Clotilde

Hanna Martin

Jana Cédric

Lydia

Dr Susanne Muschert

Dr Younes Karrout

Dr Kevimy Agossa

Mickael Maton

Hugues Florin

Muriel Deudon

En premier lieu je tiens à remercier mon directeur de thèse, le Pr Juergen Siepmann ; pour la confiance qu'il m'a accordé tout au long de ces trois années, ainsi que les précieux et nombreux conseils qu'il a su me prodiguer et qui m'ont aidé à me développer. Travailler au sein de l'unité qu'il dirige a été une aventure heureuse.

Mes remerciements vont au

Pr Jérémie Soulestin, qui me fait l'honneur de présider ce jury. Je tiens également à le remercier pour son soutien et son ouverture à la collaboration durant ces trois années à travers le projet 3D med. Je tiens également à remercier vivement le Pr. Abdul Basit, qui me fait l'honneur de juger et d'évaluer ces trois années de travail. Merci également au Pr. Yann Pellequer, qui a su dès ma première année d'étude supérieure me transmettre sa passion pour la galénique et la biopharmacie. Il y a 10 ans il me donnait mes premiers cours de pharmacie galénique, aujourd'hui il me fait l'honneur d'évaluer mes travaux de thèse, c'est une belle façon de clôturer ce dernier chapitre de ma vie étudiante. Je tiens également à remercier le Pr Florence Siepmann pour son partage, son savoir, ses qualités scientifiques et aussi humaines, ainsi que pour sa présence pendant ces trois années de thèse. Tous mes remerciements aux Dr. Sébastien Charlon, et Dr. Cédric Samuel, pour m'avoir accueilli quelques fois au sein de l'IMT Lille-Douai, et pour avoir pris le temps de partager avec moi leurs connaissances. Un grand merci également à Jérémy Vérin, pour son aide, son support, et pour toujours avoir répondu présent en cas de problème. Je garderais en mémoire les longues journées passées ensemble pour les photos SEM, mais aussi ton humour et les moments de partage autour d'un verre, d'un repas ou d'un café. Travailler pendant ces trois années au sein de l'INSERM U 1008 m'a permis de rencontrer des personnes sans lesquelles cette aventure aurait eu une saveur bien différente. Mes coéquipiers de l'ancienne vague qui m'ont vu passer par toutes sortes d'émotions : Youcef, Martin L., Coco, Adam. Et surtout mes deux co-équipières, qui avec le temps sont devenues de véritables amies, des oreilles attentives, des épaules sur lesquelles me reposer, des rire à partager : Fahima et Doha, avoir croisés vos chemins a été un véritable cadeau. Il y a également mes collègues de la nouvelle vague : Yanis, Fabiana, Lise-Anne, Filipa, Charline et Samuel, vous êtes chacun à votre façon venu ajouter votre touche de couleur et de bonne humeur dans cette aventure. Je vous

days exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates (the receptor compartment containing well-agitated bulk fluid). At predetermined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated). Please note that after exposure to the release medium, the implants were freeze-dried prior to analysis. Thus, caution must be paid The success of many therapies requires a minimum concentration of active pharmaceutical ingredients (API). Thus, to ensure pharmacological activity, without reaching the toxicity threshold causing adverse effects. For conventional drug delivery systems (immediate-release), plasmatic drug concentration can be kept constant by repeated administration of small doses of API. However, these frequent administrations may lead to lake of compliance and failure of the medical treatment. Moreover, these significant variations in the plasmatic drug concentration may be critical when the API has a narrow therapeutic window (1).

To optimize the therapeutic response, controlled drug delivery systems allow compensating elimination of the API by the body, by providing the API at an appropriate rate (1). Thus, optimal plasmatic drug concentration can be reached after one administration and kept constant during a certain period, as illustrated in Figure I-1.

Figure I-1. Active pharmaceutical ingredient plasmatic concentration, after repeated administration of conventional drug delivery systems (green lines) and administration from an "ideal" controlled drug delivery system (blue line).

A lot of oral controlled drug delivery systems have been developed, allowing optimum concentrations to be reached at a determined speed (2). However, these are not always suitable due to the hostile environment along the gastrointestinal tract (pH, enzymes, low absorption capacity, hepatic first-pass effect, etc.) (3,4). To overcome these limitations, parenteral control drug delivery systems are a good alternative. Their advantages are numerous including (i) release of the API at a speed adapted to in vivo needs and (ii) release of the API directly at the

I.1.2.2. Requirements and regulations

Long-acting injectables are complex formulations requiring dedicated regulations to assure their quality, safety, and efficacy as illustrated in Procedures for uniformity, sterility, and biocompatibility tests are well described for the parenteral formulations, unlike the release of the active substance (8,9).

It has been reported that the type of in vitro release method depends on the test objectives (11).

Long acting injectable

Therefore, quality control and biorelevant methods used to predict in vivo performance should be developed separately (11)(12)(13). Due to the complexity of parenteral formulations the in vitro drug release test methods should be developed case-by-case (14). In fact, instrument and release medium should be chosen carefully as the release profile is directly impacted by: type of instrument, release medium composition, and temperature (3,(15)(16)(17)(18). To do so, key product attributes need to be known as well as, the release mechanism, and in vivo environment that influences drug release to successfully develop in vitro methods (3). The development of the in vitro method may require the use of modified compendial or non-compendial equipment (Varying volume, medium, agitation, etc.) (14,19). The sample and separate method could be used with a different container, from small recipients, if small release medium is needed (16,(21)(22)(23)(24)(25), to United State Pharmacopeia apparatus 1 or 2 This method is easy to implement, but it should be taken into consideration that, aggregation if the formulation or potential loss of formulation pieces after degradation (if the formulation is biodegradable) could occur during the separation step leading to erroneous release profiles (13).

I

I.1.3.1.2. Continuous flow through method

The continuous flow-through method referred to as apparatus IV in the United State pharmacopeia is illustrated in In this method, the formulation is contained in the cell and the release medium is pump from the reservoir through the cell. The apparatus can work either in an open or close loop and present many advantages such as: do not need a sample and separate and easy adjustment of the release medium volume (13,19).

I.1.3.1.3. Dialysis membrane technics

In the dialysis-based technics, the formulation is dispersed in the release medium in a dialysis bag, immerged in the release medium. In some cases, it has been reported that the formulation is dispersed in the chamber, and the dialysis bag contains a release medium as illustrated in In the case of the classic dialysis-based technic, the main disadvantages are (i) the potential aggregation of the formulation in the dialysis bag due to lack of agitation, and (ii) the potential violation of sink condition in the dialysis bag. The reversed dialysis-based technics have been developed to overcome the limitation of the sink condition in the dialysis bag (13).

In some cases, dialysis membranes have been used for mimicking in vivo compartments using United States pharmacopeia referred apparatus such as apparatus I or IV (26,27).

The environment of the site of administration plays a key role in the in vivo drug release from long-acting injectables. Among the factors impacting drug release, some are of particular importance: barriers to drug diffusion (viscosity); drug partitioning at the site (uptake into

Classic Reversed

Formulation Media Media Formulation Sampling Sampling tissue); the fluid volume available at the site; and in the case of intramuscular injection muscle movement may also be an important factor (13). Reported studies that investigated the variation of these parameters have been reported in Table I

I.1.3.2. Development of a suitable in vitro

-2.
If instruments based have been demonstrated to play a key role in the development of an in vitro drug release testing method (28)(29)(30)(31), the composition and the choice of the release medium are also of great importance.

Table I-2.

Examples of reported studies on the impact of instrument-based parameters on in vitro drug release.

Investigated parameter Effect on drug release Explanation Ref

Flow rate

It was observed that flow rate did not have a significant effect on the release of triamcinolone acetonide from the microspheres.

In this case, the flow rate was not a key parameter impacting in vitro drug release mechanisms.

(

) 28

Dialysis

Comparison of release profile obtained from microparticles with USP II, IV apparatus, and dialysis method, shows that all the profiles had a sigmoidal shape, and complete drug release was achieved after 6 days for all methods. However, for the dialysis method, the lag phase could not be distinguished from the erosion phase.

It might be due to a decrease in drug diffusion of the API through the membrane in comparison with the diffusion from the microparticles.

(29)

Recipient

Drug release was studied in Eppendorf tubes, and glass flasks, and customized flow-through cells. The impact of the different methods on drug release was moderate.

In the investigated field, the type of recipient did not seem to be a key parameter controlling the drug release from microparticles.

(

) 30

Agitation

The drug release rate was decreased in tubes without agitation compared to agitated tubes in a sample and separate method.

In the case of "non-agitated" tubes convective bulk fluid transport through the implant matrix can be considered negligible (30) Greater agitation provided by USP II apparatus compared to shaking bath lead to differences in drug release rate from microparticles.

Agitation may prevent aggregation by continually dispersing microparticles in the release medium, resulting in a faster release

I.1.3.2.2. Medium based parameters

Depending on the in vivo site of administration, some release medium parameters can be adjusted such as:

-Osmolarity, Reported studies that investigated the impact of these parameters have been reported in Table I-3. Development of a suitable in vitro drug release method by modulating instrument-based and medium based parameters to be as close as possible to the parameters controlling the release rate in vivo is of great importance (16,17,(28)(29)(30)(31)(32)(33)(34). Especially for poly (lactic-co-glycolic) acidbased devices, due to the numerous mechanisms involved in the drug release from such a polymer (35) as described in the following sections.

Buffer concentration

Drug release rate slightly decreased when increasing the buffer concentration Might be due to an increase in the concentration of bases in the release medium, leading to potential diminution au autocatalytic effect of PLGA (17) When increasing buffer concentration, release rate significantly increases Increasing ionic strength of the release medium, reduce interaction between API end polymer end-groups (32) When the buffer concentration was decreased from 0.02 M to 0.01 M a significant increase in drug release rate was observed. The release profile was also changed from tri-phasic to bi-phasic.

Might be due to the difference in osmotic pressure leading to different water uptake. ( 29)

pH

A similar trend in drug release during burst and lag phases at pH 2 and pH 4.

Release rate during erosion phase was faster at pH 2 than pH4

It indicates that burst and lag phases are not dependent on pH while the erosion phase is.

(

) 29

Temperature

Release rate significantly increases when temperature increases from 37°C to 65°C Probably due to the increase of the mobility (drug mobility, but also polymer chains mobility) with temperature

Release rate decreases when temperature decreases from 37°C to 5°C Due to decrease in mobility, water uptake, and degradation rate of the polymer (16,33) The time to achieve complete peptide release was reduced from 18 to 6 days by increasing the temperature from 40°C to 45°C.

Within the first two phases (burst and lag), the drug release is accelerated by diffusion as a higher temperature softens polymer chains and increases their mobility, while during the third phase, such a temperature enhances hydration and polymer erosion.

(

) 29
By increasing temperature from 35°C to 39°C drug release rate decreases Might be due to polymer plasticization, leading to drug entrapment (28)

Surfactants

With decreasing sodium dodecyl sulfate concentrations, the cumulative triamcinolone acetonide release decreased

The triamcinolone release observed in different decreasing sodium dodecyl sulfatecontaining media correlated well with the solubility of the drug in the respective media

Viscosity

Release rate during the 3 rd phase of poly(lactic-co-glycolic) acid-based implant and microparticles, is decreased when the systems are exposed to agarose gel instead of bulk fluid

The viscosity of the agarose gel limits the swelling and the water uptake of the implant during the 3 rd release phase.

(

) 34
The release rate is increased when microparticles are exposed to agarose gel compared to phosphate buffer pH 7.4

It might be explained as follows: when surrounded by bulk fluid, the acids generated by the hydrolysis of the polyester are more rapidly neutralized than in the case of the agarose gel technique, in which rapid convective mass transport is effectively suppressed.

I.2. Poly(lactic-co-glycolic) acid

I.2.1. General characteristics and synthesis

Poly (lactic-co-glycolic acid) (PLGA) is a copolymer composed of lactic acid (LA; levogyre or dextrogyre) and glycolic acid (GA) (36,37). PLGA can be synthesized either by polycondensation or ring-opening polymerization. As described in Different PLGA grades can be obtained by controlling polymerization conditions such as temperature, time, type and concentration of catalyst, nature, and quantity of monomer. These different grades are distinguished by the choice of the monomer ratio, the molecular weight, and the nature of the end-group of the copolymer (36,38,37).

These characteristics influence the physicochemical properties of the copolymer and therefore the mechanical properties, thermoplastic behavior, rate of degradation, and release of the associated API (37).

I.2.2. Molecular weight

When the Mw increase, the polymer has longer polymeric chains and so higher intrinsic viscosity. This leads to a lower degradation rate (39). As described in section I.2.1 depending on PLGA synthesis, different PLGA Mw could be obtained. x y

Poly (lactic-co-glycolic acid)

Δ T, catalyst
x y

14

PLGA Mw is known to have an impact on drug release, from devices composed of this polymer. Some studies reported that with higher Mw not only a longer release is reached but also a lower burst effect (40,41). Moreover, it has been recently reported that drug release depends on both the molecular weight and polydispersity of the polymer (42). They observed an increase of the initial burst release when increasing the amount of low molecular weight PLGA in microparticles.

However, Kohno et al., previously reported that in some cases Mw is not a critical formulation parameter regarding drug release compared to other parameters such as glass transition temperature, or API-polymer interactions (43)

I.2.3. Crystallinity and glass transition temperatures

In some cases, PLGA copolymers composed of less than 85% of GA have been reported to be amorphous (37). In fact, due to the coexistence of two enantiomeric isomers of LA (D or L, depending on the position of the methyl group on the alpha carbon), the crystalline organization is made difficult. While GA doesn't have the methyl group, this makes it highly subject to crystallinity (44). P(L)LGA composed from 75:25 to 30:70 (LA:GA) and P(D,L)LGA composed from 100:0 to 30:70 (LA:GA) have been reported to be amorphous copolymers (45,44).

Crystalline, semi-crystalline, and amorphous materials have different properties as illustrated in Figure I-8. The amorphous grades of PLGA are characterized by their glass transition temperature (Tg). Tgs are defined as a range of temperature, in which the polymer goes from a glassy to a rubbery state. Tgs are impacted by both the ratio between LA and GA, and the Mw of the PLGA (37,44). It has been reported that depending on PLGA composition and Mw, their Tgs range between 40°C and 60°C (44).

On the other side, some substances, of low molecular weight, called plasticizers have the property of being interposed between the polymer chains, enhancing their mobility, and their Tgs. Water, but also some API have been reported to have a plasticizing effect on PLGA (46)(47)(48)(49)24).

I.2.4. Hydrophilic properties and solubility

By modulating the ratio between its two constitutive monomers: LA and GA, a more or less hydrophilic polymer can be obtained. LA methyl group makes it more hydrophobic than GA, thus increasing the proportion of LA leads to a less hydrophilic copolymer (40,50). Due to this ability to varying PLGA hydrophilicity, it is possible to find PLGA grade that has affinity either with hydrophobic or hydrophilic API (39).

Many different solvents including chlorinated solvents, tetrahydrofuran, acetone, or ethyl acetate can dissolve PLGA, whatever the ratio between its constituting monomers (51,39).

I.2.5. Biodegradation

PLGA degradation occurs through hydrolysis of the polyester backbones. Biodegradation takes place by chemical hydrolysis and induces erosion of the polymer matrix. The water molecules diffuse into the polymer matrix and break the ester bonds of the main chain, leading to the production of oligomers, and its constituting monomers: LA and GA. Hydrolysis is a reaction catalyzed either by acidic or basic products (52). The acid catalyst can come either from a strong acid in the medium, or the carboxylic acid end groups of the PLGA. In the latter case, it is called autocatalytic reaction, as illustrated in The degradation products present free carboxyl groups which decrease the microenvironmental pH (25,53,54). However, the chemical hydrolysis reaction is catalyzed by the presence of protons (25,36). This phenomenon is important because it determines the degradation and erosion regime of the polymer matrix (52,55). This lowering of pH can also influence the stability of a formulation, in particular when it is composed of API sensitive to acidic media or proteins (56).

After exposition to the release medium, porosity and size of the system will impact the degradation. Higher degradation rates have been reported for the largest devices (57)(58)(59)(60)(61)[START_REF] Klose | How porosity and size affect the drug release mechanisms from PLGA-based microparticles[END_REF][START_REF] Kapoor | PLGA: a unique polymer for drug delivery[END_REF][START_REF] Chen | Effect of Particle Size on Drug Loading and Release Kinetics of Gefitinib-Loaded PLGA Microspheres[END_REF].

Indeed, Figure I-10 illustrates the size-dependent autocatalytic effect. For larger devices, there is a region (illustrated in dark orange) in which acidic degradation products can freely diffuse through the polymer matrix (blue arrow). Thus, leading to an acceleration of the hydrolysis reaction in the light orange region, due to autocatalysis (52,[START_REF] Kapoor | PLGA: a unique polymer for drug delivery[END_REF][START_REF] Klose | PLGA-based drug delivery systems: importance of the type of drug and device geometry[END_REF]. The degradation rate is also directly influenced by the ratio of monomers. As mentioned above, GA is more hydrophilic than LA, making it more prone to hydrolysis than LA. But several publications reported that the rate of degradation is highest when the ratio LA:GA is 50:50 (36,44).

x n

+ H 2 O + H + x m + x n-m + H +
Other factors must also be considerate such as the pH, the addition of API with plasticizing or anti-plasticizing effect, the addition of acidic or basic API, mechanical constraints, manufacturing process, and sterilization of the system (49,[START_REF] Hamoudi-Ben Yelles | PLGA implants: How Poloxamer/PEO addition slows down or accelerates polymer degradation and drug release[END_REF][START_REF] Casalini | A Systematic Experimental and Computational Analysis of Commercially Available Aliphatic Polyesters[END_REF].

I.2.6. Biocompatibility

Food and Drug Administration (FDA) use the definition of biocompatibility gave by J. Black in 2006 as a reference, where it is defined as the ability of a device material to perform with an appropriate host response in a specific situation [START_REF] Black | Biological Performance of Materials: Fundamentals of Biocompatibility[END_REF][START_REF] Health | Use of International Standard ISO 10993-1[END_REF].

PLGA has been reported to have excellent biocompatibility [START_REF] Nair | Biodegradable polymers as biomaterials[END_REF][START_REF] Kumari | Biodegradable polymeric nanoparticles based drug delivery systems[END_REF][START_REF] Grund | Polymers in Drug Delivery-State of the Art and Future Trends[END_REF][START_REF] Zhao | Poly(lactic-co-glycolic acid)-based composite bone-substitute materials[END_REF]. Indeed, as illustrated in Thus, leads to a very low systemic toxicity, and makes PLGA an excellent polymer for drug delivery and biomaterial application [START_REF] Maurus | Bioabsorbable implant material review[END_REF][START_REF] Kumari | Biodegradable polymeric nanoparticles based drug delivery systems[END_REF].

I.2.7. Use in control drug delivery

PLGA is frequently used as a polymer matrix in controlled release systems such as microparticles, nanofibers, or implants [START_REF] Kapoor | PLGA: a unique polymer for drug delivery[END_REF][START_REF] Grund | Polymers in Drug Delivery-State of the Art and Future Trends[END_REF][START_REF] Anderson | Biodegradation and biocompatibility of PLA and PLGA microspheres[END_REF]. These offer many advantages: a minimal inflammatory response, an appropriate release rate depending on the application, an in-situ degradation compatible with the healing and/or regeneration process [START_REF] Kapoor | PLGA: a unique polymer for drug delivery[END_REF][START_REF] Grund | Polymers in Drug Delivery-State of the Art and Future Trends[END_REF][START_REF] Anderson | Biodegradation and biocompatibility of PLA and PLGA microspheres[END_REF].

Since 1989, the first approval of a product containing PLGA by the FDA, many specialties of parenteral products composed of PLGA have been approved for human use, such as microspheres, implants, or in situ forming implants. These products, the API delivered, the dosage form, the duration of drug release, and the company are summarized in Table I-4. Indeed, the interpretation of the observed results is not straightforward, and the underlying mass transport phenomena controlling drug release are often difficult to elucidate (35).

I.2.8. Release mechanisms and release pattern

I.2.8.1. Definitions

Understanding the mechanisms that control the release of API from PLGA based systems is far from trivial. Indeed, some hypotheses are described in the literature but many physicochemical phenomena are involved and remain to be elucidated.

API release mechanisms could be defined either as: (i) how the API is released from the system, or (ii) the phenomenon that controls its release rate [START_REF] Sansdrap | In vitro evaluation of the hydrolytic degradation of dispersed and aggregated poly(dl-lactide-co-glycolide) microspheres[END_REF][START_REF] Kranz | Physicomechanical properties of biodegradable poly(D,L-lactide) and poly(D,L-lactide-co-glycolide) films in the dry and wet states[END_REF]. So far, several phenomena controlling the release rate of an API from PLGA devices have been reported such as [START_REF] Shah | A biodegradable injectable implant for delivering micro and macromolecules using poly (lactic-coglycolic) acid (PLGA) copolymers[END_REF][START_REF] Park | Degradation of poly(lactic-co-glycolic acid) microspheres: effect of copolymer composition[END_REF][START_REF] Gasper | Formulation of L-asparaginase-loaded poly(lactide-co-glycolide) nanoparticles: influence of polymer properties on enzyme loading, activity and in vitro release[END_REF][START_REF] Zhu | Stabilization of proteins encapsulated in injectable poly (lactide-co-glycolide)[END_REF][START_REF] Jonnalagadda | A bioresorbable, polylactide reservoir for diffusional and osmotically controlled drug delivery[END_REF][START_REF] Wong | In Vitro Sustained Release of Human Immunoglobulin G from Biodegradable Microspheres[END_REF][START_REF] Friess | Release mechanisms from gentamicin loaded poly(lactic-coglycolic acid) (PLGA) microparticles[END_REF][START_REF] Bishara | PLA stereocomplexes for controlled release of somatostatin analogue[END_REF][START_REF] Matsumoto | Drug release characteristics of multireservoir type microspheres with poly(dl-lactide-co-glycolide) and poly(dl-lactide)[END_REF][START_REF] Kim | Biodegradable polymeric microspheres with 'open/closed' pores for sustained release of human growth hormone[END_REF][START_REF] Kang | Pore closing and opening in biodegradable polymers and their effect on the controlled release of proteins[END_REF][START_REF] Sun | Synchronic release of two hormonal contraceptives for about one month from the PLGA microspheres: in vitro and in vivo studies[END_REF][START_REF] Mochizuki | Controlled release of argatroban from PLA film-Effect of hydroxylesters as additives on enhancement of drug release[END_REF][START_REF] Gagliardi | Combined drug release from biodegradable bilayer coating for endovascular stents[END_REF]35,[START_REF] Siepmann | Mathematical modeling of drug delivery[END_REF]48,49,[START_REF] Gasmi | Towards a better understanding of the different release phases from PLGA microparticles: Dexamethasone-loaded systems[END_REF][START_REF] Bode | In-situ forming PLGA implants for intraocular dexamethasone delivery[END_REF]33,25,54,16) -API dissolution, -API diffusion through water-filled pores, -API diffusion through the polymer matrix, -Polymer hydrolysis,

I.2.8.2. Diffusion throughout a porous network

This mechanism describes the first release step, before the erosion of the polymer begins. It is very dependent on the porous structure of the device and is therefore dependent on the processes that promote pore formation [START_REF] Zidan | Formulation of anastrozole microparticles as biodegradable anticancer drug carriers[END_REF]. These must be continuous and large enough for the API to pass through (35). Transport through water-filled pores can be achieved either by diffusion when the concentration gradient controls transport, and/or convection when osmotic pressure controls the transport (35).

I.2.8.3. Diffusion throughout the polymer matrix

This mechanism concerns low molecular weight hydrophobic API [START_REF] Wischke | Principles of encapsulating hydrophobic drugs in PLA/PLGA microparticles[END_REF]35). The diffusion rate is very dependent on the physical state of the polymer. It can increase as the polymer changes from a glassy state to a rubbery state and unlike diffusion through water-filled pores, it is not

(A) (B) (C)
dependent on the pore structure (46). Diffusion is often higher in low molecular weight polymers, due to higher mobility of polymer chains [START_REF] Faisant | PLGA-based microparticles: elucidation of mechanisms and a new, simple mathematical model quantifying drug release[END_REF].

I.2.8.4. Degradation of polymer matrix

Matrix degradation/erosion is the process of cleavage of polymer chains into oligomers and monomers [START_REF] Göpferich | Mechanisms of polymer degradation and erosion[END_REF]. In the first case, erosion occurs only at the surface, due to a weak diffusion of water into the polymer [START_REF] Li | Further investigations on the hydrolytic degradation of poly (DLlactide)[END_REF]. In contrast, in the second case, degradation occurs more slowly and throughout the system. PLGA-based devices are generally considered to be systems undergoing erosion mass [START_REF] Klose | How porosity and size affect the drug release mechanisms from PLGA-based microparticles[END_REF][START_REF] Grund | Polymers in Drug Delivery-State of the Art and Future Trends[END_REF].

I.2.8.5. Release pattern

Different types of release profiles can be observed from PLGA devices. It is not always obvious to know which are the mechanisms of release which dominate during the phases previously described (35,100). Indeed, the release profiles are often due to the combination of several of the mechanisms described above.

I.2.8.5.1. Monophasic release pattern

This profile consists of, one single release phase, sought after in most applications to avoid the "burst" effect that can cause toxic effects. In most cases the release of API is constant and is only controlled by a diffusion process. Wang et al. have shown that zero-order kinetics can be obtained from microparticles prepared by the simple solvent evaporation emulsion technique (101). Monophasic release of API from PLGA devices is rare. It is mainly done in a biphasic or triphasic way (35).

I.2.8.5.2. Bi-phasic release pattern

This profile consists of two release phases. In most of the cases it is divided in: (i) the first phase called the "burst" effect, (ii) followed by a plateau, or in (i) zero order release, follow by a (ii) second phase of rapid release (102,103). For many formulations, an initial rapid release of API is observed before the release rate reaches a plateau (102). This phenomenon is generally called the "burst" effect. This can be defined as the amount of API released by the device before polymer erosion begins (103).

-Phase II:

Then follows a phase where the release rate of API is constant, called a plateau, during which the slowing of the release due to the increase in diffusion paths is compensated by the erosion of the polymer. The water which has penetrated inside the matrix induces the hydrolysis of the polymer chains into oligomers and water-soluble monomers (35,39,[START_REF] Faisant | PLGA-based microparticles: elucidation of mechanisms and a new, simple mathematical model quantifying drug release[END_REF].

I.2.8.5.3. Tri-phasic release pattern

This release profile also exhibits a "burst" effect followed by a second phase which corresponds to slow diffusion of the API through the polymer matrix and the pores. The third phase is the result of the degradation and/or the swelling of the polymer matrix which results in a rapid release of the API (48,49,33,25,16).

-Phase I:

As described above, the "burst" effect is mainly attributed to the diffusion of the SA which is adsorbed on the surface of the particle, or to the diffusion of the API through pores filled with water in contact. directly with the surface of the latter (33). Previous work suggests that the "burst" period ends when the pores close [START_REF] Kim | Biodegradable polymeric microspheres with 'open/closed' pores for sustained release of human growth hormone[END_REF][START_REF] Kang | Pore closing and opening in biodegradable polymers and their effect on the controlled release of proteins[END_REF]104).

-Phase II:

This phase, also called the lag phase, is characterized by slow diffusion of the API through the polymer matrix and the few pores filled with water (105). During this period, hydration of the device as well as degradation of the polymer takes place. This phase may also be due to pore closure and polymer-API interactions, which could limit the release of API. Several factors have been reported to induce pore closure, among them: degradation of the polymer, plasticizers, and high temperatures (106,107). In some systems, this lag phase is negligible due to the rapid degradation of the PLGA. The duration of this phase depends mainly on the characteristics of the polymer, the size, and the geometry of the device (57,60,[START_REF] Klose | How porosity and size affect the drug release mechanisms from PLGA-based microparticles[END_REF][START_REF] Chen | Effect of Particle Size on Drug Loading and Release Kinetics of Gefitinib-Loaded PLGA Microspheres[END_REF].

Gasmi et al. showed that this second phase of release depended mainly on the initial SA content.

They explained that at high levels, part of the API which does not have direct access to the surface of the device is trapped by the PLGA and takes time to diffuse through the polymer matrix [START_REF] Gasmi | Towards a better understanding of the different release phases from PLGA microparticles: Dexamethasone-loaded systems[END_REF]. In other studies, API is released by the swelling of the device due to the diffusion of water. Indeed, water diffuses inside the device causing the dissolution of the API and its diffusion into the release medium (108).

-Phase III:

This phase is characterized by a faster release of AS. The release during this phase is mainly due to massive erosion of the polymer and swelling and/or deformation of PM (35). The beginning of this phase is when an entirely continuous porous network is formed inside the particle. The large swelling observed at the end of the second phase can result from the osmotic pressure built up in the system and begins as soon as the polymeric structure reaches a sufficiently molecular weight low (48,49,25,54,58,35,33,16) In this thesis, the work will mainly focus on PLGA based implants for subcutaneous delivery.

Several techniques for manufacturing implants have been reported, such as compression, solvent casting, injection molding, hot-melt extrusion, and more recently, three-dimensional (3D) printing (109). I -5(115,121-125). Operating mode, advantages and disadvantages of the ME technologies have been summarized

I.3. Three

in Table I-7.

I.3.2.4. Material jetting

The Material Jetting (MJ) 3D printing technique is often compared to the standard 2D inkjet process. The use of photopolymers, metals, or waxes that solidify when exposed to light or heat (like SLA) leads to the manufacturing of 3D objects layer by layer. MJ, allows to 3D prints different materials(120,121,142-147).

The printer simultaneously projects material and another so-called "support" soluble material in the form of micro-droplets. These could be cured by UV light and then form the first layer.

The operation is repeated to form the entire object as illustrated in Only the parts created from the building material remain.

Suitable for metals and ceramics.

A limited number of materials.

Slow process, high cost.

(147) The main advantages and disadvantages of BJ 3D printing technologies are summarized in

I.3.2.5. Binder jetting

As illustrated in

I.3.2.6. Direct energy deposition

Direct Energy Deposition (DED) are 3D printing technologies in which an energy source (usually a laser beam) is focused on an area where material deposition is occurring simultaneously. This material, which may be in the form of a filament or a powder, is then melted and deposited at the same time as illustrated in As to take place in a sealed chamber filled with argon gas.

Used technology

Results Ref

Photopolymerization

SLA

SLA printing was used to prepare ibuprofen-loaded hydrogels of cross-linked polyethylene glycol diacrylate. Dissolution profiles showed that drug release rates were dependent on water content, with higher water content hydrogels releasing drug faster.

(156)

SLA

SLA was used to create custom tablet geometries using a novel biocompatible photochemistry consisting of ascorbic acid encapsulated in a poly(ethylene glycol) dimethacrylatebased polymer network and polymerized using riboflavin as a photoinitiator. Different geometries were printed with different surface area to volume ratios. After 1 h of release, honeycomb and coaxial annulus tablet gels exhibit higher release rates at approximately 80%.

