Poly Lactic-Co-Glycolic Acid Synthesis Essay

1. Introduction

A considerable amount of research has been conducted on drug delivery by biodegradable polymers since their introduction as bioresorbable surgical devices about three decades ago. Amongst all the biomaterials, application of the biodegradable polymer poly lactic-co-glycolic acid (PLGA) has shown immense potential as a drug delivery carrier and as scaffolds for tissue engineering. PLGA are a family of FDA-approved biodegradable polymers that are physically strong and highly biocompatible and have been extensively studied as delivery vehicles for drugs, proteins and various other macromolecules such as DNA, RNA and peptides [1–3]. PLGA is most popular among the various available biodegradable polymers because of its long clinical experience, favorable degradation characteristics and possibilities for sustained drug delivery. Recent literature has shown that degradation of PLGA can be employed for sustained drug release at desirable doses by implantation without surgical procedures. Additionally, it is possible to tune the overall physical properties of the polymer-drug matrix by controlling the relevant parameters such as polymer molecular weight, ratio of lactide to glycolide and drug concentration to achieve a desired dosage and release interval depending upon the drug type [4–6]. However the potential toxicity from dose dumping, inconsistent release and drug-polymer interactions require detailed evaluation. Here we present a review on the PLGA primarily as a delivery vehicle for various drugs, proteins and other macromolecules in commercial use and in research. We also present possible directions for future uses of PLGA in drug delivery applications.

2. Biodegradable Polymers

Biodegradable materials are natural or synthetic in origin and are degraded in vivo, either enzymatically or non-enzymatically or both, to produce biocompatible, toxicologically safe by-products which are further eliminated by the normal metabolic pathways. The number of such materials that are used in or as adjuncts in controlled drug delivery has increased dramatically over the past decade. The basic category of biomaterials used in drug delivery can be broadly classified as (1) synthetic biodegradable polymers, which includes relatively hydrophobic materials such as the α-hydroxy acids (a family that includes poly lactic-co-glycolic acid, PLGA), polyanhydrides, and others, and (2) naturally occurring polymers, such as complex sugars (hyaluronan, chitosan) and inorganics (hydroxyapatite) [7–9]. The breath of materials used in drug delivery arises from the multiplicity of diseases, dosage range and special requirements that may apply. Biocompatibility is clearly important, although it is important to note that biocompatibility is not an intrinsic property of a material, but depends on the biological environment and the tolerability that exists with respect to specific drug-polymer-tissue interactions [9].

2.1. Poly Lactic-co-Glycolic Acid (PLGA)

Polyester PLGA is a copolymer of poly lactic acid (PLA) and poly glycolic acid (PGA). It is the best defined biomaterial available for drug delivery with respect to design and performance. Poly lactic acid contains an asymmetric α-carbon which is typically described as the D or L form in classical stereochemical terms and sometimes as R and S form, respectively. The enantiomeric forms of the polymer PLA are poly D-lactic acid (PDLA) and poly L-lactic acid (PLLA). PLGA is generally an acronym for poly D,L-lactic-co-glycolic acid where D- and L- lactic acid forms are in equal ratio.

