisciplinary and Innovative Engineering

FABRICATION OF 3D PRINTED TITANIUM SCAFFOLDS AND

CHARACTERIZATION OF PHYSICOCHEMICAL AND

MECHANICAL PROPERTIES

Report and Annexes

Author: Patricia Veloz Torres

Directors: Elisa Rupérez – Daniel Rodríguez

Call: September 2021

Fabrication of 3D printed Titanium scaffolds and characterization of physicochemical and mechanical properties

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ABSTRACT

In this work, different geometries of 3D printed porous titanium structures were developed and

tested in order to determine if they are suitable for bone implants in humans.

Commercial pure Titanium (99.5%) and an hydrogel were mixed in order to obtain the ink to

print through 3D printing technique direct ink writing. After the printing process, samples were

subjected to thermal treatments in order to remove the hydrogel and obtain pieces with a

mechanical behavior suitable for use on human bone.

Physicochemical and mechanical characterization was performed to the samples to characterize

them.

SEM images were taken in order to observe the general microstructure of the samples,

micropores and macropores. Micro-CT technique was used to obtain porosity and density.

Mercury Intrusion Porosimetry was performed and along with Micro-CT results obtained,

allowing us to characterized micropores. Surface topographic characteristics (roughness) were

studied through Chromatic Confocal Microscopy equipment.

Young modulus was obtained using compressive strength results in order to compare them with

human bone. Vickers hardness results allowed us to determine hardness of the samples printed,

that is, the resistance of a material to being penetrated.

Results obtained on assays performed, demonstrated that all samples have physicochemical and

mechanical properties suitable for fluid transport and cellular proliferation if used as bone

scaffolds.


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RESUMEN

En este trabajo, se desarrollaron y ensayaron diferentes geometrías de estructuras porosas de

titanio impresas en 3D con el fin de determinar si son adecuadas para implantes óseos en

humanos.

Se utilizó una mezcla de titanio comercialmente puro (99,5%) y un hidrogel con el fin de obtener

la tinta para imprimir mediante la técnica de impresión 3D Direct Ink Writing. Después del

proceso de impresión, las muestras fueron sometidas a tratamientos térmicos con el fin de

eliminar el hidrogel y obtener piezas con un comportamiento mecánico adecuado para su uso

en hueso humano.

A las muestras se les realizaron una serie de ensayos con el fin de obtener sus características

fisicoquímica y mecánica.

Se tomaron imágenes SEM con el fin de observar la microestructura general de las muestras,

microporos y macroporos. Se utilizó la técnica de Micro-CT para obtener la porosidad y densidad

de las estructuras impresas. Se realizó el ensayo de Mercury Intrusion Porosimetry y junto con

los resultados de Micro-CT obtenidos, nos permitieron caracterizar los microporos. Las

características topográficas de la superficie (rugosidad) fueron estudiadas mediante un equipo

de Microscopía Cromática Confocal.

El módulo de Young se obtuvo utilizando los resultados de la resistencia a la compresión para

compararlos con el del hueso humano. Los resultados de dureza Vickers nos permitieron

determinar la dureza de las muestras impresas, es decir, la resistencia de un material a ser

penetrado.

Los resultados obtenidos en los distintos ensayos realizados, demostraron que todas las

muestran tienen propiedades fisicoquímicas y mecánicas apropiadas para permitir el transporte

de fluidos en ellas y la proliferación celular, si son usadas como material para implante en hueso

humano.


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RESUM

En aquest treball, es van desenvolupar i van assajar diferents geometries d'estructures poroses

de titani impreses en 3D per tal de determinar si són adients per a implants ossis en humans.

Es va utilitzar una mescla de titani comercialment pur (99,5%) i un hidrogel per tal d'obtenir una

tinta per imprimir mitjançant la tècnica d'impressió 3D Direct Ink Writing. Després del procés

d'impressió, les mostres van ser sotmeses a tractaments tèrmics per tal d'eliminar el hidrogel i

sinteritzar, obtenint peces amb un comportament mecànic adequat per al seu ús en teixit dur.

A les mostres se'ls van realitzar una sèrie d'assajos per tal d'obtenir les seves característiques

fisicoquímiques i mecàniques.

Es van prendre imatges SEM per tal d'observar la microestructura general de les mostres,

microporositat i macroporositat. Es va utilitzar la tècnica de Micro-CT per a mesurar la porositat

i la densitat. Es van realitzar assaigs de Mercury Intrusion Porosimetry i juntament amb els

resultats de Micro-CT obtinguts, van permetre caracteritzar els microporus de les mostres. Les

característiques topogràfiques de la superfície (rugositat) es van mesurar mitjançant un equip

de Microscopia Cromàtica Confocal.

El mòdul de Young es va obtenir utilitzant assaigs de resistència a la compressió per comparar-

los amb el del os humà. Els resultats de duresa Vickers van permetre determinar la duresa de

les mostres impreses, és a dir, la resistència d'un material a ser deformat plàsticament.

Els resultats obtinguts en els diferents assajos realitzats, van demostrar que totes les mostren

tenen propietats fisicoquímiques i mecàniques apropiades per permetre el transport de fluids

en elles i la proliferació cel·lular, si són usades com a material per a implant en os humà.


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Contents

1. OBJECTIVES .................................................................................................................................. 7

1.1. Project Scope ....................................................................................................................... 7

1.2. Previous Projects ................................................................................................................. 7

2. INTRODUCTION ............................................................................................................................ 8

2.1. Titanium ............................................................................................................................... 8

2.2. 3D Printing ......................................................................................................................... 17

2.2.1.Binder Jetting ....................................................................................................... 20

2.2.2.Direct Energy Deposition ..................................................................................... 20

2.2.3.Materials jetting ................................................................................................... 20

2.2.4.Materials Extrusion .............................................................................................. 21

2.2.5.Powder Bed Fusion (PBF) ..................................................................................... 22

2.2.6.Sheet Lamination ................................................................................................. 22

2.2.7.Vat Photopolymerization ..................................................................................... 22

3. MATERIALS AND METHODS ....................................................................................................... 24

3.1. Ink fabrication .................................................................................................................... 24

3.2. Printing parameters ........................................................................................................... 25

3.3. Binder elimination and sintering process ......................................................................... 27

3.4. Physicochemical Characterization .................................................................................... 29

3.4.1.SEM ....................................................................................................................... 29

3.4.2.Micro-CT ............................................................................................................... 29

3.4.3.Mercury Intrusion Porosimetry (MIP) ................................................................. 30

3.4.4.Roughness ............................................................................................................ 31

3.5. Mechanical Characterization ............................................................................................. 33

3.5.1.Compressive Strength .......................................................................................... 33

3.5.2.Vickers Hardness .................................................................................................. 34

4. RESULTS ...................................................................................................................................... 35

4.1. Characterization of structures........................................................................................... 35

4.2. Physicochemical Characterization .................................................................................... 38

4.2.1.Dimensional changes ........................................................................................... 38

4.2.2.SEM ....................................................................................................................... 39

4.2.3.Porosity ................................................................................................................. 42

4.2.4.Roughness ............................................................................................................ 43

4.3. Mechanical Characterization ............................................................................................. 44


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4.3.1.Young Modulus .................................................................................................... 44

4.3.2.Vickers Hardness (HV) .......................................................................................... 46

5. DISCUSSION............................................................................................................................ 47

6. CONCLUSIONS ........................................................................................................................ 48

7. BIBLIOGRAPHY ....................................................................................................................... 49

8. ANNEXES ................................................................................................................................. 54


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1. OBJECTIVES

1.1. Project Scope

The objective of this project is to obtain samples of 3D printed titanium ink, with two patterns

and two different densities in order to analyze them and compare their mechanical properties

with previous projects with similar structures.

Other objectives are:

- Define the parameter range for feasible printing structures

- Compare the printed structures in function of their design.

- Characterize their physical-chemical properties.

1.2. Previous Projects

Elia Vidal (Vidal E, 2020) performed titanium ink 3D printed samples and analyzed their physical-

chemical properties and cell response.


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2. INTRODUCTION

2.1. Titanium

In order to cover the high demand of materials for prostheses by an aging population with an

increasing average weight, biomaterials is a field that is constantly expanding. Biomaterials are

artificial or natural materials that are used to restore or replace a biological structure's form and

function after it has lost or failed in order to improve the quality and duration of human life.

(Saenz & Fuentes, 2013).

The highest number of implants is for spinal, hip and knee replacements. It is estimated that by

the end of 2030, only in the United States of America, the number of total hip replacements will

rise by 174% (572,000 procedures) and total knee arthroplasties are projected to grow by 673%

from the present rate (3.48 million procedures) (Kurtz, Ong, Jau, Mowart, & Halpern, 2007).

Figure nº1: Schematic diagram of artificial hip joint (left) and knee implant (Saenz & Fuentes,

2013)

This is because degenerative illnesses such as osteoarthritis (inflammation of the bone joints),

osteoporosis (weakening of the bones), and trauma cause pain and loss of function in human

joints. Due to excessive stress or the absence of a normal biological self-healing mechanism,

degenerative disorders cause the mechanical characteristics of the bone to deteriorate. (Saenz

& Fuentes, 2013)

Artificial biomaterials are the answers to these challenges, and surgical implantation of

appropriate artificial biomaterials helps restore the function of these human structures which

are otherwise functionally compromised. (Saenz & Fuentes, 2013).

Titanium (Ti) and its alloys, which were originally developed for aeronautics applications, have

gained significant attention in the biomedical sector because of their outstanding

characteristics, which include a moderate elastic modulus of approximately 110 GPa, a good


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corrosion resistance and a low density (approx. 4700 kg m−3). (Navarro, Michiardi, Castaño, &

Planell, 2008).

The investigation of dental and surgical applications of Ti alloys began when Branemark (Breine,

et al., 1964)discovered what he named the osseointegration phenomena for Ti implants.

Titanium and its alloys are able to form a strong attachment with bone. This characteristic

improves the long-term performance of implanted devices by lowering the risks of loosening

and failure.

To establish effective metallic/bone bonding, mechanical and biological specific requirements

must be fulfilled. (Ottria, et al., 2018)

The biocompatibility, mechanical properties, and surface properties of the material, as well as

its chemical properties and failure properties, must all be considered in the design of these

implants so that the implant closely mimics the biomechanical properties of bone and

incorporates with the native tissue while preserving its integrity for the required period. (Kim,

Wang See, Xiaochun, & Donghui, 2020).