(157)

DLP

Silk fiboin-based microneedles were fabricated using DLP 3D printing using riboflavin, a vitamin that acts as a photoinitiator. Delivery of molecules to the skin is verified using rhodamine B fluorescence dye.

(126)

CLIP

CLIP technology is used to fabricate biocompatible and drug-loaded devices with controlled release properties, using liquid resins containing rhodamine B-base, docetaxel and dexamethasone-acetate. Formulations were shown to be biocompatible over the course of 175 days of in vitro degradation and the clinically-relevant drugs could be encapsulated and released in a controlled fashion.

(

) 158
PBF SLS SLS printing is used for manufacturing medicines with two thermoplastic pharmaceutical grade polymer, and three different drug loading of paracetamol. One of the polymers showed, showed pH-independent release characteristics, while the other showed pH-dependent, modified-release profiles independent of drug loading, with complete release being achieved over 12h.

(159)

SLM

SLM was employed to manufacture Ti-6Al-4 V samples, with internal reservoirs and releasing Micro-channels (MCs) to simulate what could be a drug-delivering orthopedic or dental implant.

(160)

SLM

Drug release of Vancomycin is enabled by manufacturing of integrated permeable structures possessing high porosity through application of selective laser melting technology. (161) SLM Titanium components fabricated using SLM were functionalized with Paracetamol using phosphonic acid based self-assembled monolayers. Thus, the proposed method has the potential to incorporate drugs to titanium coated surfaces and improve their biocompatibility and reduce post-implant complications.

(162)

ME PE/SE

Bi-layered tablets containing Guaifenesin, Polyacrylic acid, microcrystalline cellulose, and sodium starch glycolate were printed using semi solid extrusion. Drug was release following a Fickian diffusion release pattern through hydrated polymer gel layer.

(163)

PE/SE

Multi-active tablets (nifedipine, glipizide,captopril) mixed hydroxypropyl methylcellulose were printed using semisolid extrusion. Both nifedipine and glipizide showed sustained-release.

(

) 164
FDM/FFF FDM/FFF 3DP technology allow to manufacture tablet shapes of different geometries, that would be challenging to manufacture by conventional pharmaceutical technologies. Drug release from the tablets was not dependent on the surface area but on surface area to volume ratio.

(

) 135
FDM/FFF 3D scanning was used to construct 3D models of a nose and ear to provide the opportunity to customize shape and size of a wound dressing to an individual patient.

(165) DPE 3D printing by avoiding the need for preparation of filaments by hot melt extrusion. This novel single-step technology could revolutionize the preparation of amorphous solid dispersions as final formulations and it may be especially suited for preclinical studies, where the quantity of drugs is limited and without the need of using traditional hot melt extrusion.

(136)

DDM

DDM can be used to manufacture drug delivery devices of varying geometries, densities and surface areas to control the drug release. This work presents a new opportunity to increase the release of Dapivirine.

(138)

MJ DOD

The functions of the bioprinter have been demonstrated by using a high-throughput drug-delivery model; and by bioprinting micro-tissues using a variety of different cell types, DOD demonstrates a promising platform for generating many types of tissues and drug-delivery models.

(166)

Polyjet

The tablets produced consist of a cross-linked poly(ethylene glycol diacrylate) hydrogel matrix containing the drug, photoinitiated in a low oxygen environment using an aqueous solution.

(167)

BJ

ZipDose ®

A high dose rapidly dispersing three-dimensionally printed dosage form comprising a high dose of levetiracetam in a porous matrix that disperses in water within a period of less than about 10 seconds.

(168,1 69)

BJ

Fabrication of donut shaped tablets based on mixture of paracetamol and polymer powder by BJ technology. A monophasic release kinetic was obtained (170)

I.3.3.2. Interest in control release

3D printing technologies offer high flexibility in the preparation of dosage forms such as combining different APIs or reaching complex release profiles (171). This is making 3D

printing promising technologies for revolutionizing controlled drug delivery systems (172). As

I.3.3.2.2. Effect of the internal structure

Examples of the effect of the internal structure on drug release are illustrated in Figure I-28. In these studies, the authors investigated the impact of porosity (percentage, and shape of the pore) on drug release rate. They exhibited that higher porosity leads to a quicker drug release rate by modulating the surface volume ratio (138,181,182).

Five different tablet geometries (cube, pyramid, cylinder, sphere and torus) were printed. Drug release from the tablets was not dependent on the surface area but instead on surface area to volume ratio. 3D printing was used to prepare hollow bullet-shaped implant with different porous surface for controlled cytoxan release

(A)
The impact of geometrical parameters of 3D printed constructs (i.e., porosity, pore shape) on drug release was investigated → Porosity modulates the surface/volume ratio, and higher porosities lead to quicker drug release rates

(B)
10% density vaginal ring exhibited a 7 times fold increase in daprivine release compared to 100% density Multicompartmental capsules were 3D printed → allow to modulate drug release.

(C)

One compartment capsule

Two compartments capsule Three compartments capsule

I.3.3.3. Poly(lactic-co-glycolic) acid

Among all the studies conducted on the suitability of 3D printing technologies in developing new controlled drug delivery devices, only a few of them focus on PLGA as a potential carrier.

Some studies have been published on the investigation of the interest of the use of PLGA and 3D printing technologies for bones repair (186-189) .

I.4. Research objectives

PLGA based devices are a first category choice for polymer-based long-acting injectables due to their good biocompatibility, biodegradability, and use in several FDA-approved products.

However, establishing an in vitro in vivo correlation regarding drug release of such devices is challenging as many different mechanisms are involved in drug release. Thus, it is of great importance to properly characterize the devices, to identify the release mechanism by which the API is released from the device.

In the past few years, 3D printing technologies have revolutionized the development of new controlled drug release devices. These technologies allow to precisely control the size, shape, internal and external porosity of the device, and so the release rate of the API.

In this context the research objectives of this thesis were dived into two parts:

(i) Development of experimental set-ups for in vitro drug release measurement for implants administrated to the subcutis, and investigation on the impact of several implants attribute on in vitro drug release;

(ii) Development of PLGA based controlled drug release implants using 3D printing technologies, and investigation of the impact of the structure of the implant on in vitro drug release.

The research results of this Ph.D. thesis will be described in seven parts corresponding to scientific papers that have either been submitted in scientific journals or plan to be:

1. How hydrogels surrounding PLGA implants limit swelling and slow down drug release 2. Hot-melt extruded ibuprofen-loaded PLGA implants: Importance of heat exposure The aim of this study was to better understand to which extent the presence of an agarose gel (mimicking living tissue) around a PLGA [poly(lactic-co-glycolic acid)] implant affects the resulting drug release kinetics. Ibuprofen-loaded implants were prepared by hot melt extrusion.

Drug release was measured upon exposure to phosphate buffer pH 7.4 in Eppendorf tubes and upon inclusion into an agarose gel which was exposed to phosphate buffer pH 7.4 in Eppendorf tubes or in transwell plates. Dynamic changes in the implants' dry and wet mass as well as dimensions were monitored by optical macroscopy and gravimetrically. SEM and GPC were used to follow system erosion and polymer degradation. Different pH indicators were applied to measure pH changes in the bulk fluids, gels and within the implants during drug release.

Ibuprofen release was bi-phasic in all cases: A zero order release phase (~20 % of the dose) was followed by a more rapid, final drug release phase. Interestingly, the presence of the hydrogel delayed the onset of the 2nd release phase. This could be attributed to the sterical hindrance of implant swelling: After a certain lag time, the degrading PLGA matrix becomes sufficiently hydrophilic and mechanically instable to allow for the penetration of substantial amounts of water into the system. This fundamentally changes the conditions for drug release:

The latter becomes much more mobile and is more rapidly released. A gel surrounding the implant mechanically hinders system swelling and, thus, slows down drug release. These observations strengthen the hypothesis of the "orchestrating" role of PLGA swelling for the control of drug release and can help developing more realistic in vitro release set-ups.

Key words: PLGA implant; release mechanism; swelling; ibuprofen; hydrogel

II.2. Introduction

Poly(lactic-co-glycolic acid) (PLGA) offers an interesting potential as matrix former for controlled parenteral drug delivery (1-3). In particular, PLGA microparticles and implants have been proposed in the literature (4)(5)(6).Several drug products are on the market since decades.

The great success of this polymer for this type of applications can be attributed to its good biocompatibility (7), complete biodegradability (avoiding the removal of empty remnants upon drug exhaust) and the possibility to provide desired drug release kinetics during flexible periods of time (8)(9)(10)(11). The resulting drug release kinetics depend on a variety of factors, including the type of PLGA (e.g., average polymer molecular weight, type of end groups), composition of the system as well as the type of manufacturing method and parameters used during processing (12)(13)(14).

Despite the great practical importance of PLGA-based drug delivery systems, yet the underlying mass transport mechanisms controlling drug release are generally not fully understood. This can be explained by the potential complexity of the involved physico-chemical processes (15,16), including for example water penetration into the systems, polymer degradation (17), drug dissolution, drug diffusion, polymerdrug interactions (18), waterpolymer interactions (19), the creation of a local acidic micro-environment within the dosage form (20)(21)(22) causing autocatalytic effects (23), the closure of surface pores (24), substantial system swelling (25), limited drug solubility effects and system disintegration. The relative importance of these phenomena can strongly depend on the specific composition and inner & outer structure of the delivery system. For instance, the extent of local drops in the micro-pH can be altered by the addition of basic excipients or by varying the initial device porosity (determining at which rate acids and bases can diffuse into and out of the system).

Often, mono-, bi-, or tr-phasic drug release profiles are observed from PLGA-based drug delivery systems, irrespective of their geometry and size (e.g., implants, microparticles and films) (26-28). In the case of tri-phasic drug release, the following phases can generally be distinguished: An initial burst release phase (frequently limited to the first few hours or 1-2 days) is followed by a "zero order drug release phase" (with an about constant release rate), and a final, again rapid release phase leading to complete drug exhaust. It has recently been proposed that PLGA swelling might play a decisive, "orchestrating" role for these different release phases (29)(30)(31), although PLGA swelling is often neglected in the literature for the explanation of drug release patterns. The initial burst release might be attributable to the release of drug particles which come into direct contact with water once the system is exposed to an aqueous fluid, because they are located directly at the system's surface, or very close to it.

Eventually, this contact is assured via tiny pores with direct surface access or via an interconnected drug particle network. Since drug dissolution is rapid (as well as diffusion through short, water-filled pores), this release phase often ends during the first 1-2 days.

Potentially, (limited) PLGA swelling can also close surface pores, terminating the initial burst release phase (32).

It is well known that water penetration into PLGA-based drug delivery systems is relatively rapid, so that the entire device is wetted and polymer degradation occurs throughout the system ("bulk erosion") (33). The possible root causes for the second (zero order release) phase are often less well understood. It has been suggested that in the case of microparticles consisting of a PLGA matrix into which tiny diprophylline crystals were dispersed, the continuous growth of a highly swollen surface layer plays a crucial role (34): Since the PLGA at the system's surface is in contact with very high amounts of water, polymer degradation can be expected to be accelerated in this region. Upon ester bond cleavage shorter chain acids and alcohols are generated, rendering especially the matrix in the outmost layer more and more hydrophilic. At a certain time point, the latter undergoes substantial swelling. With time its thickness increases, because also "deeper" polymer layers get exposed to very high water concentrations. As long as a diprophylline particle is surrounded by a dense PLGA matrix, it cannot dissolve (lack of water) and diffuse out (lack of mobility). However, once the steadily growing, highly swollen surface layer reaches the particle, the latter can dissolve and dissolved drug molecules can rather rapidly move through the highly swollen PLGA gel. This phenomenon was evidenced for single microparticles, occurring at random time points: Each particle has its own structure and its own way to release the drug. Since the drug was homogeneously distributed throughout the system, the numerous occasional individual drug release events summed up to an about constant drug release rate.

The final, again rapid drug release phase (3 rd phase) observed with certain PLGA microparticles and implants could be attributed to the onset of substantial system swelling occurring after a lag time. Due to PLGA degradation, the entanglement of the polymer chains decreases with time and the concentration of water-soluble degradation products increases (creating a continuously increasing osmotic pressure inside the device). In addition, the system becomes more and more hydrophilic (due to the newly created -COOH and -OH end groups). At a certain time point, the mechanical stability of the macromolecular network becomes insufficient and substantial amounts of water penetrate into the entire device, allowing for drug particle dissolution and significantly increased drug mobility.

In certain cases, only bi-or mono-phasic drug release patterns are observed from PLGA-based drug delivery systems. This might be due to the fact that no drug has direct initial access to the system's surface (absence of a burst release); or all drug is already released before the substantial entire system swelling sets on (no 3 rd drug release phase).

If the hypothesis of an "orchestrating role" of PLGA swelling for the control of drug release is correct, the presence of a hydrogel surrounding the device might have a non-negligible effect on drug release, since it might hinder the onset of substantial system swelling. In vivo, the living tissue surround the implant can be expected to have a similar effect. However, in most cases experimental set-ups used for in vitro drug release measurements from PLGA-based drug delivery systems expose the dosage form directly to a bulk fluid. In the literature only relatively few studies address the potential impact of the presence of a gel. For example, in an interesting recent report, the group of Lamprecht measured the release of flurbiprofen, lidocaine and risperidone from ethylcellulose or PLGA-based films, microparticles and cylindrical implants in the presence and absence of a surrounding gel (35). In many cases, the presence of the gel led to slower drug release from PLGA-based devices. However, in the case of certain flurbiprofen-loaded films the release rate became faster during most of the release period, the gel hindering the films to deform. And for some other systems, the impact of the presence of a surrounding gel was negligible. The same group proposed the use of components of muscle tissue to better mimic intramuscular environments and addressed the potential role of lipids for drug release from a variety of controlled release microparticles (36). Allababidi and Sha (37) investigated the release of cefazolin from glycerol monostearate-based implants into an agar gel or phosphate buffer pH 7.4 bulk fluid, and did not observe major differences. Furthermore, the release of profiles of ciprofloxacin hydrochloride and vancomycin hydrochloride from different types of hydroxyapatite implants functionalized with hydroxypropyl--cyclodextrin were measured into agarose gels or well agitated phosphate buffer pH 7.4. Drug release was much faster in the agitated bulk fluid, which was at least in part attributed to accelerated matrix erosion. Hydrogels have also been proposed by the group of Ostergaard as surrogates for subcutaneous tissue when studying controlled release implants (38,39) The aim of this study was to investigate ibuprofen-loaded PLGA implants using

II.3.2. Implant preparation

PLGA was milled for 4 x 30 s in a grinder (Valentin, Seb, Ecully, France). Appropriate amounts of PLGA and drug powders were blended for 5 min at 20 rpm in a Turbula T2C Shaker-Mixer (Willy A Bachofen, Basel, Switzerland). Three hundred mg mixture were filled into a 1 mL syringe (Henke Sass Wolf, Tuttlingen, Germany), followed by heating at 105 °C for 15 min in an oven (FP115, Binder, Tuttlingen, Germany). The molten blend was manually extruded using the syringe. The obtained extrudate was cut with a hot scalpel into cylindrical implants of approximately 5 mm length.

II.3.3. Practical drug loading

Implants were dissolved in acetonitrile (1 implant in 5 mL), followed by filtering (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000 Series HPLC, equipped with a LPG 3400 SD/RS pump, an auto sampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversed phase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used.

The mobile phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v). The detection wavelength was 264 nm, the flow rate was set at 0.5 mL/min. Ten microliter samples were injected.

II.3.4. In vitro drug release

Ibuprofen release from the PLGA implants was measured using 3 different experimental setups:

II.3.4.1. In well agitated bulk fluids

Implants were placed in metal baskets in 5 mL Eppendorf tubes (1 implant per basket/tube), filled with 5 mL phosphate buffer pH 7.4 USP 42 (Figure II-1 A). Optionally, one of the following pH indicators was added: bromocresol green, bromothyol blue, bromophenol blue or phenol red (0.0025 % w:v). The tubes were placed in a horizontal shaker (80 rpm, 37°C; GFL 3033; Gesellschaft fuer Labortechnik, Burgwedel, Germany). At predetermined time points, the entire bulk fluid was replaced by fresh (pre-heated) release medium. The withdrawn samples were filtered (PVDF syringe filter, 0.45 µm; Agilent, Santa Clara, California, USA) and analyzed for their ibuprofen contents by HPLC-UV, as described in section II.3.3.

II.3.4.2. In agarose gels in Eppendorf tubes

Implants were embedded in agarose gels in 5 mL Eppendorf tubes, as illustrated in

II.3.4.3. In agarose gels in transwell pates

Implants were embedded in agarose gels in transwell plates (1 implant per insert, 1 mL gel, membranes: 1.13 cm², 11 µm, 0.4 µm pore size; Nunc, Roskilde, Denmark), as illustrated in

Figure II-1 C.
The agarose gels were prepared as described above, and the implants included accordingly (placed between 2 "layers" of 0.5 mL gel). The well plates were filled with 4 mL phosphate buffer pH 7.4 USP 42 (optionally containing 0.0025 % pH indicator, as described above), covered with lids and Parafilm to minimize evaporation, and placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the entire bulk fluid was replaced by fresh (pre-heated) release medium. The withdrawn samples were treated as in the case of drug release measurements in well agitated bulk fluids.

In all cases, the pH of the release medium was measured at pre-determined time points using a pH meter (InoLab pH Level 1; WTW, Weilheim, Germany). Furthermore, in all cases, sink conditions were provided throughout the experiments in all agitated bulk fluids. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

II.3.5. Implant swelling

Implants were treated as for the in vitro drug release measurements described in section II.3.4.

At pre-determined time points:

(i) Pictures of the implants were taken with an a SZN-6 trinocular stereo zoom macroscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vison Lite 2.1 software). Cross-sections were obtained by cutting with a scalpel.

The lengths and diameters of the implants were determined using the ImageJ software (US National Institutes of Health). Dynamic changes in the systems' volume were calculated considering cylindrical geometry.

(ii) Implant samples were withdrawn and excess water was carefully removed using Kimtech precision wipes (Kimberly-Clark, Rouen, France). The samples were weighed [wet mass (t)], and the change in wet mass (%) (t) was calculated as follows:

𝑐ℎ𝑎𝑛𝑔𝑒 𝑖𝑛 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (%)(𝑡) = 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (𝑡) -𝑚𝑎𝑠𝑠 (𝑡 = 0) 𝑚𝑎𝑠𝑠 (𝑡 = 0) × 100 %
where mass (t = 0) denotes the implant mass before exposure to the release medium.

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

II.3.6. Implant erosion and PLGA degradation

Implant samples were treated as for the in vitro drug release studies described in section II. 3

II.3.8. Scanning electronic microscopy (SEM)

The internal and external morphology of the implants before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxfordshire, UK). Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome layer.

In the case of implants which had been exposed to the release medium, the systems were treated as described for the in vitro release studies in section II. 3.4. At predetermined time points, implant samples were withdrawn, optionally cut using a scalpel and freeze-dried (as described in section II. 3.6).

II.4. Results and discussion

The image in Table II-1 shows an optical macroscopy picture of an ibuprofen-loaded PLGA implant before exposure to the release medium. The practical drug loading was 6.6 +/-0.3 %.

The glass transition temperature (Tg) of the implants was determined to be 34.5 +/-0.3 °C, while the Tg of the PLGA raw material was 47.1 +/-0.1 °C. This indicates that ibuprofen acts as a plasticizer for PLGA.

II.4.1. In vitro release set-ups

The schemes in measured. Thus, to be "detected as released", drug molecules/ions released from the implant also have to cross the gel. To evaluate the impact of the presence of this additional mass obtained results, indicating that the impact of the presence of the gel was similar to the impact of the gel in the "gel -Eppendorf" set-up. This is fully sound, since the compositions of the gels are identical and the distances to be overcome similar. Hence, again, the observed delay in drug release due to ibuprofen transport through the agarose gel is relatively small compared to the much longer release periods from the investigated PLGA implants.

Importantly, all 3 experimental set-ups guaranteed sink conditions throughout the experiments in this study in the agitated bulk fluids. Interestingly, all release profiles were "bi-phasic": A zero order release phase (with an about constant release rate) was followed by a more rapid drug release phase, leading to complete drug exhaust. No noteworthy "burst release" was observed, irrespective of the experimental setup. This can probably be explained by the fact that the implants were prepared by hot melt extrusion, leading to a non-porous surface, as also evidenced by SEM (e.g. pictures at the top on the left hand side of Figure II-3. Furthermore, in contrast to small PLGA microparticles, the release of minor absolute amounts of drug from surface-near regions of a "large" implant at can be seen that the onset of important system swelling coincided with the onset of the final, rapid drug release phase in all cases. The observed ranking orders for the "lag time" for substantial system swelling and for the onset of the final rapid drug release phase were the same: "bulk fluid" < "gel -transwell" ≈ "gel -Eppendorf" set-up.

The difference in the swelling kinetics of the implants upon exposure to a "bulk fluid" versus "gel" can probably be attributed to the sterical hindrance caused by the agarose matrix: Once the PLGA implants are in contact with the bulk fluid or gel, water penetrates into the system and the entire implant is rather rapidly wetted. Since the investigated PLGA is relatively hydrophobic, the amounts of water diffusing into the implants at early time points remain limited. But these "low" amounts of water are sufficient to initiate polymer degradation throughout the device ("bulk erosion"). Consequently, the macromolecules become shorter and less entangled with time. Also, since the newly created end groups (upon ester bond hydrolysis) are hydrophilic (-COOH and -OH), the polymer matrix becomes more and more hydrophilic.

In addition, the concentration of water-soluble degradation products (short chain acids) is steadily increasing, generating a continuously increasing osmotic pressure inside the implant. release, its role might be considered as "orchestrating" (31). Please note that the dissolution of drug particles is probably not playing a role in the investigated implants, since at a practical loading of 6.6 %, the ibuprofen is likely completely dissolved in the PLGA matrix from the beginning ("monolithic solution") (40). During drug release, the polymer can be expected to be highly swollen in surface-near regions, "gel -transwell" > "gel -Eppendorf" > "bulk fluid" set-up. This drop can at least partially be attributed to the release of short chain, water-soluble acids (as degradation products of PLGA) into the bulk fluids: As discussed above, once the implants become sufficiently hydrophilic and mechanically instable, substantial system swelling sets on. This does not only fundamentally change the conditions for drug release, but also for the release of these water-soluble acids.

Please note that this temporary drop in pH can also (in part) be attributed to the longer sampling interval (3 d "week-end gap", compared to daily sampling during the week; at each sampling time point, the entire bulk fluid was renewed). Thus, the water-soluble acids accumulated during the longer sampling interval. However, as it can be seen, the following 3 d sampling gap Implant samples were treated as for the in vitro drug release studies. After pre-determined exposure periods, optical macroscopy pictures were taken. The dotted red lines highlight the most informative images at each time point. In the case of the "gel -Eppendorf" set-up, the pH in the gel surrounding the implant remained above 6.6 during the first week, and then temporarily dropped: to pH 6.0-6.5 on day 10 and to pH 5.4-6.0 on day 14. It subsequently raised again. In the "gel -transwell" set-up a similar behavior was observed. These drops are consistent with the pH drops observed in the agitated bulk fluids used in these set-ups (which were discussed above). They can mainly be attributed to the release of short chain acids from the implants after the onset of substantial system swelling, and (in part) to the accumulation of the acids during the longer (3 d) sampling interval. Since the solubility of ibuprofen is pHdependent, local acidic environments around the implant might limit drug solubility and, hence, slow down drug release. However, Kozac et al. (35) reported slower release from PLGA-based films and microparticles surrounded by agarose gel compared to agitated bulk fluid also for the free base lidocaine, which is more soluble at acidic pH. during the first week, and then decreased, due to the release of water-soluble PLGA degradation products and drug. Importantly, the erosion rate was higher in the absence of an agarose gel.

These observations are in good agreement with the hypothesized drug release mechanism:

During the first few days, the PLGA network is still highly entangled and the small amounts of generated short chain acids and drug are poorly mobile, resulting in negligible dry mass loss.

However, once substantial amounts of water are present in the system, drug and water-soluble degradation products are much more rapidly released. Since the presence of the agarose gel delays implant swelling, also the dry mass loss of the implants is delayed compared to the "bulk fluid" set-up.

The diagram at the bottom of

III. HOT MELT EXTRUDED IBUPROFEN-LOADED PLGA IMPLANTS: IMPORTANCE OF HEAT EXPOSURE III.2. Introduction

Poly(lactic-co-glycolic acid) (PLGA) is frequently used a matrix former for controlled drug delivery systems (1-5), because it is: (i) completely biodegradable (avoiding the removal of empty remnants upon drug exhaust), (ii) biocompatible ( 6), (iii) and offers the possibility to provide a considerable range of release periods, ranging from a few hours to several months (2,7,8). Different types of dosage forms can be produced, such as cylindrical implants (9)(10)(11), spherical microparticles (12)(13)(14) and thin films (2,5). A variety of manufacturing procedures can be applied to produce these systems, for instance solvent evaporation methods (10,15), direct compression (16,17), 3D printing (18,19), and hot melt extrusion (4,11,20,21).

The release mechanisms from PLGA-based drug delivery systems can be rather complex, because a variety of physic-chemical phenomena can be involved (12,(22)(23)(24)(25). This includes for example water penetration into the system, PLGA degradation via hydrolytic ester bond cleavage, physical water-polymer interactions (e.g., plasiticizing effects) (26), drug particle dissolution, the diffusion of dissolved drug and water-soluble PLGA degradation products, the creation of local, acidic microenvironments leading to accelerated polymer degradation and drug release ("autocatalytic effects") (27)(28)(29)(30)(31), polymer swelling (32)(33)(34)(35), pore formation & closure (36)(37)(38)(39), limited solubility effects, drug-polymer interactions (e.g., plasiticizing effects) (40), as well as osmotic effects (41). The relative importance of these phenomena can strongly depend on the type of drug, type of polymer (e.g. average polymer molecular weight and type of end groups), composition of the system (8,42,43), and type of manufacturing procedure. The latter might fundamentally affect the resulting inner and outer system structure, in particular its porosity (44,45) and the physical state of the drug in the polymeric matrix (46,47). The drug might be dissolved in the PLGA (molecularly distributed throughout the polymer network), or dispersed in the form of crystalline or amorphous particles. The degree of possible drug-PLGA interactions obviously strongly depends on their physical states and degree of mixing. Thus, it is very interesting to characterize them and monitor potential changes during drug release.

Due to the frequently encountered complexity of the underlying mass transport mechanisms in PLGA-based dosage forms, it is often not fully understood how a specific device controls drug release. Consequently, unexpected tendencies might be observed when varying the systems' composition or process parameters used during manufacturing. This can render system optimization and trouble shooting during production highly challenging. As an example: It can generally be expected that the increase in dosage form dimensions leads to a decrease in the resulting relative drug release rate, if diffusional mass transport plays a major role (because the 100 lengths of the diffusion pathways to be overcome increase) (48). However, it has been shown that in the case of initially non-porous, lidocaine-loaded PLGA-based microparticles a 7-fold increase in the systems' diameter did virtually not affect the relative drug release rate (49). This could be attributed to a compensating mechanism: an increase in the importance of autocatalytic effects in the system: the pH becomes more acidic inside larger microparticles, leading to higher drug mobility. In contrast, in initially highly porous lidocaine-loaded PLGA microparticles, the generated water-soluble acids can more rapidly be neutralized, resulting in decreasing drug release rates with increasing system size (50).

The aim of this study was to prepare ibuprofen-loaded PLGA implants by hot melt extrusion, varying the exposure time to 105 °C from 3 to 15 min, and to thoroughly characterize the systems before and during drug release in phosphate buffer pH 7.4. Optical & scanning electron microscopy were used to evaluate the inner and outer system structure. DSC and X-ray diffraction were applied to better understand the physical state of the drug. Gravimetric analysis was used to measure changes in the dry and wet mass of the systems, and GPC analysis to monitor the average PLGA molecular weight.

III.3. Materials and methods

III.3.1. Materials

Poly (D,L lactic-co-glycolic acid) (PLGA, 50:50 lactic acid: glycolic acid; Resomer RG 503H;

Evonik, Darmstadt, Germany); ibuprofen (BASF, Ludwigshafen, Germany); agarose (genetic analysis grade) and tetrahydrofuran (HPLC grade) (Fisher Scientific, Illkirch, France); potassium dihydrogen orthophosphate and sodium hydroxide (Acros Organics, Geel, Belgium); acetonitrile (VWR, Fontenay-sous-Bois, France); sodium hydrogen phosphate (Na2HPO4;

Panreac Quimica, Barcelona, Spain).

III.3.2. Implant preparation

PLGA was milled for 4 x 30 s in a grinder (Valentin, Seb, Ecully, France). Appropriate amounts of polymer and drug powders were blended for 5 min at 20 rpm in a Turbula T2C Shaker-Mixer (Willy A Bachofen, Basel, Switzerland). Three hundred mg mixture were filled into a 1 mL syringe (Henke Sass Wolf, Tuttlingen, Germany), followed by heating at 105°C for 3, 6, 9, 12 where mass (t = 0) denotes the implant mass before exposure to the release medium.

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

III.3.6. Implant erosion and PLGA degradation

Implants were treated as for the in vitro drug release studies (described in section III. 3.4)

III.3.7. Differential scanning calorimetry (DSC)

DSC thermograms were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland). Samples (approximately 5 mg in the case of raw materials & their physical blends, about 10 mg in the case of implants) were heated in closed aluminum pans as follows: from -

III.4. Results and discussion

Different types of ibuprofen-loaded, PLGA-based implants were prepared by hot melt extrusion using a syringe, as illustrated in Figure 1A. The drug and PLGA powders were blended and filled into a glass syringe, which was heated to 105 °C for 3, 6, 9, 12 or 15 min. The blends were manually extruded, and the extrudates were cut with a hot scalpel into cylindrical implants of approximately 5 mm length. The mean diameter was in the range of 2.5-2.8 mm in all cases (Table III-1). The practical drug loading varied between about 13 and 10 % (w/w). Partial ibuprofen loss might be explained by sublimation at 105 °C. The average polymer molecular weight decreased from about 21 to 18 kDa when prolonging the exposure time to 105 °C from 3 to 15 min (Table III-1). The Mw of the PLGA raw material was 23.1±1.1 kDa. This indicates that the polymer chains were partially cut upon heat exposure, which is consistent with reports in the literature (19,51). Importantly, the ibuprofen powder as received showed a sharp melting peak at 79.7 +/-0.5 °C. This peak was also visible in the hot melt extruded implants, but its intensity substantially decreased with prolonged heat exposure during production. This suggests that the drug is initially present in the form of tiny crystals, which partially: (i) dissolve in the PLGA during processing (the temperature is raised to 105 °C, forming a "monolithic solution"), (ii) melt and re-solidify in the form of amorphous drug particles upon cooling (forming a "solid-in-solid dispersion": amorphous-in-amorphous), and/or (iii) melt and re-solidify in the form of crystalline drug particles upon cooling (forming a "solid-in-solid dispersion": crystalline-in-amorphous). To estimate the capacity of ibuprofen to dissolve in the investigated PLGA, the DSC thermograms of different drug-polymer blends were recorded, varying the ibuprofen content from 0 to 20 % (w/w). As it can be seen in small in all cases. Thus, most of the ibuprofen is molecularly dispersed or in an amorphous form, representing about 10 % (w/w) of the total implant mass.