2.1.1. Physico-Chemical Properties

In order to design a better controlled drug delivery device, it is essential to understand the physical, chemical and biological properties of PLGA. The physicochemical properties of optically active PDLA and PLLA are nearly the same. In general, the polymer PLA can be made in highly crystalline form (PLLA) or completely amorphous (PDLA) due to disordered polymer chains. PGA is void of any methyl side groups and shows highly crystalline structure in contrast to PLA as shown in Figure 1. PLGA can be processed into almost any shape and size, and can encapsulate molecules of virtually any size. It is soluble in wide range of common solvents including chlorinated solvents, tetrahydofuran, acetone or ethyl acetate [7,10]. In water, PLGA biodegrades by hydrolysis of its ester linkages (Figure 2). Presence of methyl side groups in PLA makes it more hydrophobic than PGA and hence lactide rich PLGA copolymers are less hydrophilic, absorb less water and subsequently degrade more slowly. Due to the hydrolysis of PLGA, parameters that are typically considered invariant descriptions of a solid formulation can change with time, such as the glass transition temperature (Tg), moisture content and molecular weight. The effect of these polymer properties on the rate of drug release from biodegradable polymeric matrices has been widely studied. The change in PLGA properties during polymer biodegradation influences the release and degradation rates of incorporated drug molecules. PLGA physical properties themselves have been shown to depend upon multiple factors, including the initial molecular weight, the ratio of lactide to glycolide, the size of the device, exposure to water (surface shape) and storage temperature [11]. Mechanical strength of the PLGA is affected by physical properties such as molecular weight and polydispersity index. These properties also affect the ability to be formulated as a drug delivery device and may control the device degradation rate and hydrolysis. Recent studies have found, however, that the type of drug also plays a role in setting the release rate [12]. Mechanical strength, swelling behavior, capacity to undergo hydrolysis and subsequently biodegradation rate of the polymer are directly influenced by the degree of crystallinity of the PLGA, which is further dependent on the type and molar ratio of the individual monomer components in the copolymer chain. Crystalline PGA, when co-polymerized with PLA, reduces the degree of crystallinity of PLGA and as a result increase the rate of hydration and hydrolysis. As a rule, higher content of PGA leads to quicker rates of degradation with an exception of 50:50 ratio of PLA/PGA, which exhibits the fastest degradation, with higher PGA content leading to increased degradation interval below 50%. Degree of crystallinity and melting point of the polymers are directly related to the molecular weight of the polymer. The Tg (glass transition temperature) of the PLGA copolymers are reported to be above the physiological temperature of 37 °C and hence are glassy in nature, thus exhibiting fairly rigid chain structure. It has been further reported that Tg of PLGAs decrease with a decrease of lactide content in the copolymer composition and with a decrease in molecular weight [13]. Commercially available PLGA polymers are usually characterized in terms of intrinsic viscosity, which is directly related to their molecular weights.

2.1.2. Pharmacokinectic and Biodistribution Profile

The drug delivery specific vehicle, i.e., PLGA, must be able to deliver its payload with appropriate duration, biodistribution and concentration for the intended therapeutic effect. Therefore, design essentials, including material, geometry and location must incorporate mechanisms of degradation and clearance of the vehicle as well as active pharmaceutical ingredients (API). Biodistribution and pharmacokinetics of PLGA follows a non-linear and dose-dependent profile [14]. Furthermore, previous studies suggest that both blood clearance and uptake by the mononuclear phagocyte system (MPS) may depend on dose and composition of PLGA carrier systems [15]. Additionally whole-body autoradiography and quantitative distribution experiments indicate that some formulations of PLGA, such as nanoparticles, accumulate rapidly in liver, bone marrow, lymph nodes, spleen and peritoneal macrophages. The degradation of the PLGA carriers is quick on the initial stage (around 30%) and slows eventually to be cleared by respiration in the lung [16]. To address these limitations, studies have investigated the role of surface modification, suggesting that incorporation of surface modifying agents can significantly increase blood circulation half-life [17].

2.2. Copolymers of PLGA

The need for better delivery formulations that incorporate a variety in drugs and methods of administration has resulted in the development of various types of block copolymers of polyesters with poly ethylene glycol (PEG). PLGA/PEG block copolymers have been processed as diblock (PLGA-PEG) [18,19] or triblock molecules with both ABA (PLGA-PEG-PLGA) [20] and BAB (PEG-PLGA-PEG) [21] types. In diblock types, PEG chains orient themselves towards the external aqueous phase in micelles, thus surrounding the encapsulated species. This layer of PEG acts as a barrier and reduces the interactions with foreign molecules by steric and hydrated repulsion, giving enhanced shelf stability [22]. However, the addition of PEG to the system also results in reduction of encapsulation efficiency for drugs and proteins, even with the most appropriate fabrication techniques. The reduced drug incorporation may be due to steric interference of drug/protein-polymer interaction by the PEG chains. The precise mechanism for this effect is unclear. Better release kinetics from formulations of diblock copolymers have been demonstrated in comparison to PLGA alone. Various mechanisms of targeted delivery of drugs from diblock nanoparticles have also been reported [18,23,24].

