Open access peer-reviewed chapter

Electrospinning of Fiber Matrices from Polyhydroxybutyrate for the Controlled Release Drug Delivery Systems

Written By

Anatoly A. Olkhov, Svetlana G. Karpova, Anna V. Bychkova, Alexandre A. Vetcher and Alexey L. Iordanskii

Submitted: 30 November 2021 Reviewed: 09 June 2022 Published: 14 July 2022

DOI: 10.5772/intechopen.105786

From the Edited Volume

Electrospinning - Material Technology of the Future

Edited by Tomasz Tański and Paweł Jarka

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The submission provides an overview of current state of the problem and authors’ experimental data on manufacturing nonwoven fibrous matrices for the controlled release drug delivery systems (CRDDS). The choice of ultrathin fibers as effective carriers is determined by their characteristics and functional behavior, for example, such as a high specific surface area, anisotropy of some physicochemical characteristics, spatial limitations of segmental mobility that are inherent in nanosized objects, controlled biodegradation, and controlled diffusion transport. The structural-dynamic approach to the study of the morphology and diffusion properties of biopolymer fibers based on polyhydroxybutyrate (PHB) is considered from several angles. In the submission, the electrospinning (ES) application to reach specific characteristics of materials for controlled release drug delivery is discussed.


  • polymer therapeutic systems
  • electrospinning
  • biopolymer fibers
  • nonwoven fibrous matrices
  • polyhydroxybutyrate
  • controlled diffusion transport
  • controlled release
  • morphology
  • segmental mobility

1. Introduction

The modern paradigm of targeted drug delivery to the organ of interest, its part, or the target cell is based on the use of biocompatible carriers, the sizes of which are in the submicron (nanometer) range. Today, functional carriers of biologically active compounds and absorbents with a high specific surface area are widely used in biomedicine in the form of bioceramic minerals, polymer therapeutic systems, framework structures for cell engineering, nanoscale hybrid means of targeted drug transport, and a number of innovative systems that simultaneously function as implants and carriers of biologically active compounds [1, 2, 3, 4, 5]. Among these functional materials, the greatest practical and commercial advantages, in the near future, will be obtained by hybrid micro- and nanoparticles [6, 7], as well as composite ultrathin fibers [8, 9].

Micro- and nanoparticles based on biodegradable materials offer a wide range of applications as components of innovative forms or independent systems for drug delivery in the implementation of anti-inflammatory and antitumor therapy, wound healing, and thrombolytic therapy. The use of magnetic nanoparticles as part of innovative systems can facilitate targeted drug delivery, reducing the required dosage of a drug, and provide a prolonging effect of the activity of biological macromolecules, for example, enzymes, and the ability to visualize drug delivery processes [10, 11, 12].

The employment of biodegradable and biocompatible particles and fibers creates additional advantages in the development of a new generation of implanted therapeutic systems. Therefore, nowadays, the main part of the work aimed to develop biodegradable compositions in medicine is devoted to the creation of materials for tissue engineering. Modern metal implants (titanium and its alloys, stainless steel, etc.) are characterized by rigidity and endurance and do not cause an immune response but do not have the properties of osteoinduction and osteoconductivity. Often their application turns to add an additional surgical intervention to remove the implant. This increases additional risks and the duration of the patient’s recovery process, and the therapy aimed at restoring the bone tissue and its environment. Therefore, inorganic materials analogous to bone tissue, such as apatite and complex composite biodegradable forms that provide the transport of biological molecules and synthetic drugs, are objects of fundamental and applied research.

One of the innovative ways to obtain ultrathin fibers and fibrillar nanomaterials is the method of electrospinning (ES) of polymer solutions and melts [13, 14]. The development of technology for the formation of nanofibers in an electrostatic field makes it possible to create materials of various shapes and morphology with a high specific surface area and porosity, with adequate mechanical properties, and a wide range of structural, dynamic, and diffusion characteristics. Fibrillar matrices and mats formed by the obtained nanofibers create favorable conditions for free migration and proliferation of cells in the three-dimensional space of frame structures, and, accordingly, provide a high integration affinity of the material for living tissues of the body [15]. They are actively used in the design of biosensors, nanofilters, for wound therapy, to immobilize enzymes, in the creation of prolonged and targeted drug delivery systems, and other areas of modern biology and medicine [16, 17, 18, 19, 20].