Figure III-3. DSC thermograms of the: (A) investigated ibuprofen-loaded PLGA implants (before exposure to the release medium), and (B) raw materials (PLGA & ibuprofen). During implant preparation, ibuprofen-PLGA blends were heated for 3 to 15 min (as indicated). The

PLGA (A) (B) dashed orange oval highlights ibuprofen melting events in the implants. Please note the different scaling of the y-axes in (A) and (B).

To further confirm the hypothesis of the presence of (some) crystalline ibuprofen in the PLGA implants, also X-ray diffraction patterns of the systems were recorded before exposure to the release medium. As it can be seen in In conclusion, based on the obtained SEM pictures of surfaces and cross sections, DSC thermograms and X-ray diffraction patterns of the implants, it can be hypothesized that major parts of the ibuprofen are dissolved in the PLGA matrix in the implants (are present in the form of individual drug molecules). And that a minor proportion is also present in the form of tiny drug crystals.

III.4.2. Implants after exposure to the release medium

The diagrams at the top of The zoom on the right-hand side at the top shows that during the burst release phase the following ranking order was observed with respect to the release rate: 3 min > 6 min > 9 min > 12 min > 15 min. This can probably be explained by the fact that more drug crystals are present in the implants when prepared with shorter heating times (as discussed above). The initial burst release is likely due to drug with direct surface access once the system gets into contact with water. This includes tiny drug crystals located directly at the implants' surface, or very close to it and with access to surface pores. If such a drug crystal gets into contact with the release medium, it rapidly dissolves. In contrast, if the drug is dissolved in the polymer matrix, the probability that a drug molecule has direct surface access is lower (most of the molecules are separated from it, even if only by minor amounts of polymer).

The onset of the final rapid drug release phase from the investigated implants is observed after

III.5. Conclusions

Hot melt extrusion offers an interesting potential for the preparation of PLGA-based implants.

However, the exposure to heat can decrease the average polymer molecular weight and alter the physical state of the drug in the polymeric matrix (and of course, thermally degrade heat sensitive drugs). The drug might be present in the fom of individual molecules/ions distributed throughout the PLGA network ("dissolved"), or dispersed in the form of crystalline or amorphous tiny particles. It is worth to monitor the state of the drug and potential transformations during implant manufacturing. For instance, increasing exposure times to heat might increase the relative amount of drug, which is dissolved in the PLGA, altering the importance of the initial "burst release". In the present study this phenomenon could be monitored, but its importance on drug release was limited. However, the impact for other drugs or other polymers might be much more pronounced and certain systems might be much less robust and "foregiving".

III.6. Acknowledgments

This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed. the success of this polymer include: (i) its good biocompatibility " (4-6), (ii) complete biodegradability into lactic acid and glycolic acid, and (iii) the possibility to provide desired drug release rates during flexible time periods, ranging from a few hours to several months (7,8).

Different manufacturing techniques can be used to prepare PLGA-based dosage forms, for example: emulsificationsolvent evaporation methods (9,10), compression (11), 3D printing (12) and hot melt extrusion (13)(14)(15). The type of manufacturing method and selected process parameters can significantly affect the resulting inner and outer structure of the dosage forms and, hence, the resulting drug release kinetics (16,17). Also, the composition of the devices can have a crucial effect, such as the drug loading, type of drug (e.g., its hydrophilicity) (18), type of PLGA (e.g., average polymer molecular weight, type of end groups) (19,20), and potential addition of further excipients (2,21,22). In this study, hot melt extrusion was chosen, since this technique offers an interesting potential for the preparation of homogeneous drug-polymer blends (14,22,23). However, care must be taken not to thermally degrade the polymer, nor the drug.

The mass transport mechanisms controlling drug release from a PLGA-based dosage form can be complex (24,25). Upon contact with aqueous media (e.g. living tissue or phosphate buffer in vitro), water penetrates into the dosage form. This process is generally relatively rapid, assuring that the entire device is wetted and hydrolytic polymer chain cleavage occurs throughout the system ("bulk erosion") (26,27). However, initially, the PLGA is rather hydrophobic and the amounts of water inside the system are limited. Importantly, once an ester bond is hydrolyzed, a new -OH and a new -COOH end group is generated. Consequently, the polymeric matrix becomes more and more hydrophilic with time. In addition, the degree of polymer chain entanglement decreases and the matrix becomes mechanically less stable. Furthermore, the water-soluble degradation products create a steadily increasing osmotic pressure in the system.

This can result in substantial system swelling (28)(29)(30)(31). The presence of pores as well as pore closure (32)(33)(34), drug dissolution and diffusion (24,35), the creation of local acidic microenvironments and subsequent autocatalytic effects (36,37), as well as drugpolymer and waterpolymer interactions (38)(39)(40) can also play a major role. The relative importance of these phenomena might strongly depend on the type of system, e.g. geometry, size, composition and manufacturing method. Often, it is not fully understood which are the dominant mass transport steps, resulting in sometimes surprising effects of formulation and processing parameters. Hence, the optimization of this type of advanced drug delivery systems might be highly cumbersome, requiring cost-intensive and time-consuming series of trial-and-error experiments.

One of the key properties of a controlled drug delivery system, which can often relatively easily be varied and which might offer the possibility to adjust desired drug release profiles is the system size (41-44): This is for example the case, if diffusion plays an important role: Varying the system size means varying the length of the diffusion pathways and, thus, the resulting release rate. However, in the case of initially non-porous PLGA-based microparticles loaded with small amounts of lidocaine, an increase in system size by a factor of 7 did hardly alter drug release (45). This could be explained by the increasing importance of autocatalytic effects inside the microparticles: Not only the diffusion pathways for the drug, but also for the generated short chain acids increases with increasing system size. In this case, this led to more pronounced drops in the micro pH inside the microparticles and, thus, accelerated polymer degradation (hydrolytic ester bon cleavage being catalyzed by protons) and increased drug mobility: Larger microparticle became much more porous inside than smaller microparticles. In contrast, these effects were much less pronounced in microparticles based on the same polymer and drug, but being highly porous from the beginning (46). In these cases, acids and bases were more mobile and pH drops less pronounced.

The aim of this study was to better understand the importance of the diameter of PLGA-based implants prepared by hot melt extrusion on ibuprofen release. Differently sized implants were exposed to phosphate buffer pH 7.4 in 3 experimental set-ups: in well agitated bulk fluid, in agarose gels exposed to phosphate buffer in Eppendorf tubes and in agarose gels exposed to phosphate buffer in transwell plates. The gels were intended to mimic living tissue more realistically than bulk fluids. Optical macroscopy, gravimetric analysis, DSC, GPC, scanning electron microscopy, and pH measurements were used to monitor the implants' key properties before and after exposure to the release medium.

IV Panreac Quimica, Barcelona, Spain).

IV.3.2. Implant preparation

PLGA was milled for 4 x 30 s in a grinder (Valentin, Seb, Ecully, France). Appropriate amounts of polymer and drug powders were blended for 5 min with a mortar and a pestle, followed by extrusion using a Nano 16 twin-screw extruder (screw diameter = 16 mm, die diameter = 0.5, 1, 2 or 3 mm, length/diameter ratio = 26.25, gravitational feeder) (Leistritz, Nuremberg, Germany). The process temperatures were kept constant at 80 -75 -70 -65 °C (die -zone 3zone 2 -zone 1). The screw speed was set at 50 rpm, the screw configuration is illustrated in Figure S1. After cooling, the hot melt extrudates were manually cut into cylinders of 5 mm length.

IV.3.3. Optical macroscopy

Pictures of implants before exposure to the release medium were taken using a SZN-6 trinocular stereo zoom macroscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vision Lite 2.1 software). The lengths and diameters of the implants were determined using the ImageJ software (US National Institutes of Health, Bethesda, Maryland, USA).

IV.3.4. Practical drug loading

Implants were dissolved in 5 mL acetonitrile, followed filtration (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000 Series HPLC, equipped with a LPG 3400 SD/RS pump, an auto sampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversed phase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used. The mobile phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v). The detection wavelength was 225 nm, and the flow rate 0.5 mL/min. Ten microliter samples were injected.

IV.3.5. Differential scanning calorimetry (DSC)

DSC thermograms of the raw materials (PLGA and ibuprofen) and implants were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland). Approximately 5 mg raw material samples and around 10 mg implant samples were heated in perforated aluminum pans as follows: from -70 to 120 °C, cooling to -70 °C, re-heating to 120 °C (heating/cooling rate = 10 °C/min). The reported glass temperatures (Tgs) were determined from the 1 st heating cycles in the case of implants (the thermal history being of interest), and from the 2 nd heating cycle in the case of the raw material (the thermal history not being of interest). All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

IV.3.6. In vitro drug release

Three experimental set-ups were used to measure ibuprofen release from the PLGA implants: Gesellschaft fuer Labortechnik, Burgwedel, Germany). At predetermined time points, the entire bulk fluid was replaced by fresh release medium. The withdrawn samples were filtered (PVDF syringe filter, 0.45 µm; Agilent, Santa Clara, California, USA) and analyzed for their ibuprofen contents by HPLC-UV, as described in section IV.3.4.

IV

IV.3.6.2. In agarose gels in Eppendorf tubes

Implants were embedded into agarose gels in 5 mL Eppendorf tubes, as illustrated in Figure

IV-1 B (1 implant per tube

). An agarose dispersion (0.5% w:v) in phosphate buffer pH 7.4

(USP 42) was heated to 100 °C under magnetic stirring (250 rpm) until a clear solution was obtained. The latter was cooled to 47°C and continuously stirred (to prevent gelation). 0.5 mL of the solution was placed into the bottom of an Eppendorf tube and cooled in a refrigerator for 5 min to allow for gelation. An implant was carefully placed on top of the gel, and covered with second layer of 0.5 mL agarose solution (47 °C), followed by cooling in a refrigerator for 5 min.

Four mL phosphate buffer pH 7.4 USP 42 were added on top of the gel, and the tubes were placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the IV. PLGA IMPLANTS FOR CONTROLLED DRUG RELEASE: IMPACT OF THE DIAMETER 131 entire bulk fluid was replaced by fresh release medium. The withdrawn samples were treated as for the drug release measurements in well agitated bulk fluids.

IV.3.6.3. In agarose gels in transwell plates

Implants were embedded in agarose gels in transwell plates (1 implant per insert, membranes: The agarose gels were prepared as described above, and the implants included accordingly (placed between 2 "layers" of 0. where mass (t = 0) denotes the implant mass before exposure to the release medium.

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

IV.3.8. Implant erosion and PLGA degradation

Implants were treated as for the in vitro drug release studies described above. At pre-determined time points, implant samples were withdrawn and freeze-dried (freezing at -45°C for 2 h 35 min, primary drying at -20 °C/0.940 mbar for 35 h 10 min and secondary drying at +20 °C/0.0050 mbar for 35 h; Christ Alpha 2-4 LSC+; Martin Christ, Osterode, Germany).

The dry mass (%) (t) was calculated as follows:

𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (%)(𝑡) = 𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (𝑡) 𝑚𝑎𝑠𝑠 (𝑡 = 0) × 100 %
where mass (t = 0) denotes the implant's mass before exposure to the release medium. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation chromatography (GPC) as follows: Freeze-dried implant samples were dissolved in tetrahydrofuran (at a concentration of 3 mg/mL

IV.3.9. Scanning electronic microscopy (SEM)

The internal and external morphology of the implants before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxfordshire, England).

Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome layer. In the case of implants which had been exposed to the release medium, the systems were treated as for the in vitro release studies described in section IV.3.6. At predetermined time points, implant samples were withdrawn, optionally cut (for cross sections) using a scalpel and freeze-dried (as described in section IV. 3.8).

IV.4. Results and discussion

The aim of this study was to better understand the importance of the diameter of PLGA implants on system performance upon exposure to phosphate buffer pH 7.4 using different experimental set-ups, in particular in the light of the resulting drug release kinetics. Cylindrical ibuprofenloaded implants were prepared by hot melt extrusion using a Leistritz Nano 16 twin-screw extruder, equipped with differently sized dies.

IV.4.1. Implants before exposure to the release medium

Figure IV-2 shows optical macroscopy pictures of the implants prepared with dies with a diameter of 0.5, 1, 2 or 3 mm (as indicated). As it can be seen, they have a homogeneous appearance and smooth surface. The diameters of the implants are indicated in Table 1, ranging from 0.7 to 2.8 mm. The practical drug loading was about 12 % and rather homogeneous in all samples. The polymer molecular weight was about 24-26 kDa, which is slightly higher compared to implants of similar composition prepared by hot melt extrusion using a syringe and heating to 105 °C for 3 to 15 min (47). This difference can probably be explained by the lower temperatures applied in this study during processing ( 80 Clearly, the PLGA was amorphous, while the ibuprofen raw material showed a sharp melting peak at 79.7 +/-0.5 °C. Importantly, no sign for ibuprofen melting was visible in any implant, only one single glass transition temperature (Tg) could be detected in the investigated temperature range: at about 35-37 °C. The Tg of the PLGA raw material was about 47.0 °C.

Thus, ibuprofen acts as a plasticizer for this polymer. As it can be seen, all surfaces and cross sections were smooth and essentially non-porous.

Some tiny cavities were visible, which might be attributable to ibuprofen sublimation during processing. As it has previously been demonstrated (47), the ibuprofen can be expected to be essentially dissolved (molecularly dispersed) in the polymeric matrix at these drug loadings (47).

IV.4.2. Implants after exposure to the release medium

The schemes in As it can be seen in Figure IV-5, the observed drug release kinetics were bi-phasic in all cases, irrespective of the implants' diameter and experimental set-up: An about zero order drug release rate (with an about constant release rate) was followed by a final rapid drug release phase (accounting for approximately 80-90 % of the total drug). An initial burst phase was (if present at all) not of noteworthy importance. This is in good agreement with the hypothesis that the initial (limited) burst release (about 3-5 % in the first 12 h) from implants of similar composition, but in which the drug was partially present in the form of crystalline particles, can mainly be attributed to the presence of pure drug particles with direct surface access (47).

Interestingly, all implants did not substantially swell during the first few days after exposure to the release medium, but substantial system swelling set on after a certain lag time : The These phenomena can probably be explained as follows: Upon contact with aqueous media, limited amounts of water penetrate into the implants, which become rapidly entirely wetted.

Thus, hydrolytic polymer chain cleavage occurs throughout the systems and starts from the beginning ("bulk erosion"). This is also consistent with the observed decrease in the average Please note that the samples had to be dried before SEM analysis, thus, artefacts have been created. The highly wizened structure which can be seen on all surfaces result from the drying of highly swollen PLGA surface layers. Since the surface polymer is in contact with very high amounts of water from the beginning, ester bond cleavage is particularly rapid in these zones.

Upon PLGA surface layer increases in thickness. Importantly, the mobility of ibuprofen molecules located in these regions fundamentally changes: They can be expected to be poorly mobile in the dense, non-swollen PLGA, but to be much more mobile in a highly swollen PLGA gel: The distances between the macromolecules has considerably increased and the polymer chains themselves are much more mobile. This can be expected to result in relatively rapid drug release from highly swollen PLGA zones. Since the rate at which the thickness of the highly swollen surface layer increases can be expected to be about constant (the conditions for the swelling of "deeper" polymer layers do not change over time), also the resulting drug release rate can be expected to be about constant. This is in good agreement with the observed zero order release Please note that the implants were freeze-dried prior to analysis, creating artefacts. The implant diameters are also indicated in the diagrams.

This local, surface-near PLGA swelling has to be distinguished from the observed substantial entire system swelling occurring after about 5-10 d: Due to the fact that the entire implants are rather rapidly wetted upon contact with the release medium, polymer chain degradation occurs throughout the systems (even if at a lower rate compared to the outmost surface layers). Hence, after a certain lag time, the entire implant is sufficiently hydrophilic and mechanically instable and exhibits an osmotic pressure, which attracts substantial amounts of water into the entire device. This is the onset of fundamental entire system swelling. The consequences for ibuprofen release are the same: It becomes much more mobile in the highly swollen PLGA gel. Looking IV. PLGA IMPLANTS FOR CONTROLLED DRUG RELEASE: IMPACT OF THE DIAMETER 145 at the optical macroscopic pictures at later time points in Figure 6, this can also be "visualized".

For example, the mobility of an ibuprofen molecule in the dense PLGA matrix in an implant with an initial diameter of 3 mm exposed to a well agitated bulk fluid can be expected to be limited. However, after 1 week, the implant has tremendously swollen and ibuprofen can much more easily be released. But not only the mobility of the drug molecules/ions fundamentally increases: also the water-soluble degradation products of the PLGA can more easily diffuse out of the systems. This causes a drop in the pH of the respective bulk fluid (3 rd row in Figure 5).

Please note that the importance of the observed pH drop can in part also be attributed to the longer (3 d week end) sampling period, resulting in short chain accumulation in the bulk fluid.

As it can be seen, the drop in pH is more pronounced with larger implants, which can be explained by the higher amounts of PLGA in these systems and, thus, higher amounts of degradation products. The facilitated release of drug and short chain acids in the highly swollen PLGA matrices also leads to the onset of implant erosion, as illustrated in the 3 rd row in Larger matrices can be expected to be mechanically more stable than smaller ones, the mass of the investigated implants varying by a factor of about 20 (Table IV-1). Interestingly, no signs for the occurrence of important autocatalytic effects were observed in this study, despite the considerable system dimensions. This might eventually be explained by the relatively rapidly occurring substantial entire system swelling, which sets on after only about 5 to 10 d (Figure

IV-5-8).
Once the PLGA is highly swollen, generated water-soluble acids can rather rapidly diffuse out and be neutralized, and bases from the release medium can diffuse into the system.

Comparing the implant swelling kinetics in the 3 experimental set-ups in Poly (D,L lactic-co-glycolic acid) (PLGA)-based implants offer an interesting potential for parenteral controlled drug delivery (1)(2)(3). This is because: (i) PLGA is biocompatible and completely biodegradable (avoiding the need to remove empty remnants after drug exhaust)

(1,2,4). (ii) Drug release can be controlled during periods ranging from a few days up to several months (5). (iii) A variety of manufacturing processes can be applied for their preparation, including for example hot melt extrusion (6), compression (7) and 3D printing (8).

The underlying mass transport mechanisms controlling drug release from PLGA implants can be very complex (5,9,10). The following are examples for phenomena which can be involved.

Their relative importance can very much depend on the composition and structure of the implant (and, hence the manufacturing procedure), as well as on the experimental conditions chosen for the in vitro release measurements: Upon implantation, water penetrates into the implants.

Generally, the entire system is rapidly wetted. However, the amounts of water at this stage are often limited, because the commonly used PLGA grades are rather hydrophobic and the polymeric matrix is dense. The presence of the water initiates ester bond cleavage throughout the device ("bulk erosion") (11). In addition, the drug can dissolve in the water and becomes mobile (although often only to a limited extent, since the amounts of water in the system are still low). Consequently, the dissolved drug molecules/ions diffuse out into the surrounding environment, due to concentration gradients (9,12). Depending on the system size and initial inner structure, the pH might locally significantly drop (13)(14)(15). This is because ester bond cleavage generates shorter chain acids, which might accumulate, especially at the center of the implant: The diffusion of these acids out of the device takes time (as well as the diffusion of bases from the environment into the system), and the rate at which the acids are generated can be higher than the rate at which they are neutralized (16). Since ester bond hydrolysis is catalyzed by protons, this can lead to autocatalytic effects (10,15,16). Consequently, polymer degradation can be accelerated and the drug release rate increases. Furthermore, limited drug solubility effects within the PLGA implants as well as within the surrounding bulk fluid might play a major role (17,18). Also, drug diffusion through water-filled pores might be of importance. If the pores exist from the beginning, this might cause a high initial drug release rate. The pores might then be closed by (limited) PLGA swelling, terminating the related burst release phase (19,20).

One of the key reasons for the potential complexity of the drug release mechanisms from PLGA-based implants is polymer degradation, which has multiple consequences over time:

(i) The macromolecular network becomes more and more hydrophilic, since the cleavage of each ester bond generates two new hydrophilic end groups: an -OH and a -COOH group.

(ii) The mechanical stability of the polymeric matrix decreases, since the macromolecular chain length decreases, resulting in less intense polymer entanglement. (iii) Water-soluble degradation products accumulate (being poorly mobile in the initially dense polymeric system, similar to the drug), creating a steadily increasing hydrostatic pressure in the implant. Thus, the implants become more and more water-loving with time, get less resistant to dimensional changes and more and more actively attract water. Consequently, at a certain time point, substantial implant swelling sets on (21)(22)(23)(24). Increases in system wet mass up to 1000 % have been reported (21,25). This tremendous device swelling fundamentally changes the conditions for drug release: The polymer network becomes much less dense, the dissolved drug molecules/ions become much more mobile. Hence, the final, rapid drug release phase sets on (9,26,27).

Unfortunately, yet no compendial method has been described for the measurement of drug release from parenteral controlled drug delivery systems. So, great care must be taken when comparing drug release kinetics obtained under different conditions. Frequently, the implants are directly exposed to a well agitated bulk fluid, such as phosphate buffer pH 7.4 (28-30). At pre-determined time points, samples are withdrawn from the release medium and often replaced by fresh bulk fluid. It is well known that the sampling frequency and volume likely affect the observed drug release patterns, if limited drug solubility effects in the bulk fluid are of importance. Also, since PLGA degradation is catalyzed by protons, temporary drops in the pH of the bulk fluid can lead to accelerated polymer chain cleavage and faster drug release (10,15,31). However, little is yet known about potential further impacts of the sampling schedule on drug release.

To more realistically mimic the conditions experienced by the dosage forms upon implantation into human tissue, hydrogels have been proposed (32)(33)(34). The idea is to minimize convective mass transport (as in well agitated bulk fluids) and place the system in a semi-solid environment. It has recently been shown that the presence of such a hydrogel can limit substantial implant swelling and, thus, slow down drug release (35). The group of Ostergaard published some very interesting reports on how drug transport in a surrounding hydrogel can be monitored using UV analysis from a controlled release implant (36)(37)(38). This type of results can be very helpful to better understand which phenomena are decisive for drug transport into the environment surrounding the implant.

V. HOW BULK FLUID RENEWAL CAN AFFECT DRUG RELEASE FROM PLGA IMPLANTS 157

The aim of this study was to better understand how sensitive the experimentally measured drug release profiles from PLGA-based implants might be on the renewal of the release medium.

Ibuprofen-loaded PLGA implants were prepared by hot melt extrusion. Three experimental setups were used: The implants were either directly exposed to well agitated phosphate buffer pH 7.4, or embedded within agarose gels, which were exposed to phosphate buffer in Eppendorf tubes or transwell plates (35). The systems were thoroughly characterized before and during drug release by optical and scanning electron microscopy, gravimetric analysis, pH and solubility measurements as well as gel permeation chromatography, to explain the observed results.

second layer of 0.5 mL agarose solution (47 °C), followed by cooling in a refrigerator for 5 min.

Four mL phosphate buffer pH 7.4 USP 42 were added on top of the gel, and the tubes were placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the entire bulk fluid, or 3 mL or 1 mL of the bulk fluid was replaced by fresh release medium. The withdrawn samples were treated as for the drug release measurements in well agitated bulk fluids.

V.3.6.3. In agarose gels in transwell pates

Implants were embedded in agarose gels in transwell plates (1 implant per insert, membranes:

1.13 cm², 11 µm, 0.4 µm pore size; Nunc, Roskilde, Denmark), as illustrated in Figure V-1C.

The agarose gels were prepared as described above, and the implants included accordingly (placed between 2 "layers" of 0.5 mL gel). The well plates were filled with 4 mL phosphate buffer pH 7.4, covered with lids and Parafilm to minimize evaporation, and placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the entire bulk fluid, or 3 mL or 1 mL of the bulk fluid was replaced by fresh (pre-heated) release medium.

The withdrawn samples were treated as for the drug release measurements in well agitated bulk fluids.

In all cases, sink conditions were provided in the agitated bulk fluids. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

Furthermore, the pH of the bulk fluids was measured at pre-determined time points using a pH meter (InoLab pH Level 1; WTW, Weilheim, Germany). Mean values +/-standard deviations are reported.

V.3.7. Implant swelling

Implants were treated as for the in vitro drug release studies described above. At pre-determined time points:

(i) Pictures of the implants were taken with a SZN-6 trinocular stereo zoom macroscope (Optika), equipped with an optical camera (Optika Vision Lite 2.1 software). The lengths and diameters of the implants were determined using the ImageJ software (US National Institutes of Health). Dynamic changes in the systems' volume were calculated considering cylindrical geometry.

(ii) Implant samples were withdrawn and excess water was carefully removed using Kimtech precision wipes (Kimberly-Clark, Rouen, France) and weighed [wet mass (t)]. The change in wet mass (%) (t) was calculated as follows:

𝑐ℎ𝑎𝑛𝑔𝑒 𝑖𝑛 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (%)(𝑡) = 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (𝑡) -𝑚𝑎𝑠𝑠 (𝑡 = 0) 𝑚𝑎𝑠𝑠 (𝑡 = 0) × 100 %
where mass (t = 0) denotes the implant mass before exposure to the release medium.

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

V.3.8. Implant erosion and PLGA degradation

Implants were treated as for the in vitro drug release studies described above. At predetermined time points, implant samples were withdrawn and freeze-dried (freezing at -45°C for 2 h 35 min, primary drying at -20 °C/0.940 mbar for 35 h 10 min and secondary drying at +20 °C/0.0050 mbar for 35 h; Christ Alpha 2-4 LSC+; Martin Christ, Osterode, Germany).

The dry mass (%) (t) was calculated as follows:

𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (%)(𝑡) = 𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (𝑡) 𝑚𝑎𝑠𝑠 (𝑡 = 0) × 100 %
where mass (t = 0) denotes the implant's mass before exposure to the release medium. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation chromatography (GPC) as follows: Freeze-dried implant samples were dissolved in tetrahydrofuran (at a concentration of 3 mg/mL). One hundred µL samples were injected into an Alliance GPC (refractometer detector: 2414 RI, separation module e2695, Empower GPC software; Waters, Milford, USA), equipped with a PLgel 5 µm MIXED-D column (kept at 35°C, 7.8 × 300 mm; Agilent). Tetrahydrofuran was the mobile phase (flow rate: 1 mL/min).

Polystyrene standards with molecular weights between 1,480 and 70,950 Da (Polymer Laboratories, Varian, Les Ulis, France) were used to prepare the calibration curve. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

V.3.9. Scanning electronic microscopy (SEM)

The internal and external morphology of the implants before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxfordshire, UK). Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome (Cr) layer. In the case of implants which had been exposed to the release medium before observation, the systems were treated as described for the in vitro release studies described above. At predetermined time points, implant samples were withdrawn, optionally cut with a scalpel and freeze-dried (as described in section V.3.8 Implant erosion and PLGA degradation).

V.4. Results and discussion

During in vitro drug release measurements from PLGA implants, often parts of the release medium are withdrawn at pre-determined time points and replaced with fresh medium. This procedure might affect the dosage form and alter the release rate. The aim of this study was to investigate the potential importance of the replaced bulk fluid volume, using the 3 experimental set-ups illustrated in Figure V-1. The bulk fluid (phosphate buffer pH 7.4) was either in direct contact with the implants, or in contact with a hydrogel in which the implants were embedded.

The bulk fluid was entirely replaced (100 %), to a major extent (60-75 %), or to a minor extent (20-25 %). Ibuprofen-loaded implants were prepared by hot melt extrusion. The drug loading was about 8 %, the implants were cylindrical in shape (5.7 mm length, 2.7 mm diameter) and had a homogeneous appearance (optical macroscopy picture in Table V-1). The bulk fluids were completely renewed at each sampling time point (4 or 5 mL), or only 3 or 1 mL of it (green, orange and blue curves). As it can be seen, the volume of the renewed bulk fluid did not have a noteworthy impact on drug release from the implants embedded within an agarose gel. In contrast, in the case of direct exposure to the phosphate buffer (standard set-up), ibuprofen release was faster when the entire release medium was renewed, compared to only The onset of this final rapid release phase coincided with the onset of substantial implant swelling, as it can be seen in the diagrams in the middle and at the bottom of Figure V-2. It has previously been suggested that initially only limited amounts of water are able to penetrate into PLGA-based drug delivery systems, so that drug mobility remains limited (21). Nevertheless, the entire dosage forms are rather rapidly wetted (by small amounts of water), initiating PLGA degradation throughout the system ("bulk erosion"). Upon hydrolytic cleavage of an ester bond of a PLGA chain, 2 shorter chains are created 2 new hydrophilic end groups: an -OH and a -COOH end group. Thus, with time the polymeric system becomes more and more hydrophilic.

Table V-1. Key properties of the investigated ibuprofen

In addition, the PLGA chains get shorter and, hence, less entangled. Furthermore, water-soluble degradation products create a steadily increasing hydrostatic pressure inside the implant. At a certain time point, these changes result in the penetration of considerable amounts of water into the system, e.g. the volume of the investigated PLGA implants increased by about 500 % after 10 d upon direct exposure to phosphate buffer pH 7.4 (Figure V-2, diagram on the left hand side in the middle row). This dramatically changes the conditions for drug release: Ibuprofen is much more mobile in a highly swollen PLGA gel compared to the initially "dry" and dense polymer matrix. Consequently, the drug release rate increases.

It has previously been shown that the presence of a hydrogel surrounding a PLGA implant of similar composition substantially limits implant swelling, resulting in slower drug release (35).

This tendency has been confirmed in this study: As it can be seen in the diagrams middle and at the bottom of Figure V-2, system swelling was less pronounced, if the implants were embedded in a gel. The optical macroscopy pictures in Figure V-3 illustrate these phenomena:

The implants underwent substantial swelling after about 7-10 d. This process was particularly pronounced upon direct exposure to the bulk fluid (compared to the agarose gel set-ups). The SEM pictures at the top of Figure V-4 show surfaces and cross-sections of ibuprofen-loaded implants before exposure to the release medium. As it can be seen, the surface was smooth and non-porous (due to the manufacturing process), which can in part explain the absence of any noteworthy burst release. The other SEM pictures in Figure V-4 show implants which had been directly exposed to the release medium, or embedded within agarose gels for 8 d. Please note that the samples had been freeze-dried before analysis, so artefacts have been created. The pictures of surfaces show highly porous and "folded" structures. These likely result from the drying of highly swollen PLGA gels (35,39,40). The cross-sections indicate that in addition to these highly swollen surface-near regions, much less swollen regions exist, irrespective of the type of set-up at this time point. Please note that the pictures were not systematically taken at similar positions. Importantly, there was no sign for a significant impact of the volume of the replaced bulk fluid on the dynamic changes of the inner and outer implant morphology.