Triblock copolymers of both ABA and BAB type can act as a thermogel with an A-block covalently coupled with a B-block via ester link. The copolymer is usually a free flowing solution at low temperature and can form a high viscosity gel at body temperature. These temperature-responsive copolymers, PLGA-PEG-PLGA or PEG-PLGA-PEG, are a kind of block copolymers composed of hydrophobic PLGA segments and hydrophilic PEG segments. The hydrophobic PLGA segments form associative crosslinks and the hydrophilic PEG segments allow the copolymer molecules to stay in solution. At lower temperatures, hydrogen bonding between hydrophilic PEG segments and water molecules dominates the aqueous solution, resulting in their dissolution in water. As the temperature increases, the hydrogen bonding becomes weaker, while hydrophobic forces among the PLGA segments are strengthened, leading to solution-gel transition. The ease of handling during fabrication, formulation, filtration and filling makes such thermoresponsive polymers attractive candidates. Drug and/or protein release from both ABA and BAB copolymers occurs by two principal mechanisms: (i) drug diffusion from the hydrogel during the initial release phase; and (ii) release of drug by the erosion of the hydrogel matrix during the later phase. During the degradation of a PEG-PLGA-PEG gel, there is a preferential mass loss of PEG-rich components. Therefore, the remaining gel becomes more hydrophobic in an aqueous environment, resulting in less water content [20,25–28]. This motif can also be applied to other co-polymer combinations, including but not limited to various copolymers of PLGA and polycaprolactone [29,30].

3. Fabrication Techniques for PLGA Carriers

Drugs and proteins are the most rapidly growing class of pharmaceuticals for which controlled or targeted release is used to increase specificity, lower toxicity and decrease the risk associated with treatment. However, the stability and delivery challenges associated with these agents have limited the number of marketed products. Maintaining adequate shelf-life of peptide and protein drugs often requires solid-state formulation to limit hydrolytic degradation reactions [31]. Drug delivery of peptides and proteins may also require parenteral formulations to avoid degradation in the digestive tract and first pass metabolism, while the short circulating half-lives of peptides and proteins contribute to the need for parenteral formulations that will reduce dosing frequency. In order to avoid the inconvenient surgical insertion of large implants, injectable biodegradable and biocompatible PLGA particles (microspheres, microcapsules, nanocapsules, nanospheres) could be employed for controlled-release dosage forms. Drugs formulated in such polymeric devices are released either by diffusion through the polymer barrier, or by erosion of the polymer material, or by a combination of both diffusion and erosion mechanisms. In addition to its biocompatibility, drug compatibility, suitable biodegradation kinetics and mechanical properties, PLGA can be easily processed and fabricated in various forms and sizes. This section describes various fabrication techniques of PLGA controlled drug delivery devices [9].

3.1. Microparticle Preparation Techniques

3.1.1. Solvent Evaporation Method

  • Single emulsion process

    Oil-in-water emulsification processes are examples of single emulsion processes. Polymer in the appropriate amount is first dissolved in a water immiscible, volatile organic solvent (e.g., dichloromethane (DCM)) in order to prepare a single phase solution. The drug of particle size around 20–30 μm is added to the solution to produce a dispersion in the solution. This polymer dissolved drug dispersed solution is then emulsified in large volume of water in presence of emulsifier (polyvinyl alcohol (PVA) etc.) in appropriate temperature with stirring. The organic solvent is then allowed to evaporate or extracted to harden the oil droplets under applicable conditions. In former case, the emulsion is maintained at reduced or atmospheric pressure with controlling the stir rate as solvent evaporates. In the latter case, the emulsion is transferred to a large quantity of water (with or without surfactant) or other quench medium to diffuse out the solvent associated with the oil droplets. The resultant solid microspheres are then washed and dried under appropriate conditions to give a final injectable microsphere formulation [32–35].

  • Double (Multiple) emulsion process

    Water-in-oil-in-water emulsion methods are best suited to encapsulate water-soluble drugs like peptides, proteins, and vaccines, unlike single emulsion methods which is ideal for water-insoluble drugs like steroids. First, an appropriate amount of drug is dissolved in aqueous phase (deionised water) and then this drug solution is added to organic phase consisting of PLGA and/or PLA solution in DCM or chloroform with vigorous stirring to yield a water-in-oil emulsion. Next, the water-in-oil primary emulsion is added to PVA aqueous solution and further emulsified for around a minute at appropriate stress mixing conditions. The organic solvent is then allowed to evaporate or is extracted in the same manner as oil-in-water emulsion techniques. In double emulsion processes, choice of solvents and stirring rate predominantly affects the encapsulation efficiency and final particle size [32,36,37].

