Patent Application
20240000719
The invention relates to the field of pharmaceutical compositions comprising proteins as therapeutic agents. More particularly, it is directed to hot melt extrusion-produced antibody-containing filaments, implantable drug delivery devices made from these filaments and to methods of producing such filaments and devices. The hot melt extrusion-produced antibody-containing filaments and the devices obtained from the filaments according to the invention allow the delivery of the antibody over a certain period of time.
The hot melt extrusion (HME) is widely described and implemented in the pharmaceutical field to produce drug-loaded printable filaments (Goyanes et al., 2015; Tiwari et al., 2016). HME is based on the melting of polymeric material that is extruded through a die to obtain a homogeneous drug-loaded filament. HME is a free-solvent process which may be easily scaled-up. However, this technique is based on the use of relatively high temperatures. Such temperatures may be usually reduced by adding a plasticizer, allowing the decrease of the glass transition temperature of the polymer. Another alternative to decrease the temperature of extrusion could be the use of thermoplastic polymers characterized by a low molecular weight (Fredenberg et al., 2011). HME was already investigated to develop protein-based formulations which were characterized by a controlled-release of the loaded active ingredient over time (Cossé et al., 2016; Duque et al., 2018; Ghalanbor et al., 2010).
HME can be used in combination with 3D printing (3DP) process, such as fused deposition modelling (FDM™). FDM process is currently an integrant part of the pharmaceutical field (Jamroz et al, 2018; Azad et al., 2020). This technology is an extrusion-based 3DP method which uses heat to melt a thermoplastic polymer filament to build an object in a layer-wise manner. The use of 3DP allows the production of any kind of shapes starting from a digital design (Norman et al., 2017). The main drawback remains the lack of pharmaceutical grade polymers that are available to be used in FDM, although Poly(lactic acid) (PLA) and polyvinyl alcohol (PVA) are commonly used as thermoplastic polymers to make drug-loaded printable filaments (Jamróz et al., 2018).
Poly(lactide-co-glycolide) (PLGA) is a well-known pharmaceutical grade polymeric material that is usually used to make injectable/implantable sustained-release DDS. PLGA could be extruded at low temperature, making it a good candidate for both HME and FDM processes. Protein-loaded PLGA implants have already been described using macromolecules such as ovalbumin (Duque et al., 2018), bovine serum albumin (Cossé et al., 2016) and lysozyme (Ghalanbor et al., 2010). The major challenge remains the stabilization of the protein during the extrusion.
It was shown that the solid state of the protein could be more advantageous to promote a higher stability as well as to make easier its addition into the polymeric matrix using HME process (Cossé et al., 2016; Mensink et al., 2017). However, the protein compounds usually used as models (i.e. OVA, BSA, lysozyme) to produce protein-loaded implants are characterized by low molecular weights in comparison with immunoglobulins for instance.
Therefore, there is still a need for further filaments and implantable drug delivery devices comprising large proteins, more particularly antibodies, with sustained-release properties, improving stability of antibodies (e.g. limiting antibody degradation during the production of the filament and then of the implantable drug delivery device), while keeping their activity (i.e. without impacting drastically their biological activity).
In a first aspect, the present invention provides a filament for preparing an implantable drug delivery device, wherein the filament comprises or consists of at least one polymeric material, a plasticizer and an active ingredient, wherein said active ingredient is an antibody. The filament may further comprise at least one stabilizer, a buffering agent and/or a surfactant.
In a second aspect, the present invention relates to an implantable drug delivery device comprising or consisting of one or more layers made from a filament comprising or consisting of at least one polymeric material, a plasticizer and an active ingredient, wherein said active ingredient is an antibody. The filament may further comprise at least one stabilizer, a buffering agent and/or a surfactant.
In a third aspect, the present invention describes a 3D printed implantable drug delivery device obtained by 3D printing filaments comprising or consisting of at least one polymeric material, a plasticizer and an active ingredient, wherein said active ingredient is an antibody. The filament may further comprise at least one stabilizer, a buffering agent and/or a surfactant.
In a fourth aspect, the present invention provides a process for producing a filament for preparing an implantable drug delivery device, the process comprising the steps of:
In a fifth aspect, the present invention relates to a process for producing an implantable drug delivery device, the process comprising the steps of:
The term “polymeric material” refers to polymeric components able to support high temperatures during hot melt extrusion (HME) and 3D printing. Therefore, the preferred polymeric materials according to the invention are thermoplastic polymers or thermoresistant polymers. Examples of such thermoplastic polymers that are commonly used for 3D printing are for instance are Polyvinylpyrrolidone (PVP), acrylonitrile butadiene styrene (ABS), the poly(lactic acid) (PLA), Poly(lactic-co-glycolic acid) (PLGA), the polyvinyl alcohol (PVA), poly(ε-caprolactone) (PCL), ethylene vinyl acetate (EVA). Preferably they are biodegradable or bioeliminable for more convenience to the patients. Other thermoresistant polymeric material are for instance hydroxypropyl cellulose (HPC), Hydroxypropyl methylcellulose (HPMC), Poly(Ethylene Glycol) (PEG), Eudragit derivatives (E, RS, RL, EPO), Polyvinyl caprolactam-polyvinyl acetate-polyethylene glycol graft co-polymer (Soluplus®), thermoplastic polyurethane (TPU). Suitable polymeric materials are also herein described.
Based on advantages of Hot Melt Extrusion (HME) and/or Fused Deposition Modelling (FDM) technologies, the inventors have developed antibody-loaded filaments that can then be used to obtain implantable devices, such as via 3D-printing using FDM technology. The present invention is based on the surprising finding that it has been possible to produce filaments comprising an antibody, said filaments having a high antibody load (at 15% and higher). The filaments could then be used to obtain implantable drug delivery devices (obtained by moulding or 3D printing for instance), from which the antibody was released in a control manner over time. Further, not only the antibody was released in a timely manner, but it was still able to bind its target. It was necessary to judiciously select the type of thermoplastic polymer to be used and to optimize the manufacturing parameters of HME, or of both HME and FDM, to obtain first a filament and then an implantable drug delivery device which could allow the stability and the affinity of the loaded antibody to be maintained.
