Prozess- und Produktentwicklung von funktionellen Suprapartikeln für die biomedizinische Additive Fertigung

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Process- and Product Development of Functional Supraparticles for

Biomedical Additive Manufacturing

Prozess- und Produktentwicklung von funktionellen Suprapartikeln für

die biomedizinische Additive Fertigung

Der Technischen Fakultät der

Friedrich- Alexander-Universität Erlangen-Nürnberg zur Erlangung des Doktorgrades

D O K T O R - I N G E N I E U R

vorgelegt von

Herbert Wilhelm Canziani

aus Pegnitz


Als Dissertation genehmigt von der Technischen Fakultät der

Friedrich-Alexander-Universität Erlangen-Nürnberg

Tag der mündlichen Prüfung: 03.08.2021

Vorsitzender des Promotionsorgans: Prof. Dr.-Ing. Knut Graichen

Gutachter: Prof. Dr. Nicolas Vogel

Prof. Dr. Karl Mandel


Man muss einfach reden, aber kompliziert denken.

Nicht umgekehrt.

- Franz Josef Strauß



This PhD thesis was carried out from March 2018 to April 2021 at the Institute of Particle Tech- nology (LFG) at the Friedrich-Alexander-Universität Erlangen-Nürnberg and the University of To- ronto under the supervision of Professor Nicolas Vogel. A PhD thesis is like a marathon, the road is long and not always even. You have to work hard, learn new things, overcome obstacles, and always keep your eyes on the goal. It takes a lot of perseverance, motivation, and support to reach the finishing line. At this point, I would like to thank all the people, who contributed to this thesis in one way or another.

My deepest gratitude goes to my girlfriend Nefi and my family without whose tremendous support and encouragement, I would not have been able to write this thesis. You supported me during my studies. You always had encouraging words and an open ear for me. You always brought me back into daily life, away from all measurements, calculations, and supraparticles. With your help and support, I managed to overcome even the most stressful times.

Importantly, I would also like to express my sincere gratitude to my supervisor Professor Nicolas Vogel for his excellent supervision and for the opportunity to work on this future-oriented topic.

Thank you, Niki, for your trust and for the freedom you provided me to tailor my PhD topic. During our collaboration, I greatly appreciated your approachability, your kindness, your willingness to discuss, and your ability to listen actively, which distinguished you very positively from other pro- fessors.

Furthermore, I would like to thank Professor Karl Mandel for kindly taking on the role of the sec- ond examiner for this PhD thesis. Moreover, I would like to express my sincere gratitude to Mr.

Patrik Stenner and Dr. Heike Mühlenweg for their great support, constructive discussions, fruitful suggestions, and their constant interest in the progress of the thesis.

A big thanks goes to Professor Benjamin Hatton from the University of Toronto. Thank you, Ben, for enabling me to work in your lab and thus expanding my topic in a valuable way. I have very pleasant memories of my time in Toronto in 2020, apart from the beginning of the Corona pan- demic. Many thanks to Cameron Stewart and Brian De La Franier for the warm welcome and great support during my stay.

I would also like to thank my diligent and motivated miniproject, bachelor and master students Florentin Riedel, Frederik Bever, Moritz Fröschel, Robert Becker, Krupansh Desai and Benedikt Hanschmann for their strong support. I would like to especially mention Frederik Bever, Florentin Riedel, and Benedikt Hanschmann for their very valuable contributions. It was great fun to work with you in the lab and to accompany you on your way.


Furthermore, I want to express my gratitude to my colleagues Alexander Sommereyns (Institute of Photonic Technologies) and Thomas Distler (Institute Biomaterials) for their tremendous sup- port in the powder bed fusion and biocompatibility experiments. I highly enjoyed our enthusiastic discussions and I have benefited greatly from your extensive explanations.

At last, I would like to thank all the staff members of the Institute of Particle Technology (LFG) for the great working atmosphere. In particular, I would like to thank Nicolas Hesse, Björn Düsenberg and Laura Unger for the numerous lunch breaks with “Leberkäsbrötchen” and the very entertain- ing, scientific coffee breaks. My gratitude also goes to Salvatore Chiera, who supported me in the early days of my PhD project and was open to all my questions in the lab. Furthermore, I would like to thank Dr. Jochen Schmidt, Michael Auer, and Bettina Winzer for their help with scientific matters. Moreover, I would like to thank our always cheerful team from the secretary's office, Paula Hoppe and Julia Seekatz, for their competent help with organizational matters. A very big thank you goes to our team from the workshop for their quick and solution-oriented help in case of problems with the spray dryer.

Ein herzliches Vergelt’s Gott an alle Beteiligten

Erlangen, April 2021 Herbert Canziani



In the last decades, additive manufacturing -commonly referred to as 3D printing- has gained growing attention driven by increasing digitalization in industry and the biomedical sector. The expression additive manufacturing is used to describe a variety of manufacturing technologies that produce construction parts in a layer-by-layer approach based on CAD files. In comparison to traditional subtractive manufacturing processes, additive manufacturing allows the cheap pro- duction of highly complex, individualized specimens in small product series in a short time. These advantages are particularly attractive for the biomedical sector to produce surgical models and patient-specific implants.

In this context, the powder bed fusion laser beam melting process of polymers (PBF-LB/P) is very capable to produce geometrically complex and porous specimens for their use as bone implants.

However, the choice of suitable powder materials is very limited up to now. The available materi- als can be divided into two categories of unfilled polymeric powders, which are almost exclusively based on polyamide 12, and a few filled composite powders. These composite powders are phys- ical mixtures on the micrometer scale of polyamide 12 and additives such as carbon fibers or glass beads. This limited material selection results from the high material requirements imposed by the powder bed fusion process (PBF-LB/P). A suitable powder material needs to provide a wide ther- mal process window, a well-developed adsorption behavior in the wavelength of the CO2 laser, and high powder flowability.

In this PhD thesis, a colloid-based bottom-up process chain is presented to produce tailored su- praparticles for the powder fusion process (PBF-LB/P). In a first step, polymeric and additive pri- mary particle dispersions are prepared. These dispersions are then spray dried either as a pure polymeric dispersion to produce polymer supraparticles or as a dispersion mixture to produce composite supraparticles. The dispersion droplet serves as confinement for the self-assembly of the primary particles. The final supraparticle design can be precisely adjusted via the spray drying process conditions and the primary particle dispersions.

In the first part of the thesis, a system of polymethyl methacrylate (PMMA) and silica (SiO2) is investigated with a special focus on supraparticle formation. The produced supraparticles could be of interest for dental applications. The first aim was to identify suitable spray drying process parameters to obtain spherical supraparticles with good flowability. Then, the powder flowability of the supraparticles is further improved by adjusting the supraparticle roughness.


Subsequently, the structure formation of PMMA-SiO2 composite supraparticles is studied, based on different dispersion mass mixing ratios and primary particle diameter ratios. Furthermore, the drug release from PMMA composite supraparticles is investigated, comprising mesoporous drug- loaded silica (MSiO2) primary particles. Finally, the tailored PMMA and PMMA-SiO2 supraparticles with optimized product properties are applied in the powder bed fusion process (PBF-LB/P).