3.1.2. Phase Separation (Coacervation)

Coacervation is a process focused on preparation of micrometer sized biodegradable polymer encapsulation formulations via liquid-liquid phase separation techniques. The process yields two liquid phases (phase separation) including the polymer containing coacervate phase and the supernatant phase depleted in polymer. The drug which is dispersed/dissolved in the polymer solution is coated by the coacervate. Thus, the coacervation process includes the following three steps as reported in literature [38–40]

  • Phase separation of the coating polymer solution,

  • Adsorption of the coacervate around the drug particles, and

  • Quenching of the microspheres.

Solutions are prepared by mixing polymer and solvent in appropriate ratios. Hydrophilic drugs like peptides and proteins are dissolved in water and dispersed in polymer solution (water-in-oil emulsion). Hydrophobic drugs like steroids are either solubilized or dispersed in the polymer solution (oil-in-water emulsion). Gradual addition of organic medium to the polymer-drug-solvent phase while stirring, extracts the polymer solvent resulting in phase separation of polymer by forming a soft coacervate of drug containing droplets. The size of these droplets can be controlled by varying stirring rate and temperature of the system. The system is then quickly dipped into a medium in which it is not soluble (both organic or aqueous) to quench these microdroplets. The soaking time in the quenching bath controls the coarsening and hardness of the droplets. The final form of the microspheres is collected by washing, sieving, filtration, centrifugation or freeze drying. The processing parameters including polymer concentration, quenching temperature, quenching time and solvent composition affect the morphology and size of the microspheres [41–43].

3.1.3. Spray Drying

Emulsion techniques require precise control of processing parameters for higher encapsulation efficiency, and phase separation techniques tend to produce agglomerated particles and also require removal of large quantities of the organic phase from the microspheres. This makes the process difficult for mass production. Alternatively, spray drying is very rapid, convenient and has very few processing parameters, making it suitable for industrial scalable processing. In this process, drug/protein/peptide loaded microspheres are prepared by spraying a solid-in-oil dispersion or water-in-oil emulsion in a stream of heated air. The type of drug (hydrophobic or hydrophilic) decides the choice of solvent to be used in the process. The nature of solvent used, temperature of the solvent evaporation and feed rate affects the morphology of the microspheres. The main disadvantage of this process is the adhesion of the microparticles to the inner walls of the spray-dryer. Various spray drying techniques have been reported [44–49]. This method is known to encapsulate all kinds of drugs/peptides/proteins into microparticles without significant loss in their biological activity. Recently, coaxial capillary flows have become a preferable technique to produce monodispersed micro/nanoparticles with either simple or core-shell structure because of their precise control on mean particle size [50,51]. Using these techniques, processing parameters such as orientation of jets, material flow rates, and rate of solvent extraction can be controlled to create uniform and well-centered double-walled microspheres exhibiting a controllable shell thickness [52]. Additionally, microfluidic devices can incorporate the use of electrostatic forces to control the size and shape of particles for increased tuning of release characteristics [53].

3.2. Nanoparticle Preparation Techniques

Various groups have also reported successful preparation of PLGA nanoparticles. All the above described microparticle techniques can be employed for manufacturing PLGA nanoparticles (nanospheres and nanocapsules) by adjusting the processing parameters. These parameters usually use a small dispersed phase ratio and rate of stirring. The most common method used for the preparation of solid, polymeric nanoparticles is the emulsification-solvent evaporation technique. However, this method is primarily used in encapsulation of hydrophobic drugs. A modification on this procedure called the double or multiple emulsion technique has become the favored protocol for encapsulating hydrophilic compounds and proteins [37]. Nanoparticles can also be synthesized by nanoprecipitation methods. Polymer and drug are dissolved in acetone and added to an aqueous solution containing Pluronic F68. The acetone is evaporated at appropriate temperatures and reduced pressures leaving behind the polymer encapsulated nanoparticles with drug [54]. Salting out is another method in which a water-in-oil emulsion is first formed containing polymer, solvent (usually non chlorinated like acetone), salt (e.g., magnesium acetate tetrahydrate) and stabilizer. Water is then added to the solution until the volume is sufficient to diffuse acetone into the water, resulting in nanoparticle formulations [55–58].