The diamond representation may be used to describe the cardinal needs of bone tissue

engineering (figure nº2), which describes the following critical elements for effective bone

healing employing bone tissue engineering: An osteoconductive scaffold amenable to bone

growth, a healthy population of osteogenic cells to allow bone regrowth, growth factors to effect

cellular events to promote healing, and a good mechanical atmosphere capable of providing

sufficient enough stability for healing while still mimicking the native tissue's mechanical

characteristics. (Giannouidis, Einhorn, & Marsh, 2007)

Figure nº2: (A) Diamond concept of bone healing. (N) Natural healing process of bone: (1)

hematoma formation, (2) soft callus formation, (3) hard callus formation, and (4) remodeling.

(Giannouidis, Einhorn, & Marsh, 2007)


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The mechanism of implant-assisted bone regeneration is close to that of non-implant-assisted

bone regeneration. The implant is stabilized by the bone that extends around it, and this bone

development is dependent largely on the surface chemistry and topography of the

implanted device. Blood is one of the first tissues to come into contact with implants when is

introduced in human body, resulting in the presence and activation of platelets on the implant

surface, and therefore a subsequent inflammatory response generated by neutrophils and

macrophages, which leads to the development of hematomas. (Kuzyk & Schemitsch, 2011)

These platelets and leukocytes help to develop an environment that encourages mesenchymal

stem cells recruitment (MSCs) to the implant's surface, where their differentiate into

osteoblasts and start to generate new bone in the direction of the bone edges in a process

known as contact osteogenesis. (Davies, 2003) (Kuzyk & Schemitsch, 2011).

In a process known as distance osteogenesis, bone development happens in the opposite

direction, from the edges of the bone to the implant. (Kuzyk & Schemitsch, 2011).

Both contact and distance osteogenesis result in the development of young woven bone, which

fills the areas between the bone and the implant. (Davies, 2003) (Kuzyk & Schemitsch,

2011). The immature bone in the peri-implant bone is replaced with mature bone, which

improves the connection between the bone and the implant (remodeling). (Kuzyk & Schemitsch,

2011).

The linkages between both the bone and the implant should ideally stay stable while remodeling

proceeds; nevertheless, if somehow the peri-implant bone is not adequately stimulated, such as

due to disuse or stress shielding by the implant, the peri-implant bone can deteriorate and lead

to implant failure. (Kim, Wang See, Xiaochun, & Donghui, 2020)

On large bones of human body, this tissue is composed by two bone architectures: compact or

cortical bone and trabecular or cancellous bone. The inner part of the bone corresponds to

cancellous bone and is surrounded by cortical bone. Cortical bone's elastic modulus has been

estimated to be between 7.0 and 18.9 GPa and total porosity between 30%-90%. (Wu, Liu,

Yeung, Liu, & Yang, 2014)

Commercially pure titanium (cp Ti) is classified as biologically inert biomaterial, meaning that

when it is implanted into human bodies it does not present changes on their conformation, the

immune system identifies it as outsider but is tolerated. This action does not cause adverse

effects and are well tolerated by human tissues. (Aldani & Dominguez, 2012)

Initial and ongoing interactions between the material and the host body regulate

biocompatibility, such as molecule adsorption, protein adsorption, cell adhesion, bacterial

adhesion, macrophage activation, tissue development and inflammation. (Takao, 2019)

The formation of a strong and stable layer of oxide (TiO2) when in contact with air, the formation

limits oxygen diffusion from the environment, resulting in corrosion resistance.. (Aldani &

Dominguez, 2012)

There are four grades of cp-Ti (table nº1), classified from 1 through 4 depending on the yield

strength and allowable levels of iron, carbon, nitrogen and oxygen.


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Components (%) Grade 1 Grade 2 Grade 3 Grade 4

Iron max 0.20 0.30 0.30 0.50

Carbon max 0.10 0.10 0.10 0.10

Oxygen max 0.18 0.25 0.35 0.40

Nitrogen max 0.03 0.03 0.005 0.05

Hydrogen max 0.015 0.015 0.015 0.015

Ti (%) Balance Balance Balance Balance

Table nº1: Components of commercially pure titanium grades 1-4. (ASTM-Standard, 2005)

The study of Leonardo (Leonardo, 2013) showed that titanium has a great greed for carbon,

nitrogen, hydrogen and oxygen and the effect of these interstitials in the mechanical properties

of traction and fatigue are notable. Is possible to determine the amount of oxygen, nitrogen,

carbon and hydrogen present by instrumental analysis.

cp-Ti Grade

Young Modulus

(GPa)

Ultimate Tensile

Strength (MPa)

Yield Strength (MPa)

Elongation (%)

Density (g/cc)

1 102 240 170 24 4.5

2 102 345 275 20 4.5

3 102 450 380 18 4.5

4 104 550 483 15 4.5

Table nº2: Mechanical properties of cp-Ti grades 1-4. Source: (ASTM-Standard, 2005)

In figure nº3 we can observe the effect of oxygen in the traction resistance and creep resistance

of cp-Ti.

Figure nº3: Influence of the content of oxygen in the mechanical properties (traction) of cp- Ti.

Source: (Leonardo, 2013)


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In high temperatures, Titanium reacts with gaseous elements such as nitrogen, oxygen, and

hydrogen, therefore it must be handled in a controlled atmosphere; otherwise, the risk of

forming a thick layer of oxides, known as the "alpha case," might weaken the structure's

resistance and ductility. (Pala, et al., 2013).

Vidal (Vidal E, 2020) in her work details that in order to avoid Titanium oxidation, treatment

temperatures should be under 300ºC under air.

Titanium is a biomaterial with a high surface energy, which leads to a positive body reaction that

promotes mineral apposition on the bone-titanium interface and improves titanium

osseointegration. (Acero, Calderón, Salmeron, Verdaguer, & Consejo, 1999)

Because of their good biocompatibility, reduced in vivo corrosion, high specific strength, high

fatigue and wear resistance, ductility, and great combination of high strength and low modulus,

open-cell porous titanium scaffolds are a promising alternative for orthopedic applications.

(Lewis, 2013) (Lei Zhang, 2018) (Diana Lopes, 2018)

On the other hand, this titanium scaffold has a significant risk of loosening the implant due to

three factors: stress shielding (bone and titanium have different elastic and young modulus),

weak interfacial bonding implant/bone, and a deficiency of biological anchoring for cell in-

growth. (Lewis, 2013)

Wang & Tang (Wang & Tang, 2019) identified four main challenges can be identified in the

prevalence of implant-related infections:

(i) Insufficient mechanical stability can cause detachment of the implant surface coating, altering

the antimicrobial ability of functionalized surfaces.

(ii) Regarding drug-loaded coatings, a stable drug release profile is of vital importance for

achieving effective bactericidal effect locally; however, burst release of the loaded antibacterial

agents remains common.

(iii) Although many coatings and modified surfaces provide superior antibacterial action, such

functionalization of surfaces sometimes has a detrimental effect on tissue biocompatibility,

impairing the integration of the implants into the surrounding tissue.

(iv) Biofilm eradication at the implant surface remains particularly challenging.

Despite latest improvements in biotechnology, microorganisms can still infect the surface of

implants. Other procedures to avoid this problem includes applying a thicker oxide coating to

the implant's surface, electrochemical techniques, and plasma spraying.. (Ribeiro, Montero, &

Ferraz, 2012)

Regardless that surface coating or microstructuring can substantially reduce bacterial

adherence, it is impossible to fully remove this adhesion, and some germs may still adhere to

the implant surface. Once germs have adhered to the implant surface, a biofilm is formed, which

is difficult to eliminate. As a result, a secondary line of defense must be established to handle

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with germs that are resistant to the surface treatment's antiadhesion activity. Contact killing is

a pathogen-eradication method that aims to eliminate all attached pathogens. (Wang & Tang,

2019)

This sort of antiinfection technique usually entails immobilizing bactericidal substances on the

implant's surface, resulting in a bactericidal-capable functioning surface. The immobilized

antimicrobial agent is more permanent than coatings, which decreases the possibility of surface

morphological changes induced by coating degradation and therefore the risk of further

infection. Moreover, After binding to and killing bacteria attached to the implant surface, the

immobilized agents can recover their antibacterial activity when removed from the debris of

deceased bacteria, resulting in long-term infection control. (Costa, Carvalho, Montelaro, &

Gomes, 2011).

Bacterial biofilm will form if bacterial adhesion is not prevented and attached bacteria are not

eliminated. Treatment becomes much more difficult after a mature biofilm has formed, for two

primary reasons. First, bacteria's extracellular polymeric material acts as a barrier, preventing

chemicals from being transported. (Including antibacterial agents) through the biofilm. Second,

metabolic states of bacteria in biofilms are always substantially changed, reducing their

susceptibility to antibiotic therapy. (Costerton, Stewart, & Greenberg, 1999)

Hardness, tensile strength, Young's modulus, and elongation are the most significant mechanical

characteristics that determine the kind of material to be used as scaffold. A biomechanical

mismatch is linked to an implant fracture caused by mechanical failure. (Saenz & Fuentes, 2013)

Specifically in the mechanical aspect, Titanium elastic modulus is 110GPa and even if alloys can

go down to 55GPa, it is significantly higher than bone´s modulus (that fluctuates between 4GPa

to 30 GPa) (Magda Dziaduszewska, Andrzej Zielinski, 2019).

An important challenge for Ti scaffolds is to achieve an apparent elastic modulus similar to

nature bone to avoid mechanical issues mentioned below.

This stress shielding phenomenon is caused by the human body's predisposition to decrease or

destroy its own components when they are not in use. Even though mechanical strength is

a significant factor for scaffolds, Ti implants also need to present low stiffness to minimize bone

stress shielding. When the stresses imposed on bone by a prosthesis differ from the forces

exerted by a natural anatomical component, stress shielding occurs. (Elias, Jogaib, Moura de

Souza, Dos Santos, & de Biasi, 2019)

This discrepancy causes osteopenia (loss of bone density), which leads to bone atrophy. Lower

elastic modulus materials distribute stress better at the implant-bone contact, resulting in

reduced bone atrophy. (Niinomi & Nakai, 2011) (Schmidutz, Agarwal, Müller, & Gueorgui, 2014)

Methods to achieve a better match between bone and implant include: modification of the

geometric of the implant, modification of its material properties, or a combination of both

material and geometrical modifications. (Sajad, Burnett, & Michael, 2016)

Due to a combination of outstanding properties such as high strength, low density, high

corrosion resistance, complete inertness to the biological environment, elevated compatibility,


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low Young's modulus, and high capacity to join with bone or other tissues, titanium alloys are

rapidly emerging as the first choice for the majority of applications in biomedicine. In compared

to traditional stainless steels and cobalt-based alloys, titanium have a lower Young's modulus,

higher biocompatibility, and better corrosion resistance, making titanium solutions scaffolds an

excellent choice for bio-applications. (Liu, Chu, & Ding, 2004)

When metals and alloys enter under contact with body fluid, they are subjected to corrode since

body environment is extremely hostile due to the presence of chloride ions and proteins. On the

surface of an implanted alloy, a number of chemical reactions take place. The alloy's metallic

components are reduced to their ionic forms and dissolved oxygen results into hydroxide ions.

