formulation

Abstract

The present invention relates to the field of pharmaceutical compositions containing proteins as therapeutic agents. More particularly, it relates to antibody-containing filaments produced by hot-melt extrusion, implantable drug delivery devices made from these filaments, and methods for producing such filaments and devices. The antibody-containing filaments produced by hot-melt extrusion and devices obtained from the filaments according to the invention allow for the delivery of antibodies over a period of time.

Description

Technical Field

[0001] The present invention relates to the field of pharmaceutical compositions containing proteins as therapeutic agents. More specifically, it relates to antibody-containing filaments produced by hot melt extrusion, implantable drug delivery devices made from these filaments, and methods for manufacturing such filaments and devices. The antibody-containing filaments produced by hot melt extrusion and the devices obtained from the filaments according to the present invention enable the delivery of antibodies over a period of time.

Background Art

[0002] Hot melt extrusion (HME) has been widely described and implemented in the pharmaceutical field for manufacturing printable filaments loaded with drugs (Goyanes et al., 2015; Tiwari et al., 2016). HME is based on the melting of a polymeric material that is extruded through a die to obtain a uniform drug-loaded filament. HME is a free-solvent process that can be easily scaled up. However, this technique is based on the use of relatively high temperatures. Such temperatures can usually be reduced by adding plasticizers, enabling a decrease in the glass transition temperature of the polymer. Another alternative for reducing the extrusion temperature can be the use of thermoplastic polymers characterized by low molecular weight (Fredenberg et al., 2011). HME has already been studied for developing protein-based formulations characterized by controlled release of the loaded active ingredient over time (Cosse et al., 2016; Duque et al., 2018; Ghalanbor et al., 2010).

[0003] HME can be used in combination with 3D printing (3DP) processes such as molten metal deposition (FDM®). FDM processes are currently an integrated part of the pharmaceutical sector (Jamroz et al., 2018; Azad et al., 2020). This technology is an extrusion-based 3DP method that uses heat to melt thermoplastic polymer filaments to build objects layer by layer. The use of 3DP enables the manufacture of any kind of shape, starting from digital design (Norman et al., 2017). A major drawback is the lack of pharmaceutical-grade polymers available for use in FDM, although poly(lactic acid) (PLA) and polyvinyl alcohol (PVA) are commonly used as thermoplastic polymers to produce drug-loaded printable filaments (Jamroz et al., 2018).

[0004] Poly(lactide-co-glycolide) (PLGA) is a well-known pharmaceutical-grade polymer material commonly used to create injectable / implantable sustained-release drug delivery systems (DDS). PLGA can be extruded at low temperatures and is a good candidate for both HME and FDM processes. Protein-supported PLGA implants have already been described using polymers such as ovalbumin (Duque et al., 2018), bovine serum albumin (Cosse et al., 2016), and lysozyme (Ghalanbor et al., 2010). The main challenge remains the stabilization of the protein during extrusion.

[0005] The solid state of proteins has been shown to promote greater stability and may be more advantageous for facilitating their addition to polymer matrices using HME processes (Cosse et al., 2016; Mensink et al., 2017). However, protein compounds commonly used as models for fabricating protein-supported implants (i.e., OVA, BSA, lysozyme) are characterized by low molecular weight compared to, for example, immunoglobulins.

[0006] Therefore, there is still a need for larger proteins, more specifically antibodies, in filaments and implantable drug delivery devices that maintain antibody activity (i.e., without dramatically affecting their biological activity), improve antibody stability (e.g., limit antibody degradation during filament production and subsequent implantable drug delivery device production), and possess sustained-release properties. [Overview of the project] [Means for solving the problem]

[0007] 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 polymer material, a plasticizer, and an active ingredient, the active ingredient being an antibody. The filament may further comprise at least one stabilizer, buffer, and / or surfactant.

[0008] In a second aspect, the present invention relates to an implantable drug delivery device comprising or comprising one or more layers made from filaments comprising at least one polymer material, a plasticizer and an active ingredient, or the active ingredient being an antibody. The filaments may further comprise at least one stabilizer, buffer and / or surfactant.

[0009] In a third aspect, the present invention relates to a 3D-printed implantable drug delivery device obtained by 3D printing a filament comprising or consisting of at least one polymer material, a plasticizer, and an active ingredient, wherein the active ingredient is an antibody. The filament may further comprise at least one stabilizer, buffer, and / or surfactant.

[0010] In a fourth aspect, the present invention relates to a method for manufacturing a filament for preparing an implantable drug delivery device, a. A step of preparing a liquid formulation containing or comprising an active ingredient, wherein the liquid formulation may further contain at least one stabilizer, buffer and / or surfactant, and the active ingredient is an antibody. b. A step of freeze-drying or spray-drying the liquid formulation from step a to obtain dried fine particles. c. A step of uniformly dispersing the dried fine particles from step b together with a plasticizer and at least one polymer material. The present invention provides a method comprising the step of obtaining a filament by extruding the dispersion from step c by hot melt extrusion (HME).

[0011] In a fifth aspect, the present invention relates to a method for manufacturing an implantable drug delivery device, a. The process of loading the filament described herein into the print head of a 3D printer using a temperature higher than the glass transition temperature. b. A step of heating the construction platform at a temperature lower than the glass transition temperature of the polymer matrix, c. The method comprises the step of depositing the heated filament through a nozzle to construct a device from at least a first layer to a final upper layer.

[0012] definition The term "dried microparticles" (and its plural form, "dried microparticles") refers to dry "particles" of very small size (typically about 20 μm or less) (or they are called "microparticles" or "microspheres"). Preferably, the dried microparticles contain less than about 10% by weight of water, usually less than 5% by weight, and even less than 3% by weight. Dried microparticles can typically be obtained by spray-drying and / or freeze-drying an aqueous solution or aqueous emulsion. Alternatively, the term "dried powder" can be used.

[0013] The term "lyophilization," also known as "freeze-drying," refers to a process for obtaining dried fine particles that includes at least three main steps: 1) lowering the temperature of the product to be freeze-dried to below its freezing point (typically between -40°C and -80°C; the freezing step), 2) applying high-pressure vacuum (typically between 30 and 300 mTorr; the first drying step), and 3) raising the temperature (typically between 20°C and 40°C; the second drying step).

[0014] The term "spray drying" refers to a process for obtaining dry fine particles, comprising at least two main steps: 1) atomizing a liquid feed into fine droplets, and 2) evaporating a solvent or water with a high-temperature drying gas.

[0015] The term “sustained release” (also referred to herein as “continuous release”) refers to the delivery of an active ingredient over several days, weeks, months, or years. A typical sustained release profile for protein-supported polymer microparticles is triphasic, consisting of (i) an initial burst release (i.e., an initial large release of the active ingredient), (ii) a delay phase (i.e., a phase in which very little or no product is released), and (iii) an release phase (i.e., a phase in which the release rate is stable) (Diwan et al., 2001 and White et al., 2013). Preferably, an initial burst release of about 40% or less of the total amount of active ingredient is considered acceptable. An initial burst release of 30% or less is called a “limited burst release.” The release of antibody molecules should also be as complete as possible (i.e., a total release as close as possible to 100% of the encapsulated antibody), preferably at least more than 60%. One advantage of such sustained-release compositions is that the composition is not administered to patients too frequently.

[0016] As used herein, the term “stability” refers to the physical, chemical, and conformational stability (including maintenance of biological efficacy) of the active ingredient (antibody, as herein) in the filaments and drug delivery devices according to the present invention. Antibody instability may be caused, for example, by chemical degradation or aggregation of the antibody to form higher-order polymers, deglycosylation, modification of glycosylation, oxidation, or any other structural modification that reduces the biological activity of the formulated antibody. The term “stable” refers to a filament or drug delivery device in which the active ingredient (antibody, as herein) essentially retains its physical, chemical, and / or biological properties during manufacturing and storage. Various analytical methods for measuring antibody stability in formulations are well within the knowledge of those skilled in the art (see also the Examples section). Various parameters can be measured to determine the stability of the filament or 3DP device (compared to initial data), for example (but not limited to): 1) a change of about 15% or less in the monomeric form of the antibody, or 2) a change of 15% or less in high molecular weight species (HMW or HMWS; also referred to herein as aggregates).

[0017] As used herein, the terms “buffer” or “buffering agent” refer to a solution of a compound known to be safe in pharmaceutical formulations and having the effect of maintaining or controlling the pH of the formulation within a pH range desirable for the formulation. Acceptable buffers for controlling pH from a moderately acidic pH to a moderately basic pH include, but are not limited to, phosphates, acetates, citrates, arginine, histidine buffers, TRIS (2-amino-2-hydroxymethyl-1,3,-propanediol), and any pharmacopositically acceptable salts thereof.

[0018] As used herein, the term “surfactant” refers to a soluble compound that can be used to increase the water solubility of hydrophobic oily substances or to increase the miscibility of two substances that otherwise have different hydrophobic properties. Surfactants are commonly used in formulations, particularly to modify the absorption of drugs or their delivery to target tissues. Well-known surfactants include polysorbates (polyoxyethylene derivatives; Tween) and poloxamers (i.e., copolymers based on ethylene oxide and propylene oxide, also known as Pluronics®).

