Microbeads encapsulating GEF in two lipid formulations were prepared using drop-on-demand deposition printing for potential intraperitoneal (IP) administration, the preparation of these microbeads is presented in Fig. 1.

Preparation of microbeads by mixing alginate hydrogel with liposomes encapsulating GEF and 3D printing the formulation

Liposomes encapsulating the drug were prepared using the thin-film method. After hydration, excessive GEF was removed from liposomes using SEC. The main fraction containing the liposomes was used for further processing (Supporting information, Fig. S1). Liposomes were characterized in terms of their EE%. While size and PDI could have been more closely controlled by producing SUVs following F/T cycles and extrusion, we decided to formulate MLVs because of the several advantageous characteristics: due to the hydrophobic nature of the drug, low drug-to-lipid ratios yield high EE% as MLVs offer a higher hydrophobic area to accommodate the lipophilic drug within their lamellas; the large size of the MLVs (> 200 nm) is potentially beneficial considering the entrapment of the particles within the hydrogel; finally, with the hydrogel having an average mesh size ranging from 10 to 100 nm [12], particles with larger diameters should be entrapped tighter within the gel. With the aim of comparing the effect of unsaturated phospholipids on drug release, MLVs were not only formulated with DPPC but also with S80, an essential phospholipid extract from soybean containing 73–79% phosphatidylcholine. The reasons for choosing S80 were dual: its cost-effectiveness and scalability compared to DPPC being it derived from natural phospholipids [41], as well as its fibrosis-resolving features, which may be advantageous, particularly following abdominal surgery [42]. We examined various drug-to-lipid ratios for S80 and DPPC liposomes. Notably, lower drug-to-lipid ratios exhibited improved encapsulation efficiency (EE%) of the hydrophobic agent within the liposomes, as depicted in Fig. 2A. The results further revealed that S80 lipids demonstrated superior drug encapsulation efficiency (EE%) compared to other lipids, particularly at lower drug-to-lipid ratios. This suggests that the drug-to-lipid ratio played a more significant role in the encapsulation of GEF compared to the specific choice of lipids. Notably, S80 MLVs achieved a remarkable 90% encapsulation efficiency at a 1:30 ratio, surpassing the ~ 75% efficiency observed with DPPC MLVs. While using smaller drug-to-lipid ratios than 1:30 would likely result in even higher encapsulation of GEF, it is important to consider the scaling potential. To maintain GEF concentration while decreasing the drug-to-lipid ratio, substantial amounts of lipids would be required. This could pose challenges in liposome formation and lead to difficulties in the printing process due to the resulting high viscosity.

Encapsulation and entrapment efficiency. A Encapsulation efficiency (EE%) of GEF in DPPC or S80 MLVs with different drug-to-lipid ratios (mean ± SD, n = 4). B Entrapment efficiency of GEF-loaded beads (mean ± SD, n = 3)

The rheological characteristics of the ink, such as its viscosity at rest and its G’ values (both serve as parameters for assessing the structural strength of the material), are used as criterion to discriminate between printable and unprintable ink. All formulations exhibit shear-thinning behaviour, regardless of liposome presence (Fig. 3B and Supplementary Information Fig. S6). Moreover, the viscosity at rest increases with either rising alginate concentration or liposome addition (Fig. 3A). Inks with viscosity < 100 mPa proved excessively fluid for drop-on-demand manufacturing, resulting in continuous lines or large, uncontrollable drops instead of microbeads. Conversely, inks with viscosities > 250 mPa were too thick for extrusion at reasonable pressure, causing tear-like structures or fragmented shrapnel upon increased pressure. As for the viscosity, also G’ values (Fig. 3C) increase with rising alginate concentration and a printability window ranges from 0.046 to 0.141 Pa. All inks were viscous liquid across the measuring range, as indicated by the dominance of the viscous component (G’’) over the elastic one (G’) and the absence of a flow point (Fig. 3D).