3.3. Implant Preparation Techniques

3.3.1. Solvent-Casting and Compression Molding

Solvent casting is a method to fabricate a macroscopic millimeter size formulation which can be implanted or inserted for long term medication [59]. Large size, macroscopic formulations act as a reservoir for drug that can be delivered over a longer interval. In this method, a polymer and drug mixture is dissolved in a common solvent (e.g., acetone) in the desirable proportion, and the solvent is cast at around 60 °C until complete evaporation. Their resultant structure is a composite material of the drug together with the polymer. The solvent cast material is then compression molded into its desired geometry at around 80 °C and 25,000 psi to final density of 1 g/cc. This implant can be subcutaneously delivered in the body. The main advantage of this approach over micro/nanospheres is related to the ability to manage adverse events, since implants retain a degree of reversibility which is not available in depot mechanisms [12,59].

3.3.2. Extrusion

  • [0001]

    The invention belongs to the technical field of biodegradable medical materials, and relates to a process for synthesizing biodegradable medical polylactic-glycolic acid (lactic acid-glycolic acid copolymer) featuring a high biological safety via copolycondensation, wherein a bionic creatinine (a metabolite of arginine in human body) is used as catalyst, lactic acid and glycolic acid is used as raw materials.

  • [0002]

    Polylactic-glycolic acid (PLGA) is an important biodegradable medical material, featuring good biocompatibility, bioabsorbability and biodegradability. Since the lactic acid-glycolic acid copolymer is formed from lactic acid and glycolic acid, it combines the advantages of two homopolymer polyester materials (polylactic acid (PLA), polyglycolic acid (PGA)). The polylactic-glycolic acid has good biocompatibility, In addition, its material strength, degradation rate, mechanical properties and the like can be modulated by changing the composition and molecular weight of the copolymer. Hence, it is a biodegradable medical material featuring a wide range of practical value. The polylactic-glycolic acid has been extensively applied in several aspects of biomedical science such as implantable hard tissue-repairing materials, surgical sutures, and the carrier for targeting drugs and controlled release drugs. It is required that the degradable materials applied in the field of biomedicine should exhibit high biological safety and not contain any toxic metal and other toxic ingredients. Currently, the production of commercially available polylactic-glycolic acid is performed via stannous octoate catalyzed ring-opening polymerization or stannous chloride catalyzed polycondensation. The recent studies throughout the world have definitely proved that divalent tin salts (stannous octoate and stannous chloride) exhibit cytotoxicity. Since the tin salt catalyst used cannot be completely removed from the synthetic polymer after polymerization reaction, the safety issue of polylactic-glycolic acid synthesized by using divalent tin-containing compound as catalyst for use as medical materials for human has been generally questioned by scientists all over the world. Thus, the exploration for efficient, non-toxic, and metal-free green catalysts for synthesizing polylactic-glycolic acid has become the challenging issue in the field of degradable biomedical materials.

  • [0003]

    The object of the present invention is to solve the safety problem which may be present in the existing polycondensation, wherein stannous chloride catalyst is used to synthesize polylactic-glycolic acid applied in the field of human medicine and pharmacy, and to provide a process for synthesizing polylactic-glycolic acid via direct copolycondensation by using a bionic creatinine as catalyst.

  • [0004]

    For the first time, the present invention develops a new process for synthesizing biodegradable medical polylactic-glycolic acid featuring a high biological safety via direct copolycondensation, wherein non-toxic, metal-free biomass creatinine (a metabolite of arginine in human body) is used as catalyst, lactic acid (LA, 85% aqueous solution) and glycolic acid (GA, 95%) are used as comonomers.