[0019] As used herein, the term “stabilizer” or “stabilizer” refers to a compound that is physiologically acceptable and imparts appropriate stability / tonicity to a formulation. During freeze-drying (freeze-vacuum drying) or spray-drying processes, stabilizers are also effective as protective agents. Compounds such as glycerin are commonly used for this purpose. Other suitable stabilizers include, but are not limited to, amino acids or proteins (e.g., glycine or albumin), salts (e.g., sodium chloride), and sugars (e.g., dextrose, mannitol, sucrose, trehalose, and lactose), as described within the scope of this disclosure.

[0020] The term "polymer material" refers to polymer elements that can withstand high temperatures during hot-melt extrusion (HME) and 3D printing. Therefore, preferred polymer materials according to the present invention are thermoplastic polymers or heat-resistant polymers. Examples of such thermoplastic polymers commonly used in 3D printing include, for example, polyvinylpyrrolidone (PVP), acrylonitrile butadiene styrene (ABS), poly(lactic acid) (PLA), poly(lactic acid-co-glycolic acid) (PLGA), polyvinyl alcohol (PVA), poly(ε-caprolactone) (PCL), and ethylene vinyl acetate (EVA). Preferably, they are biodegradable or bioeliminable for greater convenience for the patient. Other heat-resistant polymer materials include, for example, hydroxypropyl cellulose (HPC), hydroxypropyl methylcellulose (HPMC), poly(ethylene glycol) (PEG), eudragit derivatives (E, RS, RL, EPO), polyvinylcaprolactam-polyvinyl acetate-polyethylene glycol graft copolymer (Soluplus®), and thermoplastic polyurethane (TPU). Suitable polymer materials are also described herein.

[0021] The term "plasticizer" refers to compounds that can be combined with thermoplastic polymers, for example, to increase their plasticity or to decrease their viscosity. They can also help lower the glass transition temperature (Tg) of the polymer. Examples of such plasticizers that can be used in the pharmaceutical industry include, for example, bioplasticizers such as alkyl citrates (e.g., acetyl triethyl citrate (ATEC), triethyl citrate (TEC)), triacetin (TA), methyl ricinoleate, epoxidized vegetable oils, or even polyethylene glycol (PEG) (depending on its molecular weight, PEG can act as either a polymer matrix or a plasticizer), castor oil, vitamin E TPGS (D-α-tocopheryl polyethylene glycol 1000 succinate), fatty acid esters (butyl stearate, glycerol monostearate, stearyl alcohol), pressurized carbon dioxide, and surfactants (polysorbate 80) (see, for example, Crowley 2007). Preferred plasticizers are also described herein.

[0022] As used herein, the term “antibody” includes, but is not limited to, monoclonal antibodies, polyclonal antibodies, and recombinant antibodies produced by recombinant technologies known in the art. “Antibody” includes antibodies of any species, particularly mammalian species; for example, any isotype of human antibodies including IgG1, IgG2a, IgG2b, IgG3, IgG4, IgE, IgD, and antibodies produced as dimers of this basic structure including IgGA1, IgGA2, or pentamers such as IgM and their modified variants; non-human primate antibodies, for example, from chimpanzees, baboons, rhesus monkeys, or cynomolgus macaques; rodent antibodies, for example, from mice or rats; antibodies of rabbits, goats, or horses; camelid animal antibodies (for example, from camels or llamas such as Nanobodies®) and their derivatives; avian antibodies such as chicken antibodies; or fish species antibodies such as shark antibodies. The term “antibody” also refers to a “chimeric” antibody in which at least one heavy-chain and / or light-chain antibody sequence has a first portion derived from a first species and a second portion derived from a second species. The chimeric antibodies of interest as used herein include “primatized” antibodies containing variable domain antigen-binding sequences derived from non-human primates (e.g., Old World monkeys such as baboons, rhesus monkeys, or cynomolgus monkeys) and human constant-region sequences. “Humanized” antibodies are chimeric antibodies containing sequences derived from non-human antibodies. In most cases, a humanized antibody is a human antibody (recipient antibody) in which residues from the recipient’s hypervariable region are replaced with residues from the hypervariable region [or complementarity-determining region (CDR)] of a non-human species (donor antibody), such as mouse, rat, rabbit, chicken, or non-human primate, having the desired specificity, affinity, and activity. In most cases, residues outside the CDR of the human (recipient) antibody; i.e., in the framework region (FR), are further replaced with corresponding non-human residues. Furthermore, humanized antibodies may contain residues not found in recipient or donor antibodies. These modifications are made to further improve antibody properties. Humanization reduces the immunogenicity of non-human antibodies in humans, thus facilitating the application of antibodies to the treatment of human diseases. Humanized antibodies and several different techniques for producing them are well known in the art.The term “antibody” also refers to human antibodies that can be produced as an alternative to humanization. For example, immunization makes it possible to produce transgenic animals (e.g., mice) that can produce a complete repertoire of human antibodies in the absence of endogenous mouse antibody production. Other methods for obtaining human antibodies / antibody fragments in vitro are based on display techniques such as phage display or ribosome display techniques, and use recombinant DNA libraries that are at least partially artificially produced or generated from the donor’s immunoglobulin variable (V) domain gene repertoire. Phage and ribosome display techniques for producing human antibodies are well known in the art. Human antibodies can also be produced from isolated human B cells that have been ex vivo immunized with the antigen of interest, subsequently fused to produce hybridomas, which can then be screened for optimal human antibodies. The term “antibody” refers to both glycosylated and non-glycosylated antibodies. Furthermore, as used herein, the term “antibody” refers not only to full-length antibodies but also to antibody fragments, more specifically antigen-binding fragments. The antibody fragment comprises at least one heavy or light chain immunoglobulin domain known in the art and binds to one or more antigens. Examples of antibody fragments according to the present invention include Fab, modified Fab, Fab', modified Fab', F(ab')2, Fv, Fab-Fv, Fab-dsFv, Fab-Fv-Fv, scFv, and Bis-scFv fragments. The fragment may also be a diabody, tribody, triabody, tetrabody, minibody, single-domain antibody (dAb), such as sdAb, VL, VH, VHH, or camelid antibody (e.g., camel or llama-derived, such as Nanobody®), and VNAR fragment. The antigen-binding fragment according to the present invention may also include a Fab linked to one or two scFv or dsscFv, each scFv or dsscFv binding to the same or different targets (e.g., one scFv or dsscFv that binds to a therapeutic target and another scFv or dsscFv that extends its half-life by binding to albumin, for example).Examples of such antibody fragments are FabdsscFv (also called BYbe®), or Fab-(dsscFv)2 (also called TrYbe®, see for example WO 2015 / 197772). Antibody fragments as defined above are known in the art.

[0023] Unless otherwise specified, value percentages (%) refer to weight percentages (alternatively called wt% or %w / w).

[0024] Detailed Description of the Invention Based on the advantages of hot melt extrusion (HME) and / or fused deposition modeling (FDM) techniques, the inventors then developed an antibody - loaded filament that can be used to obtain an implantable device via 3D printing using the FDM technique, etc. The present invention is based on the surprising discovery that filaments containing antibodies can be produced and that the filaments have a high antibody loading rate (15% or more). Then, an implantable drug delivery device (e.g., obtained by molding or 3D printing) can be obtained using the filaments, from which the antibody is released in a controlled manner over time. Furthermore, the antibody was not only released in a timely manner but still able to bind to its target. It was necessary to carefully select the type of thermoplastic polymer used and optimize the manufacturing parameters of HME or both HME and FDM to first obtain the filaments and then an implantable drug delivery device that can maintain the stability and affinity of the loaded antibody.

[0025] 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 polymer material, a plasticizer, and an active ingredient, the active ingredient being an antibody. The filament may further comprise at least one stabilizer, buffer, and / or surfactant. In such cases, and as an example, the filament according to the present invention may, as a whole, comprise or consist of at least one polymer material, a plasticizer, an antibody, and at least one stabilizer. As a further example, the filament according to the present invention may comprise or consist of at least one polymer material, a plasticizer, an antibody, at least one stabilizer, and a buffer. The filament can be molded or used with a 3D printer to obtain an implantable drug delivery device of any desired shape.

[0026] The present invention further provides an implantable drug delivery device comprising, or comprising, one or more layers made from filaments comprising at least one polymer material, a plasticizer, and an active ingredient, wherein the active ingredient is an antibody, and the filaments may further comprise at least one stabilizer, buffer, and / or surfactant.

[0027] 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 polymer material, a plasticizer, and an active ingredient, wherein the active ingredient is an antibody. The filament may further comprise at least one stabilizer, buffer, and / or surfactant.

[0028] Before adding the active ingredient to a polymer material to form a filament, and then forming an implantable drug delivery device, the active ingredient needs to be spray-dried or freeze-dried. To do this, a preliminary liquid formulation is prepared in which the formulation contains or consists of the active ingredient, and the active ingredient is an antibody. The liquid formulation may further contain at least one stabilizer, buffer and / or surfactant. The liquid formulation is then spray-dried or freeze-dried according to a standard method to obtain dried microparticles. In the form of dried microparticles, the active ingredient is uniformly dispersed in at least one polymer matrix and a plasticizer. They form an active ingredient-supported solid dispersion, such as an antibody-supported solid dispersion.