In the second part of the thesis, a system consisting of polylactide (PLA) and calcium-containing inorganic primary particles is investigated for additively manufactured bone implants. In a first step, PLA primary particles are synthesized via the miniemulsion solvent evaporation process, while binary calcium-silica (Ca-SiO2) and nanohydroxyapatite (HAP) primary particles are pro- duced via sol-gel processes. Subsequently, these colloidal dispersions are spray dried to form tai- lored supraparticles. The thermal properties and the flowability of the powder material are char- acterized in detail. Additionally, biocompatibility and bioactivity are tested. Furthermore, the me- chanical properties of composite specimens (resulting from the supraparticles) are tested and evaluated towards their use as bone replacement materials. Finally, the produced PLA supraparti- cle powders are used in the powder bed fusion process (PBF-LB/P) to fabricate multilayered square specimens, which are later tested towards their biodegradability in biologically relevant media.

In the last chapter of the thesis, the bottom-up process chain is extended to another interesting biodegradable polymer, polycaprolactone (PCL), which also finds applications in bone replace- ment. To this end, the thermal properties and the powder flowability are investigated and finally, the powders are processed in the powder bed fusion process (PBF-LB/P). Moreover, all polymeric composite supraparticles of PCL and PLA are produced and investigated towards their thermal and mechanical properties as a fully biodegradable bone replacement material.



In den letzten Jahrzehnten gewann die Additive Fertigung, oft als 3D Druck bezeichnet, zuneh- mend mehr Aufmerksamkeit angetrieben durch die steigende Digitalisierung in der Industrie und im biomedizinischen Sektor. Unter dem Begriff additive Fertigung versteht man eine Vielzahl von Fertigungstechnologien, welche Schicht für Schicht ein gewünschtes Bauteil basierend auf einem CAD Modell erstellen. Im Vergleich zu traditionellen subtraktiven Fertigungsverfahren können durch die Additive Fertigung hochkomplexe, individualisierte Geometrien, preisgünstig in kleinen Produktserien in kurzer Zeit hergestellt werden. Diese Vorteile der Additiven Fertigung sind be- sonders attraktiv für den biomedizinischen Sektor zur Herstellung von chirurgischen Modellen und patientenspezifischen Implantaten.

Der Pulverbett Laserstrahlschmelzprozess (PBF-LB/P) eignet sich dabei sehr gut zur Herstellung von geometrisch komplexen und porösen Bauteilen für Knochenimplantate aus polymerischen Materialien. Allerdings ist die Materialauswahl zum derzeitigen Zeitpunkt sehr eingeschränkt. Die verfügbaren Pulver unterteilen sich in ungefüllte Polymer Pulver, welche fast ausschließlich auf Polyamid 12 basieren, und gefüllte Kompositpulver, die wiederum physikalische Mischungen auf der Mikrometerskala von Polyamid 12 und Additiven wie Kohlefasern und Glaskugeln darstellen.

Diese eingeschränkte Materialauswahl ergibt sich aus den hohen Materialanforderungen durch den Pulverbett Laserstrahlschmelzprozess. Das Pulvermaterial muss neben einer hohe Pulver- fließfähigkeit, ein breites thermisches Prozessfenster und ein gut ausgeprägtes Adsorptionsver- halten im Bereich der CO2 Laserwellenläge aufweisen.

In der vorliegenden Doktorarbeit wird eine kolloidbasierte bottom-up Prozesskette zur Herstel- lung von maßgeschneiderten Suprapartikeln für das Pulverbett Laserstrahlschmelzen (PBF- LB/P) untersucht. Dabei werden in einem ersten Schritt polymerische und additive Primärparti- keldispersionen hergestellt. Anschließend werden diese entweder als reine polymerische Disper- sion zur Herstellung von Polymer Suprapartikeln oder als eine Dispersionsmischung zur Herstel- lung von Komposit Suprapartikeln in einem Sprühtrocknungsprozess ausgesprüht. Der Dispersi- onstropfen dient dabei als Konfinement zur Steuerung der Selbstanordnung der Primärpartikel.

Das Suprapartikel Produktdesign lässt sich dabei über die Prozessbedingungen im Sprühtrock- nungsprozess und die Primärpartikeldispersionen präzise einstellen.

Im ersten Teil der Doktorarbeit wird ein für dentale Anwendungen relevantes Materialsystem be- stehend aus Polymethylmethacrylat (PMMA) und Silica (SiO2) betrachtet; mit besonderem Fokus auf die Strukturbildung. Dabei werden zuerst geeignete Sprühtrocknungsprozessparameter erar- beitet, welche zu sphärischen und gut fließfähigen Suprapartikeln führen.


Danach wird die Pulver Fließfähigkeit dieser Suprapartikel durch Einstellung der Suprapartikel- rauigkeit weiter verbessert. Im Anschluss daran wird die Strukturbildung von PMMA-SiO2 Kom- posit Suprapartikeln basierend auf verschiedenen Dispersion Massenmischungsverhältnissen und Primärpartikel Durchmesserverhältnissen untersucht. Ferner wird die Wirkstofffreisetzung aus PMMA Komposit Suprapartikeln mit mesoporösen wirkstoffbeladenen Silica (MSiO2) Primär- partikeln beleuchtet. Zum Abschluss werden die maßgeschneiderten PMMA und PMMA-SiO2 Sup- rapartikel mit optimierten Produkteigenschaften im Pulverbett Laserstrahlschmelzprozess (PBF- LB/P) zu mehrschichtigen Bauteilen verarbeitet.

Im zweiten Teil der Doktorarbeit wird ein Materialsystem basierend auf Polylactid (PLA) und Cal- cium beladenen anorganischen Primärpartikeln für additiv gefertigte Knochenimplantate unter- sucht. Hierbei werden zuerst PLA Primärpartikel über das Miniemulsions Lösemittelverdamp- fungsverfahren und binäre Calcium-Silica (Ca-SiO2), sowie Nanohydroxyapatit (HAP) Primärpar- tikel über Sol-Gel Prozesse hergestellt werden und diese anschließend zu maßgeschneiderten Suprapartikeln versprüht werden. Das resultierende Suprapartikel Pulvermaterial wird, dabei auf seinen thermischen Eigenschaften und seine Pulverfließfähigkeit untersucht. Besonderer Fokus wird zudem auf dessen Biokompatibilität und Bioaktivität gelegt. Dazu wird auch bestimmt, ob Kompositmaterialien basierend auf den Suprapartikeln eine ausreichende mechanische Festig- keit aufweisen für deren Anwendung als Knochenersatzmaterial und ferner eine Biodegradier- barkeit in biologisch relevanten Medien zeigen. Abschließend werden die produzierten PLA Sup- rapartikel Pulver im Pulverbett Laserstrahlprozess (PBF-LB/P) verwendet werden, um mehrla- gige Baukörper herzustellen.