It should be noted that in contrast to a great availability published data on the nanoparticles applications, the data on biodegradable and bioresorbable nanofibes is rather rare [21, 22, 23]. About a third of publications analyze the peculiarities of the ES technology, and the rest considers their characteristics, functional behavior in vitro and in vivo, as well as preclinical and clinical trials.


2. Fields of application of biopolymer forms for therapeutic purposes

Biodegradable compositions are also applied in solving the problem of wound healing, which may include a set of subtasks—stopping bleeding, preventing inflammation, proliferation, and tissue reconstruction. Chronic wounds are characterized by high protease activity, infection, inflammation, and hypoxia [24, 25, 26, 27]. A wide range of wound healing materials is currently being developed. These are unfilled materials based on siloxysiloxane, dextran, urethane, collagen, etc.; materials including stem cells or vitamin E based on hyaluronic acid; materials based on extracellular matrix proteins (fibrin, fibronectin, collagen, etc.) for tissue reconstruction and angiogenesis; and materials including antibiotics and other drugs [28]. Materials in the form of films, hydrogels, foams, multilayer compositions, etc., have already received clinical use. Chitosan, known for its hemostatic and antibacterial activity, is used in some currently developed composite forms for accelerated wound healing and drug delivery for matrix formation and tissue reconstruction. In [29], chemically stable composite materials for wound coverage were based on two polysaccharides—cellulose and chitosan combining mechanical strength (cellulose) and the ability to stop bleeding, cleanse and heal wounds, deliver drugs, and overall bactericidal properties (chitosan).

It is known that the regeneration of cartilage tissue is directly related to the number of chondrocytes per unit of its mass, and the introduction of a suspension of chondrocytes into the damaged areas is currently considered a promising therapeutic approach. Foreign chondrocytes, being introduced into the joint cavity, do not cause a rejection reaction, because they have limited immunogenic activity. Much attention is now being paid to the development of composite hydrogels for injections—research is being carried out using photopolymerization, chemical crosslinking of molecules in the composition of the composite with carbodiimide, glutaraldehyde, genipin, and adipic dihydrazide. In the first case, the use of a photosensitizer and radiation is required, which limits the applicability of the approach, and in the case of crosslinkers, the main problem is their toxicity toward cells. In [30], a new representative of composite hydrogels was obtained based on chitosan and oxidized hyaluronic acid. The formation of the gel proceeded through the formation of a Schiff base between the amino and aldehyde groups of polysaccharide derivatives, N-succinyl-chitosan (S-CS) and hyaluronic acid aldehyde, and did not involve the use of a chemical crosslinking agent. The potential of using such composite hydrogels as scaffold structures for injections has been shown using the example of cartilage tissue cells in joints. In this study, the high content of chitosan led to a decrease in the rate of degradation of the composite.

Let us consider several examples of the use of composite particles based on biodegradable polymers. In [31], hydrogels based on chitosan and gelatin are used for long-term administration of the antiglaucoma drug timolol maleate and can reduce the side effects of its use. Hydrogels based on sulfonated chitosan and heparin-like chitosan (containing carboxymethyl and sulfate groups) increase blood clotting time [32], prevent protein and platelet adsorption on the membrane intended for blood purification [33], and prevent complement activation [34]. Hydrogels consisting of chitosan, heparin, and poly (γ-glutamic acid) with different ratios of components and loaded with superoxide dismutase were used to create a wound dressing with antioxidant properties [35], which had the potential to treat chronic trauma in diabetes and proved to be a promising wound healing agent. Heparin-loaded hydrogel based on photosensitive hydroxyethylchitosan promoted a long-term effect of lowering intraocular pressure after surgery to eliminate glaucoma [36]. While the possibilities of transporting the anticoagulant heparin as a constituent element of particles are being actively studied, a much smaller number of works are devoted to the transport of proteins associated with the functioning of fibrinolytic and anticoagulant systems—tissue plasminogen activator (tPA), streptokinase, and urokinase. The formation of complexes between enzymes and biodegradable polymers proved to be an effective way to overcome the short biological half-life of enzymes in the bloodstream, and the inclusion of magnetic nuclei in the particles revealed the potential of using biodegradable materials for targeted drug transport and diagnostics. Magnetic nuclei delivered on their envelope (for example, the composition of polyethylene glycol (PEG) and chitosan [37]) tPA or streptokinase promote thrombolysis controlled by an external magnetic field. Hydrogels based on chitosan and alginate with incorporated magnetic particles were investigated as carriers of the anticancer drug matrine for oral administration [38, 39]. The obtained systems provide pH-sensitive release and targeted delivery under the influence of weak magnetic fields. The inclusion of magnetic particles in systems based on hyaluronic acid allows to detect the processes of tissue regeneration and biological delivery of drugs with the participation of constructed systems [40].