The main object of the present invention is a filament for preparing an implantable drug delivery device, wherein the filament comprises or consists of at least one polymeric material, a plasticizer, and an active ingredient, wherein said active ingredient is an antibody. The filament may further comprise at least one stabilizer, a buffering agent and/or a surfactant. In such a case, and as an example, the filament according to the invention as a whole, can comprise or consist of at least one polymeric material, a plasticizer, an antibody and at least one stabilizer. As a further example, the filament according to the invention may comprise or consist of at least one polymeric material, a plasticizer, an antibody, at least one stabilizer and a buffering agent. The filament can be moulded or used in a 3D printer in order to obtain an implantable drug delivery device of any desired shape.
The invention further provides an implantable drug delivery device comprising or consisting of one or more layer(s) made from a filament comprising or consisting of at least one polymeric material, a plasticizer and an active ingredient, wherein said active ingredient is an antibody and wherein said filament may further comprise at least one stabilizer, a buffering agent and/or a surfactant.
A further object of the present invention is a 3D printed implantable drug delivery device obtained by 3D printing a filament comprising or consisting of at least one polymeric material, a plasticizer and an active ingredient, wherein said active ingredient is an antibody. Said filament can further comprise at least one stabilizer, a buffering agent and/or a surfactant.
Before being added to the polymeric material to form the filament and then the implantable drug delivery device, the active ingredient has to be spray-dried or freeze-dried. To do so, a preliminary liquid formulation is prepared wherein said formulation comprises or consists of the active ingredient, wherein said active ingredient is an antibody. Said liquid formulation may further comprise at least one stabilizer, a buffering agent and/or a surfactant. The liquid formulation is then spray-dried or freeze-dried according to standard methods to obtain dry microparticles. Once in the form of dried microparticles, the active ingredient is homogeneously dispersed into the at least one polymeric matrix and the plasticizer. They form an active ingredient-loaded solid dispersion such as an antibody-loaded solid dispersion.
Therefore, herein also provided is a process for producing a filament according to the invention, the process comprising the steps of:
The filament according to this invention can be used for producing an implantable drug delivery device. Said device can be either cut to a desired length, pelletized, moulded or 3D printed. The advantage of using a 3D printer is to enable the design and manufacture of novel and customized implantable drug delivery device that are not possible using traditional processes. Thanks to 3DP technology, the structure, shape or composition of the device can be customized and adapted to the patient on a case by case basis. Another advantage of using a 3D printer is to provide devices on demand.
3D printing is part of a technology called additive layer manufacturing (ALM). ALM can be based on liquid solidification or on solid material extrusion. Liquid solidification technologies include for instance Drop-on-powder deposition (DoP, or binder jetting), drop-on-drop deposition (DOD), whereas solid material extrusion technologies includes Pressure-assisted micro syringe (PAM) deposition, or yet Fused Filament Fabrication (FFF), also known as Fused Deposition Modelling™ (FDM®) technology. In a DoP or DoD system, two-dimensional layers are repeatedly printed until a three-dimensional object is formed. For example, inkjet or polyjet printing of dosage forms as disclosed herein can use additive manufacturing. The PAM technology involves the deposition of soft material (semi-solid or viscous) through a syringe-based print head. The syringe is typically loaded with the material which is then extruded using pneumatic pressure, plunger or a screw. The FDM technology is based on the extrusion of thermoplastic polymer which is driven by a gear system through a heated nozzle tip. The print head is composed of the pinch roller mechanism, a liquefier block, a nozzle and a gantry system that manages the x-y directions. The filament is fed and melt in the liquefier, turning the solid into a softened state. The solid part of the filament is used as a plunger to push the melt through the nozzle tip (Sadia et al., 2016). Once a layer of thermoplastic melt is deposited, the build platform is lowered, and the process is repeated to build the structure in a layer-wise manner.
Also encompassed by the invention is a process for producing an implantable drug delivery device, and in particular a 3D printed implantable drug delivery device, wherein the process comprises the steps of:
In the context of the invention as a whole, the active ingredient is an antibody. Said antibody can be any antibody as defined in the above definitions section. The antibody is preferably present in the preliminary liquid formulation, before drying, at a concentration of or of about 50 mg/ml to or to about 300 mg/mL, preferably of or of about 65 mg/ml to or to about 250 mg/mL, even preferably of or of about 80 mg/mL to or to about 200 mg/ml such as 80, 85, 90, 95, 100, 105, 110, 115, 120, 125, 130, 135, 140, 145, 150, 155, 160, 165, 170, 175, 180, 185, 190, 195 or 200 mg/mL. Alternatively, the antibody is present in the preliminary liquid formulation, before drying, at a concentration of or of about 5 to or to about 30% w/v, or preferably at a concentration of or of about 6.5 to or to about 25% w/v, even preferably of or of about 8 to about 20% such as 8, 8.5, 9, 9.5, 10, 10.5, 11, 11.5, 12, 12.5, 13, 13.5, 14, 14.5, 15, 15.5, 16, 16.5, 17, 17.5, 18, 18.5, 19, 19.5 or 20% w/v. The antibody loading in the filament, and thus in the final implantable drug delivery device, is preferably in an amount of about 15 to 40% (w/w), or in an amount of about 15 to 35% (w/w), such as 15, 16, 17, 18, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34 or 35% (w/w).