Im letzten Kapitel der Doktorarbeit wird die bottom-up Prozesskette auf Polycaprolacton (PCL), ein weiteres interessantes biodegradierbares Polymer für Knochenersatz, ausgeweitet. Hierbei werden zunächst die thermischen Eigenschaften und die Pulverfließeigenschaften untersucht werden. Danach werden die PCL Suprapartikelpulver im Pulverbett Laserstrahlschmelzprozess (PBF-LB/P) verwendet. Abschließend werden vollständig polymerische Komposit-Suprapartikel aus PCL und PLA hergestellt. Diese werden dabei auf ihre thermischen und mechanischen Eigen- schaften als vollständig biologisch abbaubares Knochenersatzmaterial untersucht


Table of Contents

1 Motivation and Objective 12

2 Theoretical Background 14

2.1 Additive Manufacturing in Biomedical Field 14

2.2 Biomaterials for Bone Grafting 17

2.3 Fundamentals on Colloids and Emulsions 22

2.4 Spray Drying of Nanoparticle Dispersions 26

2.5 Factors Influencing Powder Flowability 32

3 Experimental Methods and Materials 35

3.1 Synthesis of Primary Particles 35

3.1.1 Surfactant Free Emulsion Polymerization 35

3.1.2 Stöber Process and Binary Ca-SiO2 Nanoparticles 36 3.1.3 Synthesis of Mesoporous OCT Loaded SiO2 Nanoparticles 37

3.1.4 Miniemulsion Solvent Evaporation Technique 38

3.1.5 Precipitation of Hydroxyapatite Nanoparticles 39

3.2 Production of Supraparticles via Spray Drying 39

3.3 Post Processing of Supraparticles 40

3.4 Fabrication of Square Specimens via Powder Bed Fusion Process 40

3.5 Characterization Methods 41

3.5.1 Determination of Particle Size Distribution 41

3.5.2 Scanning Electron Microscopy & Energy Dispersive X-Ray Spectroscopy 42

3.5.3 Pore Size Analysis via Mercury Porosimetry 44

3.5.4 Characterization of Powder Flowability 45

3.5.5 Thermal Characterization of Supraparticles 46

3.5.6 Determination of Molecular Weight Distribution via GPC 48 3.5.7 Quantification of Octenidine Drug Release via UV-vis Spectroscopy 49 3.5.8 Quantification of Calcium Release via ICP-OES 49

3.5.9 Mechanical Testing of Cylindrical Specimens 50

3.5.10 Minimum Inhibitory Concentration of OCT for Streptococcus mutans 50 3.5.11 Biocompatibility Testing with Human Osteoblast-Like MG-63 Cells 51 3.5.12 Investigation of Biodegradation for PLA Specimens 52 3.5.13 Investigation of Hydroxyapatite Formation in Simulated Body Fluid 53


4 Results and Discussion 54 4.1 Bottom-up Approach for Nanoscale Composite Supraparticles 54 4.2 PMMA & PMMA-SiO2 Supraparticles for Dental Applications 56 4.2.1 Process Parameters Influencing Supraparticles Formation 57 4.2.2 Tailoring Supraparticle Roughness to Control Powder Flowability 63 4.2.3 Structure Formation of PMMA-SiO2 Composite Supraparticles 68 4.2.4 Powder Bed Fusion of PMMA and PMMA-SiO2 Supraparticles 78 4.2.5 Drug Release from OCT Loaded SiO2-PMMA Composite Supraparticles 83

4.3 Polylactide (PLA) and PLA-SiO2 Supraparticles for Bone Grafting 86 4.3.1 PLA Primary Particles via Miniemulsion Solvent Evaporation Technique 88 4.3.2 Binary Ca-SiO2 and Hydroxyapatite Primary Particles 92 4.3.3 Poly(L-lactide) (PLA) Supraparticles Containing Different Emulsifiers 95 4.3.4 PLA-SiO2 Composite Supraparticles for Improved Mechanical Strength 101 4.3.5 Biocompatibility of PLA and PLA-CaSiO2 Supraparticles 109 4.3.6 Fabrication of Square Specimens via Powder Bed Fusion Process 112 4.3.7 Degradation Behavior and Hydroxyapatite Formation 119

4.4 Process Transfer to Polycaprolactone (PCL) Supraparticles 122 4.4.1 Synthesis of Primary Particles and Formation of PCL Supraparticles 123

4.4.2 Polymeric PCL-PLA Composite Supraparticles 127

5 Conclusion and Outlook 132

6 List of Abbreviations 135

7 Curriculum Vitae 139

8 Appendix 140

8.1 Publications and Patents 140

8.2 Supervised Student Thesis 140

8.3 Supplementary Data 141

9 References 142


1 Motivation and Objective

Bottom-up design and self-assembly are the oldest principles to produce functional materials and complex organisms in nature.1,2 Life on earth is based predominately on the elements hydrogen (H), oxygen (O), carbon (C), nitrogen (N), and phosphorus (P).3,4 These elements “assemble”

through different interactions into more complex molecules like polysaccharides, lipids, nucleic acids, and proteins, the fundamental building blocks of life.5,6 In a higher level of complexity, cells like erythrocytes, neurons, and osteoblasts are obtained through the arrangement and interaction of those molecules.7 At last, complex multicellular organisms with different tissues and organs carrying out specialized tasks arise through the organization of those different cells.8,9

The human body represents an example of such a multicellular organism with various specialized organs and complex tissues.9–11 In the lifetime of a human being, accidents or diseases can lead to partial damage or complete failure of these organs and associated tissues, resulting in the partial or full loss of important physiological functions.10,12,13 Implants are often used to replace damaged tissues or organs and restore their function.11,13,14 However, traditional subtractive manufacturing methods have many tool-related design limitations when producing complex geometries for the tissue and organ implants.15–17

A newly emerging field that could provide complex structures and intricate geometrical shapes is additive manufacturing.15,16 The expression additive manufacturing -commonly known as 3D printing- designates a layer-by-layer approach, where products can be obtained from a computer- aided design without the use of any tools.18,19 The powder bed fusion process for polymers (PBF- LB/P) is very promising for the biomedical sector.15,20 Among other additive manufacturing pro- cesses it holds the promise of high design flexibility, short production times, a high degree of cus- tomization, and cheap production of small product series.15,16 This generative manufacturing ap- proach can be used for the fabrication of surgical models, personalized prosthetics and implants, as well as personalized drug delivery systems for the pharmaceutical industry.15,21,22

In particular, the fabrication of patient-specific bone implants from computer tomographic or magnetic resonance imaging combined with biodegradable polymer composite materials pro- vides a sustainable and progressive manufacturing approach.15,23 Despite all the advantages of the powder bed fusion process of polymers (PBF-LB/P), the material selection is of today still quite limited.15,24 Polyamide powders make up 90 % of the market share and only a few composite pow- der materials are available since the material requirements for the PBF-LB-P process are quite high.15,16,25


It is obvious that the currently available commercial powder materials cannot fully meet the re- quirements for biomedical applications.15,20 On the other hand, nanotechnology has provided in the last decade a broad range of different synthesis strategies for polymeric and inorganic colloi- dal nanoparticles.26–30 These colloidal nanoparticles could be assembled into more complex enti- ties and provide a functional powder material.31,32

In this PhD thesis, a novel colloid-based process chain is applied to produce tailored polymeric and composite supraparticles for powder bed based biomedical additive manufacturing. The su- praparticles are obtained from colloidal dispersions via spray drying, where the dispersion drop- let serves as the external confinement for the self-assembly of the nanoparticles.32 These supra- particles are tested towards their suitability for the PBF-LB/P process and their possible future application for bone grafting and tooth restoration. Different polymeric materials like polymethyl methacrylate (PMMA), polylactide (PLA), and polycaprolactone (PCL) are investigated together with functional additives like binary calcium silica nanoparticles (Ca-SiO2), hydroxyapatite nano- particles (HAP), and mesoporous silica (MSiO2) nanoparticles loaded with an antimicrobial drug.

