University of Groningen Biodegradable implants for the biphasic pulsatile delivery of antigens Beugeling, Max

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University of Groningen

Biodegradable implants for the biphasic pulsatile delivery of antigens

Beugeling, Max

DOI:

10.33612/diss.134204081

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Publication date: 2020

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Beugeling, M. (2020). Biodegradable implants for the biphasic pulsatile delivery of antigens: Toward single-injection vaccines. University of Groningen. https://doi.org/10.33612/diss.134204081

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General discussion & Directions for future research

Max Beugeling

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GENERAL DISCUSSION

Currently, most vaccines require the administration of multiple doses in order to optimally protect the vaccinee. Hence, most vaccines are administered by using a multi-injection boost regime. For several reasons, this multi-injection prime-boost regime contributes to worldwide under-vaccination. The development of a biodegradable device that optimally protects the vaccinee after receiving a single shot, also known as a single-injection vaccine, holds great potential to decrease worldwide under-vaccination. A majority of the work focusing on the development of such a single-injection vaccine is based on poly(DL-lactic-co-glycolic acid) (PLGA) systems that exhibit continuous antigen release [1,2]. However, such a continuous release profile does not mimic the currently applied prime-boost regimes and presents several potential disadvantages. First, it has been suggested that a continuous release of antigen may induce immune tolerance [2–5]. Second, a single-injection vaccine with continuous release kinetics may not be ideal from a regulatory approval standpoint as these kinetics do not mimic the currently approved prime-boost regimes [2]. To overcome these issues and to mimic the currently applied prime-boost regimes, a biodegradable implant with a biphasic pulsatile release could be developed. Therefore, the overall aim of the research described in this thesis was to achieve such a biphasic pulsatile release of antigens from biodegradable implants. To achieve this aim, prototypes based on two different concepts were investigated in this thesis. The first prototype was based on the direct compaction of a physical mixture, while the second prototype was based on a core-shell construct. The following sections present a general discussion of the obtained results.

The physical mixture prototype

Direct compaction of a physical mixture of PLGA with a lactic:glycolic acid ratio of 50:50 (PLGA 50:50) or poly(DL-lactic acid) (PLA), both with an intrinsic viscosity of 0.2 dl/g, and theophylline resulted in a monolithic compact that exhibited the desired biphasic pulsatile release profile in vitro at 37 °C [6]. Our investigations showed that this specific release profile could be explained by the following mechanism. After compaction, a porous compact was obtained, because the applied polymers had a glass transition temperature above room temperature. Once immersed into the release medium, a part of the incorporated theophylline was instantly released from this porous compact by diffusion through the pores. However, after immersion into the release medium, the glass transition temperature of the applied polymers decreased to a value below 37°C. As a consequence, the pores gradually closed by viscous flow of the polymer and further theophylline release was inhibited. After a period of time, however, the polymeric matrix ruptured, possibly due to a build-up in osmotic pressure within the

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compact caused by the accumulation of dissolved polymeric degradation products. This rupture of the polymeric matrix resulted in a pulsatile release of the remaining amount of theophylline [6]. The lag time prior to the delayed pulsatile release of theophylline could be increased from 18 days for PLGA 50:50-based compacts to 50 days for PLA-based compacts. This increased lag time can be ascribed to the fact that PLA is more hydrophobic than PLGA 50:50, leading to a slower degradation rate. Thus, the lag time can be tailored by changing the lactic:glycolic acid ratio of the polymer. The ability to tailor the lag time is of great relevance, as different antigens have different time periods between administration of the primer and the booster dose of the vaccine [7,8]. Subsequently, this concept was evaluated for its suitability for the biphasic pulsatile delivery of bacterial polysaccharide-based antigens. To this end, blue dextran (BD), as a model polysaccharide for polysaccharide-based antigens with either a molecular weight of 70 or 2000 kDa, was incorporated into a physically mixed PLGA 50:50-based compact. The in vitro release study with both, 70 and 2000 kDa BD-containing compacts showed biphasic pulsatile release profiles similar to that of theophylline-containing compacts, which indicates that the molecular weight of the incorporated polysaccharide-based antigen does not influence the release. Collectively, these results showed that the physical mixture prototype may be suitable for the biphasic pulsatile delivery of bacterial polysaccharide-based antigens [6]. However, this should further be investigated by incorporating a clinically relevant bacterial polysaccharide-based antigen into the prototype and confirming its biphasic pulsatile release. In addition, the intactness of the released polysaccharide-based antigen should be investigated. As the current prototype is too big (approximately 6 x 2 mm) for clinical application, miniaturization is required. Miniaturization may be achieved by using production techniques such as hot-melt extrusion. Unfortunately, the prototype was found to be unsuitable for the biphasic pulsatile delivery of protein-based antigens, which are mainly used as viral vaccines. Incorporation of ovalbumin (OVA) as a model protein for protein-based antigens into the prototype resulted in a compact that released approximately 60% of the total dose (within 1 day) as a primer, similar to theophylline- and BD-containing compacts. However, after the release of this primer dose, no delayed OVA release was observed (data not shown in the thesis). We hypothesized that one of the reasons for the absence of this delayed release phase for OVA-containing compacts was the intimate contact of the protein with the polymer. Therefore, we investigated the suitability of a prototype based on a different concept to obtain a delayed release with protein-based antigens.