Should at least one stabilizer be used in the context of the present invention as a whole, it is preferably a disaccharide (such as sucrose or trehalose), a cyclic oligosaccharide (such as hydroxypropyl-β-cyclodextrin), a polysaccharide (such as inulin), a polyol (such as sorbitol), or an amino acid (such as L-arginine, L-leucine, L-phenylalanine or L-proline) or any combinations thereof. Should more than one stabilizer be used, the combinations of stabilizers can be for instance (without any limitation) one disaccharide with one amino acid or a polyol with an amino acid. As an example, a combination of two stabilizers can be used, wherein one stabilizer is either sucrose or trehalose and the other stabilizer is L-arginine, L-leucine, L-phenylalanine or L-proline. The at least one stabilizer is preferably present in the preliminary liquid formulation, before drying, at a concentration of or of about 10 mg/mL to or to about 100 mg/mL, preferably of or of about 20 mg/ml to or to about 75 mg/mL, or even preferably of or of about 30 mg/mL to or to about 50 mg/ml such as 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49 and 50 mg/mL. Alternatively, the stabilizer is present in the preliminary liquid formulation, before drying, at a concentration of or of about 1 to or to about 10% w/v, or preferably at a concentration of or of about 2 to or to about 7.5% w/v, or even preferably of or of about 3 to or to about 5% such as 3.0, 3.1, 3.2, 3.3, 3.4, 3.5, 3.6, 3.7, 3.8, 3.9, 4.0, 4.1, 4.2, 4.3, 4.4, 4.5, 4.6, 4.7, 4.8, 4.9, or 5.0% w/v.
According to the present invention in its entirety, when at least one stabilizer is present, the ratio (w/w) antibody: stabilizer(s) (alternatively referred to ratio (w/w) antibody: at least one stabilizer) in the filament, and in the implantable drug delivery device, is preferably between about 1:1 and about 5:1 (weight/weight, i.e. w/w), more preferably between about 1.2:1 and about 4:1, even more preferably between about 1.25:1 to 3:1, such as 1.25:1, 1.5:1, 1.75:1, 2.0:1, 2.25:1 and 2.5:1 (w/w). According to the present invention in its entirety, should a buffering agent be present, said buffering agent can be selected from the group comprising or consisting of (but not limited to) phosphate, acetate, citrate, arginine, trisaminomethane (TRIS), and histidine. Said buffering agent is preferably present in the preliminary liquid formulation, before drying, in an amount of from about 5 mM to about 100 mM of the buffering agent, and even preferably from about 10 mM to about 50 mM, such as about 10, 15, 20, 25, 30, 35, 40, 45 or 50 mM.
In the context of the whole disclosure, a surfactant may also be present. Said surfactant can be for instance (but without being limited to) Polysorbate 20 (PS20) or Polysorbate 80 (PS80). When present, the surfactant is preferably added in the preliminary liquid formulation, i.e. before the drying step. Said surfactant is preferably present in the preliminary liquid formulation, before drying, in an amount of or of about 0.01 to or to about 5 mg/mL, more preferably of or of about 0.01 to or to about 1 mg/mL, more particularly of or of about 0.1 to or to about 0.6 mg/ml, such as 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 0.55 or 0.6 mg/mL. Alternatively, the polysorbate surfactant is preferably present in the preliminary liquid formulation, before drying, in an amount expressed in term of % weight per 100 mL (% w/v). In such a case, the polysorbate surfactant comprised in the formulations according to the present invention as a whole can be present in an amount of 0.001 to 0.5% w/v, preferably from 0.01 to 0.1% w/v, or even preferably from 0.01 to 0.06% w/v such as 0.01, 0.015, 0.02, 0.025, 0.03, 0.035, 0.04, 0.045, 0.05, 0.055 or 0.06% w/v.
In the context of the present invention, and in particular when referring to filaments or final implantable drug delivery devices, the optional at least one stabilizer, buffering agent and surfactant are regrouped under the collective name of excipients. When present, the excipients are preferably present in the filament, and thus in the final implantable drug delivery device, in a total amount of or of about 3 to or to about 20% w/w, preferably in a total amount of or of about 5 to 15% w/w, such as about 5, 5.5, 6, 6.5, 7, 7.5, 8, 8.5, 9, 9.5, 10, 10.5, 11, 11.5, 12, 12.5, 13, 13.5, 14, 14.5 or 15 wt %.
In the context of the invention as a whole, the at least one polymeric material is preferably a biodegradable, biocompatible and/or bioeliminable thermoplastic polymer such as polyurethane (TPU), polyvinylpyrrolidone (PVP), polyvinyl alcohol (PVA), poly(ε-caprolactone) (PCL), poly(lactic acid) (PLA), polydioxanone, polyglycolide, polytrimethylene carbonate, hydroxypropyl cellulose (HPC), Hydroxypropyl methylcellulose (HPMC) or combinations thereof such as, but not limited to, ethylene vinyl acetate (EVA), poly(lactic-co-glycolic acid) (PLGA), poly(L-lactide-co-caprolactone-co-glycolide)(PLGA-PCL). Polymeric materials can have a controlled size of about 200Da to about 50 kDa, preferably about 500 Da to about 40 kDa even preferably about 1 kDa to about 20 kDa, such as about 1, 2, 5, 10, 15 or 20 kDa. Alternatively, instead of having a given size (±), the polymeric materials can be a mix of polymers of different sizes, e.g. 5 kDa to 20 kDa or 7 kDa to 17 kDa. For instance, some commercially available polymers are a mix of polymers of different sizes such as Resomer® RG502 having a mix of polymers ranged between 7 and 17 kDa. Preferably said polymeric material is present in the filament, and thus in the final implantable drug delivery device, in an amount of about 50 to 75% (w/w), or in an amount of about 55 to 70% (w/w), such as 55, 56, 57, 58, 59, 60, 61, 62, 63, 64, 65, 66, 67, 68, 69 or 70%.