The areas of possible application of biodegradable forms are not limited to the aforementioned examples. New designs of cardio stents with biodegradable or inert coatings are capable of providing targeted transport of such cardiac drugs as paclitaxel, sirolimus, tacrolimus, etc. These drugs are responsible for suppressing cell proliferation; therefore, their prolonged delivery significantly reduces the number of restenosis incidents as compared to the first generation of spring metal stents. Moreover, a nano-level modified stent surface for controlled cell adhesion is desirable. In urology, a separate group of polymer implants consists of representatives of natural biodegradable polymers—poly-α-hydroxy acids (polylactides and their copolymers) and poly-β-hydroxyalkanoates (polyhydroxybutyrate (PHB), and its copolymers with oxyvalerate, oxyhexanoate, etc.), and the base for neurological implants are biodegradable polymers such as polylactides or polyhydroxyalkanoates [41]. For urological applications, a prolonged release of bactericidal drugs is required, whereas for nerve tissue regeneration, a drug activates the development of a nerve impulse.

Modern works also include the creation of systems with a reverse response, an example of which is the form that releases insulin depending on the concentration of glucose in the blood [42]. The possibility of visualizing the area of development of the pathological process and monitoring the course of the treatment process is presented with the use of isotopes, fluorescent dyes, quantum dots, magnetic nanoparticles, and other markers [43]. Biodegradable innovative materials, thus, have received a wide range of possible applications in the treatment of various diseases (Figure 1) in the form of various medical forms—macroobjects, films, micro-, and nanoparticles, etc. A wide variety of required compositions and morphology of matrices of medical forms contributes to the fundamental and applied issues associated with the use of natural and synthetic materials for the delivery of low- and high-molecular-weight biologically active substances (BASs). More details on the possible chemical composition of matrices for drug transport and methods of drug inclusion in the matrix will be presented in the next section.

Figure 1.

The main areas of pharmaceutical forms’ application for the delivery of low- and high-molecular-weight pharmaceuticals based on biodegradable macromolecules.


3. The main components of biodegradable innovative forms of drug delivery

Let us consider the main components of biodegradable materials that make up innovative forms using the example of collagen, hyaluronic acid, carboxymethyl cellulose (CMC), chitosan, and other natural and synthetic macromolecules, the structural formulas of which are shown in Figure 2, most often used for bone tissue regeneration [43].

Figure 2.

Structural formulas of some biodegradable macromolecules used for targeted transport of low- and high-molecular pharmaceuticals.

Chitosan is a polysaccharide of D-glucosamine and N-acetyl-D-glucosamine, linked by β (1 → 4) -glycosidic bonds. Chitosan is obtained by cleavage of the acetyl group from chitin. The deacetylation reaction can also be accompanied by the rupture of the glycosidic bonds of the polymer. The degree of deacetylation of chitosan and its molecular weight predetermines its reactivity and properties, causing the structural heterogeneity of macromolecules. The solubility of chitosan depends on the pH (it is insoluble at neutral values), and its salts, for example, chitosan hydrochloride and chitosan glutamate, are soluble at any pH.

The hydrophilicity of chitosan and its positive charge facilitate its reactions with negatively charged macromolecules and polyanions, which are both components of the biodegradable composite matrix and functional molecules.