At first, the influence of spray drying process parameters is studied towards particle size distri- bution as well as the morphology of PMMA supraparticles. Here, emphasis is placed on the influ- ence of the interplay of the supraparticle size distribution and the supraparticle surface roughness on the powder flowability. Then, the structure formation is investigated for PMMA-SiO2 composite supraparticles resulting from different mass mixing and diameter ratios. Moreover, PMMA com- posite supraparticles comprising mesoporous silica (MSiO2) primary particles loaded with the an- timicrobial drug octenidine (OCT) are analyzed towards their drug release and antimicrobial ac- tivity.33 At last, optimized PMMA and PMMA-SiO2 supraparticles are used in the powder bed fusion process (PBF-LB/P).

Secondly, polymeric polylactide (PLA) and polycaprolactone (PCL) primary particles are prepared in a scaled-up miniemulsion solvent evaporation technique.34 Calcium containing primary parti- cles (Ca-SiO2, HAP) are synthesized via precipitation and sol-gel approaches.35,36 Then, the thermal properties and powder flowability of the PLA composite supraparticles are characterized. Addi- tionally, the biocompatibility of the produced PLA composite supraparticle powders is investi- gated using human osteoblast-like cells (MG-63). Moreover, the mechanical properties for cylin- drical specimens, prepared by sintering the PLA composite supraparticles, are analyzed. Subse- quently, the PLA composite supraparticle powders are processed in the powder bed fusion pro- cess to fabricate multilayered square specimens, which are further on investigated towards their biodegradation and bioactivity.


2 Theoretical Background

In this chapter, the state of the art for additive manufacturing technologies in the biomedical field and biomaterials for bone grafting applications are explained. Moreover, the fundamental aspects of colloidal stability, spray drying, and powder flowability are described.

2.1 Additive Manufacturing in Biomedical Field

The term additive manufacturing refers to a set of generative technologies, in which objects are produced from computer-aided design (CAD) in a layer-by-layer approach.15,16 These technologies can be classified into material extrusion (MEX), material jetting (MJT), binder jetting (BJT), sheet lamination (SHL), vat polymerization (VPP), powder bed fusion (PBF), and direct energy deposi- tion (DED).15,18 Among these techniques, powder bed fusion laser beam melting process of poly- mers (PBF-LB/P) provides several attractive features for the fabrication of personalized bone im- plants.15,19 With this process complex intricate geometries and shapes with high porosities can be produced without the need for any supporting structures in comparison to the fused deposition modeling process (FDM).16,37 The minimum resolution in the PBF-LB/P process is within the range of 80-100 µm depends on the particle size and laser beam size.15,38 It should be pointed out that the mechanical properties of the resulting parts are superior compared to other polymer-based additive manufacturing technologies.39,40 Additionally, no potentially toxic binder materials are used in the PBF-LB/P process.15,16

Powder Bed Fusion Laser Beam Melting Process of Polymers

In the PBF-LB/P process thermoplastic, semicrystalline polymers are generally used.41,42 The powder bed fusion process consists of three steps that are repeated until the desired object is obtained (Figure 1).15,24 In the first process step, a thin layer of powder material is applied in the building chamber using a blade or a roller with layer heights between 100-200 µm (Figure 1, step 1).15,43 Then, the powder material is heated in the building chamber to a temperature, slightly be- low the melting temperature of the polymer (Figure 1, step 2).44,45 In a subsequent step, the pow- der material is molten in desired, spatially confined regions with the energy input of a CO2 laser (Figure 1, step 3).46,47 After that the building chamber is lowered, and a new layer of fresh powder is distributed. This process is repeated until the desired specimen is built.24,47 After the completion of the printing job, the temperature of the building chamber is slowly lowered to room tempera- ture to allow for gradual crystallization of the polymer.15,16 This slow cooling rate prevents curling of the specimen, which negatively affects the dimensional accuracy.48,49


Figure 1: Schematic representation of the powder bed fusion process for polymers (PBF-LB/P). In the first step the building chamber platform is lowered, and a homogenous powder layer is applied with a roller. In the second step, the powder bed is heated to the chamber temperature within the sintering window (temperature range between the onset of the crystallization peak and the onset of the melting peak) and in the last step, the powder particles are spatially resolved molten with additional laser energy. Own illustration based on literature.15,16,44

The powder materials need to meet certain characteristics regarding their powder flowability, their thermal and optical properties in order to be applicable in the PBF-LB/P process.15,44 In the first step of the PBF-LB/P process it is important that the powder materials create a dense and homogenous powder bed.15,24 For this reason, the powder needs to exhibit good powder spread ability and powder flowability.50,51 Furthermore, the powder material should have a high bulk powder density and a mean particle size in the range of 20-80 µm.16,41,52

In addition, the polymer particles must exhibit a broad sintering window to allow for stabile pro- cess control.18,42 The sintering window is defined as the temperature range between the onset of the crystallization and the onset of the melting peak determined via differential scanning calorim- etry (Figure 1, green area).18,42 Commercially available polyamide 12 (PA12) powders have a sintering window (Sw) of ΔT = 27.2 °C.49 Moreover, the melt viscosity and the melt surface tension of the polymeric material need to allow for sufficient coalescence of the powder particles after the energy input of the CO2 laser.41,48 Furthermore, the polymeric material should be stable enough to withstand extended periods close to the melting temperature without degrading.16,53


Also, the optical properties are of great importance, since the powder material needs to provide sufficient absorbance of the laser energy in the wavelength of 10.6 µm to melt.54,55 For polymeric particles this is possible in this spectral range due to the aliphatic carbon hydrogen bonds that can be excited with IR light into vibrations causing heat formation.15,16 In literature it has been demon- strated that the optical properties of powders can be enhanced via the addition of carbon black.56

Powder Production and Materials for PBF-LB/P Process

Currently, the commercial material selection is limited due to the high powder material require- ments of the PBF-LB/P process and about 90 % of the market share is dominated by polyam- ide 12.25,57 The available powder materials can be classified into unfilled and filled powders.15,16 Unfilled powder material represents the pure polymeric particles, which, in the case of polyamide 12 particles, are produced via precipitation.58,59 Furthermore, a few filled powder materials exist, which are physical mixtures in the microscale of polyamide 12 powder and additives like glass beads or carbon fibers.15,16