[0029] Therefore, in this specification, a method for manufacturing a filament according to the present invention is described, a. A step of preparing a liquid formulation containing or comprising an active ingredient, wherein the liquid formulation may further contain at least one stabilizer, buffer and / or surfactant, and the active ingredient is an antibody. b. A step of freeze-drying or spray-drying the liquid formulation from step a to obtain dried fine particles. c. A step of uniformly dispersing the dried fine particles from step b together with a plasticizer and at least one polymer material (also referred to herein as an active ingredient-supported solid dispersion), d. A step of obtaining a filament by extruding the dispersion from step c by hot melt extrusion (HME). This provides a method that includes [something].

[0030] The filament according to the present invention can be used to manufacture implantable drug delivery devices. These devices can be cut to desired lengths, pelletized, molded, or 3D printed. The advantage of using a 3D printer is that it enables the design and manufacture of novel and customized implantable drug delivery devices that would be impossible using conventional processes. Thanks to 3DP technology, the structure, shape, or configuration of the device can be customized and, in some cases, adapted to the patient. Another advantage of using a 3D printer is the ability to provide devices on demand.

[0031] 3D printing is a part of a technique called additive manufacturing (ALM). ALM can be based on liquid solidification or solid material extrusion. Liquid solidification techniques include, for example, drop-on powder deposition (DoP, or binder injection) and drop-on drop deposition (DOD), while solid material extrusion techniques include pressure-assisted microsyringe (PAM) deposition, or molten filament manufacturing (FFF), also known as molten material deposition (FDM®) technique. In DoP or DoD systems, two-dimensional layers are repeatedly printed until a three-dimensional object is formed. For example, inkjet printing or polyjet printing of dosage forms such as those disclosed herein can use additive manufacturing. PAM technique involves the deposition of a soft material (semi-solid or viscous) via a syringe-based print head. The syringe is typically loaded with material, which is then extruded using pneumatic pressure, a plunger, or a screw. FDM technique is based on the extrusion of thermoplastic polymers driven by a gear system through a heated nozzle tip. The printhead consists of a pinch roller mechanism, a liquefaction unit block, a nozzle, and a gantry system that controls the x and y directions. The filament is supplied, melted in the liquefaction unit, and softened into a solid state. The solid portion of the filament is used as a plunger to extrude the molten material through the nozzle tip (Sadia et al., 2016). As layers of thermoplastic molten material are deposited, the construction platform is lowered, and the process is repeated to build the structure layer by layer.

[0032] Furthermore, the present invention also encompasses methods for manufacturing implantable drug delivery devices, particularly 3D-printed implantable drug delivery devices, and this method is a. The process of loading the filament described herein into the print head of a 3D printer using a temperature higher than the glass transition temperature. b. A step of heating the construction platform at a temperature lower than the glass transition temperature of the polymer matrix. c. A step of depositing heated filament through a nozzle to construct a device from at least a first layer to a final top layer.

[0033] In relation to the present invention, the active ingredient as a whole is an antibody. The antibody may be any antibody as defined in the definition section above. The antibody is preferably present in the preliminary liquid formulation before drying at a concentration of 50 mg / mL or about 50 mg / mL to 300 mg / mL or about 300 mg / mL, preferably 65 mg / mL or about 65 mg / mL to 250 mg / mL or about 250 mg / mL, more preferably 80 mg / mL or about 80 mg / mL to 200 mg / mL or about 200 mg / mL, for example, 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 may be present in the preliminary liquid formulation before drying at a concentration of approximately 5 to approximately 30% w / v, preferably approximately 6.5 to approximately 25% w / v, or more preferably approximately 8 to approximately 20%, for example, 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 therefore in the final implantable drug delivery device, is preferably about 15–40% (w / w), or about 15–35% (w / w), for example, 15, 16, 17, 18, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, or 35% (w / w).

[0034] When at least one stabilizer is used as a whole in connection with the present invention, 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 (e.g., L-arginine, L-leucine, L-phenylalanine, or L-proline), or any combination thereof. When two or more stabilizers are used, the combination of stabilizers may be, for example (but not limited to), one amino acid and one disaccharide, or an amino acid and a polyol. As an example, a combination of two stabilizers may be used, one of which is either sucrose or trehalose, and the other stabilizer may be L-arginine, L-leucine, L-phenylalanine, or L-proline. At least one stabilizer is preferably present in the preliminary liquid formulation before drying at a concentration of about 10 mg / mL to about 100 mg / mL, preferably about 20 mg / mL to about 75 mg / mL, more preferably about 30 mg / mL to about 50 mg / mL, for example, 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 may be present in the preliminary liquid formulation before drying at a concentration of about 1 to about 10% w / v, preferably about 2 to about 7.5% w / v, or more preferably about 3 to about 5%, for example, 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.

[0035] According to the present invention, in the overall presence of at least one stabilizer, the antibody:stabilizer(s)(w / w) ratio (or alternatively referenced ratio(w / w) antibody:at least one stabilizer) in the filament and implantable drug delivery device is preferably about 1:1 to about 5:1 (weight / weight, i.e., w / w), more preferably about 1.2:1 to about 4:1, and even more preferably about 1.25:1 to 3:1, for example 1.25:1, 1.5:1, 1.75:1, 2.0:1, 2.25:1 and 2.5:1 (w / w).

[0036] According to the present invention, if a buffer is present throughout the formulation, the buffer may include, but is not limited to, phosphates, acetates, citrates, arginine, trisaminomethane (TRIS), and histidine. The buffer is preferably present in the preliminary liquid formulation before drying in an amount of about 5 mM to about 100 mM, more preferably about 10 mM to about 50 mM, for example, about 10, 15, 20, 25, 30, 35, 40, 45, or 50 mM.

[0037] In connection with the entire disclosure, surfactants may also be present. Such surfactants may be, for example (but not limited to), polysorbate 20 (PS20) or polysorbate 80 (PS80). If present, the surfactant is preferably added in the preliminary liquid formulation, i.e., before the drying step. The surfactant is preferably present in the preliminary liquid formulation before drying in an amount of about 0.01 to about 5 mg / mL, more preferably about 0.01 to about 1 mg / mL, and more specifically about 0.1 to about 0.6 mg / mL, for example, 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 as weight % (%w / v) per 100 mL. In such cases, the polysorbate surfactant contained in the formulation according to the present invention as a whole may be present in an amount of 0.001 to 0.5% w / v, preferably 0.01 to 0.1% w / v, more preferably 0.01 to 0.06% w / v, for example, 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.

[0038] In connection with the present invention, particularly when referring to filaments or the final implantable drug delivery device, any at least one stabilizer, buffer and surfactant is reclassified as a collective term for excipients. Where present, excipients are preferably present in the filament and therefore in the final implantable drug delivery device in a total amount of about 3 to about 20% w / w or about w / w, preferably about 5 to 15% w / w, for example, 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% by weight.

[0039] In relation to the present invention, the polymer material as a whole is preferably a biodegradable, biocompatible, and / or bioexcludable 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 a combination thereof, such as ethylene vinyl acetate (EVA), poly(lactic acid-co-glycolic acid) (PLGA), or poly(L-lactide-co-caprolactone-co-glycolide) (PLGA-PCL), but is not limited thereto. The polymer material can have a controlled size of about 200 Da to about 50 kDa, preferably about 500 Da to about 40 kDa, and more preferably about 1 kDa to about 20 kDa, for example, about 1, 2, 5, 10, 15, or 20 kDa. Alternatively, instead of having a given size (±), the polymer material may be a mixture of polymers of different sizes, for example, 5 kDa to 20 kDa or 7 kDa to 17 kDa. For example, some commercially available polymers are mixtures of polymers of different sizes, such as Resomer® RG502, which has a mixture of polymers in the range of 7 to 17 kDa. Preferably, the polymer material is present in the filament and therefore in the final implantable drug delivery device in an amount of about 50-75% (w / w), or about 55-70% (w / w), for example, 55, 56, 57, 58, 59, 60, 61, 62, 63, 64, 65, 66, 67, 68, 69, or 70%.

[0040] In connection with the present invention, the plasticizer as a whole is preferably polyethylene glycol (PEG) or a PEG compound, for example, but not limited to maleimide monomethoxyPEG, activated PEG polypropylene glycol, or methoxypoly(ethylene glycol) polymer. The PEG compound according to the present invention may also be of the following types of charged polymers or neutral polymers: dextran, colomic acid, or other carbohydrate polymers, amino acid polymers, and biotin and other affinity reagent derivatives. In connection with the present invention, PEG or a PEG compound may be linear or branched. In connection with the present invention, PEG or a PEG compound may have a size of about 200 Da to about 50 kDa, preferably about 500 Da to about 40 kDa, and more preferably about 1 kDa to about 20 kDa, for example, about 1, 2, 5, 10, 15, or 20 kDa. Preferably, the plasticizer is present in the filament and therefore in the final implantable drug delivery device in an amount of about 2-20% (w / w), or preferably about 5-15% (w / w), for example, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, or 15% (w / w).