This makes it possible to use sol-gel processes for drug binding. The positive charge of chitosan also promotes adhesion to the mucous membranes of the body, which makes it possible to use it for the transport of drugs through the mucous membrane. Thus, the features of chitosan are pH-dependent solubility (at pH <5 and basic pH) and insignificant toxicity due to ▬NH2 groups (it undergoes enzymatic degradation in vivo, and its degradation products are involved in the metabolic cycle). CMC can be used together with cellulose and is similar in structure to chitosan. The sodium salt of CMC is a water-soluble polymer. Cellulose is a polysaccharide with the formula (C6H10O5)n, in which the D-glucose residues are linked by β (1 → 4)-glycosidic bonds. Cellulose undergoes dissolution under the action of ionic liquids (organic salts, liquid at room temperature). In [29], such a solvent was used to obtain a composite material for wound dressings based on two polysaccharides – cellulose and chitosan.

Fibrinogen and fibrin are the next examples of protein molecules that can be components of micro- and nanoscale scaffolds. These proteins are present in the blood and are the main participants in the process of coagulation, which determines their high ability to interact with damaged tissues and cells. To solve problems associated with tissue regeneration, proteins can be modified or introduced into more complex systems, including, for example, growth factors or other proteins, as well as stem cells [44]. Fibrin gel formation can be initiated directly at the site of injury using particles on the surface of which thrombin is fixed [45].

The main protein of silkworm silk, fibroin, and the skeleton silk of the spider web, spidroin, have crystalline parts responsible for high tensile strength and amorphous parts that provide elasticity, which makes them versatile materials for use in tissue engineering, pharmacy, and medicine, regardless of the type of construction. The breakdown products of silk fibroin and spidroin are amino acids, which act as an additional building material for tissue regeneration [46].

Hyaluronic acid is a hydrophilic, non-immunogenic, biodegradable glucosaminoglycan—a polymer consisting of D-glucuronic acid and D-N-acetylglucosamine residues, linked alternately by β-1,4- and β-1,3-glycosidic bonds, which, in combination with other osteoconductive molecules, promotes bone growth. It has a high content in extracellular matrices, is a component of articular cartilage, and is part of the skin. Amphiphilic derivatives of hyaluronic acid promote its self-assembly in a core-shell nanogel for the transport of hydrophobic pharmaceuticals, e.g. anticancer.

For the transport of low- and high-molecular-weight therapeutic substances, “depot carriers” are used based on sodium alginate and its compositions with pectin, sodium hyaluronate, etc. Salts of alginic acid (alginates), in contrast to the polysaccharide acid itself, form colloidal solutions in water and have antimicrobial and hemostatic action.

Unlike natural macromolecules, synthetic biodegradable molecules such as poly-α-hydroxy acids such as hydrophobic polylactides (PLAs) and more hydrophilic polyglycolides (PGAs), as well as poly-β-hydroxyalkanoates (PHBs and its derivatives) do not elicit a biological response and are widely distributed as components of innovative forms for drug delivery. Degradation of polymers proceeds through hydrolysis of ether groups, and the rate and products of degradation are determined by the composition, structure, and molecular weight of molecules, as well as the content of ions and enzymes in biological fluids. Polymeric materials of the class of poly-β-hydroxyalkanoates or poly-α-hydroxy acids degrade to nontoxic products—CO2 and H2O [41]. Also commonly used are aliphatic polyesters, that is, polycaprolactones (PCLs). Blends of polymers or copolymers are often the basis for materials intended for the restoration of bone and cartilage tissue. The most popular copolymer is polylactide-co-glycolide (PLGA), the ratio of the components of which affects the hydrophilicity and biodegradability characteristics. During copolymerization, the degree of crystallinity decreases, as a result of which hydrolytic destruction proceeds faster [47]. Other common synthetic macromolecules for bone tissue regeneration are polypropylene fumarate (PPF), polyanhydrides, and polyphosphazenes [26]. Oligo(polyethylene glycol fumarate) (OPF) based on PEG is able to biodegrade, while PEG is not [48]. Degradation of PPF and OPF is due to the fumaric acid residue in the polymers. Copolymerization of macromolecules, for example, polyanhydrides, can be carried out to increase the hydrophobicity of materials and reduce the rate of their biodegradation. The combination of natural and synthetic materials in the compositions provides a wide range of properties of the systems and the possibilities of their use. An example of a natural material used for the transport of various drugs in the composition of composite forms is gelatin [49].