Rheological properties of liposomal ink inks. A Viscosity at rest for inks containing alginate spanning from 1 to 5% and either liposomes (S80 or DPPC; 15 mM) or hydration buffer. B Flow curve of inks used in the study (15 mM S80; 3% Alg or 15 mM DPPC; 3% Alg or 3% Alg). C G’ for assessed inks as for viscosity at rest. D Loss/storage modulus of DP ink used for in vitro study (15 mM S80; 3% Alg). Rectangle highlights values used for determining G’

To address the limitations of traditional 3D-printing methods for spherical structures, an EMD printhead was employed for the synthesis of microbeads. This printhead (illustrated in Supplementary information Fig. S5) utilizes an electromagnetically controlled valve to generate droplets ranging from micro- to nanolitres. The pressurized DP ink is jetted directly into a crosslinker, with the dispensing volume determined by parameters like nozzle diameter, valve speed, and actuation. Adjusting variables such as pressure, cycle time, and printhead height can influence microbead characteristics during printing. Highly viscous inks might require careful adjustment of process parameters to ensure reproducibility of the microbead printing.

A DoE strategy was used to screen either the range of formulation or process parameters that have been demonstrated to provide a printable drug product ink and a robust printing process resulting in microbeads (Supplementary information Table S1).

A main effects plot, as seen in Fig. 4, visually represents the average impact of independent variables, such as e.g. the alginate levels 1–5%, on the dependent variable (printability) by illustrating the factor level’s response (Supplementary information Sect. 1.1). In this context, it serves to identify how independent process parameters influence the printability of the DP ink. Thereby, the plot simplifies the relationship between independent process parameters and printability, offering a visual representation of the influence of certain process parameters on the overall printing process. The liposome-to-alginate ratio emerges as a critical factor in optimizing the printability of the DP ink. This ratio is defined as the lipid concentration divided by the alginate concentration within a specific DP ink. Throughout the study, the lipid concentration remained constant at 15 mM to ensure the delivery of the desired dose, while the alginate concentration varied from 1 to 5% (w/v), resulting in liposome-to-alginate ratios ranging from 8.8 for 1% (w/v) alginate to 39.0 for 5% (w/v). For clarity, the main effects plot displays the alginate concentrations.

Main effects plot for printability in a multifactorial experimental design (120 runs). This plot illustrates the impact of the considered factors on the printability response. Higher response values denote a more desirable printability for each factor level. The slope of the lines connecting the factor levels serves as an indicator of the significance of their influence on the printability response

After conducting the DoE with different alginate concentrations, a final concentration of 3% (w/v) was identified as the most suitable for 3D printing when combined with drug-loaded liposomes. This determination is supported by the local maximum for printability (0.29) of the alginate factor in Fig. 4. The liposome-to-alginate ratio of 13.0 (15 mM lipid; 3% alginate (w/v)), exhibited favourable printability in terms of shape retention and stability. Subsequent rheological studies revealed that a reduction in lipid content resulted in a decrease in the viscosity of the DP ink, preventing the consistent formation of droplets (Fig. 3).

For the crosslinking process, the use of CaCl2 was found to be most effective at 135 mM. Higher concentrations were considered favourable as they resulted in stronger crosslinking, potentially leading to a more sustained drug release profile [27]. At lower concentrations, the microbeads formed disk shapes. However, as indicated in Fig. 4, the line between the considered levels for crosslinker has a relatively low slope which indicates a smaller impact on the printability compared to other factors. In contrast, the steep slope for pressure indicates that lower pressure values are beneficial for printability. Maintaining the pressure within the range of 25–60 kPa proved to be advantageous, as excessive pressure led to the formation of tears in the printed structures. In terms of the EMD printhead, the cycle time was fine-tuned to an opening duration of less than 2 ms and a closing duration exceeding 200 ms. Extending the opening duration of the valve resulted in a hanging drop of DP ink on the printhead due to the continuous flow through the nozzle. The identified cycle time (1/200) ensured precise droplet formation and deposition during the printing process and are in accordance with the finding reported by Lu et al. [31]. Additionally, the height of the EMD print head should not exceed 10 mm from the surface of the crosslinker to avoid the formation of skewed spheres with increasing printhead height. The process parameters identified to be most suitable for drop-on-demand printing were derived from the DoE and are outlined in Table 1.