Different powder fabrication routes for different polymeric powder materials are reported in the literature.15 However, most of them require various post-processing steps to yield particles suit- able for the powder bed fusion process (PBF-LB/P).60–62 Cryogenic grinding can be used to fabri- cate powder particles from polymer granulates.63 Nevertheless, the resulting particles exhibit ir- regular shapes and rough edges with poor flowability.64,65 The flowability of those particles can be improved via sieving, thermal rounding, and particle dry coating.50,62,63 Another possibility is the production of spherical particles via coextrusion of an organic polymer along with a water-soluble polymer.66 Still, the recovery of the matrix water soluble polymer remains difficult for continuous processing.16 To expand the variety of polymeric materials for the PBF-LB/P process other poly- mers like polyamide 11 (PA11),59 polyethylene (PE),67 polypropylene (PP)68, polybutylene tereph- thalate (PBT),69 polycarbonate (PC)70, and polystyrene (PS)71 have been investigated to cover dif- ferent application areas. In particular for biomedical applications polylactide (PLA)72,73 and poly- caprolactone (PCL),74,75 were investigated as unfilled powders.


2.2 Biomaterials for Bone Grafting

In addition to the requirements of the PBF-LB/P process, the powder material must also meet the prerequisites as a biomaterial for bone grafting application.17,20 A biomaterial is defined as a ma- terial that interacts with a biological system and treats or replaces any tissue or organ of the body.14 Moreover, the biomaterial should not trigger any harmful immune response.76 Further- more, it is important that degradation products of the biomaterial do not cause any damaging side reactions to the body.11 Biomaterials, in general, can be classified into biotolerant, bioinert and bioactive materials.11,14 In the special case of bone implants, it has to be considered that the im- plant material should match the mechanical and healing properties of the bone tissue.77 Bone im- plants are employed after fractures, bone cancer, osteoporosis, and bone infections.78–80

Human Bone Tissue and Bone Remodeling

The human body comprises more than 200 bones that are responsible for stability and move- ment.11 Furthermore, the bones serve as a reservoir for phosphate and calcium and the formation of new blood cells via hematopoiesis.10 The bone tissue consists of two different structures; the cancellous and the cortical bone.10,23 The cortical bone tissue constitutes the outer layer of the bone and has a porosity of approximately 10 vol%, whereas the inner cancellous bone tissue has a porosity of 50-90 vol%.23 The compressive strength for human cancellous bone was reported in the range of σc = 2-12 MPa and for cortical bone from σc = 100-230 MPa.81–83 The bone matrix consists of 50-70% inorganic calcium phosphate in the form of substituted hydroxyapatite and 30 -50 % of collagen, osteocytes, lipids, and fibroblasts.10,11

Moreover, the bone undergoes constantly remodeling to adjust to mechanical loads, mobilize cal- cium, and repair small fractures.10,84 The process of new bone formation is called ossification and it can either occur endochondral or desmal.10,85 Desmal ossification is the direct bone formation from mesenchymal progenitor cells that differentiate into osteoblasts and produce osteoid, which forms the bone matrix and is subsequently mineralized.11 Endochondral bone formation refers to the formation of bone material from cartilage via primary ossification centers and is the most dominant mechanism in the human body.10,86 The bone resorption during the bone remodeling is done by the osteoclasts, which can resorb the inorganic hydroxyapatite matrix.84 The inorganic bone matrix is acid soluble and the osteoclasts attach to the bone, lower the pH and resorb the bone matrix.10,87 During the bone remodeling process the ratio between the bone building osteo- blasts to the bone resorbing osteoclasts is shifted to the osteoblasts in a healthy body.11,88


Bone Grafting Materials

The bone remodeling might be sufficient to repair smaller fractures, nevertheless for bigger bone fractures or bone defects caused by osteoporosis, osteosarcoma, or osteomyelitis, bone replace- ment implants are needed. 78–80 These bone replacement materials can be of biological or synthetic origin.11,89,90 Biological bone replacement materials can be autologous, allogenous, or xenoge- neic.91 Autologous bone replacement material would be taken from the patient, whereas al- logenous material would be obtained from another human individual and xenogeneic material would be derived from animals.91,92 Autologous bone replacement from the iliac crest represents the most suitable replacement since no immune reactions to the implant material are expected.93 However, this method depends strongly on the size of the bone defect and requires additional surgery.10,89,92 Allogenous and xenogeneic materials require careful sterilization and can cause im- mune reactions and secondary infections.90,91

In terms of synthetic materials for bone implants, metals and ceramic are commonly used.11,77 For hip and knee implants combinations of titan alloys (Ti-6Al-4V) or cobalt-chrome alloys (Co-Cr) with ceramics like aluminum oxide (Al2O3) are used.77,94,95 To provide sufficient hold in the bone matrix, the synthetic implants can be cemented with polymethyl methacrylate.96,97 Polymethyl methacrylate (PMMA) is chosen as a fixation material in bone grafting applications and as a filler material in dentistry due to its good biocompatibility and mechanical properties.96,98 However, these synthetic bone implants suffer one disadvantage.11,81 According to Wolff’s law, the bone will increase in mechanical strength when it is exposed to mechanical load but also decrease in me- chanical strength when there is less or no mechanical load applied.10,88 Therefore, it is crucial that the bone implant exhibits similar mechanical properties to the existing bone, otherwise the im- plant can break out of the bone matrix. 11,77

The presence of long-term biocompatibility issues and the challenges to match the mechanical properties of the bone have raised the interest in biodegradable implant materials or combina- tions of such with inorganic filler materials for small bone defects.11,99 Biopolyesters such as pol- ylactide (PLA) have been applied as bone replacements in form of screws or plates.15,100 These polymers can biodegrade in the human body through hydrolytic cleavage of the ester bonds in the polymer backbone.100,101 The mechanical properties of the bone are further mimicked by more complex composite materials of biopolyesters with calcium containing ceramics and glasses.102,103 The biopolymers represent the elastic quality of the bone, whereas the inorganic calcium contain- ing ceramics or glasses contribute to the mechanical stability and stimulate bone growth through the release of ions.74,103


Synthesis and Properties of Polylactide

Polylactide (PLA) represents a biopolyester that is derived from lactic acid.100,101 Lactic acid (Fig- ure 2 a) can be obtained via fermentation of renewable sources like cornstarch or sugarcane and exists in two different enantiomers L-lactic acid and D-lactic acid. 101,104 PLA can be synthesized via polycondensation, ring-opening polymerization, azeotropic dehydration, or enzymatic polymerization.101 In industry the most common route to synthesize PLA is via ring-opening polymerization (ROP) due to milder reaction conditions and shorter reaction times.105 In general, lactide, the cyclic dimer of lactic acid (Figure 2 b), is used as a monomer together with stannous octoate Sn(Oct)2 as the catalyst (Figure 2 c) for the ring-opening polymerization (ROP).106

Figure 2: Raw materials for the synthesis of polylactide (PLA) via ring-opening polymerization (ROP) using stannous octoate. a) Lactic acid which is converted via dimerization to the monomer lactide with its stereoisomers (b) and the catalyst stannous octoate for the ring-opening polymerization (c). Own illustration based on literature.101,104