To better understand the sensitivity of drug release on the sampling volume in the case of direct exposure to the release medium and insensitivity in the case of implants which were embedded into agarose gels, also the dynamic changes in the PLGA polymer molecular weight, implants' dry mass and pH of the bulk fluid were studied. (the receptor compartment containing well-agitated bulk fluid). At pre-determined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated). Please note that after exposure to the release medium, the implants were freeze-dried prior to analysis. Thus, caution must be paid due to artefact creation.

V.4.2. PLGA degradation and implant erosion

The diagrams at the top of However, in this study, sink conditions were provided in all bulk fluids in all set-ups during the entire observation periods. Importantly, the sampling volume did not affect the dynamic changes in the implants' dry mass and PLGA polymer molecular weight during the first 8 d upon direct exposure to phosphate buffer pH 7.4 or agarose gels (Figure V-5, diagrams in the middle and at the bottom). However, the implants became very fragile with time: The highly swollen implants were difficult to handle. During bulk fluid removal, the systems were rather well protected when embedded in an agarose gel and the risk of accidental damage during sampling was limited. In contrast, in the case of implants which were directly exposed to the phosphate buffer in Eppendorf tubes, the bulk fluid removal became more and more delicate with time: Especially, if the entire release medium was to be replaced. This accelerated implant disintegration (visual observation)

and was the reason why most of the measurements had to be stopped after 8 d. The partial disintegration of the highly swollen PLGA gel led to a decrease in the length of the diffusion pathways to be overcome by the drug to be released. Consequently, ibuprofen release was accelerated. In contrast, the agarose gel effectively protected the implants. In addition, the presence of the agarose gels slowed down implant swelling (Figure V-2, diagrams in the middle and bottom rows). So, the observed sensitivity of drug release from the investigated PLGA implants to the sampling volume in the case of direct exposure to the release medium can likely be attributed to the mechanical stress experienced by the systems during complete bulk fluid renewal. (8)(9)(10)(11)(12), material extrusion (13)(14)(15)(16)(17)(18), material jetting (19,20), binder jetting (21,22).

VI. 3D PRINTING OF IBUPROFEN

3D printing technologies have demonstrated high flexibility in the preparation of dosage forms such as: combining different drug substances (17), personalized devices (18), reaching complex patterns (23)… Conventional pharmaceutical technologies allow modulating the release pattern, by playing on the size of classical shape devices (24)(25)(26)(27) or the choice of the appropriate carrier (28)(29)(30). With the right choice of polymer and technology, it might be possible to precisely modulate drug release by taking advantage of the precision of 3D printing technologies to modulate the internal structure, reaching complex geometries (31), etc…

The applicability of 3D printing technologies on controlled drug delivery has been widely studied for oral dosage forms (13,16,17). Compared to oral delivery while only a few studies focus on PLGA based 3D printed controlled drug delivery devices (32)(33)(34)(35) for parenteral

delivery. Yet PLGA is widely used as a matrix former for controlled drug release, with conventional pharmaceutical technologies/ microparticles (28,29,36), implants (30,37), in situ forming implants (38)(39)(40). Due to its biocompatibility and complete biodegradability, it is widely used in different FDA-approved products (29). Drug release kinetics from such devices is mainly influenced by the type of drug, polymer grade, size of the device (24,25,28). The use of 3D printing technologies with PLGA could bring a new opportunity to modulate drug release kinetics.

Fused filament fabrication (FFF) is based on the fused deposition modeling developed and patented by L. Crump and S. Scott in 1989 (41,42). These technologies use a continuous filament of thermoplastic material as the base material. The filament is fed by a spool, via a heated, mobile printer extrusion head, often referred to as an extruder. The molten material is expelled from the extrusion nozzle and deposited first on a 3D printing platform, which can be heated to ensure good adhesion. Once the first layer is complete, the extruder and the platform move away simultaneously and the second layer can then be directly deposited on the part being manufactured (43). thermoplastic behavior of PLGA makes it a good candidate for 3D printing with FFF.

The printability of drug-polymer mixture by extrusion-based technologies can be investigated via the rheological properties of these systems (44). Investigation of the melt rheology properties can contribute to the understanding of the properties of the material that are meaningful for extrusion, injection modeling, and 3D printing (45). Proper melt rheology studies could help to understand and overcome some issues with the sometimes encountered, lack of drug-polymer miscibility, high processing temperature, and poor printability (31).

This study aimed to determine the impact of ibuprofen on viscoelastic properties of PLGA, and the impact of heating treatment by FFF on PLGA molecular weight and drug content. Moreover, ibuprofen-loaded PLGA filaments were prepared by hot-melt extrusion and used to feed a Volumic FFF 3D printer. Thus, implants with different internal structures were produced and exposed to phosphate buffer pH 7.4. Optical macroscopy, gravimetric analysis, DSC, GPC, scanning electron microscopy, and pH measurements were used to monitor the implants' key properties before and after exposure to the release medium.

VI.3. Materials and methods

VI.3.1. Materials

Poly (D,L lactic-co-glycolic acid) (PLGA, 50:50 lactic acid: glycolic acid; Resomer RG 503H; Evonik, Darmstadt; Germany); ibuprofen (BASF, Ludwigshafen, Germany); agarose (genetic analysis grade), tetrahydrofuran (HPLC grade) (Fisher Scientific, Illkirch, France); acetonitrile (VWR, Fontenoy-sous-Bois, France); sodium hydrogen phosphate (Na2HPO4, Panreac Quimica, Barcelona, Spain).

VI.3.2. Differential scanning calorimetry (DSC)

DSC thermograms of the raw materials (PLGA and ibuprofen), their physical mixture (from 5:95 to 20:80; ibuprofen:PLGA), ibuprofen loaded PLGA filaments, and 3D printed implants were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland).

Approximately 5 mg material were heated in pierced aluminum pans as follows: from -70 to 120 °C, cooling to -70 °C, re-heating to 120 °C (heating/cooling rate = 10 °C/min. The reported glass temperatures (Tgs) were determined from the 2 nd heating cycle for the raw materials and the physical mixture. The melting point and the thermograms of the ibuprofen-loaded and 3D printed implant were determined from 1 st heating cycle (thermal historic is of interest). All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

VI.3.3. Thermogravimetric analysis (TGA)

The thermal stability of PLGA, ibuprofen, and their physical mixture (from 5:95 to 20:80; ibuprofen:PLGA), was measured using a (GA 2 system (GF, Mettler Toledo, Greinfensee, Switzerland). The mass loss upon heating was obtained for 20-30 mg of samples at 10 ° C / min in the temperature range 25 to 750 ° C.

VI.3.4. Disc preparation

Appropriate amounts of polymer (milled for 4 x 30 s with a Seb Valentin grinder) and drug were mixed for 5 min at 20 rpm in a Turbula Shaker-Mixer (Sausheim, France). The blends were then mixed with a mortar and a pestle for 2 min. The obtained mixture was filled in aluminium pans (ref 0452155, Grosseron, Couëron, France) of 43 mm diameter, before heating with a MR Hei-Tec hot plate (442-1358, Avantor, Radnor, Pennsylvania, USA) under vacuum to allow moisture evaporation required for rheological studies as described in Figure 1, and

VI.3.5. Melt rheology

The rheological studies were carried out using an Anton Paar rheometer (MCR 72; Anton Paar ® , Les Ulis, France). The temperature sweep experiment was performed in the linear domain, at an oscillation frequency of 10 Hz, a strain of 1%, a gap of 0.800 mm, and using a plane-plane module with a diameter of 25 mm. Data were recorded and analyzed using RheoCompass software. Complex viscosity, storage (G'), and loss (G'') modulus were determined.

VI.3.6. Hot-melt extrusion

The PLGA was milled (4 x 30 s) with a grinder (Valentin, Seb, Ecully, France). Appropriate amounts of polymer and ibuprofen were mixed and then pre-melted in a crystallizer on a film

VI.3.7. 3D printing test

The filaments obtained were used to feed the Volumic ® 30 Stream Ultra printer (Volumic, Nice, France). Parallelepiped-shaped implants with dimensions of 2 mm x 2 mm x 0.5 mm were printed with printing temperatures ranging from 115 °C to 150 °C. Parallelepiped-shaped implants with dimensions of 10 mm x 10 mm x 2.5 mm and different filling patterns were printed according to the parameters described in

VI.3.9. Gel permeation chromatography (GPC)

The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation chromatography (GPC) as follows: raw materials, samples were dissolved in tetrahydrofuran where mass (t = 0) denotes the implant mass before exposure to the release medium.

(
All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

VI.4. Results and discussion

This study aimed to better understand the impact of the heating treatment of PLGA ibuprofen mixture on the viscoelastic properties, molecular weight of the polymer, and drug content. Thus, the use of PLGA and ibuprofen as feed materials for 3D printing by FFF was investigated. The impact of the internal structure of 3D printing ibuprofen-loaded PLGA implants on system performance upon exposure to phosphate buffer pH 7.4 was investigated, in particular in light of the resulting drug release kinetics.

VI.4.1. Raw materials characterization

To assess the impact of the percentage of ibuprofen on the thermal behavior of the mixture, the thermograms of the raw materials, and the various ratios of PLGA and ibuprofen, PLGA Tg There is therefore complete amorphization of ibuprofen, as well as it is in the polymer matrix. Interestingly, a decrease of PLGA glass transition temperature with ibuprofen content is observed, indicating its plasticizing effect, which has been reported before (39,46). It could be taken as an advantage for lowering the 3D printing temperature of PLGA. However, for a mixture containing 20% of ibuprofen, it has been noticed, that a non-homogeneous amorphous mixture is in place, with the presence of two glass transition temperatures, indicating that this could be the limit of solubility of ibuprofen in PLGA. Regarding these results, it has been decided to continue the study with a special focus on ibuprofen content corresponding to 15%.

The TGA results (Figure VI-2D) show that the mass loss of ibuprofen begins at approximately 160 ° C. This is probably due to its evaporation rather than its degradation, as it has been reported to have a boiling point of 157 ° C, and its degradation to be in a range between 180 ° C and 300 ° C (47). On the other hand, if the mass loss of PLGA begins at temperatures around 250 ° C, this does not give any indication of the state of degradation (shortening of polymer chains) of PLGA at lower temperatures. The degradation being linked to a scission of the bonds of the polymer chains influences the molar mass but not necessarily the total mass. For 3D printing, the factors limiting the maximum printing temperature appear to be the evaporation temperature of ibuprofen and the state of degradation of the PLGA chains.

VI.4.2. Melt rheology studies

Pure PLGA discs and 15% ibuprofen-loaded PLGA discs have been investigated for melt rheology study. It is important to note that the setup used to obtain the polymer discs, allows manufacturing discs without any air bubbles. The process used (Figure VI-1), allows moisture evaporation (by heating under vacuum at a temperature below the Tg of the polymer), before melting the polymer. When studying rheological behavior, it is of great importance that the material is exempt from any air bubbles and moisture (45). Importantly, the experiments have been conducted in the linear range of both investigated materials (data not shown). Ibuprofen by inserting itself between the macromolecular chains and by reducing polymer-polymer interactions in favor of polymer-plasticizer interactions facilitate the polymer chains mobility. The plasticizing effect of ibuprofen induced by hydrogen bridges has been described in the literature on certain polymers such as Eudragit® RS 30D (48,49) or ethyl cellulose (50), but also PLGA Resomer 502H and 503H (39,46). By increasing the temperature, the polymer chains increase, as well as the free volume around the polymer chains. Thus, leading to easier molecular mobility and then decreases viscosity (51). Moreover, for PLGA and PLGA ibuprofen mixture, the investigation is made in a range of temperatures above the Tg value, and in each case, the mixture behaves like a solution with a loss modulus G'' value higher than storage modulus G'. 115°C to 150°C. This might be because the filament is heated for a very short period, leading to a non-significant impact on those parameters. However, as a precautionary measure, it has been chosen to continue the study, by printing the implant at 115°C, which was the lowest temperature allowing the extrusion of the filament, except for the first layer which was printed at 125°C to improve first layer adherence (Table VI-2). print these implants. In every case, the ibuprofen is completely solubilized in the polymer matrix, indicating that the second heating treatment applied to the 3D printer does not seem to affect the interactions between ibuprofen and PLGA. printed drug product (1,3). This drug product is printed by binder jetting (4,5), and four other 3D printing technologies have been investigated for use in pharmaceutics: photopolymerization (6-9), powder bed fusion (10)(11)(12)(13)(14), material extrusion (15)(16)(17)(18)(19)(20), and material jetting (21,22).

The advantages of 3D printing technologies for the controlled release of oral dosage forms have been extensively investigated in the past few years (15,21,23). However, only a few studies focus on parenteral 3D printed devices, especially poly (lactic-co-glycolic) acid (PLGA) based devices (24)(25)(26)(27)(28)(29). Even so, PLGA is a well-known and used polymer in controlled drug delivery systems obtained with conventional pharmaceutical technologies (30)(31)(32) . Thanks to its great biocompatible and biodegradable properties (33)(34)(35)(36)(37), this polyester is widely studied (38)(39)(40) and present in numerous FDA-approved drug products, such as microparticles, implants, or in situ forming implants (41). Modulating either the polymer grade, the type of devices, its porosity, or its size could lead to different drug release kinetics (42)(43)(44). The manufacturing process as well influences the release rate (45)(46)(47).

The emerging 3D printing technologies provide new expediency to precisely control drug release from PLGA based devices. Thanks to its high flexibility, the 3D printing technologies allow the manufacturing of a wild range of dosage forms: a combination of several drug substances (19), patient-specific devices (20), complex release patterns (48)… By combining the adapted polymer and the appropriate technology precisely modulating drug release should be possible. The preciseness of 3D printing technologies is a crucial advantage when it comes to controlling the internal structure of the device, reaching complex geometries (23), etc… Due to its thermoplastic properties, PLGA seems to be a good candidate for 3D printing by extrusion-based technologies (49).

The material extrusion-based 3D printing technologies can be divided into two parts: (i) extrusion of low melting point past or gel and (ii) extrusion of high melting point thermoplastic

VII.3.3. Optical microscopy

Pictures of the implants before exposure to the release medium were taken using a SZN-6 trinocular stereo zoom microscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vison Lite 2.1 software). The lengths/width and implants pore size were determined using the ImageJ software (US National Institutes of Health, Bethesda, Maryland, USA).

VII.3.4. Practical drug loading

Pellets and pieces of 3D printed implants were dissolved in 5 mL acetonitrile, followed by filtering (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000

Series HPLC, equipped with an LPG 3400 SD/RS pump, an autosampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversedphase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used. The mobile phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v). The differences observed between the feedstock materials (filament and pellets), can be explained by the differences during the production by HME. Indeed, the production of pellets does not require having a homogeneous diameter at the outlet, since the extruded material is then cut. To obtain a filament that can be used in 3D printing, the output diameter must be finely controlled by a laser sensor placed at the output of the extruder. To operate under optimal conditions, this requires that the material comes out of the extruder at a high speed, and at the lowest temperature, so this implies the following differences during the production of pellets and filaments:

(i) Pellets: extrusion using the "gravitational" mode of the extruder, where the powder mixture, contained in a funnel, flows by gravity to the extrusion screw, and an outlet temperature of 80°C

(ii) Filament: extrusion using the "feeder" mode, where the powder is pushed by a piston at the level of the extruder screws at a controlled speed, and an outlet temperature of 70 ° C.

Those differences imply a shorter residence time in the extruder for the filament (the residence time being reduced by the force applied by the feeder, which pushes the molten material). This, coupled with an outlet temperature below the melting point of ibuprofen, therefore induces the presence of greater ibuprofen crystalline residues in the filaments than in the pellets (Figure VII-5). However, it is interesting to note that this difference is found to be smoothed out after printing. Indeed, the heating undergone by the filament during printing seems to have been sufficient to amorphized the small crystalline amount of ibuprofen that remained in the filament (54). On the other hand, the use of the "feeder" mode used for the production of the filament requires pretreating the mixture by heating it (otherwise the grain size is too fine, and clogs the "feeder").

This pretreatment has the consequence of increasing the losses of Ibuprofen (Figure VII-4B).

For the filament obtained by hot extrusion, it should also be noted that its diameter is smaller and less homogeneous (1.63 ± 0.10 mm) than the recommendations of the printer manufacturer that the pH drops sharply after 10 days of exposure of the implants obtained by DDM to the release medium, before returning to its initial value on day 16. This drop coincides with the end of the third phase during which there is a rapid release of ibuprofen which is an acidic molecule (pKa = 5.3) and during which PLGA is hydrolyzed and degraded into acidic components (lactic and glycolic acid). Ibuprofen and the degradation products of PLGA are then released into the medium in large quantities and lower the pH. However, this drop in pH is not observed with implants obtained by FFF. This may be since those implants are lighter than the ones obtained by DDM, the degradation products and ibuprofen released is not in sufficient quantity to lead to a comparable drop of pH. (i)

The first phase which takes place during the first day describes a so-called "burst" effect. This is a rapid increase in release due to diffusion of ibuprofen to the surface of implants and due to diffusion of ibuprofen through pores filled with water (55,56).

(ii) When the pores formed by the release of the ibuprofen during the "burst" close due to the mobility of the polymer chains, the second phase of slow diffusion of ibuprofen through the polymer matrix begins (from day 2 to day 4) (57,58). Concerning implants obtained by FFF show a biphasic release pattern (Figure 6 A and B) with:

(i)

The first phase of release (from day 1 to day 4) shows the same "burst" effect as the implants obtained by DDM but over a longer period.

(ii)

The second phase of slower release begins on day 4 until the complete release of the ibuprofen is achieved. mechanical stress compared to FFF technology. The implants printed by DDM have a threephase in vitro release profile with a "burst" effect, then a slower release phase, and finally a final rapid release phase due to the significant swelling of the PLGA. FFF-printed implants exhibit a two-phase in vitro release profile. This difference is mainly due to the difference in structure obtained, leading to heavier implants by DDM, and to a longer diffusion pathway for the drug substance.

VIII. PLGA BASED 3D PRINTED IMPLANTS -IMPACT OF FILLING DENSITY ON DRUG RELEASE

Research article -To be reviewed by the co-authors PLGA based 3D printed implantimpact of filling density on drug release C. Bassand 1 , L. Benabed 1 , J. Verin 1 , J. Freitag 1 , S. Charlon 2 , F. Siepmann 1 , J. Soulestin 2 , J.

Siepmann depending on the filling density of the implant was also investigated. The obtained implants exhibit a triphasic release pattern: with a slight burst effect after 1 day from 5 to 10% of ibuprofen released respectively for the 100% and 10% filling density implants. This is followed by a lag phase up to day 5 or 9 respectively for the 10% and 100% filling density implants. The decrease in the relative drug release rate with increasing implant filling can be attributed to the longer diffusion pathways to be overcome.

Keywords: PLGA, 3D printing, Droplet deposition modeling, Filling density, in vitro drug release.

VIII.2. Introduction

During the past years, (3D) printing technologies revolutionized the pharmaceutical field.

Objects are built by deposition, binding, or polymerization of successive materials layers (1,2).

The desired object can be printed by: photopolymerization, powder bed fusion, material jetting, direct energy deposition, binder jetting, selective deposition lamination, and material extrusion (1,2). Lately, the FDA approved SPRITAM, the first 3D printed dosage to obtain such an approval (1,3). The SPRITAM drug product is by binder jetting (4,5), allowing to reach an extremely high porosity and fast oral dispersing product. Besides binder jetting, four 3D

printing technologies have been used for pharmaceutical applications: photopolymerization (6-9), powder bed fusion (10)(11)(12)(13)(14), material extrusion (15)(16)(17)(18)(19)(20), and material jetting (21,22).

In the past few years, oral controlled release dosage forms obtained by 3D printing have been widely studied (15,21,(23)(24)(25). Lately, an emerging interest is reached on 3D printed dosage forms intended for parenteral administration with polyester matrix (26)(27)(28), and especially poly (lactic-co-glycolic) acid (PLGA) based devices (29)(30)(31)(32)(33)(34). PLGA is a usually used polymer in controlled drug delivery systems for parenteral administration (35)(36)(37). Its excellent biocompatible and biodegradable properties allow avoiding the removal of the device after the treatment (38)(39)(40)(41)(42). Those properties conducted to an intense interest in this polyester (43)(44)(45) and led to his use in numerous FDA-approved drug products, such as microparticles, implants, or in situ forming implants for decades (46,47). From such a polymer, the drug release kinetics can be tailored either by using different polymer grades, devices, porosity, or size (48)(49)(50)(51). The manufacturing process as well influences the release rate (52)(53)(54).

VIII.3.4. Optical microscopy

Pictures of the implant before exposure to the release medium were taken using a SZN-6 trinocular stereo zoom microscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vison Lite 2.1 software). The lengths and implant pore size were determined using the ImageJ software (US National Institutes of Health, Bethesda, Maryland, USA).

VIII.3.5. Practical drug loading

Pellets and pieces of the 3D printed implant were dissolved in 5 mL acetonitrile, followed by filtering (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000

Series HPLC, equipped with an LPG 3400 SD/RS pump, an autosampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversedphase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used. The mobile phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v).

The detection wavelength was 225 nm, and the flow rate was 0.5 mL/min. Ten microliter samples were injected.

VIII.3.6. Differential scanning calorimetry (DSC)

DSC thermograms of the raw materials (PLGA and ibuprofen), pellets and 3D printed implant were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland).

Approximately 5 mg raw material, pellets, and 3D printed implant, were heated in pierced VIII -2). Please note that this decrease might not be entirely due to the ibuprofen effect, but also to the PLGA molecular weight decrease after printing. 10% of ibuprofen released during the burst for the 10% filling density and about 5% of ibuprofen released during the burst for the 100% filling density implants), and longer lag phase (up to day 5 for the 10% filling density implants, and up to day 10 for the 100% filling density implants). This relative decrease in release rate by increasing the filling density of polyesterbased 3D printed implants is consistent with the reported results recently discussed by Zhang et al (26). This could be explained like that: Droplet deposition modeling allows obtaining ibuprofen-loaded 3D printed PLGA implants with different filling densities. Thus, the variation of the filling density is an effective tool to adjust desired absolute and relative release kinetics. The obtained implants exhibit a triphasic release pattern: with a burst release from 5 to 10% during the first days of exposure to the release medium, followed by the drug, and a lag phase up to day 5 or 9 depending on the filling density of the implants. The decrease in the relative drug release rate with increasing implant filling can be attributed to the longer diffusion pathways to be overcome. The onset of the final rapid drug release phase was delayed, because the larger polymer matrices are mechanically more stable, retarding the onset of substantial entire implant swelling, which facilitates drug release.

IX.1. General conclusion

PLGA is a first category choice for polymer-based parenteral controlled drug delivery systems due to its excellent biocompatibility, biodegradability, and use in several FDA-approved products. However, establishing an in vitro in vivo correlation regarding drug release of such devices is challenging as many different mechanisms are involved in drug release. Thus, it is of great importance to properly characterize the devices, to identify the release mechanism by which the API is released from the device.

In the past few years, 3D printing technologies have revolutionized the development of new controlled drug release devices. These technologies allow to precisely control the size, shape, internal and external porosity of the device, and so the release rate of the API.

In this context the research objectives of this thesis were divided into two parts:

(i) Development of experimental set-ups for in vitro drug release measurement for implants administrated to the subcutis, and investigation on the impact of several implants attribute on in vitro drug release;

(ii) Development of PLGA based controlled drug release implants using 3D printing technologies, and investigation of the impact of the structure of the implant on in vitro drug release.

The research results of this Ph.D. thesis had been described in seven parts corresponding to scientific papers that have either been submitted in scientific journals or plan to be. Those investigated led to the following conclusions:

1. How hydrogels surrounding PLGA implants limit swelling and slow down drug release:

The presence of an agarose gel surrounding PLGA implants significantly hinders polymer swelling (sterically) and slows down drug release (due to delayed penetration of substantial amounts of water into the system). In vivo it can be expected that surrounding tissue has a similar mechanical effect. However, yet it is unknown how important the impact of mechanical stress caused by body movements (e.g., muscle contractions) for the fate of a degrading PLGA implant is. The results presented in this study can help developing more realistic in vitro drug release set-ups for parenteral drug delivery systems. They also strengthen the hypothesis that implant swelling plays an orchestrating role for the control of drug release from PLGA-based drug delivery systems.

IX. GENERAL CONCLUSION AND DISCUSSION 255

2.

Hot-melt extruded ibuprofen-loaded PLGA implants: Importance of heat exposure:

Hot melt extrusion offers an interesting potential for the preparation of PLGA-based implants.

However, the exposure to heat can decrease the average polymer molecular weight and alter the physical state of the drug in the polymeric matrix (and of course, thermally degrade heat sensitive drugs). The drug might be present in the form of individual molecules/ions distributed throughout the PLGA network ("dissolved"), or dispersed in the form of crystalline or amorphous tiny particles. It is worth to monitor the state of the drug and potential transformations during implant manufacturing. For instance, increasing exposure times to heat might increase the relative amount of drug, which is dissolved in the PLGA, altering the importance of the initial "burst release". In the present study this phenomenon could be monitored, but its importance on drug release was limited. However, the impact for other drugs or other polymers might be much more pronounced and certain systems might be much less robust and "forgiving".

3.

PLGA implants for controlled drug release: Impact of the diameter:

The variation of the diameter of a PLGA implant is an effective tool to adjust desired absolute and relative release kinetics. In the case of hot melt extruded ibuprofen-loaded implants, biphasic release patterns were observed: A zero order release phase was followed by a rapid, final drug release phase (accounting for 80-90 % drug). The decrease in the relative drug release rate with increasing system diameter can be attributed to the longer diffusion pathways to be overcome. The onset of the final rapid drug release phase was delayed, because the larger polymer matrices are mechanically more stable, retarding the onset of substantial entire implant swelling, which facilitates drug release.

How bulk fluid renewal can affect drug release from PLGA implants:

Great caution has to be paid when defining the conditions for in vitro drug release measurements from PLGA implants: The systems become highly fragile once substantial polymer swelling sets on. The renewal of important parts of the release medium can cause partial damage of the highly fragile polymeric systems, accelerating device disintegration and decreasing the length of the diffusion pathways for the drug to be overcome. Embedding implants into agarose gels and limiting the sampling volume can limit these effects. In any case, it must not be forgotten that, once substantial system swelling sets on, the devices become highly fragile. 256

3D printing of ibuprofen loaded controlled release implants with PLGA -Proof of concept and impact filling pattern:

3D printing by FFF is an interesting tool for the manufacturing of PLGA based control drug delivery devices. The polymer demonstrated viscoelastic properties compatible with the 3D printing technologies. The use of hot-melt extrusion made possible the manufacturing of Ibuprofen-loaded PLGA filament with diameters allowing 3D printing by FFF. Moreover, those filaments were printable without any substantial loss of molecular weight of the polymer nor ibuprofen up to 150°C. 3D printing of implants containing ibuprofen, with three different internal structures was achievable. This leading to the obtention of implants allowing to maintain the release of ibuprofen for up to 9 days, with a monophasic release pattern.

6.

PLGA based 3D printed meshthe impact of the 3D printed technology on drug release:

3D printing has many advantages, such as the flexibility of dosages or the possibility of creating complex structures. However, the choice of printing technology and settings can greatly influence product performance. The objective of this research was, to study and compare the impact of two printing processes on the physicochemical properties of PLGA and the in vitro performance of the implants obtained. For this, two sets of implants were printed using DDM and FFF technology. Both technologies appear to have advantages and disadvantages regarding the production of 3D printing PLGA based control drug delivery devices. DDM technology allows to easily obtain homogeneous drug-loaded implants as the printing is done from a statistical mixture of pellets. Another advantage of this technology is that it allows avoiding the challenging development needed with FFF for the obtention of filament homogeneous in diameter. However, this technology exposes the polymer and the drug substance to high mechanical stress compared to FFF technology. The implants printed by DDM have a threephase in vitro release profile with a "burst" effect, then a slower release phase, and finally a final rapid release phase due to the significant swelling of the PLGA. FFF-printed implants exhibit a two-phase in vitro release profile. This difference is mainly due to the difference in structure obtained, leading to heavier implants by DDM, and to a longer diffusion pathway for the drug substance.

7.

PLGA based 3D printed meshthe impact of filling density on drug release:

Droplet deposition modeling allows obtaining ibuprofen-loaded 3D printed PLGA implants with different filling densities. Thus, the variation of the filling density is an effective tool to IX. GENERAL CONCLUSION AND DISCUSSION 257 adjust desired absolute and relative release kinetics. The obtained implants exhibit a triphasic release pattern: with a burst release from 5 to 10% during the first days of exposure to the release medium, followed by the drug, and a lag phase up to day 5 or 9 depending on the filling density of the implants. The decrease in the relative drug release rate with increasing implant filling can be attributed to the longer diffusion pathways to be overcome. The onset of the final rapid drug release phase was delayed, because the larger polymer matrices are mechanically more stable, retarding the onset of substantial entire implant swelling, which facilitates drug release.

IX.2. Future perspectives

In this context, multiples perspectives could be imagined to this project. Some of them would be listed below, regarding the two main objectives of this thesis: -L'augmentation du temps d'exposition à la chaleur pourrait augmenter la quantité relative de SA, qui est dissoute dans la matrice de PLGA, ce qui modifierait l'importance de l'effet « burst » initial.

-La variation du diamètre d'un implant PLGA est un outil efficace pour ajuster la cinétique de libération absolue et relative désirée.

-La définition des conditions pour les mesures de libération de substance active in vitro à partir d'implants PLGA doit être faite avec attention. -La variation de la densité de remplissage est un outil efficace pour ajuster la cinétique de libération de la SA. Les implants qui ont été imprimés avec différente densités de remplissage, présentent un schéma de libération triphasique : avec une libération « burst » relativement plus élevée et une phase de plateau plus courte pour les implants de densité de remplissage la plus faible. La diminution du taux relatif de libération de SA avec l'augmentation de la densité de remplissage de l'implant est due à des chemins de diffusion allongés.

Mots-clés : PLGA, dispositif d'évaluation de la libération de substance active, gel d'agarose, gonflement, impression 3D, fabrication de filaments fusionnés, dépôt de gouttelettes.

Figure I- 1 . 3 Figure I- 2 . 5 Figure I- 3 . 6 Figure I- 4 . 7 Figure I- 5 . 7 Figure I- 6 . 8 Figure I- 7 . 13 Figure I- 8 . 15 Figure I- 9 . 16 Figure I- 10 . 16 Figure I- 11 .