[0041] In either case, it is understood that the sum of the proportions of all components of the filament will therefore reach 100% in the final implantable drug delivery device.

[0042] In relation to the entire disclosure, implantable drug delivery devices are printed using layer thicknesses of approximately 50 μm to approximately 500 μm, preferably approximately 100 μm to approximately 400 μm, for example, 100, 125, 150, 175, 200, 225, 250, 275, 300, 325, 350, 375, or 400 μm. Implantable drug delivery devices can be designed with fillers ranging from 0 (hollow body) to 100% (completely solid body). In one embodiment, the implantable drug delivery device includes at least one internal hollow cavity. In another embodiment, the implantable drug delivery device is a completely solid body.

[0043] In further embodiments, the present invention relates to a method for manufacturing an implantable drug delivery device according to the present invention, the method being: i. Cutting the filament described herein to an appropriate length. ii. Forming the filaments described herein into a suitable delivery device, iii. Pelleting the filaments described herein into a suitable delivery device, or iv. This includes grinding the filaments described herein to obtain a powder having an appropriate particle size distribution. If necessary, this powder can be coated in the future to modify its wettability and better control the release rate of the active ingredient. The obtained powder may be compressed or incorporated into classic drug formulations such as capsules.

[0044] A non-limiting exemplary filament according to the present invention comprises about 15.5% w / w of antibody (e.g., a molecule containing a full-length monoclonal antibody or Fab fragment), about 7.5% w / w of excipient, about 69.5% w / w of polymer material (e.g., RG502), and about 7.5% w / w of plasticizer (e.g., PEG), wherein the excipient comprises or consists of histidine (used as a buffer in the initial liquid formulation) and one disaccharide (either sucrose or trehalose) as a stabilizer. Another non-limiting exemplary filament according to the present invention comprises about 15.5% w / w of antibody (e.g., a molecule containing a full-length monoclonal antibody or Fab fragment), about 7.5% w / w of excipient, about 69.5% w / w of polymer material (RG502), and about 7.5% w / w of plasticizer (PEG), wherein the excipient comprises or consists of histidine (used as a buffer in the initial liquid formulation), one disaccharide (either sucrose or trehalose) as a stabilizer, and one amino acid (L-leucine).

[0045] Preferably, the filament or device of the present invention retains at least 60% of its antibody biological activity at the time of formulation and / or packaging for several weeks after implantation into the subject to be treated. The activity may be measured, preferably during preliminary experiments, as described in the following section “Examples,” or by any other standard technique.

[0046] The present invention also provides a product for pharmaceutical or veterinary use, comprising a container containing either the filament or implantable drug delivery device described above. Packaging materials providing instructions for use are also described.

[0047] The filament or implantable drug delivery device of the present invention can be stored for at least about 12 to about 24 months before use. Under preferred storage conditions, before first use, the formulation should be kept away from bright light (preferably in the dark) at a temperature of about 2 to 18°C, for example, 18°C, 15°C, or 2 to 8°C. Those skilled in the art will understand that, depending on the Tg of the polymer, the storage temperature may be higher than 18°C, for example, up to 25°C (for example, 20°C, 22°C, or 25°C).

[0048] The present invention provides single-use filaments and implantable drug delivery devices suitable for pharmaceutical or veterinary applications. [Brief explanation of the drawing]

[0049] [Figure 1] This figure shows the process for obtaining filaments and 3DP devices from preliquid (BE) and spray-dried (SD) compositions. [Figure 2] Comparison of HMWS levels of mAb1 formulations containing sucrose (SUC), trehalose (TRE), hydroxypropyl-beta-cyclodextrin (HP-β-CD), sorbitol (SOR), and inulin (INU) (mAb:stabilizer ratio 2.0:1) after buffer exchange (BE), spray drying (SD), and hot melt extrusion (HME). [Figure 3] Comparison of HMWS levels of mAb1 formulations containing sucrose (SUC), trehalose (TRE), sucrose-leucine association (SUC-LEU), and trehalose-leucine association (TRE-LEU) (mAb:stabilizer ratio 2.0:1) after buffer exchange (BE), spray drying (SD), hot melt extrusion (HME), and 3D printing (3DP). [Figure 4](a) The dissolution profile of a 3DP device containing mAb1 stabilized with TRE-LEU (3DP_7; solid line) and the change in the in vitro pH value of the surrounding medium over dissolution time are shown in the dissolution chart (dashed line). (b) Degradation of PLGA contained in the 3DP device over 10 weeks in a dissolution medium at 37°C. [Figure 5] Comparison of monomer, HMWS, and LMWS levels (%) of mAb1 released from 3DP_7 during in vitro dissolution tests. mAb1 reference was characterized at 97.4±0.4% (monomer), 2.6±0.4% (HMWS), and without LMWS. [Figure 6] Comparison of the binding ability of mAb1 released from 3DP_7 24 hours, 5 weeks, 10 weeks, and 15 weeks after dissolution. [Figure 7] In vitro release profiles of 3DP devices containing mAb1 stabilized by TRE-LEU meetings (3DP_42 (10% filler), 3DP_43 (50% filler), 3DP_44 (100% filler)). Devices were printed with a layer thickness of 0.3 mm. [Figure 8] Comparison of HMWS levels of fAb2 formulations (Fab:stabilizer ratio 2.0:1) containing SUC, SUC-LEU, TRE, and TRE-LEU after SD, HME, and 3DP. The fAb2 reference was characterized by monomer content and HMWS level of 99.6±0.2% and 0.4±0.2%, respectively. [Figure 9] Dissolution profiles of 3DP DDS containing fAb2 stabilized with SUC (3DP_F1), SUC-LEU (3DP_F2), TRE (3DP_F3), and TRE-LEU formulations (3DP_F4). [Figure 10] Comparison over time (8 weeks) of monomer content (a) and HMWS level (b) of fAb2 released from 3DP_F1, 3DP_F2, 3DP_F3, and 3DP_F4. [Figure 11] The binding ability of fAb2 released from 3DP_F1, 3DP_F2, 3DP_F3, and 3DP_F4 24 hours after dissolution.

[0050] example Abbreviation: HMWS = High molecular weight species; LMWS = Low molecular weight species; SD = Spray-drying or spray-dried; Hot melt extrusion 3DP = 3D printing or 3D printing; BE: Buffer exchange; DDS: Drug delivery system; DDD: Drug delivery device; DSC: Differential scanning calorimetry; FDM: Molten deposition modeling; HME: Hot melt extrusion; LEU: L-leucine; Mw: Molecular weight; mAb: Full-length monoclonal antibody; fAb: Fab fragment of antibody; PBS: Phosphate buffer; 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.

[0051] 1.Material mAb1 is IgG4, with a molecular weight (MW) of approximately 150 kDa and a pI of approximately 6.0–6.3. fAb2 is the Fab portion of the antibody. fAb2 has a MW of approximately 50 kDa and a pI of approximately 9.3–9.6.

[0052] 2. Method 2.1. Spray drying The antibody-containing solutions were spray-dried using a laboratory-scale Spray-Dryer B-290 (Buhi Labertechnik) equipped with a 0.7 mm nozzle. Settings were kept constant for all formulations, following standard procedures. Solutions for spray-drying were pre-prepared in 15 mM histidine buffer at pH 5.6, along with other excipients as needed. Tables 1 and 8 summarize the composition, concentration, and mAb:stabilizer ratio of the mAb1 and fAb2 solutions. All powders were sealed in polypropylene containers and stored in a desiccator under vacuum.

[0053] 2.2. Hot Melt Extrusion The printable filament was prepared from a physical mixture of raw materials PLGA, PEG 2kDa, and mAb1 or fAb2-containing spray-dried (SD) powder, pre-blended using a Turbula® mixer (Willy A. Bachofen AG). The mixture was manually fed into an 11mm twin-screw extruder (Process-11, Thermo Fischer Scientific) equipped with a modular screw (L / D ratio 40:1) and a 1.6mm diameter round die. The barrel was heated using a temperature gradient controlled by eight thermocouples. A water circulator was used to maintain the supply zone at room temperature. The three first segments were set to 20, 40, and 80°C, respectively. The intermediate segments of the fourth and sixth thermocouples were set to 90°C. The last thermocouple, located directly before the die, was set to 85°C, and the die itself was set to 75°C. For all experiments, the screw speed was set to 40 rpm during feeding and 60 rpm when manually winding the filament. These parameters were kept constant (see Table 1).

[0054] 2.3. 3D Printing of Antibody-Loaded Devices The device design was drawn using 3D computer-aided design (CAD) software ThinkerCAD (AutoDesk® Inc.) and exported to software for slicing. The device dimensions are 178.43 mm² in volume. 3 The dimensions were 20 × 5 × 2 mm (length, width, height). mAb1 and fAb2-supported devices were printed using a Hyrel 3D System 30M printer (GA) equipped with a 0.5 mm MK2-250 hot extruder. It was not necessary to control the temperature of the construction platform. The printing temperature was set to 105 ± 2°C. The printing speed was 1 mm / s for the first layer and 10 mm / s for the subsequent layers. The device layer thickness was set to 0.1 mm and 0.3 mm to evaluate the potential decomposition of the supported mAb1 and its impact on its emission profile. Device printing was performed using 100% (v / v) filler unless otherwise specified in the following examples.