The entrapment efficiency of GEF-loaded liposomes in alginate beads is close to 100% for both, S80 and DPPC, as displayed in Fig. 2B. This indicates that only a neglectable amount of GEF was lost during the production process. As expected, the entrapment efficiency of the free drug (FD) in alginate (57.6%) is lower than in liposomal carriers (96.3% DPPC; 103.5% S80), confirming the importance of a suitable carrier to enable the delivery of the hydrophobic agent. Liposomes characterized and used on cells were always prepared freshly; nevertheless, the stability of the lipids, the drug, and its encapsulation over a period of 7 days was investigated. Regarding lipid stability for DPPC and S80, no degradation was identified over the time of observation (Supporting information, Fig. S2). Additionally, no drug degradation nor loss in GEF encapsulation was detected (Supporting information Fig. S2).

Microbeads were characterized for their size, shape, and appearance. Microscopic measurements of the microbead’s diameter revealed the size distribution as visualized in Fig. 5 and Table 2. The distribution suggests homogeneous distribution during the production, promising aspects for future scale-up of the process. The robust process of the EMD printhead contributes significantly to the improved consistency of shape and size observed in microbeads produced through drop-on-demand manufacturing, as compared to conventional extrusion methods. Nevertheless, likely due to the lipids’ influence on the rheological properties of the ink, the presence of different lipids had an influence on the printability and, hence, on the size of the microbeads. Further, the surface-to-volume ratio (SVR) provides a measure of surface per amount of volume. In systems with a large SVR, like S80 microbeads, a higher surface area implies the potential for faster drug release. Conversely, systems with lower SVR, such as DPPC microbeads, have the ability to sustain drug release for longer periods. This distinction arises from the exposed surface area and the distance the drug must traverse through diffusion before reaching the release media.

Violin plot depicting the distribution of microbead diameters across three non-consecutive batches per formulation (median ± quartile, n = 60). Considering intra- and inter-batch variation, this representation serves as an indicator of the robustness in the production of liposome-laden microbeads. While the liposome-laden microbeads manifest a similar distribution, the microbeads containing FD result in a larger process variation

The morphological difference of the prepared formulation is observed in Fig. 6. It is apparent that S80 (A) and DPPC (C) microbeads have a spherical shape while the shape of the FD (D) microbeads has minor misshaped areas. This is most likely due to the difference in viscosity of the ink during the printing process, caused by the absence of liposomes. Also, for a minor amount of microbeads, air bubbles were entrapped into the beads as seen in Fig. 6C. All irregularities on the surface of the microbeads influence the SVR and, thereby, potentially the release of the drug since a different amount of surface is exposed to the release media. The EMD printhead facilitated the production of beads within the upper micrometres range, specifically around 700 µm in size.

Images of microbeads after 3D printing. A S80 microbeads (15 mM S80; 3% Alg; 0.5 mM GEF). B S80 microbeads (15 mM S80; 3% Alg 0.5 mM GEF) after injection through an 18G needle. C DPPC microbeads (15 mM DPPC; 3% Alg; 0.5 mM GEF). D FD microbeads (3% Alg; 0.5 mM GEF). SEM images (E, F, G) of dried microbead (15 mM S80; 3% Alg; 0.5 mM GEF). E Surface image of microbeads at a macroscale. F Microscale of surface area shows homogeneous distribution of pores over surface area. G Pores with a diameter of ~ 100 nm on the surface indicated by rectangles

Microbeads were further assessed visually for injectability through an 18G needle within which no differences or irregularities were observed indicating that the system is suitable for IP administration (Fig. 6B). Beyond the advantage of injectability, this system offers the potential for alternative administration methods. Specifically, the microbeads could be administered through a peritoneal port, a technique frequently employed in corresponding surgeries [43]. For such routes of administration, the shear stress experienced by the microbeads would be drastically reduced compared to an 18G needle (potential shear stress reduction > 60-fold). As a result, the shear stress might be estimated as it can be considered negligible. Patient-centric peritoneal delivery of large systems should be considered and, in their sum, can play a vital role in ensuring treatment success.