 132536475768713815916101611 Figures list

Figure I- 12 . 19 Figure I- 13 . 20 Figure I- 14 . 21 Figure I- 15 . 22 Figure I- 16 . 25 Figure I- 17 . 26 Figure I- 18 . 28 Figure I- 19 . 30 Figure I- 20 . 31 Figure I- 21 . 32 Figure I- 22 . 34 ivFigure I- 23 . 35 Figure I- 24 . 36 Figure I- 25 . 38 Figure I- 26 . 40 Figure I- 27 . 41 Figure I- 28 . 42 Figure I- 29 . 43 Figure II- 1 . 77 Figure II- 2 . 78 Figure II- 3 . 80 Figure II- 4 . 82 Figure II- 5 . 83 vFigure II- 6 . 85 Figure II- 7 . 86 Figure II- 8 . 88 Figure III- 1 .Figure III- 2 .Figure III- 3 .Figure III- 4 .Figure III- 5 .Figure III- 6 .Figure III- 7 .Figure III- 8 .Figure IV- 1 .Figure IV- 2 .Figure IV- 3 .Figure IV- 4 .Figure IV- 5 .Figure IV- 6 .Figure IV- 7 .Figure IV- 8 .Figure V- 1 .Figure V- 2 .Figure V- 3 .Figure V- 4 .

 12191320142115221625172618281930203121322234233524362538264027412842294317727838048258368578688812345678123456781234 Figure I-12. The number of publications referring to PLGA per year from 1965 (1 publication) to 2020 (1127 publications). ......................................................................................................................................................... 19 Figure I-13. Schematic representation of the active pharmaceutical ingredient release mechanisms from a poly (lactic-co-glycolic acid) device: (A) diffusion through water-filled pores, (B) diffusion through the polymer matrix, and (C) degradation or erosion of the polymer matrix. Adapted from Fredenberg et al. 2011. ............................ 20 Figure I-14. Change of dry mass and molecular weight of the polymer in the case of (A) surface erosion and (B) bulk erosion. Adapted from Grund et al. 2011. ..................................................................................................... 21 Figure I-15. Schematic representation of the different release patterns of a drug substance from a poly (lactic-coglycolic acid) device. ............................................................................................................................................ 22 Figure I-16. Timeline of three-dimensional printing technologies. ..................................................................... 25 Figure I-17. Schematic representation of photopolymerization-based three-dimensional printing technologies. 26 Figure I-18. Schematic representation of powder bed fusion-based three-dimensional printing technologies. ... 28 Figure I-19. Schematic representation of fused deposition modeling or fused filament fabrication threedimensional printing technologies. ....................................................................................................................... 30 Figure I-20. Schematic representation of "Arburg Plastic Freeforming ® " technology -a proprietary droplet deposition modeling process. ................................................................................................................................ 31 Figure I-21. Schematic representation of material jetting-based three-dimensional printing technologies. ........ 32 Figure I-22. Schematic representation of binder jetting-based three-dimensional printing technologies. ........... 34

Figure V- 5 .Figure V- 6 .Figure VI- 1 .Figure VI- 2 .Figure VI- 3 .Figure VI- 4 .Figure VI- 5 .Figure VI- 6 .Figure VI- 7 .Figure VI- 8 .Figure VII- 1 .Figure VII- 2 .Figure VII- 3 .Figure VII- 4 .Figure VII- 5 .Figure VII- 8 .Figure VII- 9 .Figure VIII- 1 .Figure VIII- 2 .Figure VIII- 3 .Figure VIII- 4 .Figure VIII- 5 .Figure VIII- 6 .Figure VIII- 7 .Figure VIII- 8 .Figure VIII- 9 .

 56123456781234589123456789 Figure V-5. Dynamic changes in the pH of the well agitated bulk fluid, in the implants' dry mass and PLGA polymer molecular weight (Mw) upon exposure of ibuprofen-loaded implants to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates (the receptor compartment containing well-agitated bulk fluid). At pre-determined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated). ............................................................................................................. 169 Figure V-6. Dependence of the solubility of ibuprofen as a function of the pH at 37 °C. Phosphate buffer pH 7.4 USP 42 served as bulk fluid. Its pH was adjusted using of 10 N lactic acid or 0.05 N NaOH (final pH values are reported). ............................................................................................................................................................. 170 Figure VI-1. Schematic representation of the method used to prepare the polymer discs for melt rheology studies. ................................................................................................................................................................ 182 Figure VI-2. Thermograms of (A) the raw materials, (B) their physical mixture, (C) impact of the ibuprofen content on PLGA glass transition temperature measured by differential scanning calorimetry, and (D) PLGA, ibuprofen, and PLGA/ibuprofen physical mixture loss of mass upon heating, obtained by thermogravimetric analysis. ............................................................................................................................................................... 186 Figure VI-3. Effect of temperature on (A) PLGA and PLGA: ibuprofen (85:15) complex viscosity and (B) storage G' and loss G'' modulus. .................................................................................................................................... 188 Figure VI-4. Effect of printing temperature on (A) ibuprofen content and (B) PLGA molecular weight. ........ 189 Figure VI-5. Optical macroscopic pictures of the surface of ibuprofen PLGA-implants printed with the three different filling patterns....................................................................................................................................... 190 Figure VI-6. Thermograms of the ibuprofen loaded PLGA filaments, and the ibuprofen loaded PLGA 3D printed implants. .............................................................................................................................................................. 191 Figure VI-7. (A) Relative (%) and (B) absolute (mg) ibuprofen release from 3D printed PLGA implants, (C) pH changes of the release medium upon exposure to phosphate buffer pH 7.4, and (D) increase in wet mass (%) of the implants after exposure to the release medium. .................................................................................................. 192 Figure VI-8. Optical macroscopy pictures ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 37°C 80 RPM. ......................................................................................................................................... 193

-

  , thus allowing to reach normally undistributed sites or to limit systemic adverse effects According to O'Brien et al., "long-acting injectables are parenteral drug products designed to release an API at a controlled rate to achieve prolonged therapeutic exposure"(5). As stated by the European Directorate for Quality of Medicines & HealthCare (EDQM), polymeric based long-acting injectables include (In situ gelling systems. They have been developed for use via various administration routes such as subcutaneous, intramuscular, epidural, intra articular, and surgical insertion directly at the tissue of interest (3).

-Figure I- 2 .

 2 Figure I-2. Requirement for long-acting injectables formulations. Mandatory attributes are reported in black and desirable ones in grey. Adapted from Selmin et al. 2020.

. 1 . 3 .

 13 Testing in vitro drug release from dosage forms designed to subcutis I.1.3.1. Referenced methods I.1.3.1.1. Sample and separate method The sample and separate method consist in the for replenishment of the release medium either via filtration or centrifugation to maintain all the formulation in the release equipment as illustrated in Figure I-3.

Figure I- 3 .

 3 Figure I-3. Schematic illustration of sample and separate in vitro drug release method. Figure created using (20).

Figure I- 4 .

 4 Figure I-4. Schematic illustration of the United States Pharmacopeia apparatus (A) I: basket, and (B) II: paddle.

Figure I- 5 .

 5

Figure I- 5 .

 5 Figure I-5. Schematic illustration of the United States Pharmacopeia apparatus IV: continuous flow-through method.

Figure I- 6

 6 Figure I-6 (13).

Figure I- 6 .

 6 Figure I-6. Schematic illustration of the dialysis-based in vitro drug release method. Figure created using (20).

drug release testing method I. 1 . 3 . 2 . 1 .

 1321 Instrument based parametersAccording to the in vivo site of administration, some parameters can be adjusted depending on the in vitro setup used. Instrument based parameters can be listed as:

Figure I- 7 ,

 7 in the case of polycondensation, linear polymers of low average molecular weight (Mw) are obtained by intramolecular esterification of alpha-hydroxy acids, while ring-opening polymerization allows obtaining linear polymer of higher Mw by ring polymerization of the cyclic monomers(38). In the latter case, the use of a catalyst is necessary.

Figure I- 7 .

 7 Figure I-7. Synthesis of Poly (lactic-co-glycolic acid) by (A) polycondensation and (B) ringopening polymerization.

Figure I- 8 .

 8 Figure I-8. Summary of the main characteristics, and schematic arrangement of molecular chains in crystalline, semi-crystalline, and amorphous polymers.

Figure I- 9

 9 where n, m, and nm represent different degrees of polymerization, and H + the acidic catalyst coming from the carboxylic end group of the polymer.

Figure I- 9 .

 9 Figure I-9. The acid-catalyzed ester hydrolysis of poly (lactic-co-glycolic acid). Adapted from Ford Versypt et al. 2013.

Figure I- 10 .

 10 Figure I-10. Illustration of the size-dependent autocatalytic effect. Adapted from Ford Versypt et al. 2013.

Figure I- 11

 11 Figure I-11 the metabolites produced by hydrolysis are safely eliminated by the human body, either by the urine or directly integrated into the citric acid cycle and transformed in water and carbon dioxide.

Figure I- 11 .

 11 Figure I-11. Schematic representation of the elimination and metabolization of the poly (lactic-co-glycolic) degradation products. Adapted from Maurus et al. 2004.

Figure I- 12 .

 12 Figure I-12. The number of publications referring to PLGA per year from 1965 (1 publication) to 2020 (1127 publications).

-

  Formation of cracks, deformation of the device, -The collapse of the polymer structure. In addition, as illustrated in Figure I-13, Fredenberg et al. identified three-way for an API to be released from PLGA-based systems: diffusion through water-filled pores, diffusion through the polymer matrix, and degradation or erosion of the polymer matrix (35).

Figure I- 13 .

 13 Figure I-13. Schematic representation of the active pharmaceutical ingredient release mechanisms from a poly (lactic-co-glycolic acid) device: (A) diffusion through water-filled pores, (B) diffusion through the polymer matrix, and (C) degradation or erosion of the polymer matrix. Adapted from Fredenberg et al. 2011.

  Faisant et al. identified two types of erosion: surface erosion and mass erosion ass illustrated in Figure I-14 (72,97).

Figure I- 14 .

 14 Figure I-14. Change of dry mass and molecular weight of the polymer in the case of (A) surface erosion and (B) bulk erosion. Adapted from Grund et al. 2011.

  Figure I-15 describes the three types of release patterns: mono-phasic, bi-phasic, or tri-phasic.

Figure I- 15 .

 15 Figure I-15. Schematic representation of the different release patterns of a drug substance from a poly (lactic-co-glycolic acid) device.

Figure I- 17 .

 17 Figure I-17. Schematic representation of photopolymerization-based three-dimensional printing technologies.

Figure I- 19 .

 19 Figure I-19. Schematic representation of fused deposition modeling or fused filament fabrication three-dimensional printing technologies.

Figure I- 21 (

 21

Figure I- 21 .

 21 Figure I-21. Schematic representation of material jetting-based three-dimensional printing technologies.

Figure I- 22

 22 Binder Jetting (BJ) printers spray a binder solution or liquid formulation in thin layers of powder material (121,131,148,149). The powder is then bound by the formation of binding bridges, by dissolution or recrystallization. The unbound powder serves as a support for the printed object (121,131,148,149). When one layer is complete, the powder bed moves downward and a new layer of powder is sprayed onto the build surface. The process is repeated until all parts are completed. It is also possible to place the parts in the raw state inside an oven to obtain the sintering of the grains.

Figure I- 22 .

 22 Figure I-22. Schematic representation of binder jetting-based three-dimensional printing technologies.

Figure I- 23 (Figure I- 23 .

 2323 Figure I-23. Schematic representation of direct energy deposition-based three-dimensional printing technologies.

  Sheet Lamination (SL) 3D printing technology is different from all the previously described 3D printing strategies (121). This type of printer does not incorporate conventional 3D printing materials but sheets of paper, plastic, or metal as illustrated in Figure I-24. It consists of adding layer by layer an iron-on sheet cut by a laser or a blade. Between each cutting, the plate descends and a new sheet, glued by a binding agent, is unwound. A roller then exerts pressure to make it adhere to the surface. These steps are repeated until the 3D object is obtained. The new sheet can be bonded to the previous one by adhesive bonding, heat sealing, clamping, or ultrasound technique (121,131,154,155).

Figure I- 24 .

 24 Figure I-24. Schematic representation of sheet lamination-based three-dimensional printing technologies.

  previously described the gain of interest of 3D printing technologies in the pharmaceutical field is extensive. The rise of these new technologies opened an infinite range of possibilities for modulating drug release. The conventional pharmaceutical technologies allowed to modulate the release pattern by modulating the size of classical shape devices (spheres, cylinders, etc.), or the selection of the appropriate carriers (polymer grade, molecular weight, etc.) (42,43,60-62,64). 3D printing technologies have revolutionized the way to develop new devices for the controlled release of API. With the appropriate choice of polymer and technology, it could be possible to precisely modulate drug release by precisely modulating either external structure (innovative geometry), the internal structure of the device (infill percentage or pattern), or size and structure of the device, as illustrated in Figure I-26 (134,135,164,173-179).

Figure I- 26 .I. 3 . 3 . 2 . 1 .Figure I- 27 .

 26332127 Figure I-26. Illustration of the different possibilities to modulate the structure to control the drug release from 3D printed devices. Adapted from Patel et al 2021.

Figure I- 28 .

 28 Figure I-28. Examples of reported studies on the effect of the internal structure of 3D printed devices on drug release. Adapted from (A) Yang et al. (181), (B) Zhang et al.(182), and (C) Welsh et al. (138).

3 .

 3 Figure I-29. Modulating both the size and the internal structure of the device by 3D printing is of great interest in modulating drug release rate (183-185).

Figure I- 29 .

 29 Figure I-29. Examples of reported studies on the effect of size and structure of 3D printed devices on drug release. Adapted from (A) Xu et al. (183), (B) Tan et al.(184), and (C) Russi et al. (185).

  3D printing was employed to develop dexamethasone-loaded punctal plugs which showed sustained release of dexamethasone for up to 7 days → it was shown that size has a significant impact on release rate Customization of the release duration by tuning different tablet height with 3 different API (B)

3 .

 3 PLGA implants for controlled drug release: Impact of the diameter 4. How bulk fluid renewal can affect drug release from PLGA implants 5. 3D printing of ibuprofen loaded controlled release implants with PLGA -Proof of concept and impact filling pattern 6. PLGA based 3D printed meshthe impact of the 3D printed technology on drug release 7. PLGA based 3D printed meshthe impact of filling density on drug release II. HOW HYDROGELS SURROUNDING PLGA IMPLANTS LIMIT SWELLING AND SLOW DOWN DRUG RELEASE Research article -Submitted in Journal of Controlled Release How hydrogels surrounding PLGA implants limit swelling and slow down drug release C. Bassand, J. Verin, M. Lamatsch, F. Siepmann, J. Siepmann Univ. Lille, Inserm, CHU Lille, U1008, F-59000 Lille, France II.1. Abstract

Figure II- 1 B

 1 Figure II-1 B (1 implant per tube). The gel was prepared with 0.5% w:v agarose and 1 mL phosphate buffer pH 7.4 USP 42 (optionally containing 0.0025 % w:v bromocresol green, bromothyol blue, bromophenol blue or phenol red as pH indicator). An agarose dispersion in the respective buffer solution was heated to 100 °C under magnetic stirring (250 rpm) until a clear solution was obtained. The latter was cooled to 47°C and continuously stirred (to prevent gelation). 0.5 mL of the solution was placed into the bottom of an Eppendorf tube and cooled in a refrigerator for 5 min to allow for gelation. An implant was carefully placed on top of the gel, and covered with second layer of 0.5 mL agarose solution (47 °C), followed by cooling in a refrigerator for 5 min. Four mL phosphate buffer pH 7.4 USP 42 (optionally containing a pH indicator) were added on top of the gel, and the tube was placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the entire bulk fluid was replaced by fresh (pre-heated) release medium. The withdrawn samples were treated as in the case of drug release measurements in well agitated bulk fluids.

  Figure II-1 illustrate the 3 experimental set-ups, which were used to monitor ibuprofen release. The idea was to evaluate the potential impact of the presence of a hydrogel around the implants (mimicking living tissue) on the resulting drug release kinetics.

Figure II- 1 A

 1 Figure II-1 A shows the "Bulk fluid" set-up, in which an implant is placed into a metal basket in an Eppendorf tube filled with 5 mL phosphate buffer pH 7.4. The tubes are placed into ahorizontal shaker (80 rpm) and kept at 37 °C. At pre-determined time points, the entire bulk fluid is replaced by fresh (pre-heated) release medium. The metal basket assured that the implant did not sink to the bottom of the tube. Its meshes were sufficiently large (250 µm) to allow for convective flow and rapid medium exchange between the liquid inside and outside the basket.

Figure II- 1 B

 1 Figure II-1 B illustrates the "gel -Eppendorf" set-up: In this case, an implant is surrounded by a 0.5 % agarose gel prepared with 1 mL phosphate buffer pH 7.4. Four mL phosphate buffer pH 7.4 are carefully added on top of the gel, and the Eppendorf tube is placed into a horizontal shaper (80 rpm) at 37 °C. At pre-determined time points, the entire bulk fluid is replaced by fresh (pre-heated) phosphate buffer pH 7.4, and the drug content in the withdrawn samples is

  the observed ibuprofen release kinetics, the following reference experiment was conducted: Drug release from a 0.5 % agarose gel, which was prepared with a 200 µg/mL solution of ibuprofen in phosphate buffer pH 7.4, was measured with the same set-up. The black circles in Figure II-1 D show the obtained results: More than 50 % of the drug was released after 10 h. This is rapid compared to the ibuprofen release rate from the investigated PLGA implants (which was of the order of 10 d).The "gel -transwell" set-up is schematically shown in FigureII-1 C. In this case, an implant is placed in the donor compartment of a transwell plate, being embedded in a 0.5 % agarose gel prepared with 1 mL phosphate buffer pH 7.4. Four mL phosphate buffer pH 7.4 are placed into the acceptor compartment. The transwell plate is horizontally shaken (80 rpm) at 37 °C. At predetermined time points, the entire bulk fluid in the acceptor compartment is replaced by fresh (pre-heated) release medium. From a practical point of view, this is easier than replacing the release medium in the "gel -Eppendorf" set-up. The drug content in the withdrawn samples is determined by HPLC-UV analysis. Thus, also in this case, the presence of the agarose gel can be expected to slow down ibuprofen release to a certain degree. In addition, the presence of the membrane of the transwell plate might impact the rate at which the ibuprofen reaches the acceptor compartment. To evaluate the relative importance of these 2 phenomena, the following reference experiments were conducted: (i) The "release rate" of an ibuprofen solution in phosphate buffer pH 7.4 (200 µg/mL) from the donor compartment (free of agarose gel) into the acceptor compartment was measured. The black triangles in Figure II-1 D show that the entire drug amount was released in less than 10 h under these conditions. This is very rapid compared to the release periods from the investigated implants in this study (>= 10 d). Thus, the impact of the transwell plate membrane can likely be neglected. (ii) Ibuprofen release from a 0.5 % agarose gel, prepared with an ibuprofen solution in phosphate buffer pH 7.4 (200 µg/mL) was measured using this set-up. The open squares in Figure II-1 D illustrate the

Figure II- 1 .II. 4 . 2 .

 142 Figure II-1. Schematic presentations of the experimental set-ups used to monitor drug release from PLGA-based implants in: A) well agitated phosphate buffer pH 7.4 (in metal baskets) in Eppendorf tubes, B) agarose gels in Eppendorf tubes, the gels being exposed to well agitated phosphate buffer pH 7.4, C) agarose gels in transwell plates, the receptor compartment containing well agitated phosphate buffer pH 7.4. In all cases, sink conditions were provided throughout the experiments in the well agitated bulk fluids. Optionally, pH indicators were added to the phosphate buffer. Details are described in the text. The diagram in D) shows the rates at which an ibuprofen solution (200 µg/mL) was "released" from the donor compartment in a transwell plate free of gel, or from a gel in an Eppendorf tube or from a gel in a transwell plate.

Figure II- 2 .

 2 Figure II-2. Ibuprofen release from and swelling of PLGA implants upon exposure to phosphate buffer pH 7.4, observed using 3 experimental set-ups: In bulk fluids in Eppendorf tubes, in agarose gels exposed to the release medium in Eppendorf tubes, and in agarose gels in transwell plates (the acceptor compartment containing the release medium). Please note the different scaling of the x-axes on the left versus the right hand side.

  early time points is negligible from a relative point of view (the 100 % reference value being considerably higher).The diagrams in the middle and at the bottom of Figure II-2 show the dynamic changes in the wet mass and volume of the PLGA implants upon exposure to the release medium when using the 3 different set-ups. Clearly, all systems started to fundamentally swell after a "lag phase". Importantly, as in the case of drug release, there was a clear impact of the presence of a hydrogel surrounding the PLGA implant: The onset of substantial system swelling was delayed by several days. The optical macroscopy pictures in Figure II-4 illustrate this behavior: At the top, images of surfaces are shown, at the bottom images of cross-sections. The implants were exposed to the release medium for up to 10 d using the 3 different set-ups (as indicated). In all cases, system swelling was limited during the first few days, followed by the onset of substantial PLGA swelling. When looking at the 3 diagrams on the right hand side of Figure II-2 (showing drug release as well as changes in the systems' wet mass and volume during the first 10 d), it

Figure II- 3 .

 3 Figure II-3. SEM pictures of surfaces and cross sections of ibuprofen-loaded PLGA implants before and after exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups. The type of set-up and exposure times are indicated on the left hand side. Please note that after exposure to the release medium the implants were freeze-dried prior to analysis. Thus, caution must be paid due to artefact creation.

Figure II- 4 .

 4 Figure II-4. Optical macroscopy pictures of surfaces and cross sections of ibuprofen-loaded PLGA implants before and after exposure to phosphate buffer pH 7.4 using 3 experimental setups: In bulk fluids in Eppendorf tubes, in agarose gels exposed to the release medium in Eppendorf tubes, and in agarose gels in transwell plates (the acceptor compartment containing the release medium). The exposure times are indicated at the top, the type of set-up is indicated on the left hand side

  since the latter are in contact with high amounts of water (in contrast to regions deeper inside the implant). The presence of a highly swollen, surface-near matrix layer was also visible in the optical macroscopy pictures shown in Figure II-4, which were obtained without sample drying. The red rectangles in the SEM pictures on the right hand side of Figure II-3 highlight the two zones which can be distinguished: A highly swollen surface-near layer and a "non swollen" layer located below. The thickness of the highly swollen surface-near zone increases with time. This growth is due to the fact that high amounts of water are present in the highly swollen outmost layer and are, thus, in contact with the PLGA in the layer right below. Please note that the scheme in Figure II-5 does not reflect this phenomenon for reasons of simplicity.

Figure II- 5 .II. 4 . 3 .

 543 Figure II-5. Simplified schematic presentation of the mass transport mechanisms controlling ibuprofen release from the investigated PLGA implants. Initially, limited amounts of water diffuse into the system, leading to polyester degradation throughout the implants ("bulk erosion"). As soon as a critical polymer molecular weight is reached, substantial amounts of water penetrate into the device, facilitating subsequent drug release. Details are described in the text.

  at day 21 led to a much less important decrease in the pH of the bulk fluids. It has to be pointed out that the pH values shown in Figure II-6 were measured in the agitated bulk fluids in all set-ups. In the case of the "gel -Eppendorf" and "gel -transwell" set-ups, the implants were not in direct contact with this bulk fluid. This is why also potential dynamic changes in the pH within the agarose gels were monitored during drug release. Three pH indicators [phenol red (PR), bromothymol blue (BTB) and bromocresol green (BCG)] were added to the phosphate buffer pH 7.4, which was used for the preparation of the gels. On the left hand side of Figure II-7, the pH values are shown at which the indicators change their color.

Figure II- 6 .Figure II- 7 .

 67 Figure II-6. Dynamic changes in the pH of the well agitated bulk fluids, dry mass (%) of the implants and PLGA polymer molecular weight (Mw) upon exposure of the implants to phosphate buffer pH 7.4 in the 3 experimental set-ups: In bulk fluids in Eppendorf tubes, in agarose gels exposed to the release medium in Eppendorf tubes, and in agarose gels in transwell plates (the acceptor compartment containing the release medium). Please note the different scaling of the x-axes on the left versus the right hand side. The Asterix indicates that the average polymer molecular weight (Mw) was below 5 kDa.

Figure II- 8 .

 8 Figure II-8. Optical macroscopy pictures of surfaces and cross sections of ibuprofen-loaded PLGA implants before and after exposure to phosphate buffer pH 7.4 containing 0.0025 % phenol red (PR), bromothymol blue (BTB), bromocresol green (BCG) or bromophenol blue (BPB). Three experimental set-ups were used: Bulk fluids in Eppendorf tubes, agarose gels exposed to the release medium in Eppendorf tubes, and in agarose gels in transwell plates (the acceptor compartment containing the release medium). The type of set-up and exposure times are indicated at the top, the type of pH indicator is given on the left hand side. The Asterix indicates that the respective samples were too fragile to be cut.

  Figure shows that the decrease in polymer molecular weight of the PLGA is not significantly affected by the type of experimental set-up: In all cases, the length of the macromolecules exponentially decreased from the beginning, indicating pseudofirst order kinetics. After day 6, the values were too small to be reliably detected by the applied GPC method. These results suggest that the relatively rapid entire implant wetting and subsequent polyester bond cleavage during the first couple of days are not affected by the type of experimental set-up. This is consistent with the hypothesized drug release mechanism: The absence or presence of a hydrogel around the implant does not alter the rate at which the limited amounts of water diffuse into the system upon contact with phosphate buffer pH 7.4 to a noteworthy extent. Hence, also PLGA degradation throughout the polymer matrix is not affected during this initial phase.Interestingly, the absence of a noteworthy effect of the type of experimental set-up on PLGA degradation during this early phase strengthens the hypothesis that implant swelling (and not PLGA degradation) orchestrates drug release: While the decrease in polymer molecular weight is very similar during the first 6 d in the 3 experimental set-ups (Figure II-6 at the bottom), first indications for the hindrance of system swelling by the presence of an agarose gel are visible: as reflected by differences in the wet mass & volume changes (Figure II-2) as well as by optical macroscopy pictures (FigureII-4). These differences in system swelling affect drug mobility and drug release as well as the mobility of the water-soluble PLGA degradation products and, thus, implant erosion (dry mass loss).

  or 15 min in an oven (FP115, Binder, Tuttlingen, Germany) (FigureIII-1 A). The molten blend was manually extruded using the syringe. The extrudate was cut with a hot scalpel into cylindrical implants of approximately 5 mm length.

Figure III- 1 .III. 3 . 3 .

 133 Figure III-1. Schematic presentations of the experimental set-ups used to: (A) prepare ibuprofen-loaded PLGA implants by hot melt extrusion, and (B) monitor drug release from the implants upon exposure to phosphate buffer pH 7.4 in well agitated Eppendorf tubes. Details are described in the text.

Figure III- 2 .

 2 Figure shows SEM pictures of surfaces and cross sections of the investigated ibuprofenloaded PLGA implants before exposure to the release medium. The heating times applied during manufacturing are indicated on the left-hand side. As it can be seen, the internal and external structure was dominantly non-porous (some round shaped cavities might result from

Figure III- 2 .

 2 Figure III-2. Optical macroscopy pictures and SEM pictures of surfaces and cross sections of ibuprofen-loaded implants before exposure to release medium. The PLGA-ibuprofen blends were heated for 3 to 15 min during implant preparation (as indicated). Drug crystals are highlighted by dotted red circles.

Figure 4 ,

 4 Figure 4, the glass transition temperature (Tg) of the blends decreased with increasing ibuprofen content up to about 10-15 % drug and then seems to level off. This indicates that ibuprofen acts as a plasticizer for PLGA and, under the conditions provided during the DSC measurements, at least about 10 % (w/w) drug can dissolve in the polymer. Interestingly, the Tg values of the implants prepared by hot melt extrusion was about constant in the investigated heating time range (varying between about 33-34 °C, TableIII-1; please also see the DSC thermograms in Figure III-3). This suggests that already after 3 min heating, major parts of the ibuprofen are dissolved in the PLGA. This hypothesis is consistent with the amounts of crystalline ibuprofen determined in the PLGA implants when integrating the surfaces of the melting peaks observed during the DSC measurements (Figure III-3): The amounts of crystalline drug are relatively

Figure III- 4 A

 4 , all implants exhibited sharp diffraction peaks as well as an underlying amorphous halo. The X-ray diffraction patterns of the raw materials are shown in Figure III-4 B for reasons of comparison. Clearly, the ibuprofen (as received) was crystalline, and the PLGA amorphous (confirming the DSC data discussed above). Importantly, the positions of the diffraction peaks observed with the drug-loaded implants corresponded well to the positions of the crystalline ibuprofen raw material. The dashed orange ovals in Figure III-4 A highlight some of these peaks.

Figure III- 4 .

 4 Figure III-4. X-ray diffraction patterns of the: (A) investigated ibuprofen-loaded PLGA implants, and (B° raw materials (PLGA & ibuprofen) (for reasons of comparison). During implant preparation, ibuprofen-PLGA blends were heated for 3 to 15 min (as indicated). The

Figure III- 5

 5 show the resulting ibuprofen release kinetics from the investigated implants upon exposure to phosphate buffer pH 7.4 in well agitated Eppendorf vials at 37 °C (the set-up is schematically shown in FigureIII-1 B). The diagram on the righthand side is a zoom on the first 10 d of drug release. The heating time applied during implant manufacturing is indicated in the legends. As it can be seen, tri-phasic drug release was observed in all cases: A (limited) initial rapid drug release phase ("burst release") during the first few hours was followed by a zero-order release phase (with an about constant drug release rate), and a final, again more rapid drug release phase, leading to complete drug exhaust. The impact of the heating time (3 to 15 min) during implant preparation was only minor.

Figure III- 5 .

 5 Figure III-5. Ibuprofen release and swelling behavior of the investigated PLGA-based implants upon exposure to phosphate buffer pH 7.4. The diagram on the right-hand side at the top is a zoom on the first 10 h. The heating times of the ibuprofen-PLGA blends during implant preparation are indicated in the diagrams.

about 1

 1 week exposure to the release medium (Figure III-5). This coincides with the onset of the penetration of substantial amounts of water into the systems, as illustrated in the diagrams in the middle and at the bottom of Figure III-5: showing the dynamic changes in the volume and wet mass of the implants upon contact with the release medium. This was true for all the investigated heating times. The optical macroscopy pictures shown in Figure III-6 illustrate these considerable changes in implant size and morphology: During the first couple of days the system dimensions remain about constant. But after 1 week the implants substantially swell, irrespective of the heating time applied during manufacturing. The pictures on the right-hand side of Figure III-6 were obtained with light coming from the bottom (for the other pictures the light came from the top). As it can be seen, after 10 d exposure to the release medium, some darker regions (= probably less swollen) are visible close to the center of the implants. Figure III-7 shows SEM pictures of surfaces and cross sections of the implants prepared with 3 to 15 min heating time after 2, 6 and 8 d exposure to the release medium. Please note that the implants were freeze-dried prior to SEM analysis. Thus, the observed structures are in great part artefacts. The optical macroscopy pictures in Figure III-6 indicate that upon contact with the aqueous bulk fluid the implant surface starts to swell (well before substantial swelling of the entire system sets on after about 1 week). This is consistent with the highly wizened and porous surface structure observed by SEM with all implants (Figure III-7): Upon drying during sample preparation, the highly swollen, thin suface layer of "PLGA gel" shrinks.

