1. Introduction
Mass manufacturing of oral formulations fails to produce drug forms that adequately address the complexity of an individual’s disease state, age and genetic factors [
1,
2,
3]. This “one-size-fits-all” approach has led to dose-related adverse drug reactions (ADRs), particularly for older patients using polypharmacy [
4,
5,
6]. Recent investigations into the use of three-dimensional printing (3DP) technology as an alternative additive manufacturing method [
7,
8,
9,
10,
11,
12], shows great potential for delivering bespoke oral drug medications.
3DP encompasses numerous techniques based around the automated and sequential layering of material to construct an object, based on a digitally modelled design. A noteworthy appeal of 3DP in pharmaceutics can be attributed to the potential to generate oral dosage forms in a rapid, reproducible and flexible manner. To date, the realm of 3DP in pharmaceuticals has been dominated by techniques including inkjet/binder jetting [
13], selective laser sintering (SLS) [
14,
15], and fused deposition modelling (FDM) [
16,
17,
18].
FDM is a widely used 3DP technique across multiple disciplines that requires a solid filament to be melted at high temperatures, for extrusion through a small orifice on a layer-by-layer pattern and solidification to produce solid models. In medicine, and more recently in pharmaceutics, this technique has seen use in the production of hydrolysable and biodegradable polymer-based scaffolds [
19,
20,
21]. However, it is a technique that has been largely avoided in the 3DP of pharmaceutics due to the thermolabile nature of most drugs at temperatures required for its extrusion. Recently, FDM has gained greater interest with the development of new polymers that may be coextruded, with a drug, at lower temperatures to avoid thermal degradation during the extrusion process [
21,
22].
The oral bioavailability of lipophilic drugs can be improved with the coadministration of lipids in the form of food- or lipid-based formulations [
23,
24]. Personalising lipid-based formulations would represent an important advance in treatments with lipophilic small molecule drugs. However, despite the physicochemical properties of many lipids allowing their extrusion at low temperatures, there is a lack of research into the extrusion of lipid-based formulations in the context of 3DP [
25]. Even less is understood around the interplay between surface area-to-volume ratio, dispersion and digestion of such formulations which will dictate release of the drug and rate of absorption for permeable lipophilic compounds.
Lipid extrusion is already recognised as a pharmaceutical manufacturing technique, and creamy spheroids have been produced from the extrusion and pelletisation of a lipid-based solid self-emulsifying drug delivery system (S-SEDDS) [
26]. In 2019, Vithani et al. demonstrated a novel approach to the 3DP of thermolabile and lipophilic drugs loaded into solid lipid-based formulations [
27]. Vithani employed semi-solid extrusion (SSE) to 3D print solid oral tablets from a lipid-based solid self-micro emulsifying drug delivery system (S-SMEDDS) as the extrudate. In addition, Chatzitaki et al. also employed similar methods to produce a solid lipid-based suppository dosage form via SSE [
28]. These methods did not require high temperatures or solvent evaporation; however, the final dosage forms had poor printing fidelity which presents concern for the consistency of behaviour upon ingestion.
It has also been more than a decade since research investigating ‘polypill’ formulations has been conducted to combat ADR and adherence issues associated with polypharmacy [
29,
30]. Recent research using primarily polymer-based formulations suggest that it is possible to 3DP oral dosage forms with multiple drugs and different release profiles [
15,
31,
32,
33,
34,
35,
36]. Furthermore, Markl et al. demonstrated that a 3D-printed dissolvable polymer scaffold could be exploited to generate dual compartmented tablets containing a liquid lipid formulation [
36]. However, there is a lack of research in the viability of lipid-based 3DP ‘polypills’ with tuneable drug release especially utilising solid lipid systems.