In the first step of the polymerization, the catalyst Sn(Oct)2 reacts with protic educts like water or alcohols, which can be present as impurities within the monomer.107,108 The catalyst Sn(Oct)2 then converts into the initiator (OctSnOR) for the ring-opening polymerization (Figure 3 a).106,108 The initiator (OctSnOR) reacts subsequently with lactide and forms the first polymer chain segment (Figure 3 b).109,110 Then, the chain growth occurs via ring-opening of the following lactide mono- mers (Figure 3 c).105 Termination and inactivation of the growing polymer chains can be induced via reaction with protic components (Figure 3 d,e).111 Thus, control over the molecular weight of the PLA polymer can be achieved through the addition of protic components like long chain alcohols.106,108 Moreover, it should be mentioned that the molecular weight and the amount of D- and L-isomer within the PLA polymer strongly affect its physical properties.101 PLA polymers can exist in semicrystalline or amorphous forms depending on the stereochemistry of the used mon- omers.100,101 Pure poly-L-lactide (PLLA) and pure poly-D-lactide (PDLA) polymers are semicrys- talline, whereas poly-D,L-lactide (PDLLA) are amorphous.101,112 In the literature it is furthermore reported that semicrystalline PLA can be obtained if the content of L-lactide is higher than 90%

within the polymer.104,112 The glass transition temperature for PLA was reported as Tg = 58 °C and the melting temperature is in the range of Tm = 130-230°C, which depends on the crystallinity, the molecular weight, and the amount of D-isomer in the polymer.101


Figure 3: Synthesis of polylactide (PLA) via ring-opening polymerization (ROP). a) Reaction of stannous octoate (Sn(Oct)2) with protic compo- nents (ROH) and formation of the initiator (OctSnOR). b) Formation of the first chain segment through the reaction of the initiator with lactide and chain growth with the addition of lactide (c). Inactivation of growing polymer chains and termination via reaction with OctH. Own illus- tration based on literature.106,108–111

Fully crystalline PLA is reported to have a heat of fusion of ΔHm0 = 93.7 J g-1 and the degree of crystallinity is affected by the thermal history and the presence of nucleating agents such as talc or nanoclays.113–115 Moreover, it should be noted that PLA crystals can grow in three different structural positions, which also affects the melting temperature.101 The α-crystallite structure grows upon cold or melt crystallization, whereas the β-from develops upon mechanical stretching and the γ-form is only observed in special cases.101 The mechanical properties and the degradation behavior of PLA are strongly dependent on the molecular weight and the crystallinity of the poly- mer.101,110 The mechanical properties of semicrystalline PLA are higher compared to the amor- phous PLA due to the crystallite regions, which connect the amorphous regions and dissipate en- ergy before fracture.101 The degradation of the polymer occurs by scission of the ester bonds in the polymer backbone by water.101,116 For semicrystalline PLA implants the water first diffuses to the ester bonds in the amorphous regions and then hydrolytically cleaves them, leading to a de- crease in the molecular weight until the polymer chain fragments become water soluble.117


Then, those fragments can be introduced into the tricarboxylic acid cycle and further metabo- lized.100,117 The degradation time frame for PLA in vivo, was found to be in the range from 2 to 5 years.100,101 In addition, it was observed that the degradation of PLLA implants with a molecular weight of 60,000 g mol-1 was in the timeframe for bone healing and the implant maintained its mechanical properties during the healing process.118

Synthesis and Properties of Polycaprolactone

Polycaprolactone (PCL) is another example of a biodegradable, semicrystalline biopolyester for bone grafting applications.119 It is synthesized via ring-opening polymerization (ROP) from the monomer ε-caprolactone (Figure 4).100 The polymer exhibits a low glass transition temperature of Tg = -60 °C and a low melting temperature in the range of Tm = 55-60 °C.120 Polycaprolactone has a low tensile strength, but exhibits high flexibility at room temperature due to its low glass transition temperature.11

Figure 4: Synthesis of polycaprolactone from ε-caprolactone via ring-opening polymerization. Own illustration based on literature.11,100

Calcium Containing Biomaterials

To further increase the mechanical properties of biopolymers for bone grafting applications, com- posite materials of biopolymers with ceramics or glasses containing calcium can be applied.20,102 Calcium phosphates and bioactive glasses are of big interest for bone grafting applications.30,121,122 Calcium phosphate is the main mineral compound that makes up bones.10,11 In the human body it is present as non-stochiometric apatite phases through the integration of ions like Na+, Mg2+, or CO32- and is referred to as biological apatite.91,123 Calcium phosphates can be synthesized from calcium salts and phosphoric acid via precipitation as amorphous calcium phosphate.123,124 Cal- cium phosphates are mostly non-soluble in water; however, they can be dissolved by acids.124 In general, it can be said that calcium phosphates are osteoconductive and biocompatible.124 Another calcium containing inorganic material is Bioglass, which was first developed by L. Hench in 1969.122 Commercial Bioglass 45S5 consists of 45 wt% SiO2, 24.5 wt% Na2O, 24.5 wt% CaO, and 6% P2O5 and can form strong bonds with bone tissue via the formation of a calcium deficient car- bonated apatite surface layer.103,122 In addition, other bioactive glasses have been reported and special interest has recently been focused on nanoscale bioactive glasses.30,103


Those nanoscale bioactive glasses promise a higher specific surface area that allows faster ion release and mineralization.103 Different synthesis procedures using sol-gel approaches have been studied to incorporate calcium and phosphate into the silica network and calcinating the particles to obtain bioactive glasses.30,35 The implementation of nanoscale bioactive silica nanoparticles in polymer matrices showed better improvement of mechanical properties compared to micronized commercial Bioglass.103

2.3 Fundamentals on Colloids and Emulsions

Colloids represent dispersed systems consisting of a continuous and a dispersed phase, which can be solid or liquid.125,126 Moreover, the droplet or particle size of disperse phase ranges from a few nanometers to a few micrometers.127 Colloidal systems appear macroscopically homogenous and are microscopically dispersed.127 Furthermore, it should be noted that colloidal systems can be divided into the subgroups of molecular, association, and dispersion colloids.128 Examples of mo- lecular colloids are macromolecules such as methylcellulose (MC).126 Surfactants like sodium do- decyl sulfate (SDS), which can form micelles, represent an example for association colloids.129 In this thesis special emphasis is placed on the dispersion colloids, where particles in the nanometer to micrometer range are dispersed in a continuous liquid phase (Figure 5).