The core-shell prototype

The most challenging part of the biphasic pulsatile release of protein-based antigens is obtaining a delayed release after a certain, preferably adjustable, lag time. The

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desired delayed release may be achieved by using a core-shell construct that consists of a nonporous PLGA shell around a solid-state antigen-containing core. By acting as a barrier, the nonporous PLGA shell prevents continuous antigen release from the core. After a certain lag time, however, the polymeric shell is degraded to such an extent that it cannot act as a barrier anymore, resulting in a delayed antigen release. As a consequence of the nonporous PLGA shell, the system is only able to provide for the booster dose. To include the primer dose, the outer surface of the core-shell implant could be provided with a fast dissolving coating that contains the protein-based antigens. However, we first focused on obtaining a delayed release of protein-based antigens from a core-shell construct. We hypothesized that such a core-shell construct has two main advantages over the previously described physical mixture prototype that may aid in obtaining a delayed release with protein-based antigens. First, the protein-based antigens are not monolithically dispersed over the PLGA matrix, but physically separated. Second, because the PLGA shell of this core-shell implant is already nonporous before incubation, the system may have less initial water uptake.

To investigate whether a delayed release of protein-based antigens can indeed be obtained from a core-shell construct, relatively large (approximately 9 x 5 mm) core-shell compacts were evaluated as a proof of concept. These core-shell compacts consisted of an OVA-containing core surrounded by a nonporous PLGA shell that was produced by compacting the polymer at a temperature above the glass transition temperature. To investigate whether the lag time prior to the delayed release can be tailored by changing the lactic:glycolic acid ratio of the polymeric shell, PLGA with a lactic:glycolic acid ratio of either 50:50, 75:25, or 100:0 (= PLA) was applied as shell material [9]. In vitro release studies with these core-shell compacts showed that a delayed release of OVA could indeed be obtained. The lag time prior to the delayed OVA release could be increased from 3 weeks (PLGA 50:50) to 6 weeks (PLGA 100:0) by increasing the lactic:glycolic acid ratio of the shell material [9]. As previously described, this can be explained by the fact that an increase of lactic acid content in PLGA results in a more hydrophobic and thus slower degrading polymer [9]. To investigate the stability of OVA in the PLGA 50:50 and PLGA 100:0 core-shell compacts during the lag time, OVA was extracted just before the in vitro release was observed. Fluorescence spectroscopy showed minimal differences between unprocessed OVA and extracted OVA. These results indicate minimal conformational changes of the OVA that could be extracted from the core-shell compacts. However, not all OVA could be extracted. In addition, the in vitro release study showed an incomplete OVA release. Possible reasons for incomplete extraction and incomplete release from the core-shell compacts will be discussed below. BALB/c mice that were immunized with a subcutaneously (s.c.) inserted OVA-containing core-shell compact with either a PLGA 50:50 or a PLGA 100:0 shell showed a delayed OVA-specific IgG1 antibody response [9]. In addition, these mice showed higher OVA-specific IgG1 antibody titers than mice

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that were s.c. immunized with OVA dissolved in phosphate-buffered saline, indicating a potential adjuvant effect of PLGA on OVA-specific IgG1 antibody titers [9]. Such an adjuvant effect of PLGA on OVA-specific IgG1 antibody titers has also been described by others for OVA-containing PLGA-based microparticles [10]. Similar to the findings of the in vitro release study, the lag time prior to the delayed OVA-specific IgG1 antibody response could be increased from 4 weeks (PLGA 50:50) to 6-8 weeks (PLGA 100:0) by increasing the lactic acid content of the polymer [9]. These results indicate a clear in vitro - in vivo correlation of the OVA release from the core-shell compact. In conclusion, similar to the physical mixture prototype, the lag time can be tailored with the core-shell prototype. In addition, if the vaccine requires multiple booster doses, suitable polymers as shell material could be combined, providing multiple small devices for short and long lag times. As it was possible to obtain a delayed release of OVA in vitro and in vivo from the core-shell compact, we investigated whether similar results could be obtained with a clinically relevant protein-based antigen.