Figure III- 6 .

 6 Figure III-6. Optical macroscopy pictures of the investigated ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4. At the top, the time periods of implant exposure to the release medium are shown. On the left-hand side, the heating times of the ibuprofen-PLGA blends during implant preparation are indicated. The pictures on the right-hand side were obtained with light coming from the bottom.

Figure III- 7 .

 7 Figure III-7. SEM pictures (surfaces and cross sections) of the investigated ibuprofen-loaded PLGA implants after 2, 6 and 8 days exposure to phosphate buffer pH 7.4. The heating times (in min) of the ibuprofen-PLGA blends during implant preparation are indicated on the lefthand side. The dotted red circles highlight drug crystals, the dotted red rectangles surface-near regions including a highly swollen surface layer and the "not yet swollen" inner implant region.

Figure III- 8 .

 8 Figure III-8. Dynamic changes in the: (A) average polymer molecular weight (Mw) of the PLGA, and (B) dry mass of the investigated ibuprofen-loaded implants upon exposure of the systems to phosphate buffer pH 7.4. The heating times of the ibuprofen-PLGA blends during implant preparation are indicated in the diagrams.

. 3 . 6 . 1 .

 361 In well agitated bulk fluidsImplants were placed in 5 mL Eppendorf tubes (1 implant per tube), filled with 5 mL phosphate buffer pH 7.4 USP 42 (FigureIV-1 A). Metal baskets avoided that the implants sank to the bottom of the tubes (resulting in potentially limited contact with the bulk fluid). The mesh size (250 µm) was sufficient to allow for convective flow and rapid medium exchange "insideoutside" the basket. The tubes were placed in a horizontal shaker (80 rpm, 37°C; GFL 3033;

1 .

 1 13 cm², 11 µm, 0.4 µm pore size; Nunc, Roskilde, Denmark), as illustrated in Figure IV-1 C.

  5 mL gel). The well plates were filled with 4 mL phosphate buffer pH 7.4, covered with lids and Parafilm to minimize evaporation, and placed in a horizontal shaker (80 rpm, 37°C; GFL 3033). At predetermined time points, the entire bulk fluid was replaced by fresh (pre-heated) release medium. The withdrawn samples were treated as for the drug release measurements in well agitated bulk fluids.In all cases, sink conditions were provided throughout the experiments in the well-agitated bulk fluids. Furthermore, the pH of the bulk fluids was measured at pre-determined time points using a pH meter (InoLab pH Level 1; WTW, Weilheim, Germany). All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

Figure IV- 1 .

 1 Figure IV-1. Schematic presentations of the experimental set-ups used to monitor drug release from the PLGA-based implants: (A) In well-agitated release medium in Eppendorf tubes (with metal baskets), (B) In agarose gels in Eppendorf tubes, the gels being exposed to well-agitated release medium, (C) In agarose gels in transwell plates, the receptor compartment containing well-agitated release medium. In all cases, sink conditions were provided throughout the experiments in the well-agitated bulk fluids. Details are described in the text.

  Implants were treated as for the in vitro drug release studies described above. At pre-determined time points: (i) Pictures of the implants were taken with a SZN-6 trinocular stereo zoom macroscope (Optika), equipped with an optical camera (Optika Vision Lite 2.1 software). The lengths and diameters of the implants were determined using the ImageJ software (US National Institutes of Health). Dynamic changes in the systems' volume were calculated considering cylindrical geometry.(ii) Implant samples were withdrawn and excess water was carefully removed using Kimtech precision wipes (Kimberly-Clark, Rouen, France) and weighed [wet mass (t)]. The change in wet mass (%) (t) was calculated as follows:𝑐ℎ𝑎𝑛𝑔𝑒 𝑖𝑛 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (%)(𝑡) = 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (𝑡) -𝑚𝑎𝑠𝑠 (𝑡 = 0) 𝑚𝑎𝑠𝑠 (𝑡 = 0) × 100 %

  -75 -70 -65 °C: die -zone 3zone 2 -zone 1). PLGA is partially cut into shorter chains at elevated temperatures. This also explains the slightly higher glass transition temperatures of the implants prepared in this study (around 35-37 °C) compared to the implants of similar composition prepared at via a heat treatment at 105 °C for 3 to 15 min in a syringe (about 33-34 °C)(47).

Figure IV- 2 .

 2 Figure IV-2. Optical macroscopy pictures of implants before exposure to release medium. The diameter of the die used for hot melt extrusion is indicated at the top.

Figure IV- 3 .

 3 Figure IV-3. DSC thermograms of: (A) PLGA raw material and the investigated ibuprofenloaded implants (the diameter of the implants are indicated on the left hand-side), and (B) ibuprofen raw material.

Figure IV- 4

 4 Figure show SEM pictures of surfaces and cross sections of the differently sized implants before exposure to the release medium. The system diameters are indicated on the left handside. As it can be seen, all surfaces and cross sections were smooth and essentially non-porous.

Figure IV- 1

 1 shows the 3 experimental set-ups, which were used to monitor drug release from the investigated PLGA implants in this study: (i) Samples were exposed to well agitated phosphate buffer pH 7.4 in Eppendorf tubes (metal baskets with 250 µm meshes assuring that the systems did not sink to the bottom of the tubes). (ii) Implants were embedded into agarose gels, which were placed into Eppendorf tubes and exposed to well agitated phosphate buffer pH 7.4. (iii) Implants were embedded into agarose gels, which were placed into the donor compartments of transwell plates, the acceptor compartments containing well agitated phosphate buffer pH 7.4. The agarose gels were intended to better mimic the conditions upon subcutaneous administration. The top row in Figure IV-5 shows the relative drug release rates from the differently sized PLGA implants observed in the 3 set-ups. Clearly, in all cases, an increase in the implant diameter led to slower relative ibuprofen release. But please note that the opposite trend was observed when looking at the respective absolute drug release rates (bottom row in FigureIV-5). Normalizing the absolute ibuprofen release to the available surface area leads to about similar release rates at the beginning. This is sound, because the drug and polymer are the same, and the composition as well as the inner and outer system structures are very similar (TableIV-1, Figure IV-2 , FigureIV-4). However, the systems differ substantially in their total absolute drug loadings and length of the diffusion pathways to be overcome by the drug. This is why the smaller implants were much more rapidly depleted of ibuprofen compared to the larger implants (e.g. bottom row in FigureIV-5).

Figure IV- 5 .

 5 Figure IV-5. Implant behavior of the ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups: Relative (%) and absolute (mg) ibuprofen release kinetics (top and bottom), dynamic changes in the implant volume and pH of the well agitated bulk fluids (middle rows). The implant diameters are indicated in the diagrams.

diagrams in the 2 *Figure IV- 6 .

 26 Figure IV-6. Optical macroscopy pictures ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 in the 3 experimental setups: In bulk fluids in Eppendorf tubes, in agarose gels exposed to the release medium in Eppendorf tubes, and in agarose gels in transwell plates containing the release medium in the acceptor compartment. In the case of the bulk fluids, the implants became too fragile after 8 or 10 d to allow sample withdrawal for imaging.

Figure IV- 7 .

 7 Figure IV-7. Zoom on the first 10 d of implant exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups: Relative drug release, dynamic changes in the wet mass & dry mass of the systems and in the average polymer molecular weight of the PLGA. The implant diameters are indicated in the diagrams.

  polymer molecular weight measured by GPC (bottom row in Figure IV-7). However, most of the PLGA matrix remains dense (PLGA being hydrophobic, limiting the amounts of water within the systems) and the ibuprofen molecules are effectively entrapped. Only surface-near regions undergo substantial swelling: Figure IV-8 shows SEM pictures of surfaces and cross sections of the investigated PLGA implants after 2, 4 and 8 d exposure to the release media.

  cleavage of an ester bond, a new hydrophilic -OH and a new hydrophilic -CCOH group are generated. This renders the degrading polymer more and more water-loving. In addition, the mechanical stability of the 3-dimensional polymer network decreases (since the macromolecules become shorter and less entangled, Figure IV-7 bottom row). In addition, water-soluble short chain acids accumulate and create a continuously increasing osmotic pressure. This leads to substantial PLGA swelling once the polymer is sufficiently hydrophilic and mechanically instable. Importantly, this phenomenon is initially limited to surface-near regions, as it can be seen in the cross sections in Figure IV-8. With time these highly swollen

phases (Figure IV- 5 ,

 5 top row).

Figure IV- 8 .

 8 Figure IV-8. SEM pictures of surfaces and cross sections of ibuprofen-loaded PLGA implants after exposure for 2, 4 or 8 d to phosphate buffer pH 7.4 (as indicated on the left hand-side).

Figure IV- 7 .

 7 Figure Before, the dry mass loss is limited, because the polymer matrix is dense and poorly permeable.

Figure IV- 5

 5 Figure IV-8 (SEM pictures), it can be seen that the presence of the agarose gel surrounding the implants delays substantial device swelling. This is probably due to the steric hindrance by the gels. As a consequence, also the onset of the final rapid drug release phase is delayed in the gel

Figure V- 1 .

 1 Figure V-1. Schematic presentations of the experimental set-ups used to monitor drug release from the PLGA-based implants: (A) In well-agitated release medium in Eppendorf tubes, (B)In agarose gels in Eppendorf tubes, the gels being exposed to well-agitated release medium,

2 V. 4 . 1 .

 241 -loaded PLGA implants before exposure to the release medium (Tg: glass transition temperature, Tm: melting temperature). Mean values ± standard deviations are indicated (n=3). In vitro drug release and implant swelling The diagrams at the top of Figure V-2 show the observed ibuprofen release kinetics from the PLGA implants. The experiments were conducted in triplicate, mean values +/-standard deviations are reported. In each case, a single implant was: (i) directly exposed to 5 mL phosphate buffer pH 7.4 in an Eppendorf tube (Figure V-1A). The latter was horizontally shaken at 80 rpm at 37 °C. A metal basket avoided that the implant could sink to the bottom of the tube. Its mesh size (250 µm) was sufficiently large to allow for convective flow and rapid medium exchange between the fluid inside and the outside of the basket. (ii) embedded within an agarose gel (1 mL), which was exposed to 4 mL phosphate buffer pH 7.4 in an Eppendorf tube (Figure V-1B). The latter was horizontally shaken at 80 rpm at 37 °C. (iii) embedded within an agarose gel (1 mL), located in the donor compartment of a transwell plate (Figure V-1C). The acceptor compartment was filled with 4 mL phosphate buffer pH 7.4. The transwell plate was horizontally shaken at 80 rpm at 37 °C.

Figure V- 2 .

 2 Figure V-2. Drug release, increase in volume and change in wet mass of ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates (the receptor compartment containing well-agitated bulk fluid). At pre-determined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated).

  partial replacement (Figure V-2, diagram at the top on the left-hand side). To better understand the reasons for this set-up dependent sensitivity, the implants were thoroughly characterized, monitoring the dynamic changes in the systems' dimensions, dry and wet mass, PLGA polymer molecular weight, inner & outer morphology as well as pH of the bulk fluids. Furthermore, as it can be seen in the top row of Figure V-2, ibuprofen release was bi-phasic in all cases: irrespective of the type of experimental set-up and renewed bulk fluid volume. No noteworthy burst release was observed in any case. A zero-order release phase (with an about constant drug release rate) started from the beginning. After about 7-10 d, the final rapid release phase began (accounting for about 80 % drug release), leading to complete ibuprofen exhaust.

Figure V- 3 .

 3 Figure V-3. Optical macroscopy pictures of ibuprofen-loaded PLGA implants after different exposure times to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates (the receptor compartment containing well-agitated bulk fluid). At pre-determined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated).

  BULK FLUID RENEWAL CAN AFFECT DRUG RELEASE FROM PLGA IMPLANTS 167

Figure V- 4 .

 4 Figure V-4. SEM pictures of surfaces and cross-sections of ibuprofen-loaded PLGA implants before and after 8 days exposure to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates

Figure V- 5

 5 show the pH of the bulk fluids the implants or agarose gels were exposed to, as a function of the exposure time (Figure V-1). Green, orange and blue curves indicate complete and partial bulk fluid renewal. As it can be seen, the pH remained about constant during the first 7-10 d, and then temporarily dropped. This can be explained by BULK FLUID RENEWAL CAN AFFECT DRUG RELEASE FROM PLGA IMPLANTS 169 the release of short chain acids into the bulk fluids. As for the drug, the mobility of these watersoluble compounds fundamentally increases once substantial system swelling sets on. The fact that the temporary pH drop is less pronounced in the case of lower bulk fluid volume renewals can be explained by the lower amounts of withdrawn acids and added bases. Since ibuprofen exhibits strongly pH dependent solubility (Figure V-6), it might be that pronounced drops in the pH of the release medium lead to non-sink conditions and, thus, slower drug release.

Figure V- 5 .

 5 Figure V-5. Dynamic changes in the pH of the well agitated bulk fluid, in the implants' dry mass and PLGA polymer molecular weight (Mw) upon exposure of ibuprofen-loaded implants to phosphate buffer pH 7.4 using the 3 experimental set-ups. The implants were placed in well agitated bulk fluids in Eppendorf tubes, in agarose gels exposed to well agitated bulk fluid in Eppendorf tubes, or in agarose gels in transwell plates (the receptor compartment containing well-agitated bulk fluid). At pre-determined time points, the entire bulk fluid, or 1 or 3 mL thereof, was renewed (as indicated).

Figure V- 6 .

 6 Figure V-6. Dependence of the solubility of ibuprofen as a function of the pH at 37 °C. Phosphate buffer pH 7.4 USP 42 served as bulk fluid. Its pH was adjusted using of 10 N lactic acid or 0.05 N NaOH (final pH values are reported).

Figure VI- 1 .

 1 Figure VI-1. Schematic representation of the method used to prepare the polymer discs for melt rheology studies.

  PRINTING OF IBUPROFEN LOADED CONTROLLED RELEASE IMPLANTS WITH PLGA -PROOF OF CONCEPT AND IMPACT FILLING PATTERN 183 applicator hot plate (110 ° C, 6 min, Erichsen, Rueil-Malmaison, France), followed by extrusion using a Nano 16 twin-screw extruder (screw diameter = 16 mm, length/diameter ratio = 26.25, feeder flow regulator set at a speed between 5 and 8 cc/min; Leistritz, Nuremberg, Germany), equipped with a 4 mm diameter die. The process temperatures were kept constant at 70 -75 -70 -65 °C (die -zone 3 -zone 2 -zone 1). The screw speed was set at 50 rpm. The diameter of the filament obtained at the exit of the extruder was controlled by a laser sensor (target diameter between 1.60 and 1.80 mm, Tolerance Puller Noztek, Shoreham-by-Sea, United Kingdom).

1 Number of lower and upper layers 1 Filling 40 VI. 3 . 8 .

 114038 Practical drug loadingPieces of filaments and 3D printed object were dissolved in 5 mL acetonitrile, followed by filtering (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000 Series HPLC, equipped with a LPG 3400 SD/RS pump, an autosampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversedphase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used. The mobile phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v).The detection wavelength was 225 nm, and the flow rate was 0.5 mL/min. Ten microliter samples were injected. Each experiment was conducted in triplicate. Mean values +/-standard deviations are reported.

  variation over ibuprofen content, as well as the loss of mass obtained by thermogravimetry are presented in Figure

Figure VI- 2 .

 2 Figure VI-2. Thermograms of (A) the raw materials, (B) their physical mixture, (C) impact of the ibuprofen content on PLGA glass transition temperature measured by differential scanning calorimetry, and (D) PLGA, ibuprofen, and PLGA/ibuprofen physical mixture loss of mass upon heating, obtained by thermogravimetric analysis.

Figure VI- 3

 3 Figure VI-3 shows the complex viscosity of the pure PLGA and ibuprofen PLGA mixture overtemperature, and their viscoelastic behavior. The plasticizing effect of ibuprofen on PLGA observed in Figure VI-2C, is also confirmed with the melt rheology experiments. For the same temperatures, the addition of ibuprofen leads to a drop of the complex viscosity and storage and

Figure VI- 3 .VI. 4 . 3 .

 343 Figure VI-3. Effect of temperature on (A) PLGA and PLGA: ibuprofen (85:15) complex viscosity and (B) storage G' and loss G'' modulus.

Figure VI- 4 .Figure VI- 5 .

 45 Figure VI-4. Effect of printing temperature on (A) ibuprofen content and (B) PLGA molecular weight.

Figure VI- 6

 6 Figure VI-6 shows the DSC thermograms of the different implants and the filament used to

Figure VI- 6 .VI. 4 . 4 .

 644 Figure VI-6. Thermograms of the ibuprofen loaded PLGA filaments, and the ibuprofen loaded PLGA 3D printed implants.

Figure VI- 8 .

 8 Figure VI-8. Optical macroscopy pictures ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 37°C 80 RPM.

  Additive manufacturing, also known as three-dimension (3D) printing technologies allow to manufacture of objects either by deposition, binding, or polymerization of successive materials layers(1,2). Different classification of the 3D printing technologies has been reported, one of them classify those technologies depending on the object production: photopolymerization, powder bed fusion, material jetting, direct energy deposition, binder jetting, selective deposition lamination, and material extrusion(1,2). Lately, the use of 3D printing in the pharmaceutical field reached its highest point with the FDA approval of SPRITAM the first FDA-approved 3D

  The filaments obtained were used to feed the Volumic ® 30 Stream Ultra printer (Volumic, Nice, France). Parallelepiped-shaped implants with dimensions of 10 mm x 10 mm x 2.5 mm were printed according to the parameters described in TableVII-1. Figure VII-1 show a schematic illustration of the two-printing technology investigated here.

Figure VII- 1 .

 1 Figure VII-1. Schematic representation of (A) droplet deposition modeling by Arburg®Freeformer, and (B) fused filament fabrication by Volumic®.

Figure VII- 2 .

 2 Figure VII-2. (A) Optical macroscopic pictures of the surface of ibuprofen PLGA-implants obtained either with DDM with the Arburg Freeformer® technology, or FFF with the Volumic® printer, before exposure to the release medium and (B) schematic representation of the deposed lines of PLGA-ibuprofen loaded implants obtained either with DDM with the Arburg Freeformer® technology, or FFF with the Volumic® printer.

Figure VII- 3 .

 3 Figure VII-3. Scanning electronic microscopic pictures of the surface and cross-section of ibuprofen PLGA-implants obtained either with DDM with the Arburg Freeformer® technology, or FFF with the Volumic® printer, before exposure to the release medium.

Figure VII- 4

 4 Figure VII-4 diagrams illustrate the variety of molecular weight of the PLGA and ibuprofen content of the different feedstock materials (pellets, and filaments) and the implants obtained either by DDM or FFF.

Figure VII- 4 .

 4 Figure VII-4. Impact of the 3D printing technology used on (A) PLGA molecular weight and (B) the ibuprofen content.

  BASED 3D PRINTED IMPLANTS -IMPACT OF THE 3D PRINTED TECHNOLOGY ON DRUG RELEASE 215

Figure VII- 5 .

 5 Figure VII-5. Impact of the 3D printing technology used on the thermograms of the obtainedPLGA-ibuprofen loaded implants and intermediate products (pellets and filaments).

( 1 .VII. 4 . 2 .

 142 75 ± 0.05 mm). Regarding implants obtained with the two 3D printing techniques, the ibuprofen content of FFF implants is significantly lower and more variable than that of DDMzigzag implants (FigureVII-4B). Indeed, DDM technology makes it possible to print an implant from a statistical mixture of pellets, this promotes homogeneity, unlike the FFF technology which uses a short portion of the filament.Despite greater mechanical stress for the production of the filament than the pellets (the force applied to the molten polymer is the combination of the rotation of the screws, and the force applied by the material supplied by the "feeder") the PLGA molecular weight in the filament and pellets are equivalent (FigureVII-4A). In contrast, FFF implants have a higher PLGA molecular weight Compared to implants obtained by FFF technology. This is a consequence of the considerable mechanical stress associated with the Arburg® Freeformer printer screw[START_REF] Park | Degradation of poly(lactic-co-glycolic acid) microspheres: effect of copolymer composition[END_REF]. Characterization of the implants after exposure to the release mediumThe impact of 3D printing technology from previously produced and characterized implants on the relative and absolute in vitro release kinetics of ibuprofen and pH changes in the release medium have been illustrated in Figure

Figure VII- 6 .

 6 Figure VII-6. (A) Relative (%) and (B) absolute (mg) ibuprofen release from PLGA implants, and (C) pH changes upon exposure to phosphate buffer pH 7.4, 37°C, 80 RPM.

Figure VII- 7

 7 Figure VII-7 diagrams show the dynamical changes of the two types of implant upon exposure to the release medium. It was observed that the water uptake of implants obtained by FFF is slightly higher than that of implants obtained by DDM (Figure VII-7A). The impact of this water uptake on the change in the morphology of the implants is also visible in the photos (Figure VII-8). Implants obtained by DDM show triphasic release pattern (Figure VII-6A and B) with:

(

  Figure VII-7. (A) Change in wet mass (%), (B) dry mass (%), and (C) molecular weight (Mw) of ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4 37°C 80 RPM.

Figure VII- 8 .

 8 Figure VII-8. Optical macroscopy pictures ibuprofen-loaded PLGA implants upon exposure to phosphate buffer pH 7.4, 37°C, 80 RPM.

  BASED 3D PRINTED IMPLANTS -IMPACT OF THE 3D PRINTED TECHNOLOGY ON DRUG RELEASE 219 On the other hand, the absolute release kinetics (Figure VII-6B) of both types of implants are superposed until the complete release of ibuprofen contain in the implant obtained by FFF is reached. This suggests that the implants obtained by FFF released most of the ibuprofen before the onset of the swelling of the PLGA (Figure VII-7A), the absence of lag phase and 3 rd rapid release phase is also due to shorter diffusion pathway due to the difference in the shape of the implants (Figure VII-8). For both types of implants, the loss of dry mass and the molecular weight of the polymer over time are very similar (Figure VII-7B and C). The polymer erosionprofile is typical of bulk erosion(60). Until day 6, the first degradation products remain in the matrix, decrease the micro-surrounding pH, and cause, by an autocatalytic effect, the accelerated degradation of PLGA(61).

Figure VII- 9

 9 Figure shows SEM pictures of the surface and cross-section of the implants after 3 days of exposure to the release medium. In both cases, exposure to the release medium leads to porous and unshaped structures. For implants produced by FFF the event after 3 days exposure to a release medium at a temperature higher than the Tg of the implant (

Figure VII- 9 .

 9 Figure VII-9. SEM pictures of surfaces and cross-sections of ibuprofen-loaded PLGA implants after exposure to phosphate buffer pH 7.4, 37°C, 80 RPM. Please note that after exposure to the release medium the implants were freeze-dried before analysis. Thus, caution must be paid due to artefact creation.

Figure VIII- 1 .

 1 Figure VIII-1. Schematic presentation of the Arburg Freeformer® technology.

VIII. 4 .VIII. 4 . 1 .

 441 droplets at the nozzle tip Workpiece carrier moves the workpiece within X, Y and Z axes Results and discussion Characterization of the implants before exposure to the release mediumThis study aimed to produce ibuprofen-loaded PLGA implants by 3D printing with different filling densities to investigate the impact of the filling density on the in vitro performance.

Figure VIII- 2

 2 Figure VIII-2 show macroscopic pictures of the implants obtained by 3D printing (TableVIII-1). Please note that different layers thickness had to be defined depending on the filling density of the implant, to limit the polymer collapsing for the implants with low filling density. The implants pores size created by the different filling density, are highlighted in green dotted line in Figure VIII-2 and their size (measured in diagonal) are summarized in Table VIII-2.Apart from that and the weight of the implants, they exhibit similar properties regarding their, length, width, thickness, practical drug loading, glass transition temperature, and molecular weight of the polymer (TableVIII-2).

Figure VIII- 2 .

 2 Figure VIII-2. Optical macroscopic pictures of the surface of ibuprofen PLGA-implant obtained with the Arburg Freeformer® technology, with a filling density of either 10 %, or 100 % before exposure to the release medium.

  BASED 3D PRINTED IMPLANTS -IMPACT OF FILLING DENSITY ON DRUG RELEASE 239

  7 kDa before printing to 16.7 kDa, and 15.0 kDa respectively for the implants printed with 10% and 100 % filling density) (FigureVIII-3). A previous study reported that such a degradation of PLGA was not observed, even at a higher temperature, with fused filament fabrication confirming that the degradation observed here is mainly due to the high pressure during the process (±400 bars).

Figure VIII- 3 .

 3 Figure VIII-3. Scanning electronical microscopic pictures of the surface and cross section of ibuprofen PLGA-implant obtained with the Arburg Freeformer® technology, with a filling density of either 10 %, or 100 % before exposure to the release medium.

  mechanical stress and high pressure also have consequences on the internal structure observed by scanning electronic microscopy (FigureVIII-4). Thus, leading to pores formation (highlighted in red dotted line) due to high mechanical stress upon printing which would induce the evaporation of either ibuprofen molecules, even at lower temperatures than its boiling point, or residual water contained in the raw materials(17).

Figure VIII- 4 .

 4 Figure VIII-4. DSC thermograms of: (A) the raw materials (PLGA & ibuprofen), and (B)pellets and ibuprofen loaded PLGA implant obtained with Arburg freeformer ® technology.

Figure VIII- 5 .

 5 Figure VIII-5. PLGA polymer molecular weight as: raw material, hot-melt extruded pellets and 3D printed ibuprofen loaded PLGA implant.

Figure VIII- 6 . 4 .

 64 Figure VIII-6. (A) Relative (%), (B) absolute (mg) ibuprofen release for the 3D printed PGLA implants, and (C) increase in wet mass, upon exposure to phosphate buffer pH 7.4.

  filling density implant: during the first days (from day 1 to day 2) of exposure to the release medium, a high surface of the implants is in contact with water (due to the created porosity with the 3D printing filling density), leading to water penetration in the polymer matrix and diffusion of the dissolved molecule of ibuprofen close to the surface of the printed line. Once the ibuprofen close to the surface is released, then the length of the diffusion pathway for ibuprofen in the center of the polymer matrix starts to increase, slowing down the release leading to the lag phase. Then starting from day 5, the PLGA reaches a critical molecular weight leading to a more hydrophilic matrix and significant water uptake and swelling of the whole system (Figure VIII-6C and Figure VIII-7). Thus, letting enough space for the remaining ibuprofen entrapped to quickly release, leading to the third release phase (Figure VIII-6A and B).

  (ii)For the 100% filling density implants: during the first day of exposure to the release medium (from day 1 to day 2) a lower surface of the implants is in contact with water (as not such a porosity as for the 10% filling density implant is created by 3D printing as shown in Figure VIII-2 and Figure VIII-3), leading to a lower relative amount (%) of ibuprofen close to the surface released during burst (Figure VIII-6A). However, as the 100% filling density implants exhibit a higher weight, this leads to the same total amount (mg) of ibuprofen released than the 10% filling density implants (Figure VIII-6B). Then the diffusion pathway becomes even longer than for the 10% filling density implants, as only the parallelepiped faces are in direct VIII. PLGA BASED 3D PRINTED IMPLANTS -IMPACT OF FILLING DENSITY ON DRUG RELEASE 243 contact with the release medium. Thus, a lag phase with a slower release rate and a longer period is reached, as water penetrated more sluggishly compared to the lower filling density implants (Figure VIII-6V). Then after 7 days of exposure to the release medium, they start to significantly swell (Figure VIII-7), leading to initiation of the third phase which is then even accelerated at day 10.

Figure VIII- 7 .Figure VIII- 8 .

 78 Figure VIII-7. Optical macroscopy pictures of surfaces of ibuprofen loaded PLGA implant after exposure to phosphate buffer pH 7.4.

Figure VIII- 9

 9 Figure shows SEM pictures of the surface and cross-section of the implants after 3 days of exposure to the release medium. Unfortunately, implants with a 10% filling density were too fragile to cut and obtain a cross-section for SEM pictures. However, for the 100% filling density implants cross-section, it can be seen that there are two different phases, with a thin layer of porous and swelled PLGA close to the surface of the implants (FigureVIII-9). Regarding the surface of the implants after 3 days of exposure to the release medium, whatever the filling density, the surface is unshaped, due to the hydration of the surface of the polymer.

Figure VIII- 9 .

 9 Figure VIII-9. SEM pictures of surfaces and cross-sections of ibuprofen loaded PLGA implant after 3 days exposure to phosphate buffer pH 7.4. Please note that after exposure to the release medium the implants were freeze-dried prior to analysis. Thus, caution must be paid due to artefact creation.

  (i) Development of experimental set-ups for in vitro drug release measurement for implants administrated to the subcutis, and investigation on the impact of several implants attribute on in vitro drug release: -Investigation of biosimilar release medium, taking into consideration the protein and lipid content in the subcutis, -Taking into consideration the possible changes in vivo after administration (inflammation, etc…) and how it could be mimic in vitro, -Investigation of the impact of mechanical stress on in vitro performances. (ii) Development of PLGA based controlled drug release implants using 3D printing technologies, and investigation of the impact of the structure of the implant on in vitro drug release: -Investigate different shapes, geometry size of devices, -Improving the printing quality of PLGA by DDM, -Find a correlation between the in vitro performances and the morphological characteristics of the 3D printed implants. 264 Cette thèse s'est articulée autour de deux axes principaux : D'une part, le copolymère d'acides lactique et glycolique (PLGA) est un choix de première ligne en matière de libération contrôlée pour la voie parentérale. Sa biocompatibilité, et biodégradabilité ont menées à son utilisation dans plusieurs produits approuvés par la FDA. De nombreux mécanismes sont impliqués dans la libération de substance active (SA) à partir de ces systèmes. D'autre part, au cours des dernières années, les technologies d'impression en trois dimensions (3D) ont révolutionné le développement de nouveaux médicaments à libération contrôlée. Ces nouvelles technologies permettent de contrôler précisément la taille, la forme, la porosité des formes galéniques produites, et ainsi de contrôler le taux de libération de la SA. Dans ce contexte, les deux principaux objectifs de cette thèse étaient : (i) La mise au point d'une méthode de mesure de la libération de SA in vitro pour des implants destinés à être administrés par voie sous cutanée, et l'étude de l'impact de plusieurs attributs d'implants sur la libération de SA in vitro ; (ii) Le développement d'implants à libération contrôlée de médicaments à base de PLGA à l'aide de technologies d'impression 3D. Le tissu sous-cutané se comportant davantage comme un gel que comme un fluide, deux méthodes d'évaluation de la libération in vitro de SA impliquant du gel d'agarose pour mimer les propriétés du tissue sous-cutané ont été développés. Leur impact, sur le comportement d'implants à base d'ibuprofène obtenu par extrusion à chaud ont été étudiés. Cette étude a permis de tirer les conclusions suivantes : -La présence d'un gel d'agarose autour des implants PLGA entrave le gonflement du polymère et ralentit la libération de la SA.

  impact du traitement thermique sur le comportement rhéologique du mélange d'ibuprofène PLGA a également été étudié. Des implants à base de PLGA chargés en ibuprofène ont été imprimés par : fabrication de filaments fondu (FFF) et dépôt de gouttelettes (DDM). L'impact de la technologie d'impression 3D, du motif de remplissage et de la densité de remplissage des implants sur leurs performances in vitro a été étudié. Les conclusions suivantes ont été tirée :--Des filaments PLGA chargés d'ibuprofène ont pu être utilisés pour imprimé sans impacter de façon significative le poids moléculaire du polymère pour des température allant jusqu'à 150 °C, leur comportement viscoélastique est compatible avec l'impression 3D.-Il a été possible d'imprimer des implants contenant de l'ibuprofène, avec trois structures internes différentes. Ceux-ci permettant de maintenir la libération de l'ibuprofène jusqu'à 9 jours, avec un profile monophasique.-La technologie DDM permet d'obtenir facilement des implants chargés homogènement en SA et permet d'éviter l'étape de développement de filament nécessaire pour l'impression par FFF. Toutefois, cette technologie expose le polymère et la SA à un stress mécanique élevé par rapport à la technologie FFF.