In innovative materials, the simultaneous use of natural and synthetic macromolecules creates unique opportunities for the delivery of various pharmaceuticals and ensuring the required release profiles. As in the case of synthetic ones, when using natural hydrogels, chemical and physical crosslinking of macromolecules allow one to regulate the diffusion of therapeutic substances included in biodegradable forms, and the biodegradation of macromolecules is also determined by the degree of their crosslinking and the type of crosslinking agent. In this case, the pharmaceuticals themselves can be covalently and non-covalently associated with biodegradable compositions, in particular, the inclusion of pharmaceuticals in the composition of Ca3(PO4)2 particles is carried out physically or chemically [24], and the release of pharmaceuticals can be controlled by the chemistry of Ca3(PO4)2, the porosity of the material, the surface area of ​​the hydrogel particles, their charge surface, and crystallinity. In the case of covalent immobilization of macromolecular pharmaceuticals (for example, proteins), the release mechanism includes chemical/enzymatic cleavage of the active substance. The introduction of an antibiotic in the treatment of osteomyelitis can be provided by direct mixing of the drug powder with a bone graft or soaking the bone graft in an antibiotic solution [26]. Similarly, for the prolonged delivery of streptokinase, systems obtained by mixing the protein with chitosan were used [50]. Urokinase and streptokinase showed their activity as thrombolytics both on the surface and in the bulk of particles from chitosan and tripolyphosphate, providing delivery, and the indicators in both cases were higher than in the protein solution [51]. The introduction of pharmaceuticals into composite structures from a solution after their creation is one of the approaches that damage pharmaceuticals to the least extent. Chondrocytes in [30] were introduced into the hydrogel from a solution in which they were resuspended. Another example of a technique that provides a minimal effect on the drug structure is the approach using the state of a supercritical fluid, which provides a one-stage filling of porous matrices with medicinal substances. Approaches are known that include the formation of a special drug complex, for example, with cyclodextrins, for its use in a sol-gel process, during which biodegradable composite structures with pharmaceuticals incorporated into them are created [52]. Pharmaceuticals can also be adsorbed on the surface of implanted structures and coated with biodegradable polymeric materials to ensure their long-term release, which is realized, for example, in [25] during transfection. A hydrophobic drug can also be encapsulated in an oil core covered with a biodegradable shell, including, for example, chitosan [53]. Drug release can be divided by type into diffusion-controlled, controlled by chemical processes or matrix swelling, and controlled by external processes or devices.

Most of the listed approaches for the inclusion of pharmaceuticals in various matrices can also be implemented when nanoparticles are included in hydrogels. The particles are introduced into the compositions after the formation of the gel and at the stage of gel formation, with or without the use of covalent binding processes. Nanoparticles in the composition of hydrogels provide a change in their mechanical properties and swelling characteristics and are also able to impart magnetic, optical properties, electrical conductivity, and improved antimicrobial properties to systems [54]. Hybrid hydrogels may include, for example, carbon-based nanoparticles, inorganic particles, and nanoparticles of semiconductors, nanoparticles of metals and their oxides, polymers, and liposomes. For example, the introduction of magnetite nanoparticles into the hydrogel is provided due to their stabilization by oleic acid and the amphiphilic nature of the main component of the hydrogel-modified hyaluronic acid [40]. Pyrene ligands contribute to the formation of hydrogel particles with a “core/shell” structure. In this case, magnetic particles bound to thrombin were employed to form fibrin-based skeletal structures [45]. Solid colloidal nanoparticles can be the core for a polyelectrolyte shell deposited using layer-by-layer adsorption technologies [40], as well as be part of the shell itself, providing magnetically controllable particles. For example, the authors [55] modified hollow microcapsules obtained from dextran sulfate and poly-L-lysine with maghemite particles.

Thus, a large group of biological and synthetic macromolecules is used to create innovative forms for the drug delivery. Although a more detailed overview of the methods of drug encapsulation and biodegradable matrices composition cannot be presented within the required limits of this chapter, it is obvious that the chemical composition of materials combined with the features of medical forms created on their basis (which include characteristics such as size, porosity, the presence of covalent crosslinking, and stimulus sensitivity) and the method of drug encapsulation predetermine the different fate of biodegradable materials in biological media and different profiles of the release.