The core-shell concept described above may be interesting for the incorporation of protein-based antigens of various infectious diseases, including respiratory syncytial virus (RSV), which causes high morbidity and mortality rates among infants, young children, and the elderly worldwide [11,12]. Unfortunately, a safe and effective vaccine against RSV is still unavailable [11,12]. However, the pre-fusion (pre-F) protein of the virus has recently shown great potential as protein-based antigen [11,12]. Therefore, we developed a pre-F-containing freeze-dried powder and incorporated this powder into the core-shell compact [12]. Pre-F was freeze-dried in the presence of excipients that previously have shown to act as protein stabilizers during drying and subsequent storage, i.e. inulin, HEPES, sodium chloride, and Tween 80 [12–17]. Several analytical techniques indicated that the integrity of pre-F was fully maintained during freeze-drying and storage for at least 28 days at 37 °C and 60 °C [12]. Incorporation of the pre-F-containing powder into a core-shell compact with a PLGA 50:50 shell was feasible without affecting the integrity of pre-F [12]. Similar to OVA-containing core-shell compacts [9], the in vitro release study with pre-F-containing core-core-shell compacts showed a delayed but incomplete release of pre-F [12]. Nevertheless, BALB/c mice that were immunized with a s.c. inserted pre-F-containing core-shell compact showed a delayed RSV virus-neutralizing antibody (VNA) response, indicating that a delayed release of pre-F could be obtained in vitro and in vivo from the core-shell compact [12]. Moreover, pre-F-containing core-shell compacts were able to boost RSV VNA titers of primed mice, demonstrating the feasibility of the biphasic pulsatile release [12]. Unlike with OVA-containing core-shell compacts, PLGA did not seem to have an adjuvant effect on the RSV VNA response [12], indicating that the adjuvant effect of PLGA may depend on the incorporated protein-based antigen. In order to increase the feasibility of the use of a biodegradable implant with a biphasic pulsatile release profile in vaccination practice, we subsequently focused on miniaturization of the core-shell implant and on

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adding the primer dose onto the outer surface of the implant.

To this end, we investigated whether miniaturization could be achieved by using injectable nonporous PLGA 50:50 micro-tubes. In addition, we investigated whether a coating containing the protein-based antigens could be applied onto the outer surface of these filled PLGA micro-tubes. OVA was used as a model protein for protein-based antigens. An in vitro release study with the coated OVA-filled PLGA micro-tubes showed a rapid and total OVA release of the primer dose directly after incubation, indicating that a primer dose of OVA could successfully be added onto the outer surface of the implant by coating the outer surface of the OVA-filled PLGA micro-tubes. The rapid release of the primer OVA dose was followed by a lag time during which no OVA was released. After this lag time, the booster OVA dose was released from the PLGA micro-tube but again incompletely. Nevertheless, with the developed PLGA micro-micro-tubes an in vitro biphasic pulsatile release could be achieved. To demonstrate the biphasic pulsatile release in vivo, the device should be evaluated in an appropriate animal model.

Based on its outer diameter, the coated OVA-filled PLGA micro-tube would fit into an 18G needle (inner diameter of 838 µm), which is suitable for injection [12]. However, in our studies, we closed the micro-tubes provisionally with a plug that increased the outer diameter of the micro-tube. To overcome this issue, a biodegradable plug that fits into the PLGA tube could be developed using techniques such as micro-molding. In addition, a potential clinically acceptable PLGA micro-tube for the biphasic pulsatile delivery of antigens should be combined with a suitable administration device. To achieve consistent and reliable s.c. delivery of the implant, the implant could potentially be incorporated into a ready-to-use syringe similar to Zoladex® [18], which is a Food and Drug Administration-approved injectable PLGA-based implant for the slow release of the decapeptide goserelin acetate.