[0055] 2.4.Analysis method Differential Scanning Calorimetry (DSC): Thermal analysis of SD powder, filament, and 3DP DDS was performed by DSC using a heat flux type DSC Q2000 (TA instrument) equipped with a cooling system, following standard procedures.

[0056] Thermogravimetric Analysis (TGA): TGA was performed using a Q500 TGA (TA instrument) equipped with a balance with a sensitivity of 0.1 μg, following standard procedures. Data acquisition and analysis were performed using TA Instruments® Trios 4.5.0 software.

[0057] Molecular weight analysis of polyesters by size exclusion chromatography (SEC) in chloroform: The number-average molecular weight (Mn), weight-average molecular weight (Mw), and polydispersity index (Mw / Mn) of polyesters were measured by SEC according to standard methods. Relative molecular weight (number and weight-average) and polydispersity index were calculated by referring to a polystyrene standard calibration curve established using the same experimental conditions. The mean and standard deviation (STD) for molecular weight and polydispersity were calculated as detailed above for NMR analysis.

[0058] Antibody Stability Assessment: mAb1 monomer quantification, as well as the evaluation of both HMWS and LMWS content, was performed by size exclusion high-performance liquid chromatography (HPL). This analysis was performed on samples obtained either after dissolution testing or after extraction from printable filaments and 3DP devices. These quantifications were performed using an Agilent 1200 series LC system with a UV detector (Agilent Technologies) according to a standard protocol. The mobile phase was a 0.2 M PBS solution at pH 7.0. The calibration curve for mAb1 ranged from 20 to 2000 μg / mL. mAb1 stability was assessed using the monomer loss percentage, which corresponds to the difference in monomer percentage before and after both the HME and 3DP processes. Monomer, HMWS, and LMWS levels (%) were compared to a baseline consisting of the mAb1 solution obtained after buffer exchange. A similar method was used for fAb2 stability assessment.

[0059] Antibody extraction from polymer matrix: To evaluate the stability of mAb1 melt-encapsulated in both printable filaments and 3D printing devices, approximately 10 mg of the sample was placed in a Nanosep® equipped with a 0.2 μm Bio-Inert centrifuge (Pall) and dissolved in 0.5 mL of dichloromethane. PLGA was dissolved in a Nanosep® device by stirring at 600 rpm for 2 hours at room temperature using a Thermomixer Confort® tube mixer (Eppendorf AG). The sample was centrifuged at 12000 rpm for 10 minutes, and the medium was removed. Then, 0.5 mL of dichloromethane was added again. The sample was stirred for 5 minutes and centrifuged as described above. This process was repeated twice. After removing the dichloromethane, the Nanosep® device containing the mAb precipitate was placed under vacuum for 1 hour to remove any potential residual solvent. Then, 0.5 mL of PBS (0.2 M, pH 7.0) containing 0.02% w / w polysorbate 80 (PS80) was added to the tube to solubilize mAb1, and the mixture was stirred at 600 rpm for 2 hours. Then, the Nanosep® device was placed under vacuum for 12000 rpm (Arrighi et (Adapted from al., 2019) The cells were centrifuged for 10 minutes. mAb1 stability was evaluated by SEC (as described above). A similar method was used for fAb2 extraction.

[0060] Antibody loading after melt encapsulation: The amount of encapsulated mAb1 in the PLGA matrix was determined using colorimetric detection by a standard bicinchoninic acid (BCA) protein assay. The amount of melt-encapsulated mAb1 was determined by performing the Pierce® microplate procedure. Quantification of both standards and samples was performed at room temperature using a SpectraMax M5 microplate reader (Molecular Devices) at 562 nm. Overall, the mAb1 loading was determined as follows: mAb loading rate (%) = (Amount of melt-encapsulated mAb) / (Amount of 3DP device) × 100.

[0061] A similar method was used to support fAb2.

[0062] Dissolution Test: An in vitro dissolution test was performed to evaluate the release profile of mAb1 / fAb2 supported from the 3DP DDS. Approximately 200 mg of the 3DP device was placed in a 5 mL Eppendorf® tube filled with 5 mL of PBS (0.2 M, pH 7.0, 37°C) and mixed at 600 rpm using a Thermomixer Confort® tube mixer (Eppendorf AG) (adapted from Marquette et al., 2014). After a specified time, 5 mL of the medium was withdrawn, collected, and filtered through a 0.45 μm PVDF Acrodisc® syringe filter (Pall). A similar volume was replaced with fresh buffer (5 mL). The filtered solution was analyzed for pH using SEC analysis with a 280 nm UV detector.

[0063] PLGA degradation during dissolution: The decrease in polymer molecular weight (Mw) of PLGA during drug release was investigated using gel permeation chromatography (GPC). The protocol was the same as that used for the dissolution test. Mw was calculated using a polystyrene standard.

[0064] Enzyme-linked immunosorbent assay (ELISA): The binding ability of mAb1 / fAb2 was evaluated using the ELISA assay according to standard procedures.

[0065] Data Analysis: Unless otherwise specified, all experiments were conducted in triplicate. Statistical analysis was performed using Prism 8 software (GraphPad software). Results are expressed as mean ± standard deviation. Statistical significance was determined using ANOVA and Turkey or Dunnett post-hoc tests (recommended by Prism software) with a p-value < 0.05.

[0066] Example 1 - Preparation of printable filament and 3DP DDS supporting mAb1 mAb1 solutions were formulated with different stabilizers (see Table 1). These liquid solutions were spray-dried to produce mAb1-supported powders. In fact, mAb1 was used in a solid state to enhance its stability and facilitate handling during further processing. Next, a mixture of mAb1-supported powder, Resomer® RG502 (Evonik Industries) as a polymer material, and PEG as a plasticizer was extruded using HME to produce printable filaments. These printable filaments were fed into a 3DP printer to print devices (or, as herein referred to, drug delivery devices or implantable drug delivery devices). The optimal formulation was identified by evaluating mAb1 integrity after each manufacturing step (SD, HME, 3DP). Finally, in vitro evaluations (solubility tests and binding capacity) were performed.

[0067] Example 2 - Preliminary study on raw materials and printable filaments containing mAb1 The thermal properties of all raw materials (including their decomposition temperatures) were evaluated using TGA and DSC analysis, respectively.

[0068] The decomposition temperature of the raw material RG502 was approximately 175°C. No significant weight loss was observed at 200°C in extruded filaments supporting the raw materials PEG and mAb1. No residual moisture was observed in the RG502 and PEG raw materials. These results confirmed that all raw materials appeared stable and could be processed according to both HME and 3DP temperatures (90°C and 105°C, respectively). In fact, it was necessary to characterize only the mass loss using TGA and to describe mAb1 stability using other methods such as SEC and binding capacity.

[0069] The TGA thermogram of the SD mAb1 powder showed a slight weight loss (approximately 4% w / w) when it reached a temperature of 100°C. This loss may be due to the residual moisture content (approximately 3.4 ± 0.8%) in the SD mAb1 powder. A second weight loss was observed above 150°C in all SD mAb1-supported powders. Therefore, the mAb1-supported powders were able to ensure the stability of mAb1 during both HME and 3DP.

[0070] Next, DSC analysis was performed to determine the T of the thermoplastic polymer RG502. g The effects of adding PEG and mAb1-supported SD powder were evaluated. In fact, the purpose of this study was to develop an mAb1-supported 3DP DDS, and T g The temperature must be as low as possible to allow for temperature reduction in the different processes (HME, 3DP) and the resulting potential degradation of the biopharmaceutical.

[0071] Table 1. Composition of evaluated mAb1 formulations: liquid composition after buffer exchange and before spray drying (SD) (expressed as %w / v), solid composition of spray-dried powder (expressed as %w / w), printable filament produced using hot-melt extrusion (HME) batches (expressed as %w / w), and associated 3D printing (3DP) batches with layer thicknesses of 0.1 mm and 0.3 mm. mAb1 was present at 8% w / v in the initial liquid composition. [Table 1]

[0072] RG502 T g The temperature was found to be 38.0 ± 0.7°C, which was consistent with data already reported in the literature (Pignatello et al., 2009). PEG was characterized by a sharp endothermic peak at 52°C. The T of RG502 g When PEG and SD powder were added to HME, the temperature decreased to 21.8±0.4℃ (data not shown). In addition to the loss of the sharp melting peak of PEG, T gThis decrease demonstrated that the mAb1-supported SD powder and PEG were properly dispersed in the molten polymer matrix (Zhang et al., 2017).