4. Structure and morphology of ultrathin fibers as drug carriers

4.1 Influence of dipyridamole (DPD) on the structure and segmental mobility of a biopolymer in ultrathin fibers

Among the biopolymers used in dentistry, traumatology, orthopedics, cellular engineering, surgery, along with polylactides, the most common is, probably, PHB, the main representative of the biopolymers of the polyhydroxyalkanoate family [56]. This biopolymer is a product of microorganisms’ biosynthesis. It has high biocompatibility, the ability to rapid biosorption without the formation of toxic products, and increased resistance to oxidative degradation [57, 58].

In many previous reports, the structure formation of fibrillar materials based on PHB-containing dipyridamole (DPD) [59], chitosan [60], titanium dioxide and silicon nanoparticles [61], iron (III)-chlorporphyrin complexes [62], etc. was considered. In these works, the influence of low-molecular-weight substances on the structure of the crystalline and amorphous phases of PHB fibers was shown. This section will present a review of the structure and properties of ultrathin PHB fibers obtained by ES, containing immobilized drug – dipyridamole (DPD).

The structure of intercrystalline regions in fibers containing drug is directly related to the state of the crystalline phase of the carrier polymer by the spatial organization of the pass-through chains [63]. The latter, in the case of a high degree of crystallinity, experiences deformation and conformational spatial difficulties. Therefore, it is necessary to distinguish two alternative situations, the first: when the degree of crystallinity is low and the distance between the crystallites and polymer lamellae is large, so that a relatively low concentration of drug is evenly distributed between polymer crystals and should have little effect on the conformation of uncrystallized polymer molecules. An alternative situation arises when conditions are created in highly crystalline polymers (such as PGB, PLA, and polyamide-6) to realize the maximum degree of crystallinity, for example, as a result of solvent plasticization or as a result of temperature annealing. In this case, the proportion of polymer segments included in noncrystalline regions is relatively small, the mechanical and diffusion behavior of the polymer is determined by the state of the elongated flow chains, and the concentration of the introduced drug, related to the volume of the intercrystalline phase, may exceed its thermodynamic solubility. Then the excess of the low-molecular-weight component is displaced from the polymer volume with the formation of an independent drug phase on the fiber surface, and the drug molecules remaining in the volume are potentially capable of influencing the conformation and dynamics of polymer molecules in the intercrystalline regions of the polymer.

The structure and molecular dynamics of these regions in biopolymer fibers can be effectively investigated by the electron paramagnetic resonance (EPR) method using the microprobe technique of stable nitroxyl radicals developed at the IHF RAS (Moscow, Russian Federation) [64]. The satisfactory agreement between the calculated and experimental results demonstrated earlier indicates, first of all, the effectiveness of the selected two-phase model of the intercrystalline space of the PHB fiber, which consists of more and less dense regions with corresponding different rotational mobilities of a stable radical in these regions.

An increase in the concentration of dense regions in the presence of a drug was observed for the composite system PHB-DPD. The corresponding calculations revealeded that the volume fraction of more dense regions in the intercrystalline space of PHB fibers is much higher than the content of less dense regions. Moreover, when pharmaceuticals are added to the fiber, it continues to grow insignificantly following the sequence 0.90 (0) < 0.93 (1) < 0.94 (3) < 0.95 (5), where the numbers in parentheses indicate the percentage of DPD mass concentration. This result seems quite natural if we take into account that the specific enthalpy of melting of PHB changes in the same sequence, reflecting, as in the previous case, the degree of its crystallinity.