Therefore, in this study, we have evaluated the control over formulation dispersion and drug release using 3D-printed single- and multi-compartment polymer scaffolds filled with dispersible lipid formulations. The ‘polymer lipid hybrid’ (PLH) tablets were evaluated for the effect of exposed surface area; dissolvable vs non-dissolvable printed scaffolds; as well as multicompartment systems for asynchronous release, on the dispersion of the lipid formulation filled into the compartments of the scaffold (
Figure 1). Firstly, single-compartment polymer scaffolds with systematically varying exposed surface area were 3D-printed and filled with an S-SMEDDS formulation loaded with fenofibrate (FEN) as the model poorly soluble drug. The polymer scaffolds were prepared from either poly-lactic acid (PLA, not degradable in the gut on a physiological timescale) and polyvinyl alcohol (PVOH, dissolves on physiological timescales). Subsequently, as a proof of concept, multi-compartment scaffold systems were 3D-printed with six ‘wedges’ forming the circular scaffold (
Figure 1), and the wedges were filled individually with the six different combinations of three drugs (clofazimine, lumefantrine and halofantrine) in two different S-SMEDDS formulations. These studies serve as a proof of concept for a PLH system that may enable better control over the fidelity of the lipid formulation in a 3D-printable format and provide great versatility in release of different drugs to address the problems with polypharmacy.
Dispersion studies of the formulations from PLA and PVOH scaffolds were conducted within simulated gastric to intestinal media. Nephelometry and HPLC analytical techniques were employed to determine the rate of dispersion and drug release, respectively.
2. Materials and Methods
Poly-lactic acid filament (Ultimaker PLA, sourced from Imaginables, Melbourne, Australia), poly-vinyl alcohol filament (Ultimaker PVA, sourced from Imaginables, Melbourne), FaSSIF powder (FaSSIF/FeSSIF/FaSSGF, 58 g, 25 L size, Biorelevant.com Ltd, London, United Kingdom), clofazimine and lumefantrine (Sigma, Burlington, VT, USA), halofantrine HCl (SmithKline Beecham Pharmaceuticals, Mysore, India), Gelucire® 44/14 and Gelucire® 48/16 (Gattefossé, Saint-Priest, France), Kolliphor P 188 and NaOH (Sigma Aldrich, Burlington, USA), M4670 silicone rubber mixture (Barnes, Melbourne, Australia), sodium dihydrogen orthophosphate (NaH2PO4, AnalaR—Merck, Kenilworth, NJ, USA), NaOH pellets and acetonitrile gradient grade (Ajax Finechem (Univar), Scoresby, Australia).
2.1. 3D Printing PLA and PVOH Scaffolds
A fused deposition modelling (FDM)-based printer was used to generate physical 3D scaffolds from commercial poly-lactic acid (PLA) and poly-vinyl-alcohol (PVOH) filaments. A computer aided drawing (CAD) software tool ‘Sketchup’ was used to virtually model and export scaffolds as stereolithography files (’.stl’ extension). The stereolithography files were then processed (sliced) with Ultimaker Cura software (v4.6.1–‘Master’, 2020, Ultimaker, Utrecht, Netherlands) using custom slicing profiles. The printing of the PLA and PVOH scaffolds was with a layer height of 0.1 mm, and a print speed of 40 mm/s. These processed models were saved as a G-Code file onto a secure digital (SD) card (‘.gcode’ extension). An Ultimaker 2 3D printer (U2) equipped with a 0.6 mm diameter nozzle was prepared for printing with the desired filament. The U2 build plate was set to 60 °C for 5 min before extruding a small amount of filament from the nozzle at ~215 °C to ensure smooth and featureless filament extrusion. The G-Code file was then manually selected from the inserted SD card in an Ultimaker 2 3D printer and once printing had begun the printing speed was manually set to 80%.
Table 1,
Table 2 and
Table 3 below present the geometric properties for the 3D-printed scaffolds used in this study. The images show the virtual models designed for PLA-based scaffolds, with properties of the PVOH-based scaffolds underneath. Each of the multicompartment scaffolds (MCTs) and scaffold containers A, B and C (when covered) possess the same target internal volume of ~393 µL. The printing times have been calculated by multiplying the Ultimaker Cura printing time prediction by the print speed dilation (1.25 at 80%).