Table I - 1 .

 I1 Differences between quality control in vitro drug release testing, and biorelevant in vitro drug release testing methods. ...................................................................................................................................................... 5

Table I - 2 .

 I2 Examples of reported studies on the impact of instrument-based parameters on in vitro drug release. .............................................................................................................................................................................. 10

Table I - 3 .

 I3 Examples of reported studies on the impact of medium based parameters on in vitro drug release. ...12

Table I - 4 .

 I4 Summary of the FDA-approved products containing poly (lactic-co-glycolic acid). Adapted from Ochi et al. 2021. ............................................................................................................................................................. 18

Table I - 5 .

 I5 Principal photopolymerization based 3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 27

Table I - 6 .

 I6 Principal powder bed fusion-based 3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 29

Table I -

 I

7. Principal material extrusion based-3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 31 Table I-8. Principal material jetting based 3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 33 Table I-9. Principal binder jetting 3D printing technology advantages and disadvantages. ................................. 34 Table I-10. Principal direct energy deposition-based 3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 35 Table I-11. Principal sheet lamination-based 3D printed technologies, operating mode, advantages, and disadvantages. ....................................................................................................................................................... 37 Table I-12. Examples of 3D printing technologies in pharmaceutical field. ........................................................ 39

Table II -

 II

1. Physical key properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature). Mean values ± standard deviations are indicated (n=3). ................................................................. 75 Table III-1. Physical key properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature, Mw: average polymer molecular weight). The heating time applied during implant manufacturing is indicated in the top row. Mean values ± standard deviations are indicated (n = 3). ........................................... 105 Table IV-1. Key properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature, Mw: average polymer molecular weight). Mean values ± standard deviations are indicated (n = 3). ................ 138 Table V-1. Key properties of the investigated ibuprofen-loaded PLGA implants before exposure to the release medium (Tg: glass transition temperature, Tm: melting temperature). Mean values ± standard deviations are indicated (n=3). ................................................................................................................................................... 163 Table VI-1. Preparation method of polymer discs for melt rheology studies. .................................................... 182 Table VI-2. Printing parameters used to obtain the ibuprofen loaded PLGA-implant. ...................................... 183 Table VI-3. Key properties of the ibuprofen loaded PLGA-implants. (Tg: glass transition temperature, Mw: molecular weight). Mean values ± standard deviations are indicated (Weight, Length, Width, Practical drug loading n=6, Thickness, Tg, Mw n=3). ............................................................................................................... 190 Table VII-1. Printing parameters used to obtain the ibuprofen-loaded PLGA-implants.................................... 205 xii

Table VII - 2 .

 VII2 Key properties of the obtained ibuprofen-loaded PLGA-implants. (Tg: glass transition temperature, Mw: molecular weight). Mean values ± standard deviations are indicated (Weight, Length, Width, Implants pore size n=36, Practical drug loading n=6, Thickness, Tg, Mw n=3). ...................................................................... 213

Table VIII - 1 .

 VIII1 Printing parameters used to obtain the ibuprofen loaded PLGA-implants of different density. .. 234

Table VIII -

 VIII

2. Key properties of the obtained ibuprofen loaded PLGA-implants. (Tg: glass transition temperature, Mw: molecular weight). Mean values ± standard deviations are indicated (Weight, Length, Width, Implant pore size n=36, Practical drug loading n=6, Tg, Mw n=3). ......................................................................................... 239 I.1. Control drug delivery I.1.1. Definitions

Table I -

 I

1. Differences between quality control in vitro drug release testing, and biorelevant in

vitro drug release testing methods.

Extend drug release Negligible local adverse effects Sterile Biocompatible Biodegradable

			Possibility to
			interrupt the therapy
	Type of in vitro drug release	Quality control methods	Biorelevant methods
	method
	Aim	Discriminate changes in the drug	Establishment of an in vitro in vivo
		product characteristics	correlation
	Type of medium	Do not necessarily use a medium that	Should have the same kay
		has the same characteristics as the	characteristics as the administration
		administration site	site
	Type of apparatus	Usually, pharmacopeia is referred	Not necessarily a pharmacopeia
		apparatus	apparatus

Table I - 3 .

 I3 Examples of reported studies on the impact of medium based parameters on in vitro drug release.

	Investigated parameter	Effect on drug release	Explanation	Ref
	Osmolarity	Drug release rate decreased with increasing osmolarity of the release medium	Probably due to a decrease in water uptake rate	(17)

Table I -4.

 I Summary of the FDA-approved products containing poly (lactic-co-glycolic acid).

	Along with the commercialization of these products, in the literature, there is a constant
	Adapted from Ochi et al. 2021. increasing interest of devices based on PLGA. As illustrated in Figure I-12, the number of
	Product publications referring to PLGA per year is constantly increasing (Figure created using the Active pharmaceutical ingredient Formulation Drug released for Year of Company approval number of references per year using the key word PLGA on PubMed).
	Lupron Depot	Luprolide acetate	Microspheres	1,3,4,6 months	Abbvie
	Zoladex Depot	Gosereline acetate	Implant	1,3 months	AstraZeneca
	Sandostatin LAR	Octreotide acetate	Microspheres	1 month	Novartis
	Atridox	Doxycline hyclate	In situ forming implant	1 week	Zila Therapeutics
	Nutropin Depot	Somatotropin	Microspheres	1 month	Genentech
	Trelstar	Triptorelin pamoate	Microspheres	1,3,6 months	Allergan
	Somatuline Depot	Lanreotide	Microspheres	1 month	IPSEN
	Arestin	Minocycline HCl	Microspheres	2 weeks	OraPharma
	Eligard	Leuprolide acetate	In situ forming implant	1,3,4,6 months	Sanofi
	Risperdal Consta	Risperdone	Microspheres	2 weeks	Janssen
	Vivitrol	Naltrexone	Microspheres	1 month	Alkermes
	Ozurdex	Dexamethasone	Implant	3 months	Allergan
	Propel mini	Mometasone furoate	Implant	1 month	Intersect
	Bydureon	Xenatide	Microspheres	1 week	AstraZeneca
	Lupaneta pack	Leuprolide acetate	Microspheres	3 months	Abbvie
	Signifor LAR	Pasireotide	Microspheres	1 month	Novartis
	Zilretta	Triamcinolone	Microspheres	3 months	Flexion Therapeutics
	Sublocade	Buprenorphine	In situ forming implant	1 month	Indivior
	Perseris	Risperdone	In situ forming implant	1 month	Indivior

-dimensional printing technologies I.3.1. Generalities I.3.1.1. Definition

	3D printing technologies are additive manufacturing (AM) technics able to produce various 3D
	products (110). The final object is formed either by deposition, binding, or polymerization of
	materials in successive layers (110,111).
	I.3.1.2. Timeline
	As illustrated in Figure I-16, in 1980, the concept of 3D printing was first reported by Hideo
	Kodama which described a process of solidification of polymers using a beam of UV light. In
	1986, the first additive manufacturing technology, known as stereolithography, was patented
	by Charles Hull (112-116). Around the same time, selective laser sintering was invented by
	Carl Deckard (117). In 1989 the company Stratasys was funded by L. Crump and S. Scott, who
	filed the patent for fused deposition modeling (FDM) (116,118). In the 1990s other 3D printing
	technologies have been investigated (116). In 2005, the patent for FDM expired, leading to
	democratization and an increase of interest for the 3D printing technologies (112,116,119).
	Figure I-16. Timeline of three-dimensional printing technologies.
	The 3D printing technologies invented since the 1980s can be classified into seven categories
	further developed below: photopolymerization, powder bed fusion, material jetting, direct
	energy deposition, binder jetting, selective deposition lamination, and material extrusion
	(110,111).

.3.2. Classification I.3.2.1. Photopolymerization

	IPhotopolymerization-based 3D printing technologies work as described in Figure I-17. Indeed,
	a mobile platform serving as support is immersed in a vat of liquid photopolymer at a certain
	depth. This resin is generally a mixture of acrylate or epoxy monomers and a photoinitiator
	which will initiate the polymerization under the effect of UV or visible light directed using
	deflectors. The model gradually sinks during manufacturing and is finally cut, washed, and
	baked to harden it (115,120,121).
					2005 -patent for FDM
					technology	expired,
		1989 -foundation of	subsequently the use
		Startasys by L. Crump	of	3D	printing
		and S. Scott, and	technologies
	1987 -carl Deckard	patent	for	fused	democratized
	submits the patent	deposition modeling
	for selective laser
	sintering
	1984 -Charles Hull
	invents stereolithography
	which was patent in 1986
	1980 -Hideo Kodoma submits a
	patent application for a prototyping
	technique using a single laser beam.
	Patent application expires without
	proceeding to the later stages of the
	Japanese patent process

Table I -5. Principal

 I

			photopolymerization based 3D printed technologies, operating mode,
	advantages, and disadvantages.
	Name	Operating mode			Advantage	Disadvantages	Ref
	SLA	A concentrated beam of UV	High precision, high quality,	Not suitable for mass	(115)
		or laser beam is projected	smooth finish.		production.
		onto the surface of a tank				Limited choice of materials
		filled	with	liquid				(photopolymerization).
		photopolymer. The beam is
		focused, creating the 3D
		object layer by layer by
		crosslinking of the polymer.
	DLP	A digital projection screen	Faster print times (each layer	The layers are formed of small	(122,126)
		illuminates a single image of	is exposed all at once instead	rectangular bricks (the layer is
		each layer in a single	of being drawn).		made up of square pixels).
		operation.						Limited choice of materials
								(photopolymerization).
	CLIP	The bottom part of the tank	Print continually (up to 100	Limited choice of materials	(123,124)
		allows UV light to pass	times faster than traditional	(photopolymerization).
		through. UV beam passes	printing technologies).
		through this window, and
		illuminates a precise cross-
		section of the object, leading
		to cross-linking of the
		polymer resin. The object
		rises slowly enough to allow
		the resin to flow underneath
		and remain in contact with the
		bottom of the object.
	DPP	A Liquid crystal display	Faster	than	traditional	Limited choice of materials	(125)
		(LCD) is used to cure the	printing technologies.	(photopolymerization).
		polymer.						Need specially formulated
								natural light polymer.

I.3.2.2. Powder bed fusion

	Powder Bed Fusion (PBF) 3D printing technologies generate precision products. The powder
	particles are fused using a heat source (generally a laser beam or electron beam). Thus, layer
	by layer the 3D object is formed as illustrated in Figure I-18 (117,120,121,127-133).
	Heat source: laser, electron beams
	3D printed object
	Powder roller
	Powder bed
	Build platform
	Departure base: powder
	Figure I-18. Schematic representation of powder bed fusion-based three-dimensional
	printing technologies.
	Several techniques can be cited such as Selective Laser Sintering (SLS), Electron Beam Melting
	(EBM), Selective Heat Sintering (SHS), Direct Metal Laser Sintering (DMLS), Selective Laser
	Melting (SLM), Multi Jet Fusion (MJF). Their operating mode, advantage, and disadvantages
	are summarized in Table I-6.

Table I -6. Principal

 I

		powder bed fusion-based 3D printed technologies, operating mode,
	advantages, and disadvantages.
	Name	Operating mode	Advantage	Disadvantages	Ref
	SLS	Use laser beams to sinter, or	Several materials can be	Possible shrinkage and	(117)
		agglomerate,	powdered	used.	warping of thin parts when
		materials, layer after layer to	Powder recycling.	designing.
		create a 3D object			High power usage.
		Laser power up to 0.04 kW
	EBM	The material is fused by a	Use less energy and can	Requires post-processing	(128)
		high-energy electron beam.	produce layers faster than
				SLS.
				Powder recycling.
	SHS	Use a thermal print head to	Low cost.	High power usage	(129)
		selectively fuse material	Powder recycling
		powder
	DMLS/SLM Agglomerates powders and	Powder recycling.	Require additional work to	(130)
		is limited to alloys, including		compensate for the high
		those composed of titanium.		residual stress and limit the
		Laser power up to 1 kW		occurrence of distortion.
					Work only with metals.
	MJF	Infrared lamp acts as the	improved printing	requires two additional	(132,133)
		energy source to consolidate	efficiency, accuracy and	agents: a fusion agent and a
		powder particles		speed.	detailing agent

I.3.2.3. Material extrusion

  heated to ensure good adhesion. Once the first layer is complete, the extruder and the platform move away simultaneously and the second layer can then be directly deposited on the part being manufactured (111,120,121,135).

	(FFF), Direct Powder Extrusion (DPE), and Droplet Deposition Modeling (DDM)
	(111,120,121,135-139).
	FDM or FFF technologies are illustrated in Figure I-19. This is the technologies are based on
	the one developed and patented by L. Crump and S. Scott in 1989 (116,118). These technologies
	use a continuous filament of thermoplastic material as the base material. Filament is fed by a
	spool, via a heated, mobile printer extrusion head, often referred to as an extruder. The molten
	material is expelled from the extrusion nozzle and deposited first on a 3D printing platform,
	which can be
	Material Extrusion (ME) based 3D printing technologies can be divided into two parts: (i)
	extrusion of low melting point past or gel and (ii) extrusion of high melting point thermoplastic
	polymer (111). In both cases, it consists of dispensing molten or semi-solid material through
	the orifice of a nozzle, deposed layers by layers in order to obtain a 3D object (120). Nowadays,
	these printing technologies are considered the simplest and are the most widely used.
	When it comes to extrude a low melting point past or gel, Pneumatic or Syringe Extrusion
	(PE/SE) are the referenced technics (111,134). It consists of filling pastes into ink cartridges
	for extrusion though a nozzle. The past or gel, is then deposed layer by layer in order to form
	the desired object. The obtained 3D object is then under vacuum and heating for complete
	drying (134).
	Concerning the high melting point thermoplastic polymers, several technologies have been
	reported such as Fused Deposition Modeling (FDM) also known as Fused Filament Fabrication

Table I -

 I

			Departure base:
			Standard pellets
		Piezo cadences
		the nozzle shutter
				Plasticizing screw as
				for injection molding
	Name	Operating mode	Advantage	Disadvantages	Ref
	SE/PE	A past or gel, is extruded then deposed	Low heating, or no	Post treatment	(120,134)
		layer by layer in order to form the	heating
		desired object. The obtained 3D object
		is then under vacuum and heating for
		complete
		drying
	FDM/FFF Molten filament is extruded deposed	Low cost, easily usable. Need to develop a suitable	(120,135)
		Layer by layer to form 3D object.		filament if working with
				non-commercial polymer
	DPE	Powder are loaded into a screw similar	Avoid development of		(136)
		to hot melt extrusion, material is	filament
		extruded and deposed layer by layer.	Ingle step process
	DDM	Pellets are loaded into a screw similar	Avoid development of	Mechanical stress (up to	(138,139)
		to injection modeling, material is	filament	600 bar) can lead to
		extruded and deposed layer by layer.		polymer degradation

7. Principal material extrusion based-3D printed technologies, operating mode, advantages, and disadvantages.

Disacharge of individual droplets at the nozzle tip

Workpiece carrier moves the workpiece within X, Y and Z axes

Table I - 8 .

 I8 Principal material jetting based 3D printed technologies, operating mode, advantages, and disadvantages.

	Name	Operating mode	Advantage		Disadvantages	Ref
	DOP	Two printing jets: build	Good accuracy		A	limited	number	of	(142,143)
		material and soluble support			materials.
		material. A leveling blade			Slow process, high cost.
		skims the build area after each
		layer (flat surface before the
		next layer).
	Polyjet	Photopolymer materials are	Good	dimensional	A	limited	number	of	(144-
		sprayed in ultra-thin layers.	performance.		materials.		146)
		Each photopolymer layer is			Slow process, high cost.
		cured by UV. The gel-like
		carrier material can easily be
		removed manually or by
		water jet.
	NPJ	Liquid-containing building
		nanoparticles	or	carrier
		nanoparticles are sprayed in

layers. High temperatures inside the building area lead to evaporation of the liquid.

Table I -

 I

	Name	Advantage	Disadvantages	Ref
	BJ	Don't use heat.
		Low cost.

9.

Table I-9. Principal binder jetting 3D printing technology advantages and disadvantages.

Suitable for a wild range of materials.

Weak mechanical properties (compared to fused based technologies)

(121,148,149)

Table I -

 I

	Name	Operating mode		Advantage	Disadvantages	Ref
	LENS	Laser construct object layer	Process (argon gas chamber)
		by layer directly from	prevents	oxidation	and
		powdered metals, alloys,	ensures cleanliness.
		ceramics, or composites.

10. Principal direct energy deposition-based 3D printed technologies, operating mode, advantages, and disadvantages.

Table I -

 I

						3D printing technologies
	Name	Operating mode	Advantage		Disadvantages	Ref
	LOM Photopolymerization The bonding of the new sheet on the top of the other one is Powder Bed Fusion (PBF)	Material Extrusion (ME)	Post operating treatment Material Jetting (MJ)	(121,154) Binder Jetting (BJ)	Direct Energy Deposition (DED)	Sheet Lamination (SL)
		thermal
	UC	The bonding of the new sheet		Post operating treatment	(121,155)
	CBAM Stereolithography (SLA) Fiber-reinforced composites ultrasonic on top of the other ones are Selective Laser Sintering material handling Speed, low cost, ease of (SLS) Pneumatic or Syringe Post operating treatment Extrusion (PE/SE) Drop On Demand (DOD)	(121,131)	LENS	Laminated Object Manufacturing (LOM)
		fuse with a thermoplastic. Direct Metal Laser	Fused Deposition
	SLCOM Use of thermoplastics as base Digital Light Processing (DLP) Sintering (DMLS)/Selective Laser	Post operating treatment Modeling (FDM)/Fused Filament Fabrication Polyjet	(121,131)	Electron Beam Additive Manufacturing (EBAM)	Ultrasonic Additive Manufacturing (UC)
		material and composites of Melting (SLM)	(FFF)
		woven fibers.
		Continuous Liquid Interface Production (CLIP)	Electron Beam Melting (EBM)	Direct Powder Extrusion (DPE)	NanoParticle Jetting (NPJ)	Laser Deposition Welding (LDW)	Composite Based Additive Manufacturing (CBAM)
		Daylight Polymer Printing (DPP)	Selective Heat Sintering (SHS)	Droplet Deposition Modeling (DDM)		Selective Lamination Composite Object Manufacturing (SLCOM)
			Multi Jet Fusion (MJF)

11. Principal sheet lamination-based 3D printed technologies, operating mode, advantages, and disadvantages. I.3.3. Use in the pharmaceutical field I.3.3.1. State of art Among the 3D printing technologies described in section I.3.2, Five have been used for pharmaceutical application, either in actual product, or in research as illustrated in Figure I-25 (111): -Photopolymerization (126,156-158), -Powder bed fusion (132,159-162), -Material extrusion (135,136,138,163-165), -Material jetting (166,167), -Binder jetting (168-170). The efforts and advances of 3D printing technology in the field of pharmaceutics has recently culminated in 2015 with the US Food and Drug Administration approval of the world's first 3D printed orodispersible tablet SPRITAM ® (168,169). Some examples of the used of these technologies in the pharmaceutical field have been summarized in Figure I-25 and Table I-12.

Figure I-25. Different classes of 3D printing technologies, the ones used in pharmaceutical field have been colored in green, and the other ones in dark blue. Adapted from Lim et al 2018.

Table I - 12 .

 I12 Examples of 3D printing technologies in pharmaceutical field.

  Yang et al. investigated the potential of a quaternized chitosan grafted PLGA hydroxyapatite scaffold obtained by 3D printing to improve antimicrobial activity and osteoconductive properties. They observed that this type of scaffold exhibited an improvement in anti-infection and bone regeneration capabilities(190). PLGA based scaffold for tissues engineering obtained by 3D printing have also been investigated byRasoulianboroujeni et al. and Mironov et al. (191,192).

In 2017

Guo et

al. investigated the impact of printing parameters on printing resolution of PLGA scaffolds obtained by 3D printing with different PLGA grade (193). They built a model that correlate the predominant factors determining the printing precision. Recently, Carlier et al. developed 3D printed implantable devices loaded with monoclonal antibodies obtained by FDM

(194)

. They demonstrated the suitability of FDM for producing monoclonal antibodyloaded devices with good stability, affinity and sustained-release profiles of the monoclonal antibody with PLGA.

  Embedded in agarose gels, which are exposed to bulk fluid in Eppendorf tubes, and (iii) Embedded in agarose gels, which are exposed to bulk fluid in transwell plates. The observed drug release kinetics were to be explained based on the monitoring of dynamic changes in the systems' wet and dry mass (gravimetrically), average polymer molecular weight

	II.3. Materials and methods
	II.3.1. Materials
	Poly (D,L lactic-co-glycolic acid) (PLGA, 50:50 lactic acid: glycolic acid; Resomer RG 503H;
	(GPC), inner and outer morphology (by optical and scanning electron microscopy) and pH Evonik, Darmstadt; Germany); ibuprofen (BASF, Ludwigshafen, Germany); agarose (genetic
	measurements in the bulk fluids, gels and implants. analysis grade), bromocresol green (BCG), bromothyol blue (BTB), bromophenol blue (BPB),
	phenol red (PR) and tetrahydrofuran (HPLC grade) (Fisher Scientific, Illkirch, France);
	potassium dihydrogen orthophosphate and sodium hydroxide (Acros Organics, Geel, Belgium);
	acetonitrile (VWR, Fontenoy-sous-Bois, France); sodium hydrogen phosphate (Na2HPO4,
	Panreac Quimica, Barcelona, Spain).

3 different experimental set-ups: (i) Upon exposure to well agitated bulk fluid in Eppendorf tubes, (ii)

  .4. denotes the implant mass before exposure to the release medium. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

	𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (%)(𝑡) =	𝑑𝑟𝑦 𝑚𝑎𝑠𝑠 (𝑡) 𝑚𝑎𝑠𝑠 (𝑡 = 0)	× 100 %
	where mass (t = 0) The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation
	chromatography (GPC) as follows: Freeze-dried implant samples were dissolved in
	tetrahydrofuran (3 mg/mL). One hundred µL samples were injected into an Alliance GPC
	(refractometer detector: 2414 RI, separation module e2695, Empower GPC software; Waters,
	Milford, USA), equipped with a Phenogel 5 µm column (kept at 35°C, 7.8 × 300 mm;
	Phenomenex). Tetrahydrofuran was the mobile phase (flow rate: 1 mL/min). Polystyrene
	standards with molecular weights between 5,120 and 70,950 Da (Polymer Laboratories, Varian,
	At pre-determined time points, implant samples were withdrawn and freeze dried (freezing at -
	45°C for 2 h 35 min, primary drying at -20 °C/0.940 mbar for 35 h 10 min, secondary drying at
	+20 °C/0.0050 mbar for 35 h; Christ Alpha 2-4 LSC+; Martin Christ, Osterode, Germany).

The dry mass (%) (t) was calculated as follows:

Les Ulis, France) were used to prepare the calibration curve. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

II.3.7. Differential scanning calorimetry (DSC)

DSC thermograms of PLGA (raw material) and implants were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland). Approximately 5 mg PLGA and around 10 mg implant samples were heated in perforated aluminum pans as follows: from -70 to 120 °C, cooling to -70 °C, re-heating to 120 °C (heating/cooling rate = 10 °C/min). The reported glass temperatures (Tgs) were determined from the 1 st heating cycles in the case of implants (the thermal history being of interest), and from the 2 nd heating cycle in the case of the PLGA raw material (the thermal history not being of interest). All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

Table II -

 II

	Practical loading (%)	Weight (mg)	Length (mm)	Diameter (mm)	Tg (°C)	Picture
	6.6±0.3	33.7±4.4	5.1±0.3	2.6±0.2	34.5±0.3

1. Physical key properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature). Mean values ± standard deviations are indicated (n=3).

Table III -

 III

	Heating time	3 min	6 min	9 min	12 min	15 min
	Practical
	drug loading	12.8 ± 0.6	11.2 ± 0.3	9.7 ± 1.0	10.9 ± 1.0	11.4 ± 1.0
	(%)
	Diameter (mm)	2.8 ± 0.1	2.5 ± 0.1	2.7 ± 0.1	2.6 ± 0.1	2.5 ± 0.1
	Tg (°C)	33.9 ± 0.6	34.0 ± 0.3	34.3 ± 0.1	34.5 ± 0.4	33.3 ± 0.2
	Mw (kDa)	21.2 ± 1.2	19.7 ± 0.8	18.9 ± 0.2	19.4 ± 0.8	18.4 ± 1.2
	III.4.1.

1. Physical key properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature, Mw: average polymer molecular weight). The heating time applied during implant manufacturing is indicated in the top row. Mean values ± standard deviations are indicated (n = 3).

Implants before exposure to the release medium

  Waters, Milford, USA), equipped with a PLgel 5 µm MIXED-D column (kept at 35°C, 7.8 × 300 mm; Agilent). Tetrahydrofuran was the mobile phase (flow rate: 1 mL/min).Polystyrene standards with molecular weights between 1,480 and 70,950 Da (Polymer Laboratories, Varian, Les Ulis, France) were used to prepare the calibration curve. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

). One hundred µL samples were injected into an Alliance GPC (refractometer detector: 2414 RI, separation module e2695, Empower GPC IV. PLGA IMPLANTS FOR CONTROLLED DRUG RELEASE: IMPACT OF THE DIAMETER 133 software;

Table IV -1. Key

 IV

			Diameter, die
	Implant	0.5 mm	1 mm	2 mm	3 mm
	Diameter	0.66 ± 0.02	1.08 ± 0.02	1.98 ± 0.03	2.79 ± 0.19
	(mm)
	Weight	2.53 ± 0.10	6.84 ± 0.21	23.58 ± 1.11	46.22 ± 6.83
	(mg)
	Practical loading	12.1 ± 0.1	12.5 ± 0.1	12.3 ± 0.1	12.0 ± 0.1
	(%)
	Tg	35.6 ± 0.2	35.2 ± 0.2	36.2 ± 0.1	36.9 ± 0.3
	(°C)
	Mw (kDa)	23.9 ± 0.3	24.5 ± 0.3	25.7 ± 0.1	24.6 ± 0.4

properties of the investigated ibuprofen-loaded PLGA implants (Tg: glass transition temperature, Mw: average polymer molecular weight). Mean values ± standard deviations are indicated (n = 3).

Table VI - 1 .

 VI1 Pictures of the disc were taken using a SZN-6 trinocular stereo zoom macroscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vision Lite 2.1 software).

Table VI - 1 .

 VI1 Preparation method of polymer discs for melt rheology studies.

	Temperature (°C)	Time (min)	Vaccum
	Heating from 20 to 40	1 : 07
	40	15 : 00
			Yes
	Heating from 40 to 110	1 : 37
	110	5 : 00
	110	2 : 00	No

Table VI - 2 .

 VI2 Table VI-2. Printing parameters used to obtain the ibuprofen loaded PLGA-implant.

VI.3.10. In vitro drug release

  3 mg/mL). Fifty µL samples were injected into an Alliance GPC (refractometer detector: 2414 In all cases, at pre-determined time points, pictures were taken as described in section VI.3.4.

	VI. 3D PRINTING OF IBUPROFEN LOADED CONTROLLED RELEASE IMPLANTS
	WITH PLGA -PROOF OF CONCEPT AND IMPACT FILLING PATTERN
	experiments in all agitated bulk fluids. All experiments were conducted in triplicate. Mean
	values +/-standard deviations are reported.
	VI.3.11. Implant swelling
	Implants were treated as for the in vitro drug release measurements. At pre-determined time
	points: implant samples were withdrawn and excess water was carefully removed using
	Kimtech precision wipes (Kimberly-Clark, Rouen, France). The implants were weighed [wet
	mass (t)], and the change in wet mass (%) (t) was calculated as follows:
	𝑐ℎ𝑎𝑛𝑔𝑒 𝑖𝑛 𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (%)(𝑡) =	𝑤𝑒𝑡 𝑚𝑎𝑠𝑠 (𝑡) -𝑚𝑎𝑠𝑠 (𝑡 = 0) 𝑚𝑎𝑠𝑠 (𝑡 = 0)	× 100 %
	RI, separation module e2695, Empower GPC software; Waters, Milford, USA), equipped with
	a PLgel 5 µm MIXED-D column (kept at 35°C, 7.8 × 300 mm; Agilent). Tetrahydrofuran was
	the mobile phase (flow rate: 1 mL/min). Polystyrene standards with molecular weights between
	1,480 and 70,950 Da (Polymer Laboratories, Varian, Les Ulis, France) were used to prepare the
	calibration curve. All experiments were conducted in triplicate. Mean values +/-standard
	deviations are reported.
	Implants were placed in 50 mL tubes (352070 Corning-Falcon, New York, U.S;1 implant per
	tube), filled with 50 mL phosphate buffer pH 7.4 USP 42. The tubes were placed in a horizontal
	shaker (80 rpm, 37°C; GFL 3033; Gesellschaft fuer Labortechnik, Burgwedel, Germany). At
	predetermined time points, the entire bulk fluid was replaced by fresh release medium. The
	withdrawn samples were filtered (PVDF syringe filter, 0.45 µm; Agilent, Santa Clara,
	California, USA) and analyzed for their ibuprofen contents by HPLC-UV, as described in
	section VI.3.8.
	and the pH of the release medium was measured using a pH meter (InoLab pH Level 1; WTW,
	Weilheim, Germany). Also, in all cases, sink conditions were provided throughout the
			185

Table VI - 3 .