Stabilization of Colloidal Particles and DLVO Theory

In general colloidal systems can be stabilized via steric and electrostatic interactions.127,129 For sterically stabilized dispersion colloids, macromolecules can be used such as polyvinyl alcohol (PVA) or methylcellulose (MC).127,130 These molecules adhere to the surface of the nanoparticles, increase the distance and reduce their mobility.128 Moreover, if the nanoparticles encounter each other, the concentration of the adsorbed polymer molecules increases.131 This leads to an increase in the osmotic pressure, triggering the diffusion of solvent molecules and thus increasing the dis- tance between the nanoparticles (Figure 5 a).129

Colloidal nanoparticles can be also stabilized electrostatically.127 This can be achieved through the presence of charged functional groups originating from reactions with acid and base or through the adsorption of charged molecules like sodium dodecyl sulfate (SDS) (Figure 5 a).128 Based on the Stern and the Gouy-Chapman model the adsorption of ions from the electrolyte can be de- scribed (Figure 5 b).132,133 The Stern layer represents the first layer that forms around a charged particle in presence of an electrolyte.129 It consists of the inner and outer Helmholtz layer (Figure 5 b, II).129 The inner Helmholtz layer is composed of a fixed monolayer of the same charged ions adsorbed to the particle surface due to van der Waals forces.129


In addition, a fixed monolayer of cations, the outer Helmholtz layer, is adsorbed.127 The Stern layer is followed by the diffusive (Gouy-Chapman) layer, which is rich with cations.129 These ions de- crease with increasing distance from the particle and as a consequence the electric potential de- creases (Figure 5 b, III).130 Partially ions can escape from the diffusive layer since colloidal parti- cles are subject to the Brownian molecular movement and statistical diffusion processes.130 This measured potential at the slipping plane is referred to as the zetapotential and can be used as an indication for the colloidal stability of a system (Figure 5 b, IV).127 A nominal value of ζ = 20 mV is typically considered stable.127,129

The DLVO theory represents a commonly accepted theory to describe the interaction of colloidal particles (Figure 5 c).128 This theory rationalizes the interaction potential between two colloidal particles based on the interplay of the repulsive electrostatic (Figure 5 c, red) and attractive van der Waals forces (Figure 5 c, green) as a function of their distance (Equation 3).129,130 When two charged nanoparticles get into closer contact, the first force they experience is the repulsive elec- trostatic force (VR) which can be described via Equation 1.129

For the calculation of the electrostatic repulsive force (VR) the radius of the nanoparticle (a), the charge of the ions (z), the dielectric constant of the dispersion medium (ε), the Boltzmann constant (kB), the elementary charge (e0), the Debye length (κ), the distance between the particles (H) and a parameter (γ), which takes the surface potential (ψ0) into account, are considered (Equation 1).129 The attractive van der Waals forces (VA) only come into consideration when the energy bar- rier of the total interaction is overcome (Figure 5 c, II).127,130 The van der Waals interaction can be approximated with Equation 2, where the Hamaker constant (A) and the distance between the particles come into consideration.129

𝑉𝑅= 𝑎


𝑒02 ∙ 𝛾2𝑒−𝜅𝐻 (1)

𝑉𝐴= −𝐴

12∙ (𝑎 + 0.75𝐻

𝐻 + 2𝑙𝑛 𝐻

𝑎 + 0.75𝐻) (2)

𝑉𝑇= 𝑉𝐴+ 𝑉𝑅 (3)

The particles would be then aggregated at the first minimum (Figure 5 c, III).127 The Born repul- sion caused by the overlapping of the electron clouds hinders the particles to further approach each other.130


Figure 5: Stabilization of colloidal particles with Stern model for electrical double layer and DLVO theory. a) Electrostatic and steric stabiliza- tion of colloidal particles. b) Stern model for an electrical double layer of a colloidal particle in an electrolyte with particle surface charge (I), Stern layer (II), and diffusive (Gouy-Chapman) layer (III) with slipping plane (IV). c) Total interaction potential of two colloidal particles in dependency of the distance with electrostatic repulsion, van der Waals attraction, and Born repulsion. I) Secondary minimum, (II) energy barrier, and (III) primary minimum. Own illustration based on literature.127,129

Formulation of Emulsions

Emulsions are mixtures of at least two immiscible liquids and an emulsifier.131 One liquid serves as the inner phase and the other liquid as the outer phase.129 The volumetric phase ratio (Φ) be- tween the inner (Vinner phase)to the outer phase (Vouter phase) can be defined according to Equation 4.126,134 Binary emulsions of water and oil can form two different emulsion types depending on the used emulsifier and the volumetric phase ratio (Φ).131 An oil in water (O/W) emulsion is obtained, when the oil is the dispersed and water the continuous phase.125 In the opposite case, a water in oil (W/O) emulsion is achieved, when the water represents the inner phase and the oil the outer phase.129,130

𝜙 = 𝑉𝑖𝑛𝑛𝑒𝑟 𝑝ℎ𝑎𝑠𝑒

𝑉𝑜𝑢𝑡𝑒𝑟 𝑝ℎ𝑎𝑠𝑒 (4)

Furthermore, emulsions can be classified according to the droplet size of their inner phase into macroemulsions (100 nm - 100 µm) and miniemulsions (100 nm - 1 µm).126 To formulate two immiscible liquids to an emulsion, a suitable emulsifier and a sufficient mechanical energy input are required.126,135 Emulsifiers embody a subcategory of surfactants and certain polymers -so called hydrocolloids- that can stabilize droplets of two immiscible liquids.129,131 The emulsifiers can be classified according to their head groups into ionic, amphoteric, and non-ionic emulsifi- ers.127,131


The choice of emulsifier strongly depends on the emulsion system and the desired stabilization mechanism.125,126 To select a proper emulsifier the Bancroft rule and the HLB value, which repre- sents the hydrophilic to the lipophilic balance of the emulsifier, can be used.125,129 Emulsifiers with a low HLB value can be used for water in oil (W/O) emulsions and emulsifiers with a high HLB value for the formulation of oil in water (O/W) emulsions.126,131 Ionic emulsifiers like sodium do- decyl sulfate (SDS) stabilize the droplets of the inner phase electrostatically, whereas polymers like methylcellulose (MC) or polyvinyl alcohol (PVA) stabilize the droplets sterically.129,130 To generate the droplets of the inner phase in the continuous phase energy input is needed.125 The energy input can be provided via ultrasonic processors, high pressure homogenizers, or rotor- stator machines.131 Successful droplet breakup is achieved when the external shearing forces (σ) are larger than the internal forces (pk) holding the droplet together.126,136 The external shearing forces strongly depend on the flow regime.131 For rotor-stator machines like the Ultraturrax tur- bulent flow conditions are present.126 The surface tension (γ) holds the droplet together and can be expressed by the capillary pressure (pk) via the Laplace equation (Equation 5).136 The capillary number Ca (Equation 6) describes the ratio of the droplet deforming forces to the internal stabi- lizing forces. When a value above a critical threshold is reached, droplet breakup can occur.126,136 The extent of the deforming force (σ) can be expressed via the power (Pv) and energy density (Ev).126


𝑥 (5)

𝐶𝑎 = 𝜎 ∙ 𝑥

4 ∙ 𝛾 (6)

The power density (Pv, Equation 7) is obtained by dividing the applied mechanical power (P) by the emulsion volume (V), and the energy density (Ev, Equation 8) is obtained by multiplying the power density (Pv) with the emulsification time (t).126,136 The droplet size of the inner phase can be estimated via Equation 9, where the Sauter diameter (x32) is proportional to the energy density (Ev) and two fitting constants C and b.126