  • [0005]

    The chemical name of the non-toxic, metal-free biomass organic guanidinium compound, creatinine, as used in the present invention is 2-amino-1-methyl-2-imidazolin-4-one (the common name in English is creatinine; the abbreviation in English: CR), and the molecular structure thereof is shown as follows:

  • [0000]

  • [0006]

    The process provided in the present application for synthesizing biodegradable medical polylactic-glycolic acid via direct copolycondensation of lactic acid and glycolic acid catalyzed by creatinine has the following steps:

  • [0007]

    Aqueous solution of industrial grade lactic acid (LA) with a mass percentage of 85% and glycolic acid (GA) with a mass percentage of 95% were used as comonomers at a molar ratio of 9:1 to 1:9, to firstly synthesize an oligolactic-glycolic acid (lactic acid-glycolic acid copolymer), wherein the weight average Mw=200˜400;

  • [0008]

    Process conditions: a reactor was charged with lactic acid and glycolic acid, and then vacuumized and charged with argon for three repetitions; under an argon atmosphere at normal pressure, the reaction system was heated to 130-150° C. and subjected to dehydration for 1˜3 h; the pressure in the reactor was then reduced to 100 Torr, reacting at 130˜150° C. for 1˜3 h; finally, the pressure in the reactor was reduced to 30 Torr, reacting at 130˜150° C. for 1—3 h;

  • [0009]

  • [0010]

    The oligolactic-glycolic acid (LGA) synthesized in the first-step was used as raw material; the creatinine was used as catalyst; under a reduce pressure, the polycondensation of fused monomers was performed to synthesize biodegradable medical polylactic-glycolic acid featuring a high biological safety;

  • [0011]

    The process conditions and operation methods of synthetic reaction were described as follows: oligolactic-glycolic acid and creatinine catalyst were added into the reactor; the mass ratio between oligolactic-glycolic acid and creatinine was set as 100:1˜1000:1; the pressure in the reactor was reduced to 10 Torr, heating to 150˜190° C. for 40˜170 h, to obtain polylactie-glycolic acid.

  • [0012]

  • [0013]

    The polylactic-glycolic acid synthesized in the present invention has a weight average molecular weight within 1.8˜17.7×104, and according to the actually required molecular weight, the polymer can be synthesized by controlling the time of polymerization reaction so that the molecular weight thereof falls into the above range.

  • [0014]

    The polylactic-glycolic acid synthesized according to the present application does not contain any metal and other toxic ingredients and thus can be used as implantable hard tissue-repairing materials, surgical sutures, and the carrier for targeting drugs and controlled release drugs.

  • [0015]

    The advantages and beneficial effects of the present invention are as follows:

  • [0016]

    1. The catalyst used exhibits high biocompatibility and biological safety;

  • [0017]

    2. The synthesized product exhibits high biocompatibility and biological safety and does not contain any metal and other toxic ingredients.

  • [0018]

    3. The weight average molecular weight of synthesized product, polylactic-glycolic acid, can be regulated within the range of 1.8˜17.7×104;

  • [0019]

    4. The biodegradable medical polylactic-glycolic acid (featuring a high biological safety) is synthesized by using green catalyst and green process (no use of solvent, no occurrence of toxic products);

  • [0020]

    5. Low cost of raw materials, simple technical operation, easy for industrial practice.

  • [0021]

    A reactor was charged with 45 g of aqueous solution of industrial grade lactic acid (LA) featuring a mass percentage of 85% and 38 g of glycolic acid (GA) featuring a mass percentage of 95%, and then vacuumized and charged with argon for three repetitions. Under an argon atmosphere at normal pressure, the reaction system was then heated to 130° C. and subjected to dehydration for 3 h. The pressure in the reactor was then reduced to 100 Torr, reacting at 130° C. for 3 h, Finally, the pressure in the reactor was reduced to 30 Torr, reacting at 130° C. for 3 h, to give oligolactic-glycolic acid (OLGA) (yield: 98.0%), with a weight average molecular weight of 220.

  • [0022]

    A reactor was charged with 45 g of aqueous solution of industrial grade lactic acid (LA) featuring a mass percentage of 85% and 38 g of glycolic acid (GA) featuring a mass percentage of 95%, and then vacuumized and charged with argon for three repetitions. Under an argon atmosphere at normal pressure, the reaction system was then heated to 150° C. and subjected to dehydration for 1 h. The pressure in the reactor was then reduced to 100 Torr, reacting at 150° C. for 1 h. Finally, the pressure in the reactor was reduced to 30 Torr, reacting at 150° C. for 1 h, to give oligolactic-glycolic acid (OLGA) (yield: 98.2%), with a weight average molecular weight of 280.