SEM images of the microbeads as shown in Fig. 6 provide details of the microbeads’ surface. While the beads show minor irregularities at the macroscale (E) which might be caused during printing, the surface at a microlevel seems homogeneous (F). Zooming in on the surface area, the beads’ pores were identified with a diameter of approximately 100 nm (G) which is in line with previous reporting [44]. It is assumed that said pores are small enough to allow surrounding release media to enter the beads, further increasing the exposed surface area of the bead. But at the same time, pores of 100 nm are likely capable to entrap larger MLVs and thereby hinder the lipids to be directly released for the system [45].

A solubility study of GEF in peritoneal simulation fluid (PSF) was performed to determine the most appropriate drug loading in microbeads for the subsequent drug release study in biorelevant conditions (vide infra).

Drug release from liposomes often depends on having the right environment to dissolve the drug in their surroundings [12, 46]. To gauge the solubility of GEF, we examined it in various media. For instance, we added 20% DMSO to PSF and assessed GEF solubility at 37 °C, utilizing both kinetic and thermodynamic approaches (see Table 3). PSF serves as a mimic for peritoneal fluid in individuals with peritoneal disease, covering pH, buffer capacity, glucose, MgCl2, and CaCl2 content [19]. Interestingly, we found that the best solubilities were achieved by introducing an excessive amount of GEF relative to the PSF from the delivery system, rather than relying on excessive free GEF powder dissolution. Therefore, liposomes improve the solubility of the hydrophobic agent GEF. This improved solubility can potentially lead to enhanced therapeutic effectiveness [47].

To study the microbeads’ capability to retain the loaded liposomes, hindering thus a fast peritoneal clearance, we measured the release of MLV from the system. DiD, a dye with similar lipophilic characteristics as the evaluated drug, was encapsulated in liposomes and the fluorescently labelled liposomes mixed in the ink with alginate and printed as microbeads. Given the hydrophobic nature of DiD, the fluorescence in the release medium could be considered an indicator of MLV release. Neglectable amounts (< 0.1%) of lipid release for both assessed lipids over a period of 9 days (Supporting information, Fig. S3). Similar results were obtained with microbeads filled with 200 nm SUVs, showing a neglectable liposome release for 6 days (Supporting information, Fig S4). This observation suggests that the CaCl2 in the release medium is capable of preserving the alginate’s integrity and, in turn, securing the liposomes, thus indicating that the surrounding environment plays a role in the liposome release process [48, 49]. This finding provides further evidence to the hypothesis that the 100-nm-sized nanopores on the surface of the microbeads are capable of retaining the MLVs within the alginate beads effectively. The negligible release of liposomes from the microbeads could thus provide the necessary hydrophobic volume to accommodate GEF in the carrier and at the same time ensure a sustained release of the drug in the peritoneal space thanks to large size (> 1 µm) of the microbeads, offering overall a clear advantage with respect to the use of plain GEF-liposomes.

Drug release from microbeads was studied in PSF with 20% DMSO at physiological temperature of 37 °C and horizontal shaking to simulate the movement within the peritoneum. Microbeads entrapping GEF-loaded DPPC and S80 liposomes were tested. As a control, FD-laden microbeads were used. Both liposomes and microbeads were stable in PSF with 20% DMSO over the assessed time as preliminary studies indicated. In absence of a liposomal carrier slowing down the diffusion of GEF, FD microbeads manifested a burst release of the drug within the first 4–5 h, as shown in (Fig. 7).