𝑉 (7)

𝐸𝑉= 𝑃 ∙ 𝑡

𝑉 (8)

𝑥32= 𝐶 ∙ (𝐸𝑉)−𝑏 (9)


Miniemulsion Solvent Evaporation Technique

The miniemulsion solvent evaporation technique is a versatile process to formulate colloidal pol- ymeric nanoparticles via miniemulsions (Figure 6).27,137 In this process, the respective polymer is dissolved in an organic solvent, which is not miscible with water.34 In the next step, an aqueous phase containing an emulsifier like sodium dodecyl sulfate (SDS) is added to the organic phase and stirred to form a pre-emulsion (Figure 6, step 1).27 Then, mechanical energy in form of an ultrasonic processor or rotor-stator system is applied to induce the droplet breakup and the for- mation of a miniemulsion (Figure 6, step 2).126 In the last step, the solvent of the inner organic phase is evaporated and a polymeric colloidal dispersion is obtained (Figure 6, step 3).126,138 The miniemulsion solvent evaporation technique is a convenient method to produce colloidal disper- sions from different polymers.138 In the PhD thesis the polymeric primary particles of polylactide (PLA) and polycaprolactone (PCL) are obtained through this method. The size of the polymeric nanoparticles can be tailored via the concentration and the type of the emulsifier as well as the extent of mechanical energy input.27,126

Figure 6: Schematic representation of the miniemulsion with solvent evaporation technique. In the first step, an aqueous phase containing a surfactant and an organic phase containing the dissolved polymer is prepared and added to a reaction vessel. Then, the two miscible phases are homogenized with an energy input given by ultra-turrax stirring or ultrasonication to form an emulsion. At last, the inner organic phase of the emulsion is evaporated, and a colloidal polymeric dispersion is obtained. Own illustration based on literature.27,139

2.4 Spray Drying of Nanoparticle Dispersions

Spray drying represents a process to formulate a particulate material from a liquid feedstock.140,141 Spray drying is used in the food, pharmaceutical, and chemical industry.125,140,142 In the PhD thesis the spray drying process is employed to obtain supraparticles from the colloidal primary particle dispersions. A supraparticle is an entity resulting from a colloidal dispersion of primary particles that were assembled in a confinement.32 Primary particles in that sense are colloidal nanoparticle dispersions.32,71 In the spray drying process the dispersion droplet is used as a confinement and the process itself can be divided into the following steps (Figure 7 a):142–144


First, the liquid feed stock in form of a colloidal dispersion is introduced into the process via a pump and is subsequently atomized into small droplets via an atomization equipment (Figure 7 b).145,146 Then, the droplets enter the heated drying chamber, where the droplets get in contact with a heated drying gas and the droplet to particle conversion occurs (Figure 7 c).147–149 In a last step, the particulate product is separated from the heated gas via a cyclone and gathered in a col- lection vessel (Figure 7 a).142,144 The product quality and design can be influenced via the external process parameters as well as from the product properties of the liquid feedstock.142–144 In the following paragraphs the individual spray drying process steps will be discussed in detail as well as their effect on the product design.

Droplet Generation via Atomization Step

After the liquid feed is pumped into the spray dryer, the atomization process plays a crucial role for the product properties.143,150,151 The atomization process strongly influences the final mean particle size and particle size distribution.143,152 To break the liquid feed into small droplets, the atomization equipment can apply centrifugal force, electrostatic force, pressure or ultrasonic en- ergy.142,152 The corresponding atomization equipment used in spray dryers is fluid nozzles, rotary disks, or ultrasonic nebulizers.142,152

The droplet size of the atomized liquid feed can vary from 1-1000 µm depending on the used at- omization device and the properties of the liquid feed material.142,144,145 Additionally, the choice of the atomization equipment is strongly based on the characteristics of the used liquid feed mate- rial.140–142 In most lab-scale spray dryers like the Büchi B-290, used in this thesis, two-fluid nozzles are employed (Figure 7 b).153 To create the droplets from the continuous liquid feed an additional nitrogen gas is used in the atomization process.145,146 The size of the resulting droplets (x32) can be calculated (Equation 10) using the diameter of the two-fluid nozzle (dN), the Weber number (We, Equation 11), two fitting parameters (H,c), and the air to liquid ratio (ALR, Equa- tion 12).145,146,154

𝑥32= 𝐻 ∙ 𝑑𝑁∙ [ 𝑊𝑒 (1 + 1





𝑊𝑒 = 𝑣𝑔2∙ 𝜌𝑎,𝑔𝑎𝑠∙ 𝑑𝑁

𝛾𝑓𝑒𝑒𝑑 (11)


The air to liquid ratio (ALR) is an important parameter to tailor the size of the droplet and thus of the resulting supraparticle.142–144 The ALR (Equation 12) is defined as the ratio of the mass flow of the atomization gas (ṁa,gas) to the feed mass flow(ṁfeed).140,145,146 For higher ALR values, smaller droplet sizes are obtained, whereas for lower ALR values, larger droplets are obtained.146,155 Moreover, the Weber number (We) is calculated using the atomization gas velocity (vg), the den- sity of the atomization gas (ρa,gas), the diameter of the nozzle (dN) and the surface tension (γfeed) of the liquid feed (Equation 11).145,146,154 The atomization gas velocity (vg) is obtained via Equation 13, where the inlet temperature (Tinl), the ideal gas constant (R), the molar mass (M), the isen- tropic coefficient (κ), the inlet (pin) and the outlet pressure (pout) are used.146,154 In addition, it should be mentioned that also the viscosity of the feed dispersion and the surface tension strongly influence the droplet breakup and therefore the mean particle size.140,142,143

𝐴𝐿𝑅 = 𝑚̇𝑎,𝑔𝑎𝑠

𝑚̇𝑓𝑒𝑒𝑑 (12)

𝑣𝑔= √𝑇𝑖𝑛𝑙∙ 𝑅 𝑀𝑎,𝑔𝑎𝑠 2 ∙ 𝜅

𝜅 − 1∙ (1 − (𝑝𝑜𝑢𝑡



𝜅 ) (13)

Droplet Drying and Particle Formation

In the next step, the dispersion droplets enter the heated drying chamber of the spray dryer (Fig- ure 7 c).147–149 The carrier gas flow can be in co-current flow mode with the atomization gas or counter current.141,142 In the Büchi B-290 lab-scale spray dryer a co-current flow mode is used providing milder process conditions.153 After the dispersion droplets enter the drying chamber with the heated carrier gas, the convective drying process occurs.142,148 In general, the drying pro- cess can be divided into two drying stages for dispersions containing insoluble particles (Figure 7 c).147,156

The first drying stage starts with the heating up of the droplet to the wet bulb temperature due to the heated carrier gas (Figure 7 c, I).147,148 Then, the evaporation (ṁv) of the liquid starts from the surface (Ad) of the droplet.141,144,157 The droplet radius shrinkage over time (Equation 14) can be expressed using the density of the wet droplet (ρw,d), the initial droplet radius (Rd), and the mass flow (ṁv) of the vapor (Figure 7 c, II).147,149




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