 VI3 Key properties of the ibuprofen loaded PLGA-implants. (Tg: glass transition temperature, Mw: molecular weight). Mean values ± standard deviations are indicated (Weight, Length, Width, Practical drug loading n=6, Thickness, Tg, Mw n=3).

			Hexagonal	Cross	Zigzag
	implant	Weight (mg)	138.1 ± 13.5	136.5 ± 7.1	127.9 ± 5.9
	PLGA	Length -width (mm)	10.1 ± 0.1	10.0 ± 0.1	10.0 ± 0.1
	loaded	Thickness (mm)
	Ibuprofen	Practical drug loading (%) Tg (°C)	10.3 ± 0.6 35.7 ± 0.2	10.7 ± 0.5 35.8 ± 0.4	10.5 ± 0.3 35.8 ± 0.1
		Mw (kDa)	21.7 ± 0.3	21.5 ± 0.6	21.7 ± 0.2

Table VII - 1 .

 VII1 Printing parameters used to obtain the ibuprofen-loaded PLGA-implants.

	VII. PLGA BASED 3D PRINTED IMPLANTS -IMPACT OF THE 3D PRINTED
			TECHNOLOGY ON DRUG RELEASE
	(target diameter between 1.60 and 1.80 mm, Tolerance Puller Noztek, Shoreham-by-Sea,
	United Kingdom).
			Droplet deposition	Fused filament
	Used technology
			modeling (DDM)	fabrication (FFF)
	Number of contouring layers	0
	Number of lower and upper layer	0
	Filling pattern		ZigZag
	Filling density (%)		30
	Nozzle diameter (mm)		0.25
	Layer thickness (mm)		0.25
			zone 1 : 120
				First layer: 125
	Printing temperature (°C)	zone 2 :110
				Otherwise: 115
			zone 3 : 110
	Printing speed (mm/s)		40	50
	Température du plateau (°C)	25	30
	Ventilation (%)		Not applicable	10
	Speed of retraction (mm/s)	Not applicable	40
	VII.3.2.2.	Fused deposition modeling, Volumic ®
	The PLGA was milled (4 x 30 s) with a grinder (Valentin, Seb, Ecully, France). Appropriate
	amounts of polymer and ibuprofen were mixed and then pre-melted in a crystallizer on a film
	applicator hot plate (110 ° C, 6 min, Erichsen, Rueil-Malmaison, France), then extruded as
	described in section 2.2.1. Except that: (i) a 4 mm die was used, (ii) the die temperature was set
	at 70 ° C, (iii) and a feeder flow regulator was used at a speed between 5 and 8 cc/min. The
	diameter of the filament obtained at the exit of the extruder was controlled by a laser sensor
				207

Table VII -2).

 VII

  The study aimed to evaluate the impact of filling density of ibuprofen-loaded PLGA implants obtained by DDM on implants' physicochemical characteristics and in vitro drug release performance. To do so, ibuprofen-loaded PLGA implants with 10% and 100% filling density were printed using the DDM Arburg Freeformer® technology. Thus, both implants' physicochemical characteristics were investigated before and after exposure to phosphate buffer pH 7.4. The impact of the filling density on the practical drug loading, polymer molecular weight, morphology by optical macroscopy, and scanning electronic microscopy was investigated. The impact of the difference observed on the in vitro drug release after exposure to phosphate buffer pH 7.4, and erosion of PLGA was also investigated.

	VIII. PLGA BASED 3D PRINTED IMPLANTS -IMPACT OF FILLING DENSITY ON
	DRUG RELEASE
	VIII.3. Materials and methods
	VIII.3.1. Materials
	Poly (D,L lactic-co-glycolic acid) (PLGA, 50:50 lactic acid: glycolic acid; Resomer RG 503H;
	Evonik, Darmstadt; Germany); ibuprofen (BASF, Ludwigshafen, Germany); agarose (genetic
	analysis grade), tetrahydrofuran (HPLC grade) (Fisher Scientific, Illkirch, France); potassium
	dihydrogen orthophosphate and sodium hydroxide (Acros Organics, Geel, Belgium);
	acetonitrile (VWR, Fontenoy-sous-Bois, France); sodium hydrogen phosphate (Na2HPO4,
	Panreac Quimica, Barcelona, Spain).
	VIII.3.2. Pellets preparation
	PLGA was milled (4 x 30 s) with a grinder (Valentin, Seb, Ecully, France). Appropriate
	amounts of polymer and drug powders were blended for 5 min with a mortar and a pestle,
	followed by extrusion using a Nano 16 twin-screw extruder (screw diameter = 16 mm,
	length/diameter ratio = 26.25, gravitational feeder; Leistritz, Nuremberg, Germany), equipped
	VIII.3.3. 3D printing of the implant
	Parallelepipeds (10 mm, 10 mm, 2.5 mm) with different filling densities (10 %, and 100 %)
	were printed using the Arburg Plastic Freeforming ® technology (Arburg GmbH, Loßburg,
	The emerging 3D printing technologies provide new horizons to precisely control drug release
	from PLGA based devices. PLGA thermoplastic properties make it a good choice for use with
	extrusion-based 3D printing technologies (55). Modeling (FDM) also known as Fused Filament Table VIII-1. Printing parameters used to obtain the ibuprofen loaded PLGA-implants of
	Fabrication (FFF), Direct Powder Extrusion (DPE), and Droplet Deposition Modeling (DDM) different density.
	are extrusion-based 3D printed technologies and are nowadays the most widely used (15-
	Filling density 17,56-58). The "Arburg Plastic Freeforming ® " technology is a thermoplastic DDM 10 % 100 % Nozzle diameter (µm) 250
	technology, capable to process a wide range of granulated polymer feedstocks commonly used Printing speed (mm/s) 40
	in injection molding processes (17,58). For example, Welsh et al studied the possibility to Layer thinckness (mm) 0.18 0.33
	modulate Dapivirine, release from polyurethane matrix for vaginal delivery (17). They T° 1 st zone (°C) 120
	investigate the impact of the filling density on drug release kinetic from vaginal rings, T° 2 nd zone (°C) 110
	highlighting that by decreasing the filling density, the Dapirivine release rate increases. T° 3 rd zone (°C) 110
	T° Printing chamber (°C)
	233

with a 2 mm diameter die. The process temperatures were kept constant at 80 -75 -70 -65 °C (die -zone 3 -zone 2 -zone 1). The screw speed was set at 50 rpm. After cooling, the hot melt extrudates were manually cut into pellets of 5 mm length.

Allemagne) illustrated in

Figure VIII-1. The pellets obtained by hot melt extrusion were fed into the Freeformer ®. The manufacturing parameters were set depending on the filling density as indicated in Table VIII-1.

Table VIII - 2 .

 VIII2 Key properties of the obtained ibuprofen loaded PLGA-implants. (Tg: glass transition temperature, Mw: molecular weight). Mean values ± standard deviations are indicated (Weight, Length, Width, Implant pore size n=36, Practical drug loading n=6, Tg, Mw n=3).As previously reported (17) DDM induces high mechanical stress, leading to a decrease in PLGA molecular weight (from 27.

	Filling density	10 %	100 %
	Weight (mg)	206.5 ± 5.9	304.7 ± 11.1
	Length -width (mm)	10.4 ± 0.4	10.0 ± 0.1
	Implants pore size (mm)	1.1 ± 0.2	-
	Practical drug loading (%)	13.6 ± 0.2	13.4 ± 0.1
	Tg (°C)	33.7 ± 0.3	34.0 ± 0.8
	Mw (kDa)	16.7 ± 0.5	15.0 ± 2.0

Carlier E, Marquette S, Peerboom C, Amighi K, Goole J. Development of mAb-loaded 3D-printed (FDM) implantable devices based on PLGA. Int J Pharm. 2021 Mar 15;597:120337.

J. Siepmann Conceptualization; Methodology; Supervision; Resources; Writing -Review & Editing; Visualization; Project administration; Funding acquisition

Review & Editing; Visualization; Project administration; Funding acquisition

Remerciements

VII.6. Acknowledgments

This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed. The authors are very grateful for this support. They would also like to thank Mr. A. Fadel from the "Centre Commun de Microscopie" of the University of Lille ("Plateau techique de la Federation Chevreul CNRS FR 2638") for his valuable technical help with the SEM pictures. VII.7. Author statement C. Bassand Investigation; Methodology; Validation; Conceptualization; Writing -Original Draft; Visualization L. Benabed Investigation; Methodology; Validation S. Charlon Investigation, Visualization; Project administration J. Verin Investigation; Methodology; Validation J. Freitag Investigation F. Siepmann Conceptualization; Methodology; Supervision; Resources; Writing -Review & Editing; Visualization; Project administration; Funding acquisition J. Soulestin Conceptualization; Methodology; Supervision; Resources; Writing -Review & Editing; Visualization; Project administration; Funding acquisition J. Siepmann Conceptualization; Methodology; Supervision; Resources; Writing -Review & Editing; Visualization; Project administration; Funding acquisition VIII.6. Acknowledgements This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed. The authors are very grateful for this support. They would also like to thank Mr. A. Fadel from the "Centre Commun de Microscopie" of the University of Lille ("Plateau techique de la Federation Chevreul CNRS FR 2638") for his valuable technical help with the SEM pictures.

souhaite à tous beaucoup de réussite à venir, et espère avoir This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed. encore de belles occasions de partager des moments avec vous autour d'un « thé », d'un gâteau, d'une petite blague ou bien de la suite de « la dernière série » hollywoodienne.

II.5. Conclusion

The presence of an agarose gel surrounding PLGA implants significantly hinders polymer swelling (sterically) and slows down drug release (due to delayed penetration of substantial amounts of water into the system). In vivo it can be expected that surrounding tissue has a similar mechanical effect. However, yet it is unknown how important the impact of mechanical stress caused by body movements (e.g., muscle contractions) for the fate of a degrading PLGA implant is. The results presented in this study can help developing more realistic in vitro drug release set-ups for parenteral drug delivery systems. They also strengthen the hypothesis that implant swelling plays an orchestrating role for the control of drug release from PLGA-based drug delivery systems.

II.6. Acknowledgements

This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed.

The authors are very grateful for this support. They would also like to thank Mr. A. Fadel from the "Centre Commun de Microscopie" of the University of Lille ("Plateau techique de la Research article -To be submitted in Journal of Drug Delivery Science and technology Hot melt extruded ibuprofen-loaded PLGA implants: Importance of heat exposure C. Bassand 1 , L. Benabed 1 , J. Verin 1 , F.Danede 2 , L.A. Lefol 1 , J.F. Willart 2 , F. Siepmann 1 , J. Siepmann 1 1 Univ. Lille, Inserm, CHU Lille, U1008, F-59000 Lille, France 2 Univ. Lille, USTL UMET UMR CNRS 8207, F-59650 Villeneuve d'Ascq, France

III.1. Abstract

Hot melt extrusion offers an interesting potential for the manufacturing of poly(lactic-coglycolic acid) (PLGA)-based implants. However, the heat treatment might substantially alter the polymer, drug and degree of drug-polymer mixing. The aim of this study was to better understand the impact of varying exposure to 105 °C from 3 to 15 min in the case of ibuprofenloaded PLGA implants. In vitro drug release was measured in phosphate buffer pH 7.4, optical & scanning electron microscopy, DSC, GPC, X-ray diffraction and gravimetric analysis were used to monitor dynamic changes of the implants' morphology, dry & wet mass and average polymer molecular weight. Interestingly, increasing the heat exposure from 3 to 15 min led to a decrease in the amount of crystalline drug present in the system, resulting in a slight decrease in the intial burst release. The average PLGA molecular weight also slightly decreased during the heat treatment. In contrast, the relatively rapid penetration of water into the entire implants and subsequent polymer degradation throughout the devices was not affected to a noteworthy extent. Also the onset of substantial implant swelling after about 1 week and the subsequent onset of the final rapid drug release phase (accounting for about 80 % drug release) was not significantly altered. However, for different drugs and polymers changes in their physical state upon heat exposure during hot melt extrusion might have much more importance consequences for drug release, and the hot melt extrusion process might be much less robust. Keywords: PLGA; implant; drug release mechanism; swelling; monolithic solution; solid dispersion 70 to 120 °C, cooling to -70 °C, re-heating to 120 °C (heating and cooling rates = 10 °C/min).

The reported glass temperatures (Tgs) were determined from the 1 st heating cycles in the case of the implants (the thermal history being of interest), and from the 2 nd heating cycles in the case of the raw materials and their physical mixtures (the thermal history not being of interest).

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

III.3.8. X ray powder diffraction

X-ray powder diffraction patterns were recorded using a PANalytical X'Pert pro MPD powder diffractometer (PANalytical, Almelo, Netherlands), equipped with a Cu X-ray tube (λCuKα = 1.54 Å) and the X'celerator detector. Samples were placed in a spinning flat sample holder. The measurements were performed in Bragg-Brentano θ-θ geometry. The diffractograms were recorded from 3 to 60° (2θ) (0.0167 ° steps, 100 s/step).

III.3.9. Scanning electronic microscopy (SEM)

The internal and external morphology of the implants before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxford, England). Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome layer.

If indicated, the investigated implants had been exposed to the release medium before, as described for the in vitro release studies (please see section III. 3.4). At predetermined time points, implant samples were withdrawn and freeze-dried (as described in section III. 3.6). Cross sections were obtained by manual cutting with a scalpel, prior to freeze drying.

Research article -To be submitted European Journal of Pharmaceutics and Biopharmaceutics

PLGA implants for controlled drug release: Impact of the diameter C. Bassand, J. Freitag, L. Benabed, J. Verin, F. Siepmann, J. Siepmann Univ. Lille, Inserm, CHU Lille, U1008, F-59000 Lille, France

IV.1. Abstract

The aim of this study was to better understand the importance of the diameter of poly(lacticco-glycolic acid) (PLGA)-based implants on system performance, in particular the control drug release. Different types of ibuprofen-loaded implants were prepared by hot melt extrusion using a Leistritz Nano 16 twin-screw extruder. Drug release was measured in well agitated phosphate buffer pH 7.4 bulk fluid and in agarose gels in Eppendorf tubes and transwell plates. Dynamic changes in the implants' dry & wet mass, volume, polymer molecular weight, inner & outer morphology were monitored using optical macroscopy, gravimetric analysis and scanning electron microscopy. The physical states of the drug and polymer were determined by DSC and pH changes in the release medium investigated. Irrespective of the type of experimental set-up, the resulting absolute and relative drug release rates decreased with increasing implant diameter (0.7 to 2.8 mm). Bi-phasic drug release was observed in all cases from the monolithic solutions (the ibuprofen being completely dissolved in the polymer matrices): A zero order release phase was followed by a final, rapid drug release phase (accounting for 80-90 % drug). The decrease in the relative drug release rate with increasing system diameter can be explained by the increase in the diffusion pathway lengths to be overcome. Furthermore, the observed delay in the onset of the final rapid drug release phase with increasing implant diameter could be explained by the higher mechanical stability of the devices, delaying the onset of substantial entire system swelling after a lag phase of about 5 to 10 d. Interestingly, no signs for the occurrence of important autocatalytic effects were observed in this study, despite the considerable system dimensions.

Keywords

IV.5. Conclusions

The variation of the diameter of a PLGA implant is an effective tool to adjust desired absolute and relative release kinetics. In the case of hot melt extruded ibuprofen-loaded implants, biphasic release patterns were observed: A zero order release phase was followed by a rapid, final drug release phase (accounting for 80-90 % drug). The decrease in the relative drug release rate with increasing system diameter can be attributed to the longer diffusion pathways to be overcome. The onset of the final rapid drug release phase was delayed, because the larger polymer matrices are mechanically more stable, retarding the onset of substantial entire implant swelling, which facilitates drug release.

IV.6. Acknowledgments

This project has received funding from the Interreg 2 Seas programme 2014-2020 co-funded by the European Regional Development Fund under subsidy contract No 2S04-014 3DMed.

The Authors would also like to thank Mr. A. Fadel from the "Centre Commun de Microscopie" of the University of Lille ("Plateau technique de la Federation Chevreul CNRS FR 2638") for his valuable technical help with the SEM pictures. Sodiun hydroxide solution 10% (Sigma Aldrich, Saint-Louis, Missouris, USA).

IV

V.3.2. Implant preparation

Appropriate amounts of polymer [milled for 4 x 30 s with a grinder (Valentin, Seb, Ecully,

France)] and drug powders were blended for 5 min at 20 rpm in a Turbula T2C Shaker-Mixer (Willy A Bachofen, Basel, Switzerland). Three hundred mg of the mixture were filled into a 1 mL syringe (Henke Sass Wolf, Tuttlingen, Germany), followed by heating to 105°C for 15 min in an oven (FP115, Binder, Tuttlingen, Germany). The molten blend was manually extruded using the syringe. The extrudate was cut with a hot scalpel into cylindrical implants of approximately 5 mm length.

V.3.3. Optical macroscopy and implant dimensions

Pictures of implants before exposure to the release medium were taken using a SZN-6 trinocular stereo zoom macroscope (Optika, Ponteranica, Italy), equipped with an optical camera (Optika Vision Lite 2.1 software). The lengths and diameters of the implants were determined using the ImageJ software (US National Institutes of Health, Bethesda, Maryland, USA).

V.3.4. Practical drug loading

Implants were dissolved in 5 mL acetonitrile, followed by filtration (PVDF syringe filters, 0.45 µm; Agilent Technologies, Santa Clara, USA) and drug content determination by HPLC-UV analysis using a Thermo Fisher Scientific Ultimate 3000 Series HPLC. The latter was equipped with an LPG 3400 SD/RS pump, an autosampler (WPS-3000 SL) and a UV-Vis detector (VWD-3400RS) (Thermo Fisher Scientific, Waltham, USA). A reversed-phase column C18 (Gemini 5 µm; 110 A°; 150 x 4.6 mm; Phenomenex, Le Pecq, France) was used. The mobile V. HOW BULK FLUID RENEWAL CAN AFFECT DRUG RELEASE FROM PLGA IMPLANTS 159 phase was a mixture of 30 mM Na2HPO4 pH 7.0: acetonitrile (60:40, v:v). The detection wavelength was 264 nm, and the flow rate 0.5 mL/min. Ten microliter samples were injected.

V.3.5. Drug solubility measurements

Excess amounts of ibuprofen were exposed to 20 mL phosphate buffer pH 7.4 USP 42 in glass flasks, which were horizontally shaken (80 rpm) at 37°C in an incubator (GFL 3033;

Gesellschaft fuer Labortechnik, Burgwedel, Germany). Optionally, the pH was adjusted using aqueous solutions of 10 N lactic acid or 0.05 N NaOH (final pH values are reported). At predetermined time points, samples were withdrawn, immediately filtered (PVDF syringe filter, 0.45 µm; Agilent, Santa Clara, USA) and diluted. The drug contents of the samples were determined by HPLC-UV, as described above. Samples were withdrawn until equilibrium was reached. Each experiment was conducted in triplicate, mean values +/-standard deviations are reported.

V.3.6. In vitro drug release

Three experimental set-ups were used to measure ibuprofen release from the PLGA implants:

V.3.6.1. In well-agitated bulk fluids

Implants were placed in 5 mL Eppendorf tubes (1 implant per tube), filled with 5 mL phosphate buffer pH 7.4 USP 42 (Figure V-1A). Metal baskets avoided that the implants sank to the bottom of the tubes (resulting in potentially limited contact with the bulk fluid). The mesh size (250 µm) was sufficient to allow for convective flow and rapid medium exchange "insideoutside" the basket. The tubes were placed in a horizontal shaker (80 rpm, 37°C; GFL 3033;

Gesellschaft fuer Labortechnik, Burgwedel, Germany). At predetermined time points, the entire bulk fluid (5 mL), or 3 mL or 1 mL of the bulk fluid was replaced by fresh release medium.

The withdrawn samples were filtered (PVDF syringe filter, 0.45 µm; Agilent) and analyzed for their ibuprofen contents by HPLC-UV, as described above.

V.3.6.2. In agarose gels in Eppendorf tubes

Implants were embedded into agarose gels in 5 mL Eppendorf tubes, as illustrated in Great caution has to be paid when defining the conditions for in vitro drug release measurements from PLGA implants: The systems become highly fragile once substantial polymer swelling sets on. The renewal of important parts of the release medium can cause partial damage of the highly fragile polymeric systems, accelerating device disintegration and decreasing the length of the diffusion pathways for the drug to be overcome. Embedding implants into agarose gels and limiting the sampling volume can limit these effects. In any case, it must not be forgotten that, once substantial system swelling sets on, the devices become highly fragile.

V.6. Acknowledgments

This project has received funding from the Interreg This study aimed to determine the impact of heating treatment by fused filament fabrication (FFF) on poly (lactic-co-glycolic acid) (PLGA) properties and to investigate the impact of filling pattern on in vitro drug release from printed ibuprofen loaded PLGA implants. To do so, the interaction between the PLGA and Ibuprofen was investigated by differential scanning calorimetry and thermogravimetric analysis. The viscoelastic properties of the melt polymer and Ibuprofen PLGA mixture were investigated by rheology temperature sweep experiments.

Ibuprofen-loaded PLGA filaments were produced by hot-melt extrusion (HME) using a Leistritz Nano 16 twin-screw extruder coupled with a laser sensor Tolerance Puller Noztek ® .

Those filaments were used to investigate the impact of temperature applied by 3D printing, via fused filament fabrication (FFF) with a Volumic ® printer, on PLGA molecular weight and ibuprofen content. PLGA ibuprofen mixture was suitable for 3D printing, as their viscoelastic properties allow to (i) obtain ibuprofen loaded filaments by HME, (ii) 3D printing with temperature from 115°C to 150°C without any substantial loss of PLGA molecular weight or ibuprofen content (iii) obtain ibuprofen loaded implants with three different internal structures.

The in vitro performance of the different 3D printed implants was investigated after exposure to phosphate buffer pH 7.4 37°C 80 RPM. Absolute and relative drug release, increase in wet mass, pH in the release medium, and morphology of those implants after exposure to the release medium were investigated. The obtained ibuprofen-loaded 3D printed PLGA implants allow to maintain the release of ibuprofen for up to 9 days, with a monophasic release pattern.

Keywords: PLGA, 3D printing, melt rheology, infill pattern, fused deposition modeling, drug release Three-dimension (3D) printing has many advantages, such as the flexibility of dosages or the possibility of creating complex structures. However, the choice of printing technology and settings can greatly influence product performance. The objective of this research was, first of all, to study and compare the impact of two printing processes on the physicochemical properties of poly (lactic-co-glycolic) acid (PLGA) and the in vitro performance of ibuprofen loaded-implants obtained. To do so, two sets of implants were printed using Droplet Deposition Modeling (DDM) with the Arburg Freeformer ® technology and Fused Filament Fabrication (FFF) with a Volumic ® printer. The obtained implants were characterized regarding morphology (optical macroscopy and scanning electronic microscopy), solid-state (differential scanning calorimetry), drug loading, and molecular weight of the polymer. The implants in vitro performance was assessed after exposure to phosphate buffer pH 7.4, 37°C, 80 RPM. Absolute and relative in vitro drug release, the evolution of the pH in the release medium, increase in wet mass, and morphologically changed upon exposure to release medium were evaluated. The degradation behavior of PLGA upon exposure to release medium (loss of dry mass, and molecular weight of the polymer) with the two types of implants obtained was also investigated.

The implants printed by DDM have a three-phase in vitro release profile with a "burst" effect, then a slower release phase, and finally a final rapid release phase due to the significant swelling of the PLGA. While the FFF-printed implants exhibit a two-phase in vitro release profile.

Keywords: PLGA, 3D printing, fused filament fabrication, droplet deposition modeling, in vitro drug release.

polymer (1). In both cases, it consists of dispensing molten or semi-solid material through the orifice of a nozzle, deposed layers by layers to obtain a 3D object (50) (die -zone 3 -zone 2 -zone 1). After cooling, the hot melt extrudates were manually cut into pellets of 5 mm length.

The pellets obtained by hot extrusion were fed into the feeder of the Arburg Freeformer® printer (Freeformer, Arburg, Germany). Parallelepiped-shaped implants with dimensions of 10 mm x 10 mm x 2.5 mm were printed using a 250 μm nozzle according to the parameters described in

The detection wavelength was 225 nm, and the flow rate was 0.5 mL/min. Ten microliter samples were injected.

VII.3.5. Differential scanning calorimetry (DSC)

DSC thermograms of the raw materials (PLGA and ibuprofen), pellets, filament, and 3D printed implants were recorded using a DCS1 Star System (Mettler Toledo, Greifensee, Switzerland).

Approximately 5 mg raw material, pellets, and 3D printed implants, were heated in pierced aluminum pans as follows: from -70 to 120 °C, cooling to -70 °C, re-heating to 120 °C (heating/cooling rate = 10 °C/min. The reported glass temperatures (Tgs) were determined from the 1 st heating cycles in the case of the pellets, filaments, and 3D printed implants (the thermal history being of interest), and from the 2 nd heating cycle in the case of the raw material (the thermal history not being of interest). All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

VII.3.6. Gel permeation chromatography (GPC)

The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation chromatography (GPC) as follows: raw materials, samples were dissolved in tetrahydrofuran

where mass (t = 0) denotes the mass of the implant before exposure to the release medium.

All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

VII.3.9. Implants erosion and PLGA degradation

Implants samples were treated as for the in vitro drug release studies described above. At predetermined time points, implants samples were withdrawn and freeze-dried (freezing at -45°C for 2 h 35 min, primary drying at -20 °C/0.940 mbar for 35 h 10 min, secondary drying at +20 °C/0.0050 mbar for 35 h; Christ Alpha 2-4 LSC+; Martin Christ, Osterode, Germany).

The dry mass (%) (t) was calculated as follows:

where mass (t = 0) denotes the mass of the implant before exposure to the release medium. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

The average polymer molecular weight (Mw) of the PLGA was determined by gel permeation chromatography (GPC) as described in section VII.3.6. All experiments were conducted in triplicate. Mean values +/-standard deviations are reported.

VII.3.10. . Scanning electronic microscopy (SEM)

The internal and external morphology of the implants before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxfordshire, England).

Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome layer. In the case of implants that had been exposed to the release medium, the systems were treated as described for the in vitro release studies as described in section VII.3.7. At predetermined time points, implant samples were withdrawn, optionally cut (for cross-sections) using a scalpel, and freeze-dried (as described in sectionVII.3.9). Photos taken by SEM (Figure VII-3) confirmed this observation, as the difference in the deposed line can be seen. Both types of implants show a slight surface roughness and small cross-sectional pores (enlighten in dotted red line). Those pores are more numerous inside the polymer matrix of DDM implants. These small pores are described in the literature as being the consequence of a high pressure exceeding 400 bar at the nozzle of the Arburg® Freeformer printer which would induce the evaporation of either ibuprofen molecules, even at lower temperatures than its boiling point, or residual water contained in the raw materials (17).

aluminum pans as follows: from -70 to 120 °C, cooling to -70

VIII.3.11. . Scanning electronic microscopy (SEM)

The internal and external morphology of the implant before and after exposure to the release medium was studied using a JEOL Field Emission Scanning Electron Microscope (JSM-7800F, Japan), equipped with the Aztec 3.3 software (Oxford Instruments, Oxfordshire, England).

Samples were fixed with a ribbon carbon double-sided adhesive and covered with a fine chrome layer. In the case of implants that had been exposed to the release medium, the systems were treated as described for the in vitro release studies as described in section VIII.3.7. At predetermined time points, implant samples were withdrawn, optionally cut (for cross-sections) using a scalpel, and freeze-dried (as described in section VIII.3.10).

IX. GENERAL CONCLUSION AND

DISCUSSION

SUMMARY

SUMMARY

This thesis was driven by two main axes:

On one hand, poly (lactic-co-glycolic) acid (PLGA) is a first category choice for polymer-based parenteral controlled drug delivery systems due to its excellent biocompatibility, biodegradability, and use in several FDA-approved products for decades (microparticles, implants, in situ forming implants). However, establishing an in vitro in vivo correlation regarding drug release of such devices is challenging as many different mechanisms are involved in drug release. Thus, it is of great importance to properly characterize those kinds of devices, to identify the release mechanism by which the API is released from the device.

On the other hand, in the past few years, three -dimension (3D) printing technologies have revolutionized the development of controlled drug release devices. These technologies allow to precisely control the size, shape, porosity of the device, and so the release rate of the API.

In this context, the two objectives of this thesis were:

(i) To develop an experimental set-up for in vitro drug release measurement for implants aimed to be administrated to the subcutis, and investigate the impact of several implants attribute on in vitro drug release;

(ii) To develop PLGA based controlled drug release implants using 3D printing technologies.

As subcutaneous environment behaves more like a gel than a fluid, two in vitro set-up involving agarose gel to mimic the subcutis have been developed. The impact of those set-ups, on the behavior or hot melt extruded ibuprofen loaded PLGA based implants have been investigated.

Through this investigation, the following conclusions have been drawn:

-The presence of an agarose gel surrounding PLGA implants significantly hinders polymer swelling (sterically) and slows down drug release (due to delayed penetration of substantial amounts of water into the system).

-Increasing exposure times to heat might increase the relative amount of drug, which is dissolved in the PLGA, altering the importance of the initial "burst release".

-The variation of the diameter of a PLGA implant is an effective tool to adjust desired absolute and relative release kinetics.

-Great caution has to be paid when defining the conditions for in vitro drug release measurements from PLGA implants.

SUMMARY 261

Then, the impact of the heating treatment on PLGA ibuprofen mixture rheological behavior to access its printability have been investigated. Ibuprofen loaded PLGA based implants have been printed using: Fused Filament Fabrication (FFF), and Droplet Deposition Modeling (DDM).

The impact of the 3D printed technology, the filling pattern, and the filling density of the implants on their in vitro performances have been investigated. By investigating those parameters, the following conclusions have been drawn:

-Ibuprofen loaded PLGA filaments were printable without any substantial loss of molecular weight of the polymer nor ibuprofen up to 150°C, and their viscoelastic behavior are compatible with 3D printing.

-3D printing of implants containing ibuprofen, with three different internal structures was achievable. This leading to the obtention of implants allowing to maintain the release of ibuprofen for up to 9 days, with a monophasic release pattern.

-DDM technology allows to easily obtain homogeneous drug-loaded implants as the printing is done from a statistical mixture of pellets, allowing to avoid the challenging development of filament with constant diameter. However, this technology exposes the polymer and the drug substance to high mechanical stress compared to FFF technology.

-The variation of the filling density is an effective tool to adjust desired absolute and relative release kinetics. The obtained implants exhibit a triphasic release pattern: with relative higher burst release and shorter lag phase for the lowest filling density implants.

The decrease in the relative drug release rate with increasing implant filling can be attributed to the longer diffusion pathways to be overcome.

Keywords