(Saenz & Fuentes, 2013) .

It is well known that the protective and stable oxides (TiO2) on titanium surfaces can promote

osseointegration. The chemical structure and thickness of the film have a significant impact on

the oxide's stability. (Zhu, Chen, Scheideler, Reichl, & Geis-Gerstorfer, 2004).

(Wang, et al., 2016) performed a study where subjected different titanium samples to oxidation

treatments during 2,4 and 6 hours. The nano-scaled titanium surface grew sturdier, the rutile

ratio increased, and the oxide layer thickness rose after 6 hours of heat oxidation treatment,

resulting in improved cell adhesion and osteogenic activity in vitro. Despite the lack of a

substantial change in various cell activities, such as cell attachment, cell adhesion, and cell

shape, osteogenic activity in vitro and in vivo were both enhanced. Cells grown on the three

oxidized titanium surfaces developed faster and had more osteogenic activity than cells cultured

on the control samples.

The disintegration rate of a passive metal at a given potential is significantly lower than that of

an active metal due to the existence of an oxide layer. It is mostly determined by the passive

film's characteristics and solubility in the environment. These layers, which develop

spontaneously on the metal's surface, doesn’t allow metallic ions and/or electrons from moving

over the film. In order to conform an effective protective barrier, films must have an atomic

structure that prevents ions and/or electrons from migrating across the metal oxide–solution

interface. Additionally, must be tight and completely cover the metal surface. They must be able

to stay on the surface of these alloys even when subjected to mechanical stress or abrasion, as

is common with orthopedic devices. (Hallab, Urban, & Jacobs, 2004).

The lack of appropriate tribological properties (wear, friction & lubrication) and potential

corrosion problems have led to the development of surface treatments that highly affect and

improve near-surface strength, achieve better hardness and abrasive wear resistance, reducing

the coefficient of friction, and avoiding or reducing the transference of ions from the surface or

bulk material to the surrounding tissue. To achieve the desired biological reactions after

implantation, great effort has been put towards thickening and stabilizing surface oxides on

titanium. The biological reaction to titanium is determined by the chemical constitution of the

surface as well as the capacity of titanium oxides to absorb molecules and integrate

components. Surface topography is critical in managing cell behavior, such as cell shape,

orientation, and adhesion. (Saenz & Fuentes, 2013)


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Different alloy chemistry and thermomechanical treatments can be applied to create a large

range of microstructures. These various alloys aid in lowering Titanium's young modulus to a

comparable rate of natural human bone. (Chen Q, 2015). Niobium, Zirconia, Molybdenum, and

Tantalum are alloy ‘elements that may be used to lower the modulus of elasticity without

sacrificing strength. (Li N et al, 2019)

Healing and remodeling process are influenced by surface roughness. In vitro, osteoblastic cells

bond to rough metal surfaces effectively. (Ratner, Hoffman, Schoen, & Lemons, 2013).

Porosity is significant because it influences fatigue response of the implant. This factor is

calculated in scaffolds to find the balance between preserving the implant's mechanical strength

while this value also enables tissue ingrowth. (Cameron, Pilliar, & Macnab, 1976)

Porosity encourages cells to develop into desirable physical forms and encourages ingrown

tissue vascularization. Scaffolds must have a total open porosity of 60%-70% or higher, with a

minimum pore size of 100 µm to ensure adequate cell penetration and vascularization. (Mitra,

Tripathi, Sharma, & Basu, 2013)

The ingrowth of bone tissue into pores leads mechanical anchoring. Even when surface

treatment enhances the chemical composition, the impacts of not just the chemical composition

but also the roughness generated by the treatment appear to accelerate bone growth and bone

bonding in the majority of cases. (Takao, 2019)

Her-Hsiung, et al (Her-Hsiung, et al., 2004) conducted a research into the influence of ground Ti

surface roughness on the first adhesion of osteoblasts. For commercial Ti implants with

machined surfaces, the range of surface roughness was studied. U-2 OS cells (105 cells/3.8 cm2

in density) were cultured on the metal specimens with different surface roughness. Three of the

samples suffered surface modification trough sandpaper #120, #600, and #1500 and two

samples Al2O3 powder polishing (1μ and 0.3μm). Results of roughness are shown on table nº3

and percentage of optical density (OD) (%) on figure nº4.

Since light absorbance is proportional to the quantity of the absorbing particles in the sample,

optical density is a typical approach for quantifying the concentration of substances (Beer-

Lambert law).

#120 #600 #1500 1μm 0.3μm

Ra(μm) 1.2 ± 0.12 0.33 ± 0.005 0.15 ± 0.004 0.007 ± 0.01 0.005 ± 0.002

Table nº3: Surface roughness (Ra) results for the 5 surface modified titanium samples. Source:

(Her-Hsiung, et al., 2004)


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Figure nº4: Percentage of optical density (OD) (%) of ground Ti specimens after 2 h cell

incubation. (Her-Hsiung, et al., 2004)

Their study showed that when surface roughness increasing (figure nº6), lead into a surface in

which topography revealed a more significant groove and ridge morphology. In terms of the OD

value, the polished Ti samples (with values of Ra:0.05μm and 0.07μm) had a surface roughness

less optimal for initial cell adhesion, while the #1500 Ti samples with an Ra of 0.15μm had the

optimal cell adhesion performance. For the ground Ti specimen with an Ra higher than 0.15μm,

the OD value decreased with an increase in the roughness (Ra: 0.15μm–1.20μm). (Her-Hsiung,

et al., 2004).

After 2 hours of cell incubation (figure nº5 and figure nº6), results demonstrated that the surface

roughness of Ti is an important factor on the attachment of osteoblast. In the commented study,

the ground Ti specimen with an Ra value of 0.15 μm has the best cell adhesion and spreading

appearance as compared with either the smoother (Ra: 0.05μm and 0.07μm) or rougher (Ra:

0.33μm and 1.20μm) specimens. (Her-Hsiung, et al., 2004).

Figure nº5: Crystal violet staining micrographs a)Ra:1.2 ± 0.12 µm: , b)Ra: 0.33 ± 0.005 µm, and

c)Ra: 0.15 ± 0.004 µm, of ground Ti specimens after 2 h cell incubation (light violet: cytoplasm;

dark violet: nucleus).Source: (Her-Hsiung, et al., 2004)


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Figure nº6: Crystal violet staining micrographs a)1μm polished and b)0.3μm polished Ti

specimens after 2 h cell incubation (light violet: cytoplasm; dark violet: nucleus). Source: (Her-

Hsiung, et al., 2004)

2.2. 3D Printing

3D printing is a process for creating a diverse range of structures and challenging geometries

from three-dimensional digital models. Printing consecutive layers of materials that are

produced on top of each other constitutes the process. Charles Hull invented stereolithography

(SLA) in 1986, which was followed by advancements such as powder bed fusion, fused

deposition modeling (FDM), inkjet printing, and contour crafting (CC). Novel materials and

additive manufacturing (AM) techniques like 3D printing are constantly being explored, resulting

in the development of new applications. The expiration of previous patents, which has given

manufacturers the opportunity to build new 3D printing equipment, is one of the primary

reasons for this technology becoming more accessible. Metals, polymers, ceramics, and

concrete are just a few of the materials utilized in 3D printing today. The poorer mechanical

characteristics and anisotropic behavior of 3D printed components, on the other hand, continue

to limit the possibilities of large-scale printing. (Tuang, Kashani, Imbalzano, Nguyen, & Hui, 2017)

The precision of the printed components, on the other hand, is determined by the accuracy of

the printing process and the printing scale. Micro-scale 3D printing, for example, has problems

with resolution, surface quality, and layer bonding, requiring post-processing methods like as

sintering. (Vaezi, Seitz, & Yang, 2013)

3D printing is now extensively used all around the world. In the fields of agricultural, healthcare,

automotive, and aerospace, 3D printing technology is rapidly being used for mass customization

and manufacture of any sort of open source design. (Keles, Blevins, & Bowman)

Nowadays is possible to print in 3D a wide spectrum of structures with a large variety of

materials with different properties according to the purpose. Complex 3D printed structures

with highly accurate and specific architecture is one of the biggest advantages that additive

manufacturing breakthrough in biomaterials is bringing to the biomedical field. Most of these

geometries are impossible to manufacture with traditional techniques, thus 3D printing is

solving an important problem for complex surgical cases.The process to obtain a personalized

3D printed scaffold for bone in humans, is explained in figure nº7.


18

Figure nº7: The process stages of titanium EBM (electron bean melting 3D printing technique)

implant production. Source: (Vladimir Popov Jr., et al., 2018)

Figure nº8: Titanium 3D printed scaffolds. Source: Left: Renishaw, right: EOS.

The goal of 3D printing technology is to provide solutions to challenges like mechanical

mismatches between titanium scaffolds and natural human bone. It is possible to alter the

volume fraction of porous structures, the degree of porosity, the size and form of the pores, and

therefore modify mechanical behavior, by adjusting several parameters in the 3D printing pre-

process. (Zhao S, 2016). Furthermore, this defined porosity allows for cell growth, support for

nutrients and waste movement, and enhances angiogenesis. (Wang, 2017)

Different techniques of 3D printing have been developed during the last 20 years and have been

used to print titanium as a biomaterial for different applications on human body.

Ni et.al (Ni, et al., 2019) performed a bibliographic revision of the most important 3D printing

applications for biomaterials, summarizing their application field, objectives and benefits. (table

nº4)


19

Application field Objectives Benefits

General-purpose complex

implants

Provide complex-structure

implants for general purposes

including hip, knee, shoulder,

oral implants, and so on.

Metal implants with complex

3D internal structures can be

produced by 3D printing

technology with high speed

and high precision.

Personalized permanent

implants

Solve the problem brought by

the mismatch between

general-purpose implants and

patients with individual

differences such as poor

adaptability, intraoperative

incomplete coverage, implant

wear and loosening, and so

on.