[0073] Example 3 - Formulation screening and mAb1 stability after spray drying process Stabilizers were selected to maintain antibody integrity throughout the entire manufacturing process. The main anticipated adverse factor was the relatively high temperature used during both HME and 3DP. Unfortunately, the choice of stabilizer is not universal and must be adapted to each biotherapy drug and to the stress factors associated with the process (Le Basle et al., 2020; Wang et al., 2007). SUC, TRE, HP-β-CD, SOR, and INU are commonly used in antibody-containing formulations (Baek et al., 2017; Bowen et al., 2013; Gidwani and Vyas, 2015; Kanojia et al., 2016; Maury et al., 2005). The effect of adding stabilizers on the stability of supported mAb1 was investigated using three different mAb:stabilizer ratios (w / w) (1.5:1, 2.0:1, and 2.5:1) (see formulations in Table 1). To enhance the stability of mAb1 during the SD process, an mAb:stabilizer ratio (w / w) of 2.0:1 has been previously described (Bowen et al., 2013). To evaluate the effects of these ratios on the stability of mAb1 not only during SD, but more specifically during HME and 3DP (two processes that result in high thermal stress), the inventors also investigated higher and lower ratios.

[0074] The different liquid compositions to be evaluated were obtained by buffer exchange. No instability was observed between the mAb1 reference (before buffer exchange) and the various liquid compositions (after buffer exchange, BE). The percentage of HMWS was very similar to that observed from the mAb1 reference (2.6 ± 0.4%) (Table 2). After SD, regardless of the properties of the stabilizer (p > 0.05), no significant formation of HMWS was observed for either the 1.5:1 or 2.0:1 mAb:stabilizer ratio (Figure 2). In contrast, when using the 2.5:1 ratio, the percentage of HMWS increased, except for SUC and TRE, regardless of the properties of the stabilizer (p > 0.05) (Table 2). LMWS levels were also evaluated, and no fragmentation was observed in the raw material mAb1 solution. Similar observations were made after BE and SD, regardless of the mAb:stabilizer ratio (Table 2). Since the ratios of 1.5:1 and 2.0:1 yielded similar results, the ratio of 2.0:1, which allows for a higher proportion of mAb1 relative to the stabilizer, was selected for further investigation.

[0075] Table 2. Comparison of HMWS and LMWS levels of mAb1 formulations (mAb:stabilizer ratio: 1.5:1; 2.0:1 and 2.5:1; see Table 1 for formulations) after buffer exchange (BE), spray drying (SD), and hot melt extrusion (HME). All HMWS and LMWS values ​​are expressed as percentages. [Table 2]

[0076] Example 4 - Extrusion of mAb1-supported printable filament To obtain filaments, mAb1-supported SD powder was mixed with PLGA and PEG, and extruded (HME) to produce printable filaments (see formulations in Table 1). The filaments were successfully prepared with a diameter of 1.70 mm to 1.75 mm, as recommended for supplying to FDM 3D printers (Melocchi et al., 2015). An mAb1 loading rate of 15% (w / w) was selected.

[0077] As shown in Table 2 and Figure 2, the proportion of HMWS increased due to the use of relatively high temperatures, regardless of the properties of the stabilizer (p < 0.0001). When HP-β-CD, SOR, and INU were added to the formulation (mAb:stabilizer ratio 2.0:1), the proportion of HMWS reached 6.4 ± 0.2%, 11.2 ± 0.5%, and 4.9 ± 0.1%, respectively (see Figure 2). In contrast, SUC and TRE appeared to be the most suitable for stabilizing mAb1 during the HME process performed at 90°C. In fact, the proportion of HMWS increased to only 3.3 ± 0.3% and 3.8 ± 0.5%, respectively (Figure 2). No significant differences were highlighted for either disaccharide after the HME process (p > 0.05).

[0078] The proportion of LMWS was also evaluated after HME (see Table 2). When HP-β-CD and SOR were used as stabilizers, slight fragmentation was observed. In contrast, LMWS was not observed with SUC, TRE, and INU.

[0079] Overall, HP-β-CD, SOR, and INU were less effective than SUC and TRE in maintaining mAb1 stability in HME. Based on evaluations of HMWS and LMWS levels, TRE and SUC were used as stabilizers to ensure the integrity of mAb1 in HME.

[0080] Examples 2 and 3 showed that SUC and TRE appear to be the most suitable stabilizers for stabilizing the formulation throughout the continuous manufacturing process (after SD and HME).

[0081] Finally, prior to the printing process, the mAb1 loading rate of the printable filaments was evaluated. This showed that the actual loading rate of all filaments was similar to the target loading rate (15% w / w) with a very low standard deviation (Table 3). These results demonstrate that the manufacturing process is suitable and reproducible for producing uniform printable filaments with a uniform dispersion.

[0082] Table 3. mAb1 loading rates in printable filaments and 3DP devices obtained by BCA assay. [Table 3]

[0083] Example 5 - 3D printing of mAb1 implantable delivery devices Using slicing software, a model of an embeddable 3DP device with an embeddable shape was designed. The printing process was carried out in a room at 20°C. In fact, the physical state of the filament was its T g As previously mentioned, the temperature was approximately 22°C, which can change rapidly due to room temperature. Therefore, at 20°C, the filament's rigidity was maintained, allowing for printing. However, handling the filament induced heat transfer by conduction. This phenomenon was more pronounced when the filament was loaded into the print head. In fact, they were too soft to move along the feed gear. To limit heat transfer by conduction during printing, 3DP had to be performed using the "flexible high-temperature flow" modular print head MKE-250.

[0084] The device resolution was macroscopically evaluated, and a completely solid device was expected when the filler was set to 100%. Immediate visualization showed defects and material absence on the top of the device (data not shown). The printing process was performed at 105°C, a temperature that promotes both adhesion to the construction platform and adhesion between continuous layers. The printing speed was selected to improve the resolution of the DDS, with 1 mm / s for the first layer and 10 mm / s for subsequent layers. 3D printing with layer thicknesses of 0.1 mm and 0.3 mm was evaluated.

[0085] mAb1 extraction was performed using a 3DP device, and the proportions of both HMWS and LMWS were evaluated. The proportion of HMWS increased after 3DP, regardless of layer height and disaccharide properties (Figure 3). However, this was significantly higher when a layer thickness of 0.1 mm was used (p < 0.0001 and p < 0.0004). For example, the proportion of HMWS increased from 3.3 ± 0.1% (formulation HME_16) and 3.8 ± 0.1% (formulation HME_18) after HME, respectively, when SUC and TRE were used, to 4.7 ± 0.3% (formulation 3DP_2) and 4.8 ± 0.1% (formulation 3DP_5) after 3DP with a layer thickness of 0.3 mm, or to 6.14 ± 0.1% (formulation 3DP_1) and 6.2 ± 0.1% (formulation 3DP_4) after 3DP with a layer thickness of 0.1 mm. This was attributed to the relatively high temperature used during 3DP. This can be explained by the slower movement of the construction platform, which has resulted in an increased contact area between the printer nozzle and the printing device (Carlier et al., 2019).

[0086] Despite the addition of SUC or TRE, a significant (acceptable) increase in HMWS was observed after 3D printing. Therefore, it was hypothesized that the addition of hydrophobic amino acids such as LEU (Minne et al., 2008) could enhance the stability of the supported mAb1. Starting from a preliminary liquid formulation containing a combination of stabilizers SUC-LEU or TRE-LEU, 3DP devices were printed using a layer thickness of 0.3 mm (see Table 1).

[0087] HMWS levels were assessed after each process (from SD to 3DP, with the starting value being BE) (see Figure 3). After 3DP, these levels were 4.4±0.2% and 3.6±0.1% for 3DP_3 and 3DP_6, respectively. These levels were compared to those obtained when SUC and TRE were used alone. It was demonstrated that the addition of LEU to SUC and TRE could limit HMWS production. The reduction in HMWS was significant with TRE-LEU association (p<0.0001). The increase rate was compared as a ratio over the entire process (between 3DP and SD). After adding LEU to TRE, the percentage of HMWS increased by approximately 18% compared to 50% when TRE was formulated alone. The same trend was observed when LEU was added to SUC at higher HMWS levels (SUC-LEU: 33% increase vs. SUC: 66% increase).

[0088] LMWS levels were also investigated after 3DP. Regardless of the addition of LEU to SUC or TRE, a slight increase in LMWS (approximately 0.05 ± 0.04%) was observed (data not shown).

[0089] Finally, drug loading was evaluated using 3DP DDS, and the BCA results showed an actual loading rate close to the target loading rate of 15% (w / w) (see Table 4). These results confirm the uniform dispersion of mAb1 in the polymer matrix expressed after HME.

[0090] Table 4. mAb1 loading rates in printable filaments and 3DP devices obtained by BCA assay. [Table 4]

[0091] Therefore, our surprising discovery was that 1) it was possible to extrude a printable filament by HME starting from an SD supporting mAb1, and 2) it was possible to fabricate a 3DP device by FDM using the filament. The increase in HMWS observed mainly after HME and 3DP was directly related to thermal degradation occurring at 90°C (during HME) and 105°C (during 3DP). We further investigated the most promising formulations containing TRE-LEU and less SUC-LEU that could minimize HMWS production and promote mAb1 stability.

[0092] Example 6 - Dissolution test in a 3D printed device containing mAb1 It has been previously demonstrated and described that PLGA-based drug delivery systems (DDS; e.g., microparticles and implants) feature a triphasic release profile. Facilitating a release profile with a limited latent phase may be of greater interest. Indeed, the latent phase may lead to mAb1 degradation due to its retention in the polymer matrix and uptake into the culture medium. Furthermore, a linear release profile, which may tend toward "zero-order kinetics," should allow for constant drug release of mAb1 in the dissolution medium and a stable release concentration.