As the concentration of the dense fraction in the intercrystalline regions of PHB increases, a corresponding decrease in the rotation rate of the radical is observed, and, therefore, the segmental mobility of macromolecules slows down. For the slow component of the rotational mobility, similar changes in the values of the correlation time in the fiber are observed, namely, the value of this dynamic parameter increases with an increase in the DPD content in the sequence 6.6 × 10−9 s (0) > 7.1 × 10−9 s (1) > 8.8 × 10−9 s (3) > 9.0 × 10−9 s (5%), which indicates a slowdown in the molecular mobility of the radical and, accordingly, a decrease in the molecular mobility of PHB chains in intercrystalline areas. The correlation time for the fast component in all samples, except PHB with 5% DPD (7 × 10−10 s), was 2.4 × 10−10 s, i.e. more than an order of magnitude lower than the previous values. Consequently, a change in the crystalline phase of the polymer affects the dynamics of chains in denser regions of the intercrystalline space and practically does not affect the fast component of their mobility in less-dense intercrystalline regions. These results can be explained within the framework of the model of the heterogeneous (binary) structure of intercrystalline regions of polymers with a high degree of crystallinity [65, 66]. The perfection of the crystal structure of nanofibers depends on the conditions of electrospinning, namely, on the rate of solvent desorption from the formed fiber and the temperature regime of its cooling [63]. Indeed, fragments of the polymer chain with a predominance of the straightened conformation are mainly involved in the formation of paracrystalline regions and the recrystallization of the polymer. Therefore, probe molecules with a correlation time τ in paracrystalline regions are sensitive to changes in the degree of crystallinity of PHB. On the contrary, the molecules of the same probe, located at a considerable distance from the crystals in regions with fast segmental mobility and approaching in their dynamic characteristics to the amorphous phase of the polymer, are practically unaffected by the crystalline phase of the fiber and, within some limits, retain their constant value.

4.2 Influence of thermal heating of PHB fibers containing a dipyridamole on the molecular dynamics of polymer chains

During ES, as a result of difficult cooling and curing conditions of ultrathin fibers, their polymer structure can be far enough from the state of thermodynamic equilibrium. The imperfection of the crystalline phase and morphology is manifested in the insufficient orientation of the segments in the fiber, as well as in an atypically low degree of crystallinity [63]. Thermal annealing several tens of degrees below the melting point of PHB (annealing temperature 140°C) allows to sharply intensify segmental mobility and transfer the system to a more thermodynamically equilibrium state. Indeed, the time of thermal annealing at 2 h for PHB combinations with pharmaceuticals in the absence or low drug content for a number of samples significantly affects the rotation dynamics of the TEMPO radical. Whereas for highly crystalline fibers (with 3 and 5% DPD), this process has little effect on the rotational mobility of the probe, reflecting the segmental mobility of the chains (Figure 3).

Figure 3.

Dependences of the effective correlation time (τ) on the annealing time at 140°C: 1—PHB, 2—PHB with 1%, 3—PHB with 3%, 4—with 5% DPD.

One of the possible mechanisms of PHB recrystallization upon thermal annealing of the fiber is that a part of the chain segments with a high degree of orientation β, which was acquired during the electrospinning process, reaches values greater than the critical value h/L ∼ √2/l, where h is the distance between chain ends, L is the persistent chain length, and l is the length of a single segment [67]. As a result, the sorption capacity of the fiber due to an increase in the degree of crystallinity decreases. Indeed, measurements of the concentration of a stable radical in the fibers also show a sharp decrease in this value after annealing for 2 h. For example, in the initial PHB fibers, the concentration of the radical after absorption from vapors was 8 × 1015 spin/g, while this value after annealing for 2 h decreased to 1.4 × 1015 spin/g. A similar picture was observed for other biopolymer fibers containing a porphyrin complex.

The correspondence is observed between the content of the crystalline phase of PHB (structural characteristic) and the value of the correlation time (dynamic characteristic), so that molecular mobility decreases as the total volume of intercrystalline regions decreases due to an increase in crystallinity of PHB and, consequently, the involvement of an increasing number of mobile polymer segments in dense polymer regions with low mobility and high values of τ. Obtained results allow to conclude that, in the presence of a drug, from a PHB solution by ES, ultrathin fibers of various geometries with structures of various degrees of equilibrium and perfection are formed, which is true both for the crystalline phase and for intercrystalline regions, where redistribution between the amorphous and paracrystalline states of PHB is possible. The nonequilibrium of intercrystalline regions for PHB with the absence or low content of biologically active substances (DPD) is confirmed by a change in the rotational mobility of a stable radical and an increase in crystallinity. All results, including the effect of drug concentration on the fiber shape and its dynamic characteristics, are in good agreement with the thermophysical parameters of the system and should be directly applied if describing the directed and prolonged transport of bioactive compounds.