2.2. Preparing Lipid Filled 3DP Scaffolds
2.2.1. Preparation of Drug-Loaded SMEDDS
The polymer–lipid hybrid tablets contained a drug-loaded SMEDDS formulation, which was manually injected as molten liquid into the 3DP scaffold. The dispersion properties of the SMEDDS have been previously reported by Vithani et al. [
27]. Drug-loaded SMEDDS were prepared following a previously reported process and formulation [
27] (
Table 4 and
Table 5), where FEN was used as the model drug. In this study, FEN was used for the single compartment PLHs in SMEDDS A, while clofazimine (CLO), lumefantrine (LUM) and halofantrine (HAL) were used for the multicompartment systems, where each of the six drug/SMEDDS combinations were individually prepared and loaded separately into a ‘wedge’ in the dose form. It is important to realise that 3DP systems with FDM and SSE heads in the one unit have been reported [
37], and that the manual preparation method is sufficient to test the concept of controlling dispersion and release from varying geometries and from independent compartments in the multicompartment system [
36].
Batches of drug-loaded SMEDDS for use in single compartment studies were prepared in quantities ranging from 10 to 15 g (
Table 4.), and individual drug batches for multicompartment studies were prepared in quantities of 3 g (
Table 5).
The process for preparing a batch of drug-loaded SMEDDS is illustrated in
Figure 2. Gelucire® 44/14 (GEL 44) was weighed into a glass beaker which was then placed onto a heated hotplate stirrer set to 80 °C (melting point of FEN) and stirred using a magnetic stirrer bar. Gelucire® 48/16 was deposited into the molten GEL 44. Then, Kolliphor P 188 was deposited into the mixture in a similar manner. Lastly, the drug component was deposited into the mixture and stirred for 30 min to ensure homogeneity. The mixture was then transferred to a 50 mL polypropylene tube and immediately sonicated for 2 min in a sonication bath. This was done to remove any air bubbles from the molten SMEDDS.
2.2.2. Filling 3DP Scaffolds with Molten Lipid Formulation
Five different tablet formats possessing a range of surface area-to-volume ratios (SA:V) were constructed for the study of single-compartment systems (
Table 6). Each tablet format refers to a different scaffold (or scaffolds), printed with PLA (
Figure 3A) or PVOH (
Figure 3B), that has been filled with a molten SMEDDS formulation. SA:V values were calculated on the assumption that the theoretical cavity volume (392.6 µL) was occupied by the SMEDDS. ‘Semi-open’ and ‘Closed’ tablet formats required scaffold covers (
Table 2A,B, respectively) in addition to a scaffold container (
Table 1C), which were manually attached after the SMEDDS had solidified.
A target mass of 400 mg molten drug-loaded SMEDDS was pipetted into a polymer scaffold, whilst the scaffold remained on the weighing plate. For the filling of ‘no-scaffold’ type systems, a silicon mould made from the scaffold cavity was used instead (See
supplementary information Figure S1A). For the filling of ‘dual-face’ scaffolds, the scaffold was attached to a cold smooth surface before filling (See
supplementary information Figure S1B).
2.2.3. Filling Multicompartment Scaffolds with Molten Lipid Formulation
Multicompartment tablet systems consisted of CLO, HAL and LUM incorporated into compositions 1 and 2 (
Table 5) to uniquely fill each ‘wedge’ with a target of 66 mg of molten drug-loaded SMEDDS, as indicated in
Figure 4 below. The filled mass for each wedge is provided in the
supplementary information (Table S2).
2.2.4. Preparation of Mixed Drug and Formulation Pellets
Pellets without a scaffold were also prepared from a mixture of all the drug:lipid compositions used in the multicompartment systems as control systems (
Figure 5). The molten mixture was deposited into a silicon mould, solidified and the pellet pressed out of the mould to generate solid mixed drug/formulation pellets without a scaffold.