The incomplete protein release from the prototypes

The incomplete in vitro protein release that was observed with all the prototypes described in this thesis has also been described with many other protein-containing PLGA-based systems [19–23]. It is an extensively described issue [19–23] that is yet to be resolved unambiguously. Specific processes that may lead to an incomplete protein release from PLGA-based systems are numerous and complex [24]. Processes described in literature include: covalent and non-covalent aggregation [23,24], non-specific adsorption to the (degrading) PLGA matrix [23,24], and hydrolysis [23]. The main stress factors inducing these processes after administration are thought to be: hydration [2,25], exposure to an elevated temperature (37 °C) for weeks or months [2,25], the formation of an acidic microclimate within the PLGA matrix (due to the formation of degradation products) [2,23,25,26], and protein-PLGA interactions [23,25,26]. The following section elaborates on the incomplete protein release that was observed with

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the prototypes described in this thesis.

The absence of a delayed release phase for OVA-containing physical mixture compacts may be explained by the intimate contact of the protein with the polymer in combination with the release mechanism of the prototype, which was described elaborately above. The porous structure of the compact will allow for water penetration immediately after incubation. After some time, the pores at the surface are closed by viscous flow and water will be entrapped in the compact. This entrapped water will initiate internal PLGA degradation, as PLGA degrades via hydrolysis of its ester linkages [27,28]. The resulting degradation products, lactic and glycolic acid [27,28], will accumulate inside the matrix, as they are unable to diffuse through the closed polymer layer of the compact. Eventually, this accumulation may have created a highly acidic microclimate [29,30] inside the matrix of the compact. Previous studies with PLGA-based microspheres have indeed shown that the pH within the microspheres can drop as low as pH 1.5 upon PLGA degradation [29]. Such an acidic microclimate is known to degrade the incorporated protein [23,30] and may explain why no delayed release phase was observed from OVA-containing compacts with the prototype.

As demonstrated by the results described above, obtaining a delayed release is the most challenging part of the biphasic pulsatile delivery of protein-based antigens. We hypothesized that such a delayed release may be achieved by using a core-shell construct that consists of a nonporous PLGA shell around a solid-state antigen-containing core. With this core-shell construct, the antigens are physically separated from the polymer, which may reduce the polymer-protein contact-related stress factors described above. Indeed, the evaluated OVA-containing and pre-F-containing core-shell compacts showed the desired delayed release profile (in contrast to the physical mixture compacts). However, the release was incomplete. To investigate the in vitro water penetration into the core-shell compacts, anhydrous copper(II) sulfate, which is a white powder, was incorporated into the core. Upon hydration, anhydrous copper(II) sulfate turns into the blue pentahydrate form. The results of the in vitro study with these anhydrous copper(II) sulfate-containing core-shell compacts indicated that water penetrated into the core after approximately 2 weeks of incubation (data not shown in the thesis). In conclusion, although the desired delayed release profile was achieved with the core-shell compacts, physically separating the antigens from the polymer did only reduce, but not fully prevent the described stress factors and only a part of the protein was released from the core.

As was expected, an incomplete delayed OVA release was also observed with the micro-tubes. However, the amount that was released from the micro-tubes was approximately half of the amount that was released from the core-shell compacts. This can possibly be explained by the fact that water penetrated into the core of the micro-tubes much more rapidly (already after 5 h of incubation) than into the core of the compacts (data not shown in this thesis). This much more rapid water penetration to

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the core of the micro-tubes can be ascribed to differences between the two prototypes. First, the PLGA shell thickness of the micro-tube was approximately 100 µm, while the PLGA shell of the compact was approximately 500 µm thick (data not shown in thesis). Therefore, penetrated water reached the core inside the micro-tube sooner. Secondly, the water-soluble content relative to the amount of PLGA was higher with the micro-tubes, which may facilitate water penetration. Collectively, these results indicate the complexity related to the incomplete protein release from PLGA-based systems.

In addition, it is noteworthy that the measured in vitro release of OVA from PLGA-based systems does not necessarily reflect the in vitro release of protein in its antigenic conformation. The amount of in vitro released OVA in the studies described in this thesis was determined with a modified micro Lowry protein assay. This assay determines the total protein content in a solution, without detecting any conformational changes of the protein. By contrast, the amount of in vitro released pre-F was determined with an enzyme-linked immunosorbent assay, which only detects pre-F in its antigenic conformation. Based on the results of the in vitro release studies with OVA-containing and pre-F-containing core-shell compacts, the amount of OVA released (approximately 65% of the total dose) seems higher than the amount of pre-F released (approximately 20% of the total dose). However, this comparison is incorrect, as the measured amount of OVA also includes soluble protein that lost its native conformation, while potentially soluble pre-F that had lost its native conformation was not measured. Therefore, multiple analytical techniques should preferably be combined to investigate both, the total amount of released protein and the amount of released protein in its native conformation.