[0093] As shown in Figure 4a, the release of mAb1 from the 3D printed device was characterized by a low burst effect of 2.0 ± 0.3% within 24 hours. Sustained release occurred over time, starting with a sustained-release phase (latent period) within the first week. Weeks 1–4 actually showed low antibody release up to 10.6 ± 1.9%. This is because the culture medium struggled to penetrate the PLGA matrix, and is known to be low during the first week. Subsequently, an increase in the rate of mAb1 release was observed in the following weeks. Cumulative release accelerated, increasing from 17.3 ± 2.8% at 5 weeks to 57.8 ± 2.5% at 12 weeks. Finally, the low-release phase was observed to reach 59.7 ± 2.3% at 15 weeks. The release of mAb1 depended on water uptake, which allowed mAb1 to diffuse through the pores of the device.

[0094] The degradation of polymer RG502 was evaluated in a 3D printed device during the dissolution test (Figure 4b). Diffusion of the medium through the polymer matrix is ​​necessary to induce hydrolysis and accelerate the erosion of the DDS. Resomer® RG502, a PLGA derivative, was characterized with an initial Mw of 17867 ± 577 g / mol. RG502 hydration occurred during the first week of the dissolution test. Slight degradation of the polymer was observed, and the pH value of the surrounding medium remained constant (Figure 4a). Subsequently, degradation increased after 3 weeks with a loss of approximately 20% of its initial mass (14367 ± 462 g / mol) (loss due to hydrolytic cleavage of RG502 in the oligomer into the device). Erosion began after 3 weeks, following the decreased pH value of the surrounding medium (Figure 4a). During the first week, degradation occurred primarily, but the initiation of erosion was triggered and accelerated with decreasing pH. Thus, autocatalysis accelerated erosion and increased both PLGA degradation and mAb1 release. For example, a loss of 64% of its initial mass (5373 ± 1217 g / mol) was observed after 7 weeks of dissolution (Figure 4b). Interestingly, the pH value decreased to 6.3 ± 0.1%, which indicated the highest rate of erosion. After this major degradation, no further degradation was reported, and the polymer's Mw remained stable at approximately 6000 g / mol (Figure 4b). Furthermore, the rate of erosion decreased after 7 weeks. This description was supported by the increase in the pH value in the following week from 6.7 ± 0.1% (week 8) to 7.0 ± 0.1% (Figure 4a) after 15 weeks. Such a profile was consistent with expectations (Cosse et al., 2016; Ghalanbor et al., 2013).

[0095] The release of mAb1 was evaluated over 15 weeks. The pH value remained slightly acidic due to oligomer formation and their diffusion into the dissolution medium. After 10 weeks, no further degradation of PLGA or further release of mAb1 was observed. Since the sample generated in the dissolution medium after 10 weeks remained insoluble in chloroform, PLGA and mAb1 may form insoluble aggregates over time.

[0096] To study the stability of mAb1 during dissolution, HMWS and LMWS levels, as well as monomer content, were evaluated by dissolution tests (Figure 5). A decrease in monomer proportion was shown to be associated with an increase in either HMWS or LMWS species. The highest HMWS levels were observed between week 6 (25.4±3.6%) and week 8 (25.9±3.1%). This increase correlated with the aforementioned highest erosion rate and the decrease in pH to 6.3±0.1 at week 7. Interestingly, a slight increase in LMWS was observed during the first nine weeks of dissolution (<0.7%). LMWS levels increased to 17.0±5.7% after 10 weeks. This level remained high at 15.4±5.2% after 14 weeks. Fragmentation was observed in the delayed phase of the dissolution tests. This may be due to hydration of the PLGA-based device core that occurred after the main erosion of the matrix. Therefore, the decrease in pH, coupled with the complexity of extracting mAb1 from the core, appeared to be more detrimental than during the primary erosion process. Monomer content was 96.5±0.3% after 24 hours (burst effect), and then decreased to 74.1±3.6% and 64.6±3.3% after 6 and 12 weeks, respectively.

[0097] An ELISA assay was performed to evaluate the binding capacity of mAb1 after diffusion from the device to the lysis medium (see Figure 6). The binding capacity of mAb1 was found to be 69.0 ± 1.5% after 24 hours. A slight decrease in binding capacity was demonstrated after 5 weeks (66.2 ± 3.8%). After 10 and 15 weeks, the binding capacity decreased dramatically to 43.8 ± 6.8% and 38.8 ± 7.9%, respectively. Although the value after 24 hours was lower than expected considering the low HMWS levels and the observed high monomer content (96.5 ± 0.3%) (Figure 5), the result is very promising, indicating that despite thermal stress, mAb1 is still able to bind to its target and is therefore likely to remain active, potentially allowing for continuous release for several weeks.

[0098] Example 7: Stability study of mAb1-supported 3DP devices. Regarding temporal stability, an important aspect to consider when developing a pharmaceutical formulation, the effects of storage temperatures (T0, T1, T2, T3, and T6 months) at 5±3°C and 25±2°C for 6 months were evaluated. 3DP devices were fabricated using mAb1 stabilized with a TRE-LEU combination.

[0099] Physical state of polymer matrix: DSC analysis of 3DP devices was compared at different time points (see Table 5). As previously stated, PLGA was plasticized using 11% (w / w) PEG, and the filament Tg (before printing) was 21.8 ± 0.4°C. The Tg of the reference sample (T0) was close to this value at 20.7 ± 0.3°C. No increase in Tg was observed over 3 months according to both storage temperatures (i.e., 5°C and 25°C). However, an increase in Tg to 29.7 ± 0.3°C (T6) was observed after 6 months at 25°C. The Tg of the devices remained constant at 5°C during stability testing. Furthermore, small melting peaks were observed in samples stored at 25°C for 2 months (T2), 3 months (T3), and 6 months (T6). The melting peaks observed at Tm 45.2±1.4℃ (T2); 45.9±0.8℃ (T3) and 46.7±0.4℃ (T6) may be attributed to PEG being able to migrate at a higher temperature (i.e., 25℃) than the Tg of the polymer matrix. The enthalpy of melting for these melting peaks was recorded, showing an increase from 1.7±0.9 J / g (T2) to 7.4±0.6 J / g (T6) over several months. The increase in enthalpy of melting likely indicated phase separation with PLGA chain mobility at 25℃. After 2 and 3 months, the enthalpy of melting remained low, indicating that the plasticizing effect was effective. The increase in Tg after 6 months at 25℃ was associated with higher values ​​of enthalpy of melting consistent with phase separation between PLGA and PEG. The enthalpy of melting for net (neat) PEG was recorded at approximately 193.4 J / g (data not shown). Therefore, only small amounts of PEG tended to separate from the PLGA blend over a 6-month period. Similar observations were made during polymer aging studies. It has been reported that PEG can crystallize over time with increasing storage temperature and humidity. Crystallization of PEG may increase the rigidity of the devices and alter their mechanical and emission properties.

[0100] Table 5. Printed 3DP devices for performing stability tests identified using time points (T0, T1, T2, T3, T6) with characteristics such as Tg (°C), Tm (°C), enthalpy of fusion (J / g), and Mw (kDa). [Table 5]

[0101] The degradation of PLGA was evaluated using GPC measurements. The Mw at T0 was recorded as 17.02 ± 0.38 kDa, which was consistent with the received raw material PLGA (Mw: 17.05 ± 0.45 kDa). No degradation occurred during 3 months of storage. These results demonstrate the stability of the device when stored at 5°C and 25°C for 3 months. However, the results obtained after 6 months were obtained after 1.5 months of storage in a refrigerator, where relative humidity can affect the Mw of the polymer.

[0102] Visual evaluation of devices over storage time: Visual evaluation was performed on 3DP devices (not shown) stored at 5°C and 25°C. No difference was observed in devices stored at 5°C for 6 months. When devices were kept at 25°C, adhesive test pieces were observed. The devices adhered to the bottom of the glass vial, but no material loss was observed during the recovery process. This observation was made for all devices stored at 25°C from T1 to T6. The cross-section of device T6 at 25°C showed a highly porous network due to chain mobility. Increased device porosity was expected to result in faster release of mAb1 during dissolution studies.

[0103] Drug content and extraction from devices: The target loading rate was 15% (w / w). As shown in Table 6, the mAb1 loading rate in each device was consistent with the experimentally obtained values.

[0104] Table 6. Comparison of mAb1 loading rate (%), monomer content (%), and both HMWS and LMWS levels (%) at time points from T0 (reference) to T6 (6 months). [Table 6]

[0105] The stability of mAb1 was also evaluated at each time point (Table 6). Monomer content remained stable for 6 months at 5°C. However, a slight decrease in monomer content was observed after 6 months at 25°C. This decrease was associated with an increase in both HMWS (5.1±0.2%) and LMWS (0.08±0.01%) levels of the sample.