2.3. Dispersion Studies in Simulated Gastric to Intestinal Media
In vitro dispersion studies in sequential gastric and intestinal media were conducted as illustrated in
Figure 1 (in vitro dispersion testing). Titration vessels (Metrohm titration vessel with thermostat jacket 20–90 mL) were connected to a heated water pump, set to 37 °C (Ratek immersion heater and insulation tank), using rubber tubing. The jacketed vessels were placed on a multipoint magnetic stirrer (Daihan Scientific, Wonju, Korea) set to 300 rpm with one magnetic stirrer bar placed in each vessel (L. 30 mm, bar diam. 8 mm).
A 3D-printed PLA mount (see
supplementary information Figures S2 and S3) was used to attach to and submerge the scaffolds containing the formulations. The weight of the scaffolds, formulations and drug mass for each system are detailed in the
Supplementary Information Tables S1 and S2. For the first 30 min, the formulation was submerged in simulated gastric fluid (gastric fluid herein) followed by a 3-min transition window, then 30 min of fasted state simulated intestinal fluid (intestinal fluid herein). Gastric fluid (0.1 M HCl, pH 1.2) was deposited into the titration vessel 5 min prior to dispersion, to allow for heat transfer. Aliquots (1 mL) were taken at timepoints; 1, 3, 5, 10, 15, 20, 30, 34, 26, 38, 43, 53, 63 min after immersion (t
1–t
14). Aliquots were immediately transferred to 1.5 mL Eppendorf tubes and stored for analysis. At the end of the gastric phase, 3.2 mL of 0.95 M NaOH was added to the vessel to neutralise the gastric fluid, followed by 6.4 mL of ‘Concentrated FaSSIF solution’ (
Table 7), and finally 6.4 mL of deionised water to establish the simulated intestinal condition.
The concentrated FaSSIF solution was prepared using amounts of each component at 4.5 times the amount needed for physiological relevance to allow for dilution.
2.4. Dispersion Kinetics of SMEDDS Studied Using Nephelometry
Samples retrieved at the timepoints indicated in
Section 3.3 were vortexed for ~5 s before aliquoting 350 µL of the dispersion into the well of a 96-well UV-transparent microplate. Turbidity was determined using a Nephelostar Plus nephelometer (BMG Labtech) set to 37 °C, with a laser beam focus of 2.5 mm and laser intensity of 80%. A settling and interval time of 1 s each was allowed between measurements. Furthermore, the sample plate was agitated using the double orbital shaking option for 1 s at 400 rpm, prior to each measurement. The total formulation dispersed (%) in gastric or intestinal fluid was determined by extrapolation from a dispersed FEN:SMEDDS calibration curve.
2.5. Quantification of Drug in Dispersion Media—UPLC
2.6. Solid State Characterisation - X-ray Diffractometer (XRD)
A Shimadzu X-ray Diffractometer (XRD-7000L) was used to measure the changes in diffracted X-ray intensities of solid samples, at a rotation angle range of 10–90°. Samples were deposited and smeared into the sample tray such that the sample was flush with the sample tray height. The X-ray diffraction system was set to a horizontal gonio type and data were processed using XRD-6100/7000 software (XRD-series system, version 7.0.1.1, 2014, Shimadzu, Rydalmere, Australia).
2.7. X-ray CT Imaging Analysis
X-ray computed tomography (CT) was conducted using a Zeiss Xradia 520 Versa and image files were post-processed using Dragonfly Pro software (version 2021.1.0.997, 2021, Object Research Systems (ORS) Inc., Montreal, QC, Canada) [
38].
2.8. Cryo-TEM Imaging Analysis
Cryogenic transmission electron microscopy (Cryo-TEM) was used to analyse ~400 mg of a blank SMEDDS dispersed in 20 mL of deionised water. A total of 3.5 µL of the dispersion was spotted onto a TEM grid (EM Lacey Formvar Cu 300 mesh, 50 µm) followed by vitrification on a FEI Vitrobot Mark IV (blot force = -4 and blot time = 3). Images were collected on a 120 keV FEI Tecnai Spirit G2 TEM at a low-dose mode. Dispersion samples for cryo-TEM analysis were prepared and imaged on the same day.