In the context of the invention as a whole, the plasticizer is preferably polyethylene glycol (PEG) or a PEG compound such as, but not limited to, maleimido monomethoxy PEG, activated PEG polypropylene glycol, methoxypoly(ethyleneglycol) polymer. PEG compounds according to the invention can also be charged or neutral polymers of the following types: dextran, colominic acids, or other carbohydrate-based polymers, polymers of amino acids, and biotin and other affinity reagent derivatives. PEG or PEG compounds in the context of the invention can be linear or branched. PEG or PEG compounds in the context of the invention can have a size of about 200Da to about 50 kDa, preferably about 500 Da to about 40 kDa even preferably about 1 kDa to about 20 kDa, such as about 1, 2, 5, 10, 15 or 20 kDa. Preferably said plasticizer is present in the filament, and thus in the final implantable drug delivery device, in an amount of about 2-20% (w/w), or preferably in an amount of about 5 to 15% (w/w), such as 5, 6, 7, 8, 9, 10, 11, 12, 13, 14 or 15% (w/w).
It is understood that in any case the sum of the percentages of all the components of the filaments, and thus in the final implantable drug delivery device, reaches 100%.
In the context of the whole disclosure, the implantable drug delivery device is printed using a layer thickness from about 50 μm to about 500 μm, preferably from about 100 μm to about 400 μm such as 100, 125, 150, 175, 200, 225, 250, 275, 300, 325, 350, 375 or 400 μm. The implantable drug delivery device can be designed with an infill from 0 (hollow object) to 100% (full solid object). In an embodiment, the implantable drug delivery device comprises at least one internal hollow cavity. In an alternative embodiment, implantable drug delivery the device is a fully solid object.
In a further embodiment, the present invention relates to a process for producing an implantable drug delivery device according to the invention, the process comprising:
A non-limiting exemplary filament according to the invention comprises about 15.5% w/w of an antibody (such as a full-length monoclonal antibody or a molecule comprising a Fab fragment), about 7.5% w/w of excipients, about 69.5% w/w of a polymeric material (such as RG502), about 7.5% w/w of plasticizer (such as PEG), wherein the excipients comprise or consist of histidine (used as a buffering agent in the initial liquid formulation) and one disaccharide (either sucrose or trehalose) as the stabilizer. Another non-limiting exemplary filament according to the invention comprises about 15.5% w/w of an antibody (such as a full-length monoclonal antibody or a molecule comprising a Fab fragment), about 7.5% w/w of excipients, about 69.5% w/w of a polymeric material (RG502), about 7.5% w/w of plasticizer (PEG), wherein the excipients comprise or consist of histidine (used as a buffer in the initial liquid formulation), one disaccharide (either sucrose or trehalose) and one amino acid (L-Leucine) both acting as stabilizers.
Preferably the filaments or devices of the invention retain at least 60% of the antibody biological activity at the time of formulation and/or packaging over a period of several weeks after implantation in the subject to be treated. The activity may be measured as described in the following section “Examples” or by any other standard techniques, preferably during preliminary experiments.
The invention also provides an article of manufacture, for pharmaceutical or veterinary use, comprising a container comprising any of the above described filament or implantable drug delivery device. Also described, a packaging material providing instructions for use.
The filaments or the implantable drug delivery devices of the invention may be stored before use for at least about 12 months to about 24 months. Under preferred storage conditions, before the first use, the formulations are kept away from bright light (preferably in the dark), at temperature from about 2 to 18° C., e.g. 18° C., 15° C. or at 2-8° C. The skilled person would understand that depending on the Tg of the polymer, the temperature of storage may be higher than 18° C., such as up to 25° C. (e.g. 20° C., 22° C. or 25° C.).
The present invention provides filaments and implantable drug delivery devices, for single use, suitable for pharmaceutical or veterinary use.
HMWS=high molecular weight species; LMWS=low molecular weight species; SD=spray-drying or spray-dried; HME=Hot melt extrusion; 3DP=Three-dimensional printing or Three-dimensional printed; BE: Buffer exchange; DDS: Drug delivery system; DDD: Drug delivery device; DSC: Differential scanning calorimetry; FDM: Fused deposition modelling; HME: Hot melt extrusion; LEU: L-leucine; Mw: Molecular weight; mAb: full length monoclonal antibody; fAb: Fab fragment of an antibody; PBS: Phosphate buffer solution; gel permeation chromatography (GPC); SEC=Size exclusion chromatography; PEG: Polyethylene glycol; PLGA: Poly(lactide-co-glycolide) acid; rpm: Revolutions per minute; SUC: Sucrose; Tg: Glass transition temperature; TGA: Thermogravimetric analysis; Tm: Melting temperature; TRE: Trehalose; % (w/w): Weight percentage; Stab: stabilizer; HP-β-CD: Hydroxypropyl-β-cyclodextrin; SOR: sorbitol; INU: inulin.
mAb1 is an IgG4, has a molecular weight (MW) of about 150 kDa and a pl of about 6.0-6.3.
fAb2 is a Fab moiety of an antibody. fAb2 has a MW of about 50 kDa and a pl of about 9.3-9.6.
The antibody-containing solutions were spray-dried using a lab-scale Spray-Dryer B-290 (Büchi Labertechnik) equipped with a 0.7 mm nozzle. Settings were based on standard procedure and kept constant for all formulations. The solutions to be spray-dried were preliminary prepared in a 15 mM histidine buffer at pH 5.6, with other excipients on demand. An overview of the mAb1 and fAb2 solution compositions, concentrations and mAb: stabilizer ratios are shown in Tables 1 & 8. All powders were sealed in a polypropylene container and stored in a desiccator under vacuum.