Cumulative drug release from 3D-printed microbeads (15 mM lipid; 3% Alg; 0.5 mM GEF) in PSF (mean ± SD, n ≥ 3). Total of 100% drug was retrieved with extraction following release experiment

GEF release from liposomes in microbeads was sustained over 3 days reaching overall ~ 100% and ~ 80% drug release for DPPC and S80, respectively. The longer retention of GEF in the S80 system can be attributed to a potential preferential interaction between the drug and the S80 lipid bilayer, which aligns with the observed trend in encapsulation efficiency (EE%). Moreover, the sustained release is a result of the synergistic effect of both the alginate and the liposomes. As the liposomes are entrapped within the alginate mesh, the drug is gradually released from the lipidic environment into the alginate gel.

Surgical and locoregional treatment of peritoneal metastasis, particularly in cases originating from colorectal cancer, has gained widespread acceptance following the publication of favourable patient outcomes by various groups worldwide [10]. However, it is important to acknowledge that the peritoneum can be susceptible to damage resulting from factors such as surgical trauma, infection, or exposure to peritoneal dialysis fluid following these surgical interventions in the peritoneal cavity [50, 51]. In such situations, the attachment of fibroblasts to fibrin and subsequent collagen production can lead to the formation of adhesive fibrotic tissue, necessitating advanced treatment strategies [50, 51]. Even if the liposomes are entrapped in the microbeads, the lipids could potentially manifest their bioactive role once the microbeads undergo degradation, despite the prior release of the drug. Hence, the employment of S80 is notably beneficial, given its documented antifibrotic qualities as shown by our research [42]. Further, the high availability and low cost of S80 make it a desirable candidate for further drug development and scale-up efforts.

To test the suitability of our drug delivery system, we applied different concentrations of GEF to Huh-7 cells, an immortalized human hepatic carcinoma cell line, either as free drug in DMEM (0.5% V/V DMSO used as a vehicle) or encapsulated in MLVs printed as composite alginate microbeads. The Ca2+ (and Mg2+) ions present in DMEM medium were capable of stabilizing sodium alginate microbeads during the study, avoiding fast erosion of alginate and thereby burst release of the liposomes. Cells were treated for 24 h and their viability was measured after removal of the beads. To avoid unwanted toxicity due to contact between beads and cells, inserts were used to physically separate the beads from the cells, while still ensuring GEF to pass through the insert membrane and be absorbed by the Huh-7 cells. The diffusion of GEF through the insert membrane was validated experimentally.

Concentration-dependent Huh-7 cell death was observed with GEF, both when in its free form and when released from the microbeads (Fig. 8). Higher cell death was obtained if the drug was released from the microbeads, which can be explained due to cytotoxicity of the beads themselves, but also due to the enhanced solubility of the drug, facilitated by the delivery system, as previously discussed. The sustained release might be enough to avoid having an excessively high free GEF concentration which would lead to precipitation, whereas albumin, the main component of FBS, probably facilitated GEF transport due to its ability of binding small hydrophobic compounds [52, 53].

Cell viability of different GEF concentration on Huh-7 cells (mean ± SD, n = 3) normalized to DMEM. GEF was either administered as free drug in presence of 0.5% V/V DMSO (dark grey) or encapsulated in MLVs entrapped in microbeads (light grey), where no DMSO had to be added. Cell viability in DMEM and exposed to empty beads (light grey dotted) was measured as control. For experiments in DMEM, the solubility did not allow the testing under sink conditions

Our initial, yet encouraging, in vitro results underscore the efficacy of our formulation to induce cell death without the use of organic solvents, thus suggesting a possible pathway for future in vivo studies involving GEF-loaded microbeads. The use of GEF-loaded microbeads in peritoneal chemotherapy potentially offers several advantages. Firstly, it enables localized and targeted drug delivery, minimizing systemic side effects [14, 54]. The sustained release profile of the liposome-alginate system ensures prolonged exposure of tumour cells to the drug, potentially enhancing efficacy [5]. Additionally, the use of liposomes in peritoneal cancer treatment opens possibilities for a wider range of therapeutic agents and enables to overcome limitations associated with current treatments, such as the use of charged drugs to avoid rapid clearance [5, 54].