4. Discussion
The developments in three-dimensional (3DP) printing within the realm of pharmaceutics to generate orally administered dosage forms has, to date, been dominated by polymer-based formulations. Lipid-based systems have been proven to boost the bioavailability of lipophilic drugs, yet research into the use of lipid-based formulations in 3D printing is in its infancy. Vithani et al. developed a proof of concept for a semi-solid lipid-based formulation with self-micro emulsifying properties that could be 3D-printed via soft material extrusion [
27]. However, the printing fidelity was not sufficiently high to translate to viable printed medicines and this system required further optimisation to better control its dispersion kinetics under gastric and intestinal conditions.
To address this shortcoming, the potential of 3D-printed polymer scaffolds containing the lipid-based SMEDDS formulation to control print fidelity and exposed surface area was explored. Different polymer scaffolds (
Table 1,
Table 2,
Table 3) were printed using a fused deposition modelling 3D printer with PVOH and PLA as separate extrudates, using a process that resembles previous reports of building PVOH capsules from 3DP core shell design [
39]. This approach allowed much higher fidelity dosage units without the ‘slumping’ apparent when printing SMEDDS systems without a scaffold [
27].
The single-compartment scaffolds were filled with FEN-loaded SMEDDS to generate a range of 3D-printed PLH-based tablets with increasing SA:V ratios by modifying the exposed surface area (
Figure 3 and
Figure 4). For this study, the SMEDDS was manually filled into the scaffolds, which while not practical in a manufacturing context enabled the proof-of-concept study and determination of the relationships between exposed surface area, volume, and polymer dissolution to be investigated. There are reports in the literature of combined FDM and semi-solid extrusion printers with the two different print heads in the printer [
2,
19]. Hence, it would be possible to manufacture the PLH tablets in one operation to achieve this capability.
The dispersion rate of the solid SMEDDS into gastric media was directly proportional to the SA:V ratio when considering the amount of formulation dispersed at 20 min for each tablet (
Figure 12A). Interestingly, there was no difference between the dispersion rates for the PLA and PVOH scaffolds, indicating that dissolution of the PVOH scaffold occurred more slowly than dispersion of the formulation. Note that the closed PVOH scaffold had not released formulation at the 20 min mark but was clearly differentiated from the PLA scaffold at later timepoints (
Figure 10).
Fenofibrate is a poorly water-soluble drug, hence release of drug into the dissolution media would be expected to follow the kinetics of dispersion. The relationship between SA:V and drug release kinetics from solid dosage forms has been studied for 3D-printed polymer-based tablets [
17,
40]. In this study, a general increase in drug release as the amount of formulation dispersed increased was observed at 20 min (
Figure 12B). However, the amount of drug released was slightly lower than what would be expected for a 1:1 correlation with formulation dispersion.
Scalability of printed dose forms to suit the patient is a major advantage of 3D-printed dose forms. Of course, the large scaffolds printed here would be too large for a patient—this was merely an arbitrary size selected for easy manual handling when assembling dispersion and dissolution experiments.
The additive manufacturing approach offers an interesting opportunity to provide tailored release behaviour of multiple drugs from the one dose unit. The concept of a polypill was first coined in 2003 by Wald et al. as a combined therapy daily pill to treat cardiovascular disease [
29] and has been used to describe orally administered dose forms containing multiple active ingredients. The use of polypill approaches has been shown to increase adherence rates, improve cost effectiveness of therapies and increase the safety of medicines [
41]. There have been reports of 3DP polypills with defined release profiles [
15,
32], and 3DP dual-compartment lipid-based tablets with sequential release kinetics of two drugs [
36]. However, to our knowledge there has not been any reports of 3DP solid lipid-based polypills. Pereira et al. published a novel multicompartment 3D-printed system that shows promise for testing and delivery of fillable formulations with defined release rates [
34].