Printable filaments were prepared from physical mixtures of raw PLGA, PEG 2 kDa and mAb1- or fAb2-containing spray-dried (SD) powders which were previously blended together using a Turbula® mixer (Willy A. Bachofen AG). The blend was manually fed into a 11-mm twin screw extruder (Process-11, Thermo Fischer Scientific), equipped with modular screws (L/D-ratio 40:1), and a round die with a diameter of 1.6 mm. The barrel was heated using a gradient of temperature controlled by eight thermocouples. The feeding zone was maintained at room temperature using a water circulator. The three first segments were set at 20, 40 and 80° C., respectively. The middle segments, from the 4th to 6th thermocouple, were set at 90° C. The last thermocouple, which was located right before the die, was set at 85° C. and the die itself was set at 75° C. For all experiments, the screw speed was set at 40 rpm during the feeding and 60 rpm when the filament was manually coiled. These parameters were kept constant (see Table 1).
The design of the devices was drawn using the 3D computer-aided design (CAD) software ThinkerCAD™ (AutoDesk® Inc.) and exported into a software for slicing. The dimensions of the devices were 20×5×2 mm (length, width, height) for a volume of 178.43 mm3. An Hyrel 3D system 30M printer (GA), equipped with a 0.5 mm MK2-250 hot extruder, was used to print the mAb1- and fAb2-loaded devices. The temperature of the build platform did not need to be controlled. The printing temperature was set at 105±2° C. The printing speed was set at 1 mm/s for the first layer and 10 mm/s for the others. The layer thickness of the devices was set at 0.1 mm and 0.3 mm to evaluate its influence on the potential degradation of the loaded mAb1 as well as on its release profile. The printing of devices was performed with an infill of 100% (v/v), except specified otherwise in the examples below.
Differential scanning calorimetry (DSC): Thermal analyses of SD powders, filaments, 3DP DDS were performed, according to a standard method, via DSC using a heat-flux type DSC Q2000 (TA instruments) equipped with a cooling system.
Thermogravimetric analysis (TGA): TGA were performed on a Q500 TGA (TA instrument), equipped with a balance with a sensitivity of 0.1 μg, as per standard methods. Data collection and analysis were performed using TA Instruments® Trios 4.5.0 software.
Molecular weight analysis of polyesters by size exclusion chromatography (SEC) in chloroform: Number average molecular weight (Mn), weight average molecular weight (Mw), and polydispersity index (Mw/Mn) of polyesters were measured by SEC, as per standard methods. Relative molecular weights (number and weight average) and polydispersity index were calculated by reference to a polystyrene standard calibration curve established using the same experimental conditions. Means and standard deviations (STD) related to molecular weights and polydispersity were calculated as detailed above for NMR analysis.
Antibodies stability evaluation: The quantification of mAb1 monomer as well as the evaluation of both HMWS and LMWS contents was carried out by size exclusion high performance liquid chromatography. This analysis was conducted on samples obtained from either dissolution studies or after extraction from the printable filaments and 3DP devices. These quantifications were performed on an Agilent 1200 series LC system equipped with a UV detector (Agilent Technologies), according to standard protocols. The mobile phase was a 0.2 M PBS solution, at pH 7.0. The calibration curve of mAb1 was ranged from 20 to 2000 μg/mL. The stability of mAb1 was evaluated using the percentage of monomer loss, which corresponded to the difference in the percentage of monomers before and after both HME and 3DP processes. Monomer, HMWS and LMWS levels (%) were compared to a reference that consisted of mAb1 solution obtained after buffer exchange. Similar method was used for fAb2 stability evaluation.
Antibodies extraction from the polymeric matrix: To evaluate the stability of mAb1 that was melt-encapsulated in both printable filaments and 3D-printed devices, samples of approximatively 10 mg were placed in Nanosep® with 0.2 μm Bio-Inert centrifugal devices (Pall) and dissolved in 0.5 mL of dichloromethane. The Nanosep® devices were stirred at 600 rpm during 2 hours at room temperature to dissolve the PLGA, using a Thermomixer confort® tubes mixer (Eppendorf AG). The samples were centrifuged at 12 000 rpm during 10 min and the medium was withdrawn. Then, 0.5 mL of dichloromethane were added again. The sample was stirred for 5 minutes and centrifuged as previously mentioned. This step was repeated twice. Dichloromethane was removed and the Nanosep® devices containing the mAb's precipitate were placed 1 hour under vacuum to remove potential residual solvent. Then, 0.5 mL of PBS (0.2 M, pH 7.0), containing 0.02% w/w of polysorbate 80 (PS80), were added in the tube to solubilize mAb1 before being stirred at 600 rpm for 2 hours. Then, the Nanosep® devices were centrifuged 10 min at 12 000 rpm (Adapted from Arrighi et al., 2019). mAb1 stability was evaluated by SEC (as described above). Similar method was used for fAb2 extraction.
Antibodies loading after melt-encapsulation: The amount of encapsulated mAb1 into PLGA matrix was determined using colorimetric detection by bicinchoninic acid (BCA) protein assay according to standard methods. The Pierce™ microplate procedure was carried out to determine the amount of melt-encapsulated mAb1. The quantification of both standards and samples was performed at 562 nm on a SpectraMax M5 microplate reader (Molecular Devices) at room temperature. Overall, mAb1 loading was determined as follows:
mAb loading (%)=(amount of melt-encapsulated mAb)/(amount of 3DP device)*100.
Similar method was used for fAb2 loading.
Dissolution studies: To evaluate the release profiles of the loaded mAb1/fAb2 from 3DP DDS, in vitro dissolution studies were performed. 3DP devices (˜200 mg) were placed in 5 mL Eppendorf® tube filled with 5 mL of PBS (0.2 M, pH 7.0, 37° C.) and stirred at 600 rpm using a Thermomixer confort® tubes mixer (Eppendorf AG) (adapted from (Marquette et al., 2014)). At predetermined times, 5 mL of medium was withdrawn, collected and filtrated on 0.45 μm PVDF Acrodisc® syringe filters (Pall). Similar volume was replaced with fresh buffer (5 mL). The filtrated solutions were measured using SEC analysis equipped with UV-detector at 280 nm and analysed for pH.