3D printed metal implants can

be custom designed and

fabricated into specific

structure to meet the

individual requirements of

different patients, and thus

become a personalized

treatment.

Porous implant prosthesis

Circumvent the stress

shielding effect brought by

the dense metal implants due

to the mismatch of stiffness

and elastic modulus between

them and human bone tissue

3D printing technology is

significantly superior to

traditional machining

methods in constructing

porous metal implants

Personalized surgical tools

Reduce the long processing

cycle for traditionally

personalized surgical tools

Personalized surgical tools

fabricated by 3D printing

technology enable more

precise procedures, simplify

the operation, and increase

the operational speed and

efficiency.

In vitro medical devices

Provide in vitro medical

models and devices such as

prosthesis, hearing aids,

dental surgery template

models, and so on.

With 3D printing technology,

simulation models of human

tissues and bones can be

printed out in advance, which

allows surgeons to practice

the operation procedures,

thus enabling more accurate

and safer actual surgical

operations.

Table nº4: Application fields of 3D-printed biometals in medical devices. Source: (Ni, et al., 2019)

Binding jetting, directed energy deposition, material extrusion, material jetting, powder bed

fusion, sheet lamination, and vat photopolymerization are the seven 3D printing processes

classified by ASTM. There are no disagreements about whether equipment or technology

performs better because each has its own set of uses. 3D printing technologies are no longer

restricted to prototype but are rapidly being utilized to create a wide range of goods.

(Shahrubudin, Lee, & Ramlan, 2019)


20

2.2.1. Binder Jetting

To connect powder particles, a liquid binding agent is applied selectively. Binder jetting

technique forms a layer by spraying a chemical binder onto the dispersed powder. (Ze-Xian, et

al., 2016)

Binder jetting is used to make casting patterns, raw sintered goods, and other large-volume

products from sand. Metals, sands, polymers, hybrids, and ceramics are some of the materials

that binder jetting can print. Some materials, such as sand, may not require further processing.

Binder jetting is also easy, quick, and inexpensive since powder particles are bonded together.

Finally, binder jetting allows for the printing of very large objects. (Shahrubudin, Lee, & Ramlan,

2019)

2.2.2. Direct Energy Deposition

Thermal energy is used to melt materials (in powder or wire form) as they are placed to fuse

them. Because of the reduced precision and post-processing requirements, directed energy

deposition for metal AM is less common for whole-part additive manufacturing. (Duda & Venkat,

2016)

Directed energy deposition (DED) is a more complicated printing method that is frequently used

to repair or add material to existing components.

The idea behind directed energy deposition is similar to extrusion of material, however the

nozzle is not locked to a single axis and may move in any direction. In addition, this technique

may be utilized with ceramics and polymers, although it is most commonly employed with

metals and metal-based hybrids in the form of metals and metal-based hybrids. Wire or powder

can be used. (Shahrubudin, Lee, & Ramlan, 2019)

2.2.3. Materials jetting

A printhead distributes droplets of a photosensitive substance that hardens, layer by layer,

creating a component under ultraviolet light. (Silbernagel, 2018)

Material jetting offers multi-material printing and a wide range of materials such as polymers,

ceramics, composites, biologicals, and hybrids. Material jetting produces components with an

extremely smooth surface finish and excellent dimensional accuracy at the same time. (Syed, et

al., 2017).


21

• Direct Ink Writing

This was the 3D printing employed to print the scaffolds of this thesis. The ink is deposited onto

a surface in a predefined order creating a continuous filament writing. New sections can then

be layered on top of each other to build a multilayer 3D object. (Solis, Smirnov, Peretyagin,

Seleznev, & Peretyagin, 2020)

Figure nº9: Scheme of direct ink writing process. Ink is deposited in the syringe and is extruded

thought-out the nozzle in order to built the object. Syringe moves along axis X and Y, while built

platform performs movements along Z axis. Source: (Solis, Smirnov, Peretyagin, Seleznev, &

Peretyagin, 2020)

2.2.4. Materials Extrusion

Plastics, food, and live cells may be printed in multi-materials and multi-colors using material

extrusion-based 3D printing technology. (Muller & Karevska, 2016). This method is widely

utilized, and the costs are low. Furthermore, this method may be used to create completely

functional product pieces. (Syed, et al., 2017)

The earliest example of a material extrusion system is fused deposition modeling (FDM). Was

created in the early 1990s and is based on the usage of polymer as the primary material.

(Stansbury & Idacavage, 2016).

By heating and extruding thermoplastic filament, FDM creates parts layer by layer from the

bottom to the top. (Shahrubudin, Lee, & Ramlan, 2019)


22

2.2.5. Powder Bed Fusion (PBF)

Thermal energy selectively fuses areas of the powder bed in PBF-based technologies. (Duda &

Venkat, 2016)

The electron beam melting (EBM), selective laser sintering (SLS), and selective heat sintering

(SHS) printing techniques are all used in the powder bed fusion process. To melt or fuse the

material powder together, this approach employs an electron beam or a laser. Metals, ceramics,

polymers, composites, and hybrids are examples of materials utilized in this procedure. Powder-

based 3D printing method is most often known as selective laser sintering (SLS). In 1987, Carl

Deckard invented SLS technology. SLS is a 3D printing process that is functionally fast, accurate,

and offers a wide range of surface finishes. (Tiwari, Pande, Agrawal, & Bobade, 2015).

Figure nº10: Concept of SLM process. (i) High-power laser melts selective areas of the powder

bed. (ii) Process is repeats for successive layers. (iii) Loose powder removed and finished part

revealed. (Yap, et al., 2015)

2.2.6. Sheet Lamination

Sheet lamination has the advantages of being able to print in full color, being reasonably

affordable, ease material handling, and the ability to recycle leftover material.

Two of the most important techniques that uses this technology are Laminated Object

Manufacturing (LOM) and Ultrasound Additive Manufacturing (UAM). LOM allows for the

creation of complex geometrical elements at a cheaper cost and with less operation time. UAM

is a cutting-edge technique that combines layers of metal taken from featureless foil material

using sound. (Vikayavenkataraman, Jerry, & Wen, 2017)

2.2.7. Vat Photopolymerization

Stereolithography (SLA) and digital light processing (DLP) are two examples of

photopolymerization-based 3D printing technologies. (Stansbury & Idacavage, 2016)

A more traditional light source, such as an arc lamp with a liquid crystal display panel, is used in

the DLP. It can cover the whole surface of the photopolymer resin pot in a single step, making it

quicker than other methods. (Reddy, 2019).


23

Photopolymerization is ideal for creating a premium product with fine features and a high

surface quality. The period of exposure, wavelength, and quantity of power supply are all critical

elements in Vat Photopolymerization. Physical state at the beginning of the process are liquid,

but when subjected to UV radiation, they solidify. (Muller & Karevska, 2016)


24

3. MATERIALS AND METHODS

In order to obtain the designed geometries, is necessary to follow the process:

- Ink fabrication

- 3D impression

- Binder elimination and sintering process

3.1. Ink fabrication

Ti powder (~ 325 mesh, 99,5% purity, density: 4.5g/cm³, Alfa Aesar, MA, USA) is the titanium

powder used in this study. To form an hydrogel and mix it with the Ti Powder, Pluronic F-127®,

(density: 1.1g/cm³, Sigma-Aldrich, St Louis, MO, USA) was used. The hydrogel was prepared by

dissolving Pluronic F-127®in distilled water by means of a off-axis planetary mixer (Hauschield

SpeedMixer®, Hauschdil, GmbH, Germany), - at 2.500 rpm for 5 minutes (30% w/v). After this

mixing process, titanium powder is added to the hydrogel and mixed 2.500 rpm for 5 minutes.

The proportion of hydrogel/titanium of the ink is based on the work of Elia Vidal (Vidal E, 2020),

because this concentration has an appropriate viscosity for 3D printing (the load to extrude the

ink is supported by the printer) and the fraction of metal powder is enough to ensure mechanical

properties required. In Figure nº12 we can distinguish the first image on the left from the image

in the center, where in this last one titanium is already added and finally the texture of the mix

is a consistent paste.

Titanium Powder (%w/w) Hydrogel (%w/w)

69.0 31.0

Table nº5: Ink composition chosen to perform the samples. Source: Elia Vidal, 2020.

Figure nº11: left to right: Pluronic F127 hydrogel, final titanium ink and syringe with ink previous

printing process.


25

3.2. Printing parameters

All the samples were printed using a modified Reprap 3D printer for extruding pastes (Reprap,

Centre CIM, UPC, Barcelona). This printer was the first low-cost 3D printer and is capable of

printing some of his parts in order to replicate the printer.

This printer uses dense materials (paste) that are extruded at high pressure by a piston. It is

possible to regulate the temperature of the paste before extrusion, in order of obtain a desired

consistence on the printed paste. In the case of this study, was not necessary to heat the paste

in order to obtain a certain consistence because the material used had enough viscosity to flow.

To proceed with the printing of the samples, is necessary to adjust the parameters in order to

obtain the different geometries. Extrude parameters are the same for all the samples. A total of

20 samples of 2 different geometries (rectilinear and full honeycomb, with 45% and 50% infill,

creating a variety of 4 different type of samples) were printed using this parameter.

The program used to modify the models and obtain the final geometries of the pieces is Simplify

3D v3.0.2. (Simplify 3D, OH, USA).

One of the most important parameters that is able to be adjusted in order to obtain a feasible

sample is “Infill”, that is related to the density of the 3D printed object. It is possible to adjust

this value to obtain different patterns with the same geometry, that in this case correspond to

rectilinear and fully honeycomb. If this value is reduced, the printed piece have more hollow

space, increasing this value implies a higher density of the sample.

Velocity of the printing process itself (how fast the nozzle moves to create the geometry) is an

important factor to control because a non-appropriate value will lead on layer tinner than

expected and non-accurate shape if the velocity is too fast and ticker and distorted layers if the

velocity is too low.

Different values of infill were experimented, was not possible to increase the infill over 50%

because the samples could not resist their own weight, thus 50% the maximum value of infill

using this ink and 45% was the minimum.


26

Figure nº12: Extruder parameters for printing. Source: Simplify 3D Software

• Rectilinear, infill 45%

Figure nº13: printing parameters for rectilinear infill 45% sample, using Modify3D Software.

• Rectilinear, infill 50%

Figure nº14: printing parameters for rectilinear infill 50% sample, using Modify3D Software.


27

• Full Honeycomb, infill 45%

Figure nº15: printing parameters for full honeycomb infill 45% sample, using Modify3D

Software.