Conceptually, printing one drug as a single compartment is as challenging as printing multiple drugs when they are printed as multiple isolated compartments. It is believed, with the advent of 3D-printed products, these challenges are perhaps less of an issue than might initially be considered compared to combination dose forms where the drugs are mixed together.
The release behaviour from the polypills was different to the single-compartment systems—release of drug was approximately linear with time. This is suggestive that the formulation is gradually dispersed into the media based on the available surface area of the formulation, which is essentially dictated by the wedge shape exposed to the dispersion media. Steady erosion of the formulation from within the wedge-shaped holes then is expected to be faster than the erosion of the PVOH scaffold. This was confirmed by retrieval of multicompartment systems at different timepoints as seen in the photos in
Figure 13. SMEDDS A (composed of Gelucire 48/16 combined with other surfactants) and B (only composed of Gelucire 48/16) were used in the study, but it was not possible to discriminate release from either SMEDDS because total amount of each drug simultaneously from the two formulations was quantified together. Thus, it is possible that the ‘linear’ profiles result from a fast and slow contribution from the different SMEDDS, however the erosion evident in the photos in
Figure 13 would suggest that SMEDDS A and B dispersed at a similar rate.
X-ray CT analysis of the void area, or the gap, between the PLA scaffold and SMEDDS formulation increased with height through the tablet. This suggests that these two materials are not miscible. In addition, the increase in void area with height may be attributed to shrinking of the SMEDDS as it cools within the scaffold. Furthermore, the porosity of the PLA scaffolding is not uniform, and a particularly large difference in porosity may be observed between the foot and walls of the scaffolding (
Figure 6A). It is unclear whether porosity has influenced the dispersion and subsequent release of the SMEDDS from PLA formulations; however, it is believed to have a negligible effect considering the rapid dispersion time and hydrophobicity of PLA. This demonstrates the need for more research into miscible 3DP PLA-SMEDDS core shell designs, and to optimising the printing resolution to eliminate porosity, that may be achievable with PLH-type filament and continuous printing methods [
42].
The model drugs for the multicompartment combination study were selected based on lipophilicity, not direct application of the combination. Clofazimine, halofantrine and lumefantrine each possess high logP values of 5.2, 7.6, 8.6 and 8.6, respectively (Pubchem). In a clinical setting, all three are anti-infective drugs, with clofazimine being used for leprosy [
43], and lumefantrine and halofantrine to treat malaria [
44,
45] Multidrug combinations for anti-infective treatments are common and the demonstration of incorporating relevant combinations of drugs in the one dose form, potentially in different formulations to control dispersion and release opens new opportunities for convenient treatment of such conditions in clinically challenging settings.
The drugs appear to be partly crystalline in the solid SMEDDS at room temperature, and the X-ray diffraction data at least for fenofibrate support this assertion. Our experience with these compounds is that digestion of lipid formulations favours drug solubilisation [
46,
47]. Here, the focus was on control of dispersion of the formulation and drug and less on its solid state upon dispersion, so the effects of digestion were not studied specifically in this study. There is, however, sufficient literature evidence that digestion will stimulate solubilisation of already crystalline drug, not precipitation as might be expected for a liquid SMEDDS formulation where the drug is fully dissolved prior to dispersion [
48].
The translation of this concept into a product will require the development of processes at sufficient speed so that it becomes viable to print such systems. It should be noted that 3D printers have been reported that have an FDM and semisolid extrusion nozzle in the one printer [
49], meaning it is technically possible to prepare these types of combination dose forms without manually filling the scaffold. Furthermore, it is difficult to predict the output speed of a dosage form based on the complexity of its design alone, as the speed and resolution of deposited materials is highly dependent on the material itself and the printing method used. Therefore, the current application of such a formulation is more likely to serve a niche requirement for bespoke dose forms generated in a hospital or doctor’s surgery, rather than in large scale additive manufacturing. Nevertheless, it is possible that the continual development of 3D printing technology will realise larger scale additive manufacturing of bespoke drug forms over time.