PLGA degradation during dissolution: The decrease in polymer molecular weight (Mw) of PLGA during the drug release was carried out using gel permeation chromatography (GPC). The protocol was similar to that used for the dissolution test. Mw were calculated using polystyrene standards.
Enzyme-linked immunosorbent assay (ELISA test): The binding capacity of the mAb1/fAb2 was assessed using an ELISA test, according to standard methods.
Data analysis: All experiments were performed in triplicate, unless otherwise specified. Prism 8 software (GraphPad software) was used for statistical analysis. The results are expressed as a mean±standard deviation. Statistical significance was determined at p-value<0.05 using ANOVA and Turkey's or Dunnett's post-hoc test (as recommended by Prism software).
The mAb1 solutions were formulated with different stabilizers (see Table 1). These liquid solutions were spray-dried to produce mAb1-loaded powders. Indeed, mAb1 was used in solid state to increase its stability and to facilitate the handling during further processing. Then, a mixture of mAb1-loaded powder, Resomer® RG502 as a polymeric material (Evonik Industries) and PEG as a plasticizer was extruded using HME to produce filaments suitable for printing. These printable filaments were used to feed the 3DP printer to print the devices (alternatively herein named drug delivery device or implantable drug delivery device). Optimal formulations were identified by evaluating mAb1 integrity after each manufacturing step (SD, HME, 3DP). Finally, in vitro evaluations (dissolution test and binding capacity) were performed.
The thermal properties, including their temperature of degradation, of all raw materials were assessed using TGA and DSC analysis, respectively.
The degradation temperature of raw RG502 was around 175° C. No apparent weight loss was observed under 200° C. on raw PEG and on the extruded filaments loaded with mAb1. No residual moisture was observed in RG502 and PEG raw materials. These results confirmed that all raw materials seemed stable and may be processed according to the temperatures in both HME and 3DP (90° C. and 105° C., respectively). Indeed, only the mass loss was characterized using TGA and other methods were required to state on mAb1 stability such as SEC and binding capacity.
The TGA thermograms of the SD mAb1 powder showed a slight weight loss (˜4% w/w) when a temperature of 100° C. was reached. Such decrease could be attributed to the residual moisture content into the SD mAb1 powder (about 3.4±0.8%). A second weight loss was observed above 150° C. on all the SD mAb1-loaded powders. The mAb1-loaded powders were thus able to ensure the stability of mAb1 during both HME and 3DP.
Then, DSC analyses were carried out to evaluate the influence of the addition of PEG and mAb1-loaded SD powder on the Tg of the thermoplastic polymer RG502. Indeed, the aim of this work being to develop mAb1-loaded 3DP DDS, the Tg should be as low as possible to allow decreasing the temperature of the different processes (HME, 3DP) and the potential degradation of the biotherapeutic as a consequence.
The Tg of RG502 was found to be 38.0±0.7° C., which was consistent with data already described in literature (Pignatello et al., 2009). PEG was characterized by a sharp endothermic peak at 52° C. The Tg of RG502 decreased to 21.8±0.4° C. when PEG and SD powder were added during HME (data not shown). Such decrease of the Tg, in addition to the loss of the sharp melting peak of PEG, demonstrated that mAb1-loaded SD powder and PEG were properly dispersed in the molten polymeric matrix (Zhang et al., 2017).
Stabilizers were selected to maintain antibody integrity during all the steps of manufacturing. The main expected deleterious factor was the relatively high temperatures that were used during both HME and 3DP. Unfortunately, stabilizer selection is not universal and needs to be adapted to each biotherapeutic and in regard with the stress factors associated to the process (Le Basle et al., 2020; Wang et al., 2007). SUC, TRE, HP-β-CD, SOR and INU are commonly used in formulations comprising antibodies (Baek et al., 2017; Bowen et al., 2013; Gidwani and Vyas, 2015; Kanojia et al., 2016; Maury et al., 2005). The effect of the addition of stabilizers on the stability of the loaded mAb1 was investigated using 3 different mAb: stabilizer ratios (w/w) (1.5:1, 2.0:1 and 2.5:1) (see formulations of Table 1). A mAb: stabilizer ratio (w/w) 2.0:1 was previously described to increase the stability of mAb1 during a SD process (Bowen et al., 2013). Higher and lower ratios were also investigated to evaluate their influence on the stability of our own mAb1, not only during SD, but more especially during HME and 3DP (two steps bringing a high thermal stress).
The different liquid compositions to be assessed were obtained by buffer exchange. No instabilities were observed between mAb1 reference (before buffer exchange) and the various liquid compositions (after buffer exchange, BE). The percentages of HMWS were quite similar to that observed from mAb1 reference (2.6±0.4%) (Table 2). After SD, there was no significant formation of HMWS either for mAb: stabilizer ratios of 1.5:1 and 2.0:1, regardless of the nature of the stabilizer (p-value>0.05) (
To obtain filaments, mAb1-loaded SD powders were mixed with PLGA and PEG and extruded (HME) to make printable filaments (see formulations of Table 1). The filaments were successfully prepared with a diameter between 1.70 and 1.75 mm as recommended to feed the FDM 3D printer (Melocchi et al., 2015). mAb1 loading was chosen at 15% (w/w).
As shown in Table 2 and
The percentage of LMWS was also evaluated after HME (see Table 2). It was observed that slight fragmentation appeared when HP-β-CD and SOR were used as stabilizers. In contrast, no LMWS were observed with SUC, TRE and INU.