  • [0023]

    A reactor was charged with 45 g of aqueous solution of industrial grade lactic acid (LA) featuring a mass percentage of 85% and 38 g of glycolic acid (GA) featuring a mass percentage of 95%, and then vacuumized and charged with argon for three repetitions. Under an argon atmosphere at normal pressure, the reaction system was then heated to 140° C. and subjected to dehydration for 2 h. The pressure in the reactor was then reduced to 100 Torr, reacting at 140° C. for 2 h. Finally, the pressure in the reactor was reduced to 30 Torr, reacting at 140° C. for 2 h, to give oligolactie-glycolic acid (OLGA) (yield: 98.6%), with a weight average molecular weight of 400.

  • [0024]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 190° C. for 40 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 85.9%), with a weight average molecular weight of 1.83×104.

  • [0025]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 170° C. for 48 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactie-glycolic acid (yield: 85.0%), with a weight average molecular weight of 1.86×104.

  • [0026]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 150° C. for 54 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 87.4%), with a weight average molecular weight of 1.80×104.

  • [0027]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 190° C. for 124 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 85.1%), with a weight average molecular weight of 7.12×104.

  • [0028]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 170° C. for 132 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 85.6%), with a weight average molecular weight of 7.08×104.

  • [0029]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 150° C. for 150 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 86.2%), with a weight average molecular weight of 7.07×104.

  • [0030]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 190° C. for 154 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 84.7%), with a weight average molecular weight of 17.7×104.

  • [0031]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 170° C. for 160 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 84.5%), with a weight average molecular weight of 17.3×104.

  • [0032]

    70 g of oligoactic-glycolic acid and 265 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 150′C for 169 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 84.9%), with a weight average molecular weight of 17.0×104.

  • [0033]

    A reactor was charged with 90 g of aqueous solution of industrial grade lactic acid (LA) featuring a mass percentage of 85% and 7.6 g of glycolic acid (GA) featuring a mass percentage of 95%, and then vacuumized and charged with argon for three repetitions. Under an argon atmosphere at normal pressure, the reaction system was then heated to 130° C. and subjected to dehydration for 3 h. The pressure in the reactor was then reduced to 100 Torr, reacting at 130° C. for 3 h. Finally, the pressure in the reactor was reduced to 30 Torr, reacting at 130° C. for 3 h, to give oligolactic-glycolic acid (OLGA) (yield: 98.1%), with a weight average molecular weight of 220.

  • [0034]

    A reactor was charged with 15 g of aqueous solution of industrial grade lactic acid (LA) featuring a mass percentage of 85% and 102 g of glycolic acid (GA) featuring a mass percentage of 95%, and then vacuumized and charged with argon for three repetitions. Under an argon atmosphere at normal pressure, the reaction system was then heated to 130° C. and subjected to dehydration for 3 h. The pressure in the reactor was then reduced to 100 Torr, reacting at 130° C. for 3 h. Finally, the pressure in the reactor was reduced to 30 Torr, reacting at 130° C. for 3 h, to give oligolactic-glycolic acid (OLGA) (yield: 98.0%), with a weight average molecular weight of 220.

  • [0035]

    70 g of oligoactic-glycolic acid and 700 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 160° C. for 170 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 84.7%), with a weight average molecular weight of 17.1×104.

  • [0036]

    70 g of oligoactic-glycolic acid and 140 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 170° C. for 60 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 84.6%), with a weight average molecular weight of 1.88×104.

  • [0037]

    70 g of oligoactic-glycolic acid and 70 mg of creatinine catalyst were added into the reactor. The pressure in the reactor was reduced to 10 Torr, heating to 180° C. for 48 h. After stopping the reaction, the reactor was cooled to room temperature. The polymer was dissolved in acetone, and then poured into ethanol at 0° C., filtered at reduced pressure. The resulting product was dried under vacuum at 50° C. for 36 h, to give a white solid, i.e. polylactic-glycolic acid (yield: 85.1%), with a weight average molecular weight of 1.98×104.

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