[0106] To investigate the release pattern as a function of the device packing, lysis studies were performed on printed devices (3DP_39~3DP_44; Table 7) (Figure 7). Burst release from all devices was limited regardless of the formulation. For example, the burst release of 3DP_42 reached 6.1±0.5% after 24 hours. A triphasic profile was observed for all devices, as previously observed. These results were consistent with previous results shown for mAb1-supported 3DP devices (Figure 4a). Finally, maximum cumulative releases of 63.2±4.7% and 62.3±5.1% for devices 3DP_43 and 3DP_44 were demonstrated after 6 weeks. As can be observed, the device packing has only a slight effect on the cumulative release of the antibody.

[0107] Table 7 - HME batch starting materials and printed associated 3DP batches (n=3) with packing densities of 10, 50, or 100% (v / v) and a layer thickness of 0.3 mm. [Table 7]

[0108] Example 8-fAb2 filament and 3DP device containing fAb2. Based on the findings from Examples 1-7, we investigated the use of TRE and SUC as stabilizers for the Fab antibody fragment fAb2, regardless of whether LEU was added.

[0109] A similar approach was applied to fAb2 using a series of processing methods, including buffer exchange (BE), spray drying (SD), hot melt extrusion (HME), and 3D printing (3DP). The fAb2 formulations were prepared in LEU-containing or LEU-free TRE or SUC to compare four different formulations and determine which could stabilize the Fab against thermal stress (Table 8), as well as to identify relevant 3D printing (3DP) batches with layer thicknesses of 0.1 mm and 0.3 mm. The fAb2 content was 8% w / w in the initial liquid composition, 66.7% w / w in the SD powder, and 15.3% w / w in the filament / 3DP device.

[0110] Characteristics of extracted fAb2: The raw material fAb2 material was characterized by a high monomer content of 99.6±0.2% and a low HMWS level of 0.4±0.2%. fAb2 extracted from the PLGA matrix after HME and 3DP was compared with fAb2 extracted from the SD powder (Figure 8). The HMWS level of the formulated Fab increased slightly during the high-temperature process (e.g., HME and 3DP). For example, the HMWS level of the formulation containing TRE-LEU developed from 0.6±0.3% (SD) to 1.0±0.1% after 3DP. All results showed that none of them differed significantly from the Fab-SD powder (p > 0.05).

[0111] Dissolution Test: Dissolution tests were performed on all printed devices (DDS; 3DP-F1 to 3DP_F4) to investigate both the release pattern and stability of Fab over time (Figure 9). Burst release from all devices was limited regardless of the formulation. For example, the burst release of 3DP_F4 reached 2.4 ± 0.2% after 24 hours. As previously observed, a triphasic profile was observed in all devices. This observation was mainly dependent on the monolithic state of the devices and the difficulty of water penetration. As a result, faster release was observed between weeks 5 and 6. These results were consistent with previous results shown for mAb1-supported 3DP devices (Figures 4a and 7). Finally, a maximum cumulative release of 79.3 ± 1.7% by device 3DP_F4 was demonstrated after 8 weeks.

[0112] Stability study of fAb2 over time: During the dissolution test, both monomer content and HMWS levels were investigated over 8 weeks (Figure 10). A slight decrease in monomer content was observed after 8 weeks. In fact, the monomer content changed from 99.3±0.1% to 97.6±1.0% in 3DP_F4 (Figure 10a). After 2 weeks of dissolution, deformation of the monomer peak on the thermogram was observed (data not shown). A shoulder appeared in the monomer peak and increased over several weeks, but no fragmentation was reported after 8 weeks regardless of the formulation (data not shown). This occurrence on the chromatogram may be related to the acidic microenvironment pH in the PLGA matrix that compromised the integrity of the Fab. An increase in HMWS levels was demonstrated over the dissolution time (Figure 10b). Aggregation of fAb2 over the dissolution weeks appeared to be limited compared to previous results produced with mAb1. For example, the value of 3DP_F4 after 8 weeks of dissolution was 1.9±0.1%. No fragmentation of fAb2 was observed after an 8-week dissolution period. FAb2 appeared to be more stable under similar conditions. In fact, aggregation of fAb2 over several weeks was very low after the high-temperature process and during the dissolution test.

[0113] Binding Ability Study: The binding ability of fAb2 was evaluated to confirm that fAb2 could still bind to its target. ELISA testing demonstrated that the binding ability of fAb2 was preserved after 24 hours of release, regardless of the formulation. For example, the binding ability of 3DP_F4 was 99.5 ± 6.4% (Figure 11). These results suggest that all formulations were adapted to allow continuous drying of the Fab, produce printable filaments, and print 3DP devices with limited Fab degradation.

[0114] Conclusion: fAb2 was sequentially dried, extruded, and 3D printed using four different formulations containing LEU-containing / uncontaining SUC or TRE. 3DP DDS enabled sustained release of Fab for at least 8 weeks. HMWS levels were very low, with a maximum of 1.9 ± 0.1% (3DP_F4). Results obtained using formulations containing TRE (+ / -LEU) as a stabilizer were slightly better in terms of total release, but SUC (+ / -LEU) was also very promising.

[0115] Overall conclusion The results presented herein, surprisingly, demonstrate for the first time that not only HME but also the association of HME with FDM 3D printing, as shown herein with monoclonal antibody (mAb1) and Fab fragment (fAb2), is suitable for producing antibody-supported filaments and antibody-supported embedded devices in which the antibody can still bind to its target (and therefore is likely still active). Homogeneous solid dispersions of the antibody in PLGA matrices were achieved in both printable filaments and 3DP devices. Various stabilizers were investigated to stabilize the antibody against thermal decomposition. The most promising ones (trehalose and sucrose) promoted mAb integrity during the SD, HME, and 3DP processes using different mAb:stabilizer ratios. Further optimization of the formulation using small amounts of amino acids such as leucine improved the antibody's stability against potential thermal decomposition. Furthermore, the lysis profiles demonstrated an interesting sustained-release profile with limited burst effects, particularly with antibody fragments. Finally, it was demonstrated that the antibody binding ability remained for at least approximately 5 weeks despite the relatively high temperatures of extrusion (90°C) and printing (105°C).

[0116] Table 8. SD batch (%w / V), solid composition of SD powder (%w / w), and yield (%) of the spray drying process. The theoretical composition of the fAb2 formulation evaluated for the process yield (%) and printed at a packing density of 100% (v / v) and a layer thickness of 0.3 mm (n=3), and the HME batch with the associated 3DP batch. Printable filament (%w / w) Excip. = Excipient [Table 8]

[0117] References

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Claims

1. A filament for preparing an implantable drug delivery device, wherein the filament comprises at least one polymer material, a plasticizer, at least two stabilizers, and an active ingredient. The active ingredient is an antibody, The at least one polymer material is poly(lactic acid-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL), poly(lactic acid) (PLA), or a combination thereof, and is in the range of 50-75% (w / w). The plasticizer is polyethylene glycol and is in the range of 2 to 20% (w / w), The first stabilizer is a disaccharide, and the second stabilizer is an amino acid, and The antibody-to-stabilizer ratio is 1:1 to 5:1 (w / w). filament.

2. The filament according to claim 1, wherein the filament further comprises a buffer and / or a surfactant.

3. The filament according to claim 1 or 2, wherein the disaccharide is sucrose or trehalose.

4. The filament according to any one of claims 1 to 3, wherein the amino acid is L-arginine, L-leucine, L-phenylalanine, or L-proline.

5. The filament according to any one of claims 1 to 4, wherein the total amount of the stabilizer is in the range of 5 to 15% (w / w).

6. The filament according to any one of claims 1 to 5, wherein the active ingredient is uniformly dispersed in the polymer matrix.

7. The filament according to any one of claims 1 to 6, wherein the loading rate of the active ingredient is in the range of 15 to 35% (w / w).

8. An implantable drug delivery device comprising or consisting of one or more layers made from a filament according to any one of claims 1 to 7.

9. A 3D-printed implantable drug delivery device obtained by 3D printing a filament according to any one of claims 1 to 8.

10. The implantable drug delivery device according to claim 8 or 9, wherein the device is printed using a layer thickness of 100 μm to 400 μm.

11. The implantable drug delivery device according to any one of claims 8 to 10, wherein the device includes at least one internal hollow cavity.

12. The implantable drug delivery device according to any one of claims 8 to 10, wherein the device is a completely solid object.

13. A method for manufacturing a filament according to any one of claims 1 to 7, a. A step of preparing a liquid formulation containing the active ingredient, wherein the liquid formulation may further contain at least one stabilizer, buffer and / or surfactant. b. A step of freeze-drying or spray-drying the liquid formulation from step a to obtain dried fine particles. c. A step of uniformly dispersing the dried fine particles from step b together with a plasticizer and at least one polymer material. d. A step of obtaining a filament by extruding the dispersion from step c by hot melt extrusion (HME), Methods that include...

14. A method for manufacturing an implantable drug delivery device according to any one of claims 8 to 12, a. The process of loading filament into the print head of a 3D printer using a temperature higher than the glass transition temperature. b. A step of heating the construction platform at a temperature lower than the glass transition temperature of the polymer matrix, c. A step of depositing the heated filament through a nozzle to construct a device from at least a first layer to the final upper layer, Methods that include...

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