Overall, HP-β-CD, SOR and INU were less effective to maintain mAb1 stability during HME in comparison to SUC and TRE. Based on the evaluation of HMWS and LMWS levels, mAb1 integrity was ensured during HME using TRE and SUC as stabilizers.
Examples 2 and 3 have shown that SUC and TRE seemed to be the most suitable stabilizers to stabilize the formulations over the successive steps of production (after SD and HME).
Finally, mAb1 loading was assessed on the printable filaments before the printing process. This showed that the real loadings of all the filaments were similar to the targeted one (15% w/w) with very low standard deviations (Table 3). These results indicated that the manufacturing process was suitable and reproducible to produce uniform printable filaments with homogeneous dispersion.
A slicing software was used to design a model of implantable 3DP device with a shape that could be implantable. The printing process was performed in a room at 20° C. Indeed, physical state of the filaments may be quickly modified due to the room temperature as it was previously mentioned that their Tg was around 22° C. Therefore, at 20° C., filaments were able to be printed as their stiffness was preserved. However, the handling of the filaments induced a heat transfer by conduction. This phenomenon was greater when the filaments were loaded in the print head. Indeed, they were too soft to be travelled along the feeding gears. To limit the heat transfer by conduction during printing, 3DP had to be performed using a “flexible hot flow” modular printing head MKE-250.
The device resolution was macroscopically evaluated and, when the infill was set at 100%, a fully solid device was expected. Immediate visualization showed defects and a lack of matter at the top of the devices (data not shown). The printing step was performed at 105° C. which was the temperature where both adhesion to the build platform and between successive layers were promoted. The printing speed was selected at 1 mm/s for the first layer and 10 mm/s for the following layers to improve the resolution of DDS. 3D printings with a layer thicknesses of 0.1 mm and 0.3 mm were evaluated.
Extraction of mAb1 was performed on 3DP devices to evaluate the percentage of both HMWS and LMWS. The percentage of HMWS increased following 3DP, regardless of the layer height as well as the nature of the disaccharide (
Despite the addition of SUC or TRE, it was demonstrated that a significant (although acceptable) increase of HMWS appeared after 3D printing. Therefore, it was hypothesized that the addition of a hydrophobic amino acid such as LEU (Minne et al., 2008) could enhance the stability of the loaded mAb1. 3DP devices were printed using a layer thickness of 0.3 mm, starting with preliminary liquid formulations comprising the combination of stabilizers SUC-LEU or TRE-LEU (see Table 1).
HMWS levels were evaluated after each process (from SD to 3DP, the starting values being those for BE) (see
LMWS levels were also investigated after 3DP. It was demonstrated that a slight increase of LMWS (around 0.05±0.04%), regardless the addition of LEU to SUC or TRE (data not shown).
Finally, drug loadings were assessed on 3DP DDS and the BCA results showed real loadings closed to the targeted loading of 15% (w/w) (see Table 4). These results confirm the uniform dispersion of mAb1 in the polymeric matrix expressed after HME.
It was therefore a surprising finding from the inventors that it was possible 1) to extrude printable filaments by HME starting from mAb1-loaded SD and 2) to create 3DP devices via FDM with said filaments. The increase of HMWS observed mainly after HME and 3DP was directly related to the thermal degradation occurring at 90° C. (during HME) and 105° C. (during 3DP). The most promising formulation, containing TRE-LEU and to a lesser extend SUC-LEU, and able to minimize the production of HMWS and to promote mAb1 stability, was further investigated.
It was previously demonstrated and described that PLGA-based drug delivery system (DDS; e.g. microparticles and implants) are characterized by a triphasic release profiles. It could be more interesting to promote a release profile where a limited latent phase occurred. Indeed, the latent phase could lead to mAb1 degradation due to its retention in the polymeric matrix and the medium uptake. Moreover, a linear release profile which could tend towards a “zero order kinetic” should allow a constant drug release and a steady release concentration of the mAb1 in the dissolution medium.
As shown in
Degradation of the polymer RG502 was evaluated on the 3D printed device during the dissolution test (
The release of mAb1 was assessed over 15 weeks. The pH values remained slightly acidic due to the generation of oligomers and their diffusion to the dissolution medium. No further degradation of the PLGA nor further release of mAb1 was observed after 10 weeks. As the sample generated after 10 weeks in the dissolution medium remained insoluble in chloroform, it is likely that PLGA and mAb1 form insoluble aggregates over time.
In order to study the stability of mAb1 during release, HMWS and LMWS levels as well as the monomer content were assessed in the dissolution test (
ELISA assays were performed to evaluate the binding capacity of mAb1 after its diffusion from the devices to the dissolution medium (see
Stability over time being an important aspect to consider when a drug product is developed, the effect of storage temperatures using 5±3° C. and 25±2° C. for 6 months (T0, T1, T2, T3 and T6 months) was assessed. 3DP devices were produced using mAb1 stabilized with the TRE-LEU combination.