• Full Honeycomb, infill 50%

Figure nº16: printing parameters for full honeycomb infill 50% sample, using Modify3D

Software.

3.3. Binder elimination and sintering process

In order to reduce the contamination of the samples during binder elimination and sintering

heat treatment, clean ceramic platforms with zirconia (ZrO₂) beads were used to position them

inside the oven , as shown in figure nº17.

Figure nº17: Samples positioned in the ceramic base with zirconia previous insertion in the

vacuum tubular furnace.


28

With the aim of eliminating the binder (debinding) and avoid titanium oxidation (under 300ºC

in air atmosphere), after 6 hours of air drying at room temperature, the samples were heated in

an oven at 275ºC for 12 hours in air atmosphere. Immediately after, scaffolds were sintered in

a vacuum tubular furnace (Hobersal, Spain) at 1.200 ºC for 2 hours, at a pressure of 1·10-5 mbar.

The sintering atmosphere must fulfill the following functions: protecting the material from

oxidation during the sintering process, reducing surface oxides in order to improve the contact

between adjacent particles and preventing carburization of manufactured Titanium samples due

to incomplete binder removal.

The presence of impurities that could contaminate the samples is reduced because of the high-

quality vacuum created inside the furnace (below 10-5 mbar), that removes air and impurities

from inside the vacuum chamber constantly by means of a diffusion pump. (Peawband &

Thedsakhulwong, 2017)


29

3.4. Physicochemical Characterization

3.4.1. SEM

In order to obtain SEM images of the samples, equipment used was JEOL JSM-7001F Field

Emission Scanning Electron Microscope (JEOL, Japan).

Figure nº18: Field Emission Scanning Electron Microscope. Source: JEOL

3.4.2. Micro-CT

Micro-CT is a high-resolution X-Ray 3D imaging technology that permits the viewing of a sample's

interior structure without damaging it. It is possible to analyze the photos and create a 3D model

(reconstruction) of the item as well as a cross-section expansion using the related software, and

it is also feasible to determine porosity, structural thickness, and object volume using the images

and 3D model. In this case, was used to obtain the total porosity of the samples.

Equipment used to obtain total porosimetry was High-resolution 3D X-ray Microscopy Skyscan

1272 (Bruker, Belgium).

Figure nº19: micro-CT employed to obtain total porosity of the samples.


30

3.4.3. Mercury Intrusion Porosimetry (MIP)

This method is used to determine only the open porosity of a material by applying different

levels of pressure on a sample immersed in mercury. The amount of pressure needed to get the

mercury into the pores is inversely related to the size of the holes. This enables the entrance

size of micro pores and macro pores.

MIP was performed to samples on sample of each geometry and infill (in total 4 samples were

subjected to these assay), in order to analyze open porosity and micropore entrance size

distribution. This technique only allows you to measure diameters from 0.003 µm to 360 µm. In

the case of the printed samples in this study, macropores exceed 360 µm.

To characterize the sample, size smaller than 50µm are considered micropores and higher than

50µm are macropores.

Parameters used were evacuation pressure 50 μmHg, evacuation time 5 minutes, mercury filling

pressure 0,50 psia and equilibration time 30 seconds.

With the value of density (ρbulk, with volume obtained with micro-CT and weight of the sample),

calculations were made to obtain total porosity of the samples and knowing that density of cp-

Ti is 4.5 g/cm3 (ρ cp − Ti ), applying equation nº1:

𝑇𝑜𝑡𝑎𝑙 𝑃𝑜𝑟𝑜𝑠𝑖𝑡𝑦 (%) = 100 𝑥 (1 −(ρbulk )

(ρ cp−Ti)) Equation nº1

Is important to mention that total porosity is a value given by Micro-CT but is possible to obtain it also through equation mentioned before. The difference between the results using both methods, is due to lack of accuracy in micro-CT assay regarding the edges of the samples (real volume). Afterwards, to obtain microporosity, equation nº2 was applied: 𝑃𝑒𝑟𝑐𝑒𝑛𝑡𝑎𝑔𝑒 𝑜𝑓 𝑀𝑖𝑐𝑟𝑜𝑝𝑜𝑟𝑒𝑠 (%) = (𝑉𝑚𝑖𝑐𝑟𝑜 𝑥 ρbulk )𝑥100 Equation nº2 Where: 𝑉𝑚𝑖𝑐𝑟𝑜= volume of micropores obtained after the sum of the incremental mercury intrusion in micropores (size lower than 50 μm) In order to obtain closed porosity, we applied equation nº3:

𝐶𝑙𝑜𝑠𝑒𝑑 𝑃𝑜𝑟𝑜𝑠𝑖𝑡𝑦 (%) = 1 −(ρ skeletal )

(ρ cp−Ti)) Equation nº3

Where: ρ skeletal correspondes to apparent density, obtained with MIP.


31

3.4.4. Roughness

Chromatic Confocal Microscopy equipment "Station Micromesure 2 3D" (STIL, France) was the

equipment used to obtain the parameters of roughness on the different samples, one per

geometry.

The optical axis is “colour coded” by passing incident white light through a chromatic lens, which

produces a continuous stream of monochromatic light along the z-axis. When a single

wavelength is fixed to the surface of an item in this color field, it is reflected back to the optical

system. The backscattered beam is collected by a spectrometer after passing through a filtering

pinhole. To accurately determine the position of the surface in the measurement field, the

particular wavelength of the beam is determined. Chromatic confocal technique enables for

high-resolution dimensional measurements that are reliable, accurate, and repeatable.

Figure nº20: Scheme of the chromatic-confocal measurement principle. Source: Polytec

For each sample, an area of 0,1mm2 was considered (X=200μm, Y=500μm), and measures were

made every 2μm. Before starting with measures procedure, special attention to the inclination

of the sample in the platform of the equipment was taken, in order to obtain an horizontal

surface, maintaining the intensity of the focus in the moving Z axis.

Data obtained from these measures was analyzed using SPIP 6.6.5 software (STIL, France).

Using the investigation of Saurí (Saurí, et al., 2015) to describe the different amplitude roughness

parameters obtained:

- Sa: arithmetical mean height (mean surface roughness). Average distance to mean

surface.


32

- Sq: root mean square height of the surface. Standard deviation of the height distribution.

Higher amplitude peaks and valleys have more weight in the quadratic mean.

- Sz: defined as the average of the five highest local maximums plus the average of the

five local minimums.

- Sv: Depth of deepest valley

- Sp: Height of highest peak

Figure nº21: Surface Roughness parameters. Source: ISO 4287:1997. Geometrical Product

Specifications (GPS) — Surface texture: Profile method — Terms, definitions and surface texture

parameters.


33

3.5. Mechanical Characterization

3.5.1. Compressive Strength

Assays were made with a servohydraulic universal testing machine Bionix 858 (MTS, CA, USA),

according to ISO 13314:2011 (it specifies a test method for compressive properties of porous

and cellular metals). Speed of the press was set at 1 mm/minute.

Figure nº22 shows an example of the results of this assay, where point 3 of the curve represent

the maximum compressive strength.

In this figure is possible to distinguish two regions, the first one corresponds to a quasi-elastic

region followed by a long plastic deformation plateau.

In the case of this structures, deformation causes a rapid rise in stress in the elastoplastic area.

Stress concentration occurs in particular areas, such as sinter necks, due to the non-uniform

distribution of stress in the porous titanium structure. This results in local plastic deformation

below the porous structure's yield strength. The main contributor of deformation at this stage

is macroscopic buckling of the cell walls. (Özbilen, Liebert, Beck, & Bram, 2016)

In the Y axis is represented σ, hat corresponds to compressive force divided by the initial cross-

sectional area perpendicular to the loading direction. Compressive stress is expressed in

newtons per square millimeter. In X axis compressive strain is represented, that is the overall

compressive displacement divided by the initial height (gauge length) of the test specimen.

Additionally, “first maximum compressive strength” is defined as the first local maximum in the

stress-strain curve if there is one (ISO13314, 2011).

Figure nº 22: Stress-strain curve to determine the characteristic values from compression testing

of porous and cellular metals. Source: (ISO13314, 2011)

Young's modulus (also known as elasticity modulus) is the slope of the linear part of the stress-

strain curve for a material under tension or compression, this describes the elastic

characteristics of the material. Young’s modulus is equal to the longitudinal stress divided by the

strain.


34

3.5.2. Vickers Hardness

Vickers Hardness test was performed using Hardness Testing Machine model MVK-HO, Akashi,

Japan. Force applied was 0,2 kgf. As a reference, annealed Cp-Ti Vickers Hardness is around200

HV. (Vidal et al, 2020)

Angle of the diamond indenter 136 ° and knowing force applied was 0,2kgf, with these values

and the average length of diagonals left by the indenter (d), is possible to computer the hardness

values. A total of 8 measures were made on each geometry.

𝐻𝑉 = 1,8544 𝐹

𝑑2 [𝑘𝑔𝑓

𝑚𝑚2]

Figure nº23: left: Sample prepared and polished ready for chemical etching, right: Vickers-

Hardness picture result.


35

4. RESULTS

4.1. Characterization of structures

Initially 20 of the samples were sintered at Ames Sintered metallic components, enterprise

located in Barcelona that is dedicated to the manufacture and sale of sintered metal

components. These samples were subjected to different assays in order of obtaining a

characterization of the structures.

Geometry Sample Test

Rectilinear Infill 45%

(R-45)

R-45/1 Compression / SEM

R-45/2 Reserved

R-45/3 Compression

R-45/4 Compression

R-45/5 SEM / MIP

Rectilinear Infill 50%

(R-50)

R-50/1 Compression / SEM

R-50/2 Reserved

R-50/3 Compression

R-50/4 Compression

R-50/5 SEM / MIP

Full Honeycomb Infill 45% (FH-45)

FH-45/1 SEM / MIP

FH-45/2 Reserved

FH-45/3 Compression

FH-45/4 Compression / SEM

FH-45/5 Compression

Full Honeycomb Infill: 50%

(FH-50)

FH-50/1 SEM / MIP

FH-50/2 Compression

FH-50/3 Compression / SEM

FH-50/4 Reserved

FH-50/5 Compression

Table nº6: Geometries and test performed on each sample.

Figure nº24: Rectilinear infill 45% 3D printed titanium samples, before sintering process. Left: sample R-45/1, right: sample R-45/2.


36

Figure nº25: Rectilinear infill 50% 3D printed titanium sample R-50/1, before sintering process.