Conclusion

The desired biphasic pulsatile delivery of antigens could successfully be achieved with prototypes based on the two different concepts described in this thesis. However, for a clinically acceptable product, substantial challenges related to the stability of protein-based antigens within the PLGA device after administration still need to be overcome. When this is achieved, the successful development of a clinically acceptable product could save many lives annually, especially in low- and middle-income countries.

DIRECTIONS FOR FUTURE RESEARCH

The desired biphasic pulsatile delivery of polysaccharide-based and protein-based antigens could successfully be achieved with the physical mixture and the core-shell prototype, respectively. However, the greatest challenge, i.e. obtaining a delayed and complete release of protein in the native conformation from the core-shell implant, still remains. Therefore, future research should focus on ways to achieve a complete

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(delayed) release of protein in the native conformation from polymeric matrices. As previously described, the main stress factors leading to protein instability after administration are: hydration [2,25], exposure to an elevated temperature (37 °C) for weeks or months [2,25], the formation of an acidic microclimate within the PLGA matrix (due to degradation products) [2,23,25,26], and protein-PLGA interactions [23,25,26]. Therefore, an increase or even a complete release of protein in the native conformation may be achieved by incorporation of excipients into the implant that counteract these stress factors and/or by replacing PLGA with another biodegradable and biocompatible polymer. Both strategies will be described in the following two sections.

Incorporation of stabilizing excipients

Incorporation of stabilizing excipients into the PLGA-based system that counteract the main stress factors causing protein instability is the most commonly applied strategy to increase the stability of the protein within the polymeric matrix [2]. Several excipients that are known to stabilize proteins during drying, also have shown to be protective in a hydrated, warm environment [2]. Therefore, incorporation of sugars, buffers, surfactants, and salts [13,14] may also increase the release of protein in the native conformation from PLGA-based systems. However, in our hands, incorporation of these excipients into pre-F-containing core-shell compacts still resulted in a low (approximately 20% of the total dose) release of pre-F in the native conformation. To counteract the formation of an acidic microclimate within the PLGA matrix (due to degradation products), some studies incorporated acid neutralizing excipients, such as Mg(OH)₂, into protein-containing PLGA-based implants with good results [31,32]. Because the investigated implants were based on a continuous protein release [31,32], it is unclear whether a similar stabilizing effect can be obtained with the core-shell implant described in this thesis. In our hands, however, incorporation of Mg(OH)₂into the physical mixture prototype did not result in a stabilizing effect (data not shown in this thesis). A recently published study reported that incorporation of shellac, a pH-responsive biocompatible polymer, into OVA-containing PLGA-based implants could increase the release of OVA in the native conformation [19]. The authors of this study hypothesized that shellac, which has a low solubility at low pH, protected OVA against the acidic microclimate [19]. Unfortunately, in our hands, incorporation of shellac into the physical mixture prototype did not have a stabilizing effect (data not shown in this thesis). Although some studies have indeed shown that incorporation of certain excipients led to an increased release of protein in the native conformation from PLGA-based systems, achieving a complete (delayed) release of protein in the native conformation by incorporation of stabilizing excipients remains a great challenge. Therefore, replacing PLGA with another polymer may be a more practical approach to achieve a complete (delayed) release of protein in the native conformation from

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polymeric matrices.

Replacing PLGA with another polymer

Replacing the bulk-degrading PLGA with a surface-eroding polymer could be a promising approach to increase the release of protein in the native conformation from the core-shell implant described in this thesis. Although PLGA has many advantages, its main disadvantage remains the fact that it degrades into lactic and glycolic acid. Because PLGA is a bulk-degrading polymer, these acidic degradation products are generated throughout the implant [2], which results in an acidic microclimate within the implant. To overcome this issue, the bulk-degrading PLGA could potentially be replaced with a eroding polymer, e.g. poly(ortho esters) [33]. Because surface-eroding polymers degrade from the exterior surface [34], acidic degradation products are only generated at the exterior surface of the implant where they can quickly diffuse away and be buffered by the environment [2]. Therefore, surface-eroding polymers are unlikely to have any issues related to an acidic microclimate within the implant [2]. This makes the use of surface-eroding polymers interesting for the core-shell prototype described in this thesis.

In addition, future research could focus on the rational design of novel polymers, such as polymers that prevent water uptake and/or that do not create an acidic microclimate upon degradation [25]. These novel polymers could potentially be applied in the two different prototypes described in this thesis to finally obtain a clinically acceptable biodegradable implant suitable for the biphasic pulsatile delivery of antigens.

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