Physical state of polymeric matrix: DSC analyses of the 3DP devices were compared at different time points (see Table 5). As previously mentioned, the PLGA was plasticized using PEG at 11% (w/w) and the Tg of the filament (before printing) was 21.8±0.4° C. The Tg of the reference samples (T0) were close to this value with 20.7±0.3° C. No increase of the Tg was observed over 3 months according to both storage temperatures (i.e. 5° C. and 25° C.) However, an increase of the Tg to 29.7±0.3° C. (T6) was observed after 6 months at 25° C. The Tg of devices during the stability study remained steady at 5° C. Besides, minor melting peaks were observed on samples stored at 25° C. for 2 months (T2), 3 months (T3) and 6 months (T6). The Tm were observed at 45.2±1.4° C. (T2); 45.9±0.8° C. (T3) and 46.7±0.4° C. (T6). The melting peak could be attributed to the PEG which was able to move at temperature higher than Tg of the polymeric matrix (i.e. 25° C.). The melting enthalpy of these melting peaks were recorded and showed an increase over months from 1.7±0.9 J/g (T2) to 7.4±0.6 J/g (T6). The increase of the melting enthalpy demonstrated a probably phase separation with the PLGA chain mobility at 25° C. After 2 and 3 months, the melting enthalpy remained low and the plasticizing effect was effective. The increase of the Tg after 6 months at 25° C. was associated with a higher value of melting enthalpy which was consistent with a phase separation between PLGA and PEG. The melting enthalpy of neat PEG was recorded around 193.4 J/g (Data not shown). Therefore, only a small quantity of the PEG tended to separate from the PLGA blend over 6 months. Similar observations were observed during polymer ageing study. It was reported that PEG was able to crystallize over time due to an elevation of storage temperature and humidity. The crystallization of the PEG may increase the stiffness of the devices and modify their mechanical and release properties.
The degradation of the PLGA was assessed using GPC measurement. The Mw of T0 was recorded at 17.02±0.38 kDa which was consistent with the raw PLGA as received (Mw: 17.05±0.45 kDa). No degradation occurred over 3 months of storage. These results demonstrated the stability of the devices when they were stored at 5° C. and 25° C. for 3 months. However, the result obtained after 6 months were stored in the fridge for 1.5 months and may affect the Mw of the polymer due to the relative humidity.
Visual assessment of the devices over storage time: Visual assessment was carried out on the 3DP devices stored at 5° C. and 25° C. (not shown). No difference was observed on the devices stored at 5° C. over 6 months. Sticky specimens were observed when devices were kept at 25° C. Devices adhered to the bottom of glass vial but no loss of material was observed during the withdrawal step. This observation was performed on every device stored at 25° C. from T1 to T6. The cross-section of devices T6 at 25° C. showed high porous network due to the mobility of the chain. An increase of the device porosity was expected to have a faster release of mAb1 during the dissolution study.
Drug content and extraction from the devices: The targeted loading was 15% (w/w). As shown in Table 6, mAb1 loading in each device was consistent with the values obtained experimentally.
The stability of mAb1 was also assessed at each time point (Table 6). The monomer content remained stable over 6 months at 5° C. However, a slight decrease of the monomer percentage was observed after 6 months at 25° C. This decrease was associated to an increase of both HMWS (5.1±0.2%) and LMWS (0.08±0.01%) levels of the samples.
A dissolution study was performed on printed devices (3DP_39 to 3DP_44; Table 7) to investigate the release patterns as a function of the infill of the devices (
Based on the findings from examples 1 to 7, TRE and SUC as stabilizers, with and without the addition of LEU, were investigated on fAb2, a Fab antibody fragment.
A similar approach was applied to fAb2 with successive processing methods such as buffer exchange (BE), spray-drying (SD), Hot melt extrusion (HME) and 3D printing (3DP) at the end. The formulations of fAb2 were made of TRE or SUC with or without LEU to compare four different formulations and find out which one was able to stabilize the Fab against thermal stresses (Table 8) and associated 3D printing (3DP) batches with layer thickness of 0.1 mm and 0.3 mm. fAb2 was at 8% w/v in the initial liquid composition, 66.7% w/w in the SD powder and 15.3% w/w in the filaments/3DP devices.
Characteristics of the extracted fAb2: The raw fAb2 material was characterized with a high monomer content of 99.6±0.2% and a low HMWS level of 0.4±0.2%. The extracted fAb2 from the PLGA matrix after HME and 3DP were compared with the extracted fAb2 from SD powder (
Dissolution study: A dissolution study was performed on all printed devices (DDS; 3DP-F1 to 3DP_F4) to investigate both the release patterns and the stability of the Fab over time (
Stability study of fAb2 over time: During the dissolution test, both monomer content and HMWS level were investigated over 8 weeks (
Binding capacity study: The binding capacity of fAb2 was assessed to confirm that it was still able to bind its target. ELISA test demonstrated that the binding capacity of fAb2 was preserved after 24 h of release, regardless the formulation. For instance, the binding capacity of 3DP_F4 was 99.5±6.4% (
Conclusion: fAb2 was successively dried, extruded and 3D printed using four different formulations containing SUC or TRE with/without LEU. The 3DP DDS allowed a sustained release of the Fab over at least 8 weeks. The HMWS level remained quite low with a maximum value of 1.9±0.1% (3DP_F4). Although results obtained with formulations comprising TRE (+/−LEU) as stabilizers were slightly better in term of total release, SUC (+/−LEU) were also very promising.
The results herein presented surprisingly showed for the first time that not only HME but also the association of HME and FDM 3D printing were suitable to produce antibody-loaded filaments and antibody-loaded implantable devices in which the antibody can still bind its target (and therefore is likely still active), as herein shown with a monoclonal antibody (mAb1) and a Fab fragment (fAb2). Homogeneous solid dispersion of the antibodies in the PLGA matrix was reached in both printable filaments and 3DP devices. Different stabilizers were investigated to stabilize the antibodies against thermal degradation. The most promising ones (trehalose and sucrose) promoted the mAb integrity during SD, HME and 3DP steps using different mAb: stabilizer ratios. The further optimization of the formulation, using a low amount of an amino acid, such as Leucine, led to improved stability of the antibodies against potential thermal degradation. In addition, dissolution profile demonstrated an interesting sustained-release profile with a limited burst effect, especially with an antibody fragment. Finally, it was demonstrated that, despite the relatively high temperatures of extrusion (90° C.) and printing (105° C.), the binding capacity of the antibodies remains for about at least 5 weeks.
| Number | Date | Country | Kind |
|---|---|---|---|
| 2018889.2 | Dec 2020 | GB | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/EP2021/083401 | 11/29/2021 | WO |