Figure nº26: Full Honeycomb infill 45% 3D printed titanium sample FH-45/1, before sintering process.

Figure nº27: Full Honeycomb infill 50% 3D printed titanium sample FH-50/2, before sintering process.


37

Other 4 samples (one each geometry) were printed and sintered using the equipment available

in the University, using a modified vacuum tubular furnace at 1.200ºC for 2 hours. This samples

were used to measure the Vickers hardness.

Figure nº28: Samples after sintering process.


38

4.2. Physicochemical Characterization

4.2.1. Dimensional changes

All the samples were subjected to high temperatures to eliminate the binder and obtain a pure

titanium piece, subsequently the vacuum tubular furnace elevated the temperature of the

pieces.

In both process the samples experimented morphological changes (dimensional changes) (figure

nº24). Removal of the binder and the mass loss this implies, triggered the shrinkage of the

sample. Height variation showed a reduction of the dimensions between 24.07% and 25.82%,

both are full honeycomb geometries and the one that suffered the biggest reduction

corresponds to infill 45%. In the other hand, the one with less height reduction had 50% infill.

Measures of the dimensions of each sample were made pre and post sintering in order to

obtain the contraction of the samples (figure nº25). Sample number 3.1 (R-50) could not be

measured due to a human mistake.

Sample Height Variation (%) Diameter Variation (%)

R-45/1 to R-45/5 25.08 ± 1.18 34.36 ± 3.21

R-50/1 to R-50/5 24.64 ± 4.10 37.03 ± 1.91

FH-45/2 to FH-45/5 25.82 ± 1.68 35.59 ± 1.74

FH-50/1 to FH-50/5 24.07 ± 1.16 37.34 ± 1.94

Table nº7: Height and diameter variation of samples according to the different geometries and

infill value.

Diameter variation were mostly experimented by full honeycomb samples with 50% infil, with

37.34% of reduction. By contrast the rectilinear 45% infill sample had the lowest diameter

variation (34.36%).


39

4.2.2. SEM

Figure nº29: top to bottom: Different printed structures x12 and x50, corresponds to a) R-45,

b)R-50, c)FH-45 and d)FH-50.

In SEM images of figure nº29 is possible to observe the different amount of contact points

between filaments of the different layers on each geometry printed, where figure d), which infill

is 50% and geometry is fully honeycomb, we can observe a higher number of filaments in contact

with another filament on the different layers. This images corresponds to samples subjected to

compression assay in order to analyze their resistance.

SEM images showed on figure nº29 to nº32 to different number of pores, micropores (on the

filament, size lower than 50µm) and macropores (between filaments, size bigger than 50µm). It


40

is possible to notice that honeycomb infill 45% has a higher number of visible pores. On the

other hand, the rectilinear infill 45% samples showed the lower number of visible pores.

Figure nº30: SEM images of sample rectilinear infill 45%, R-45/5.

Figure nº31: SEM images of sample rectilinear infill 50%, R-50/5.

Figure nº32: SEM images of sample full honeycomb 45% Infill, FH-45/1


41

Figure nº33: SEM images of sample full honeycomb 50% Infill, FH-50/1


42

4.2.3. Porosity

Was possible to obtain the parameters of porosity with MIP and Micro-ct results. Considering

that for those pores with a size smaller than 50µm are considered micropores and higher than

50µm are macropores.

With the values of weight, volume and density obtained with Micro-CT and using equation nº1,

was possible to compute closed porosity.

Sample Weight

(g) Bulk Density

(g/cm3) Vmicro MIP

(cm3/g) Skeletal Density

(MIP) (g/cm3) Total Porosity

(equation nº1) (%)

R-45/5 1.770 1.845 0.013 4.251 59,1

R-50/5 1.909 1.928 0.011 4.168 57,3

FH-45/1 1.710 1.616 0.016 4.276 64,2

FH-50/1 1.815 1.718 0.015 4.253 61,9

Table nº8: Porosity parameters and total porosity result according to equation nº1.

With the aim of compute microporosity, closed porosity and microporosity and by applying

equations nº2 and nº3, was possible to generate table nº9 with the results for each sample:

Total Porosity Micro-CT (%) (a)

Microporosity (%) (equation nº2) (b)

Closed Porosity (%) (equation nº3) (c)

Macroporosity (%) (a-b-c)

61.3 2.44 0.06 58.80

56 2.17 0.08 53.75

63.6 2.60 0.05 60.94

61.9 2.63 0.06 59.21

Table nº9: Microporosity, microporosity and closed porosity results for each sample studied.

MIP results shows that Micropore size average was found between 11,78μm and 23μm. Samples

with higher open microporosity is honeycomb with 50% of infill, followed by the same geometry

with 45% infill. Rectilinear 50% infill samples is the one with lowest porosity. Honeycomb

geometry has a bigger macroporosity possibly because there are more contact points between

filaments than in the rectilinear geometry.

Results of distribution of entrance pore diameter are shown on figure nº34. In general, size of

the micropores are similar for all the samples. Macro porosity results are also similar except for

R-45/1 where the number of smallest micropores are considerably less in comparison with the

rest of the samples.


43

Figure nº34: Entrance pore diameter distribution for samples.

Micro-CT total porosity exposed that both honeycomb 45% and 50%, had more total porosity

with values of 63,6% and 61,9% respectively. The lowest values correspond to rectilinear infill,

where the lowest porosity is 56% (rectilinear 50% infill).

4.2.4. Roughness

Values obtained after measures and the processing of the data for the samples can be seen on

figure nº33.

Regarding rugosity, results showed on table with an average roughness (Sa) between 3 and 5

microns, are similar between samples.

Sample Sa (µm) Sq (µm) Sz (µm) Sv (µm) Sp (µm)

R-45/1 3.296 4.405 40.887 25.146 15.741

R-50/4 3.525 4.574 33.258 14.767 18.491

FH-45/3 4.743 6.344 60.825 34.739 26.086

FH-50/5 3.829 9.188 100.038 90.203 10.181

Table nº10: Rugosity parameters for 4 samples (one each geometry).

00,010,020,030,040,050,060,070,080,09

0,10,110,12

2,3

E-0

1

2,8

E-0

1

3,6

E-0

1

5,2

E-0

1

8,1

E-0

1

1,2

E+0

0

1,5

E+0

0

2,0

E+0

0

2,5

E+0

0

3,1

E+0

0

4,1

E+0

0

5,9

E+0

0

6,6

E+0

0

8,0

E+0

0

1,0

E+0

1

1,2

E+0

1

1,5

E+0

1

2,0

E+0

1

3,0

E+0

1

4,5

E+0

1

1,0

E+0

2

3,0

E+0

2

dV

/dlo

gD

Entrance Pore Diameter (µm)

MIP

R-45/1

R-50/5

FH-45/1

FH-50/1


44

4.3. Mechanical Characterization

4.3.1. Young Modulus

Young modulus of the samples are situated between 2.5 6GPa and 3.8GPa (figure nº37) with

standard deviations between ±5% and ±13%. The geometry that experimented higher young

modulus is rectilinear with infill 50% and both honeycombs (45% and 50% infill) performed the

lowest young modulus of the entire set of samples.

If we relate infill and young modulus, in the case of infill 45% for both geometry, Young Modulus

mean equals to 2.81 ± 0.41 GPa and for infill 50% value of the mean is 3.09 ± 0.55 GPa.

For example, to obtain young modulus of sample R-50/1 we use the values of the curve between

11 seconds and 42 seconds (figure nº35), as explained in the chapter before, we need to avoid

the mistake of consider a value of local plastic deformation below the porous structure's yield

strength.

Time 31 Seconds Time 27,2 Seconds

Force (N/mm2) 42 11

Deformation (mm) 0.06 0.05

dF (N/mm2) 31

dD (mm) 0,01

Young Modulus (GPa) 3.22E+03

Table nº11: Young modulus value and variables for sample R-50/1.

Figure nº35: Tension vs deformation graph for sample R-50/1.

1,11E+01

4,21E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

3,E-03 1,4E+00 5,4E+00 1,1E+01 1,8E+01 2,7E+01 3,9E+01 5,4E+01 7,1E+01

Ten

sio

n (

N/m

m²)

Displacement(%)

Tension vs Deformation R-50/1


45

Figure nº36: Young modulus assay´s results for every sample.

Figure nº37: Average of young modulus per geometry and infill.

3,303,00

3,22

3,82

2,34

2,712,48

3,00

0,00

0,50

1,00

1,50

2,00

2,50

3,00

3,50

4,00

4,50

R-45/1 R-45/3 R-50/1 R-50/3 FH-45/3 FH-45/4 FH-50/2 FH-50/3

You

ng

Mo

du

lus

(GP

a)

Samples

Young Modulus

3.417 3.509

2.521 2.726

0.000

0.500

1.000

1.500

2.000

2.500

3.000

3.500

4.000

4.500

You

ng

Mo

du

lus

(GP

a)

Young Modulus

R-45

R-50

FH-45

FH-50


46

4.3.2. Vickers Hardness (HV)

The results of the Vickers-hardness measures exposed that the higher hardness was for

rectilinear infill 45% with values close to 737 ± 232 HV. All the other samples of rectilinear infill

50% and both honeycombs showed similar values between 412 ± 128 and 465±139 HV.

The results of assays performed are shown on figure nº35.

Sample Vickers Hardness (HV)

R-45/6 737± 232

R-50/6 439 ± 102

FH-45/6 412 ± 128

FH-50/6 465 ± 139

Table nº12: Vickers hardness obtained.


47

5. DISCUSSION

Direct ink writing technique allowed the fabrication of the samples successfully, but is important

to mention that filling the syringe with the mix of titanium and hydrogel is a critical process

because presence of bubbles in the paste inside the syringe will lead on a non-continuous

filament on the printed sample. This is the reason why printing process must be strictly

supervised in order to be aware of this possible discontinuous segment on the sample because

this would affect the shape of the sample (due to non-continuous filament, some part of the

upper layers would not have support underneath, thus a deformation on the sample is going to

be visible immediately).

According to literature revision, dimensional changes use to be higher in the vertical axis (height

variation with respect to diameter variation) due to the action of gravity force (Vidal E, 2020). In

the case of this study, diameter presented a higher variation, this could be explained because

possibly an expansion was generated in the outer loop of the filament, and exaggerated the

measured diameter.

Is demonstrated that porosity is a key influence in bone ingrowth. Values obtained in the

porosity analysis, using micro-CT and MIP results, evidence that porosity of the samples

(between 56% and 63.6%) are in the range of bone porosity (30%-90%) and microporosity and

macroporosity values are consistent with literature revision and suitable for bone regeneration

when implanted in human body because allows body fluids circulation and cell migration and

ingrowth.

Surface analysis of the samples concerning roughness parameters, demonstrated that surface

characteristics of the samples are acceptable for a biomaterial in contact with hard tissue, since

it is a suitable roughness range for the osseointegration of titanium with bone.

Regarding the images obtained through SEM on figure nº29, number and location of contact

point of metal on each layer of the samples are important because that affects compressive

resistance and microporosity ass well. Higher number of filaments in contact means a better

distribution of the loading charges and higher infill means more material to support that load,

in consequence, higher infill demonstrated to show higher young modulus for the 50% infill in

comparison with 45% infill samples for each geometry.

Mechanical results showed that young modulus of all samples studied, was between the range

of cancellous bone (0.1 to 5 GPa), thus the material is appropriate to perform as scaffold for

human bone without risk of stress shielding.

Assays results of Vickers-Hardness showed that all the samples performed a remarkable higher

resistance to penetration than cp-Ti. This value could be caused because a presence of oxygen

during sintering process may interfere and affect the sample, in consequence affecting HV.


48

6. CONCLUSIONS

In this study, porous titanium structures printed in 3D through direct ink writing were obtained

with physicochemical and mechanic characteristics compatibles and suitable to act as implants

in human bodies.

• It was possible to successfully print in 3D titanium scaffold through direct ink writing and

post process them in order to eliminate binder and obtain samples with certain

physicochemical and mechanical properties to study them.

• Total porosity situated between 59% and 64%. Micropore size average was found

between 11,78μm and 23μm. The values presented are suitable for fluid transport and

cellular proliferation if used as bone implants on human body.

• Regarding mechanical behavior, elastic modulus situated between 2.3 and 3.8 GPa.

These results are appropriate for scaffolds in human bone.

• Results of microhardness (Vickers Hardness) demonstrated to be higher than

commercially pure titanium. Presence of oxygen during sintering process may interfere

and affect the sample hardness.


49

7. BIBLIOGRAPHY

Acero, J., Calderón, J., Salmeron, J., Verdaguer, J., & Consejo, C. (1999). The behaviour of titanium

as a biomaterial: microscopy study of plates and surrounding tissue in facial

osteosynthesis. (J. o.-m. Surgery, Ed.) pp. 117-123.

Aldani, C., & Dominguez, A. (2012, January). Titanium as a Biomaterial for Implants. In Recent

Advances in Arthroplasty. doi:10.5772/27413

ASTM-Standard. (2005). Standard specification for wrought titanium-6 aluminum-7 niobium alloy

for surgical implant applications.

Breine, U., Johansson, B., Roylance, B., Roeckert, P., Yoffey, J., & Branemarck, P. (1964).

Regeneration of bone marrow. A clinical and experimental study following removal of

bone marrow by curettage. 59, 146.

Cameron, H., Pilliar, R., & Macnab, I. (1976). The rate of bone ingrowth into porous metals. Journal

of Biomedical Materials Research. doi:DOI: 10.1002/jbm.820100210

Chen Q, T. G. (2015). Metalic Implant Biomaterials. Materials Science & Engineering, pp. 1-57. doi:

https://doi.org/10.1016/j.mser.2014.10.001

Costa, F., Carvalho, I., Montelaro, R., & Gomes, P. (2011). Covalent immobilization of antimicrobial

peptides (AMPs) onto biomaterial surfaces. Acta Biomater, 7, 1431-1440.

Costerton, J., Stewart, P., & Greenberg, E. (1999). Bacterial biofilms: a common cause of persistent

infections. Science, 284, 1318-1322.

Davies, J. (2003). Understanding peri-implant endosseous healing. Journal of Dental Education, 67,

932-949.

Diana Lopes, C. M.-C. (2018). Bone Physiology as Inspiration for Tissue Regenerative Therapies.

Biomaterials. doi:doi: https://10.1016/j.biomaterials.2018.09.028

Duda, T., & Venkat, L. (2016). 3D Metal Printing Technologies. IFAC- Papers On Line, 103-110.

Elias, C., Jogaib, D., Moura de Souza, F., Dos Santos, E., & de Biasi, S. (2019). Mechanical and clinical

properties of titanium and titanium-based alloys (Ti G2, Ti G4 cold worked nanostructured

and Ti G5) for biomedical applications. Journal of Materials Research and Technology, 8,

1060-1069.

Giannouidis, P., Einhorn, T., & Marsh, D. (2007). Fracture healing: The diamond concept. Injury, 38,

S3-S6.

Hallab, N., Urban, R., & Jacobs, J. (2004). Corrosion and Biocompatibility of Orthopedic Implants.

Biomaterials in Orthopedics, 63-92. doi:DOI:10.1201/b14227-4


50

Her-Hsiung, H., Chun-Te, H., Tzu-Hsin, L., Tien-Ling, L., Ko-Kaung, L., & Fang-Lung, C. (2004). Effect

of surface roughness of ground titanium on initial cell adhesion. Biomolecular Engineering,

93-97.

ISO13314. (2011). Mechanical testing of metals — Ductility testing — Compression test for porous

and cellular metals.

Keles, O., Blevins, C., & Bowman, K. (n.d.). “Effect of build orientation on the mechanical reliability

of 3D printed ABS. Rapid Prototyping Journal, 320-328.

Kim, T., Wang See, C., Xiaochun, L., & Donghui, Z. (2020). Orthopedic implants and devices for bone

fractures and defects: Past, present and perspective. Enginereed Regeneration, 1, 6-18.

Kurtz, S., Ong, K., Jau, E., Mowart, F., & Halpern, M. (2007). Projections of primary and revision hip

and knee arthroplasty in the United States from 2005 to 2030. Journal of Bone and Joint

Surgery – American, 89, 780-785.

Kuzyk, P., & Schemitsch, E. (2011). The basic science of peri-implant bone healing. Indian Journal

Orthopedics, 45, 108-115.

Lei Zhang, G. Y. (2018). Three-dimensional (3D) Printed Scaffold and Material Selection for Bone

Repair. Acta Biomaterialia. doi:https://doi.org/10.1016/j.actbio.2018.11.039

Leonardo, G. (2013). Trabajo de investigacion bibliográfica: Estructura y propiedades de fatiga del

titanio y sus aleaciones utilizadas como biomateriales (Vol. 17). (C. y. Universidad, Ed.)

Lewis, G. (2013). Properties of open-cell porous metals and alloys for orthopaedic applications.

Journal of Materials Science: Materials in Medicine. doi:https://doi.org/10.1007/s10856-

013-4998-y

Li N et al. (2019). Progress of additive manufacturing technology for porous metal implant

applications and triple minimal surface structures: A review. J Mater Sci Technology.

Liu, X., Chu, P., & Ding, C. (2004). Surface modification of titanium, titanium alloys, and related

materials for biomedical applications. Mater Science Engineering, 47, 49-121.

Magda Dziaduszewska, Andrzej Zielinski. (2019). Titanium Scaffolds – Hopes and Limitations.

American Journal of Biomedical Science & Research. doi:DOI:

10.34297/AJBSR.2019.04.000841

Mitra, J., Tripathi, G., Sharma, A., & Basu, B. (2013). Scaffolds for bone tissue engineering: role of

surface pattering on osteoblast response. RSC Advances.

Muller, K., & Karevska, S. (2016). How will 3D printing make your company the strongest link in the

value chain? Retrieved from http://www.ey.com/Publication/vwLUAssets/EY-3d-druck-

studie-executive-summary/$FILE/EY-3d-druck-studie-executive-summary.pdf

Navarro, M., Michiardi, A., Castaño, O., & Planell, J. (2008). Biomaterials in orthopaedics. doi:DOI:

10.1098/rsif.2008.0151


51

Ni, J., Ling, H., Zhang, S., Wang, Z., Peng, Z., Benyshek, C., . . . Khademhosseini, A. (2019). Three-

dimensional printing of metals for biomedical applications. Materials Today Bio, 3.

doi:https://doi.org/10.1016/j.mtbio.2019.100024

Niinomi, N., & Nakai, M. (2011). Titanium-based biomaterials for preventing stress shielding

between implant devices and bone. International Journal of Biomaterials, 10.

Ottria, L., Lauritano, D., Andreasi Bassi, M., Palmiera, A., Candotto, V., Tagliabue, A., & Tettamanti,

L. (2018). Mechanical, chemical and biological aspects of Titanium. Journal of Biological

Regulators and Homeostatic Agents, pp. 81-90.

Özbilen, S., Liebert, D., Beck, T., & Bram, M. (2016). Fatigue behavior of highly porous titanium

produced by powder metallurgy with temporary space holders. Material Science

Engineering, 60, pp. 446-457.

Pala, J., Barao, V., Aparecida, J., Pérez, L., Pereira, T., & Goncalves, W. (2013). Titanium in Dentistry:

Historical Development, State of the Art. Indian Prosthodont Society. doi:DOI

10.1007/s13191-012-0190-1

Pattanayak, D., Fukuda, A., Matsushita, M., Takemoto, M., Fujibayashi, S., Sasaki, K., . . . Kokubo, T.

(2011). Bioactive Ti metal analogous to human cancellous bone: fabrication by selective

laser melting and chemical treatments. Acta Biomaterials, 1398-1406.

Peawband, A., & Thedsakhulwong, A. (2017). Design and Construction of a Small Vacuum Furnace.

Journal of Physics: Conference Series. IOP Publishing.

Ratner, B., Hoffman, A., Schoen, F., & Lemons, J. (2013). Biomaterials Science-An Introduction to

Materials in Medicine. (Oxford, Ed.) U.K.: Academic Press.

Reddy, P. (2019). Digital Light Processing (DLP). Think3D.

Ribeiro, J., Barao, V., Aparecida, J., & Perez, L. (2013). Titanium in Dentistry: Historical

Development, State of the Art and Future Perspectives. The Journal of Indian Prosthodontic

Society. doi:DOI:10.1007/s13191-012-0190-1

Ribeiro, M., Montero, F., & Ferraz, M. (2012). Infection of orthopedic implants with emphasis on

bacterial adhesion process and techniques used in studying bacterial-material interactions.

Biomatter 2(4):176-194.

Saenz, V., & Fuentes, E. (2013). Titanium and Titanium Alloys as Biomaterials. IntechOpen. doi:DOI:

10.5772/55860

Sajad, A., Burnett, J., & Michael, T. (2016). Fully porous 3D printed titanium femoral stem to reduce

stress-shielding following total hip arthroplasty. Journal of Medical Virology.

Saurí, J., Suñé-Negre, J., Díaz-Marcos, J., Vilana, J., Millán, D., Ticó, J., . . . García- Montoya, E.

(2015). Relationships between surface free energy, surface texture parameters and

controlled drug release in hydrophilic matrices. International Journal of Pharmaceutics,

328-340.


52

Schmidutz, F., Agarwal, Y., Müller, P., & Gueorgui, B. (2014). Stress-shielding induced bone

remodeling in cementless shoulder resurfacing arthroplasty: a finite element analysis and

in vivo results. Journal Biomechanics, 3509-3516.

Shahrubudin, N., Lee, T., & Ramlan, R. (2019). An Overview on 3D Printing Technology:

Technological, Materials and Applications. 2nd International Conference on Sustainable

Materials Processing and Manufacturing.

Silbernagel, C. (2018). What is material jetting? Retrieved from Canada Makers: Avaliable:

http://canadamakes.ca/what-is-material-jetting/ [Accesed 2021]

Solis, N., Smirnov, A., Peretyagin, N., Seleznev, A., & Peretyagin, A. (2020). Direct Ink Writing

Technology (3D Printing) of Graphene-Based Ceramic Nanocomposites: A Review.

Standard, A. (2005). Standard specification for wrought titanium-6 aluminum-7 niobium alloy for

surgical implant applications (Vol. F1295). West Conshohocken.

Stansbury, W., & Idacavage, M. (2016). 3D Printing with polymers: Challenges among expanding

options and opportunities. Dental Materials, 54-64.

Syed, A., Elias, P., Amit, B., Susmita, B., Lisa, O., & Charitidis, C. (2017). Additive manufacturing:

scientific and technological challenges, market uptake and opportunities. Materials Today,

1, 1-16.

Takao, H. (2019). Titanium–Tissue Interface Reaction and Its Control With Surface Treatment.

Tokyo: Department of Metallic Biomaterials, Institute of Biomaterials and Bioengineering.

Tiwari, S., Pande, S., Agrawal, S., & Bobade, S. (2015). Selection of selective laser sintering materials

for different applications. Rapid Prototyping Journal, 21, 630-648.

Tuang, D., Kashani, A., Imbalzano, G., Nguyen, K., & Hui, D. (2017). Additive manufacturing (3D

printing): A review of materials, methods, applications and challenges. Composites.

Vaezi, M., Seitz, H., & Yang, S. (2013). A review on 3D micro-additive manufacturing technologies.

The International Journal of Advanced Manufacturing Technology, 67.

Vidal E, T. D.-M.-P. (2020). Titanium Scaffolds by Direct Ink Writing: Fabrication and

Functionalization to Guide Osteoblast Behavior. Metals.

doi:https://doi.org/10.3390/met10091156

Vikayavenkataraman, S., Jerry, Y., & Wen, F. (2017). 3D Printing and 3D Bioprinting in Pediatrics.

Bioengineering, 63, 1-11. doi:10.3390/bioengineering4030063

Vladimir Popov Jr., V., Muller-Kamskii, G., Kovalevsky, A., Dzhenzhera, G., Strokin, E., Kolomiets, A.,

& Ramon, J. (2018). Design and 3D-printing of titanium bone implants: brief review of

approach and clinical cases. Biomedical Engineering Letters.

Wang. (2017). Analysis of factor influencing bone ingrowth into three-dimensional printed porous

metal scaffolds: A review (Vol. 717). J Alloys Compd.


53

Wang, G., Li, J., Kaige, L., Wenjie, Z., Xun, D., Guanghzeng, Y., . . . Xinquan, J. (2016). Surface

thermal oxidation on titanium implants to enhance osteogenic activity and in vivo

osseointegration. Sci Rep, 6. doi:https://doi.org/10.1038/srep31769

Wang, M., & Tang, T. (2019). Surface treatment strategies to combat implant-related infection from

the beginning. Journal of Orthopedics Translation, 17, 42-54.

Wu, S., Liu, X., Yeung, K., Liu, C., & Yang, X. (2014). Biomimetic porous scaffolds for bone tissue.

Materials Science Engineering, 80(1), 1-36.

Yap, C., Chua, C., Dong, Z., Liu, Z., Zhang, D., Loh, L., & Sing, S. (2015). Review of selective laser

melting: Materials and applications. Applied Physics Reviews.

Ze-Xian, L., Yen, T., Ray, M., Mattia, D., Metcalfe, I., & Patterson, D. (2016). Perspective on 3D

printing of separation membranes and comparison to related unconventional fabrication

techniques. Journal of Membrane Science, 523, 596-613.

Zhao S, L. S. (2016). The influence of cell morphology on the compressive fatigue behavior of Ti-6Al-

4V meshes fabricated by electron beam melting (Vol. 59). J Mech Behav Biomed Mater.

Zhu, X., Chen, J., Scheideler, L., Reichl, R., & Geis-Gerstorfer, J. (2004). Effects of topography and

composition of titanium surface oxides on osteoblast responses. Biomaterials, 25, 4087-

4103.


54

8. ANNEXES



6,99E+00

5,19E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

2,7E-03 5,4E-02 7,7E-02 1,0E-01 1,3E-01 1,8E-01 2,3E-01 2,8E-01 3,3E-01

Ten

sio

n (

N/m

m²)

Deformation (%)


1,35E+01

5,25E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

1,40E+02

8,E-04 2,7E-02 5,3E-02 5,8E-02 8,4E-02 1,1E-01 1,4E-01 1,6E-01 1,9E-01

Ten

sio

n (

N/m

m²)

Deformation (%)



55

Figure nº38: Tension vs deformation graph for sample R-50/1

.


1,11E+01

4,21E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

3,E-03 1,4E+00 5,4E+00 1,1E+01 1,8E+01 2,7E+01 3,9E+01 5,4E+01 7,1E+01

Ten

sio

n (

N/m

m²)

Displacement(%)


1,43E+01

5,52E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

1,40E+02

1,60E+02

1,80E+02

2,4E-01 3,6E-03 -7,5E-02 2,5E-04 1,8E-01 4,4E-01 8,4E-01 1,4E+00 2,1E+00 3,E+00

Ten

sio

n (

N/m

m²)

Deformation(%)



56

Figure nº40: Tension vs deformation graph for sample FH-45/3.


1,14E+01

4,24E+01

0,00E+00

1,00E+01

2,00E+01

3,00E+01

4,00E+01

5,00E+01

6,00E+01

7,00E+01

8,00E+01

1,41E-04 8,90E-02 3,38E-01 7,48E-01 1,23E+00 1,79E+00 2,51E+00 3,39E+00

Ten

sio

n (

N/m

m²)

Deformation (%)

Tension vs Deformation FH-45/3

1,2E+01

3,9E+01

0,0E+00

1,0E+01

2,0E+01

3,0E+01

4,0E+01

5,0E+01

6,0E+01

7,0E+01

8,0E+01

9,0E+01

-1,1E-03 2,9E-02 2,3E-01 5,9E-01 1,1E+00 1,6E+00 2,4E+00 3,3E+00 4,3E+00 5,6E+00

Ten

sio

n (

N/m

m²)

Deformation (%)



57



1,03E+01

4,61E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

1,40E+02

1,60E+02

-2,6E-03 7,4E-02 7,3E-01 1,8E+00 3,5E+00 5,8E+00 8,8E+00 1,2E+01

Ten

sio

n(N

/mm

²)

Deformation (%)


1,01E+01

4,81E+01

0,00E+00

2,00E+01

4,00E+01

6,00E+01

8,00E+01

1,00E+02

1,20E+02

8,4E-05 1,9E-01 7,6E-01 1,5E+00 2,6E+00 4,E+00 5,8E+00 8,E+00

Ten

sio

n (

N/m

m²)

Deformation (%)



58

Vickers Hardness

Geometry Diameter

1 (d1) Diameter

2 (d2) d1

[um] d2

[um] d1

[mm] d2

[mm] D

[mm] HV Gpa

Rectilinear 45% Infill

112 113 28 28,25 0,028 0,028 0,028 469 4,594

70 77 17,5 19,25 0,018 0,019 0,018 1098 10,762

78 80 19,5 20 0,020 0,020 0,020 951 9,316

103 104 25,75 26 0,026 0,026 0,026 554 5,428

84 85 21 21,25 0,021 0,021 0,021 831 8,143

83 86 20,75 21,5 0,021 0,022 0,021 831 8,143

114 113 28,5 28,25 0,029 0,028 0,028 461 4,513

90 94 22,5 23,5 0,023 0,024 0,023 701 6,869

Rectilinear 50% Infill

111 115 27,75 28,75 0,028 0,029 0,028 465 4,553

103 98 25,75 24,5 0,026 0,025 0,025 587 5,756

121 119 30,25 29,75 0,030 0,030 0,030 412 4,038

103 115 25,75 28,75 0,026 0,029 0,027 499 4,894

130 123 32,5 30,75 0,033 0,031 0,032 371 3,633

109 110 27,25 27,5 0,027 0,028 0,027 495 4,849

151 160 37,75 40 0,038 0,040 0,039 245 2,405

109 123 27,25 30,75 0,027 0,031 0,029 441 4,321

Honeycomb 45% Infill

141 151 35,25 37,75 0,035 0,038 0,037 278 2,728

98 95 24,5 23,75 0,025 0,024 0,024 637 6,244

106 123 26,5 30,75 0,027 0,031 0,029 453 4,435

132 139 33 34,75 0,033 0,035 0,034 323 3,167

118 113 29,5 28,25 0,030 0,028 0,029 445 4,358

128 139 32 34,75 0,032 0,035 0,033 333 3,262

143 142 35,75 35,5 0,036 0,036 0,036 292 2,863

107 104 26,75 26 0,027 0,026 0,026 533 5,224

Honeycomb 50% Infill

113 121 28,25 30,25 0,028 0,030 0,029 433 4,247

114 119 28,5 29,75 0,029 0,030 0,029 437 4,284

97 96 24,25 24 0,024 0,024 0,024 637 6,244

137 136 34,25 34 0,034 0,034 0,034 318 3,120

135 140 33,75 35 0,034 0,035 0,034 314 3,075

130 126 32,5 31,5 0,033 0,032 0,032 362 3,549

98 111 24,5 27,75 0,025 0,028 0,026 543 5,324

93 95 23,25 23,75 0,023 0,024 0,024 671 6,580

Table nº13: Vickers-Hardness results for each sample.


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