A natural continuation of these studies is the transition from the structural characteristics of ultrathin fibers to the study of the diffusion kinetics of a drug in their bulk. In the next section, the results of the dependence of diffusion transport as the main process responsible for controlled drug release on the geometric dimensions of the fiber, its crystallinity and the porosity of fibrillar films will be presented. The obtained results will be used to consider diffusion kinetics, which, in combination with the analysis of segmental dynamics, represent two fundamental processes that determine the rate and mechanism of controlled drug release from fibrillar therapeutic systems.


5. Diffusion and controlled drug release in ultrathin PHB fibers and fibrillar films based on them

The process of controlled targeted delivery of biologically active substances, for example, anti-inflammatory and antidiabetic pharmaceuticals or growth hormone [68, 69, 70], cannot be adequately described without considering their diffusion. In this regard, despite the impressive technological advances in the creation of ultrathin polymer fibers for various applications, the solution of diffusion and enzymatic-hydrolytic problems in the systems monofilament—biologically active substance (BAS) and fibrillar matrix—BAS is found in an extremely limited number of works (e.g. [71, 72, 73]) and requires in-depth analysis both experimental and theoretical levels. Consideration of the earlier-mentioned diffusion-kinetic problem was carried out using the example of ultrathin fibers with an encapsulated drug. Combining the results of transport in fibers and fibrillar membranes with the structural and dynamic characteristics of a biopolymer carrier, a consistent controlled release model is proposed that satisfactorily describes the combination of BAS diffusion and hydrolysis kinetics in an innovative therapeutic system based on a typical biodegradable polyester (PHB).

In Figure 4, typical kinetics of the release of dipyridamole (DPD) from fibers of different geometry consisting of ellipse-like and cylindrical structures and with different drug concentrations is plotted. As in the case of the “monolithic PHB film—drug” system [74], for the fibrillar matrix formed by ultrathin PHB fibers, the kinetic release profiles have two characteristic kinetic regions of linear and nonlinear form. The bimodal nature of controlled release is especially pronounced for fibers with a higher DPD content in the fiber (3 and 5 wt%). Under our proposed model, the initial nonlinear section mainly reflects a diffusion process with a characteristic drug diffusion coefficient, while the linear region corresponds to a kinetic process reflecting fiber degradation with partial loss of its mass. Briefly, the essence of this process is determined by the onset of hydrolysis of the ester functional groups of PHB. In the process of hydrolytic destruction, the drug encapsulated in the fiber passes into the surrounding aqueous solution not only as a result of diffusion but also as a result of partial fiber disintegration by the mechanism of surface destruction or erosion [75].

Figure 4.

Typical kinetic profiles of controlled release of DPD from PHB fibers. The concentration of DPD is 1 (1), 3 (2), and 5 wt% (3). scanning electron microscopy (SEM) photomicrographs illustrate the shape of the fibers; magnification 1000×.

In kinetic measurements, the fiber sample was immersed in a 60% aqueous solution of ethanol and the optical density of DPD samples was determined with a periodic sampling. The interval of sampling depended on the fiber composition and, accordingly, on the rate of drug release; it was from 1 to 30 min. The experiments lasted from several hours to several days. Termination of an increase in the optical density of DPD exhibiting two characteristic peaks at λ = 410 nm and a more intense peak at λ = 292 nm with an extinction coefficient of 31,260 L/(mol × cm) indicated that the drug release was completed.

With this formulation of the problem, the kinetic profile of drug release is described by the following system of equations. During the time interval Δt, the cumulative mass of the drug released from the fibrillar film by the time t (ΔMd(t)) is the sum of two terms: the amount of drug that entered the solution volume by the diffusion mechanism (ΔMD) and the amount of the drug that passed into solution as a result of partial loss of fiber mass ΔMf, which contained ΔMk grams of drug immobilized by polymer molecules and incapable of diffusion in a polymer medium. The above reflects a simple balance of the change in the mass of the active component in the fiber, which reflects the Eq. (1);


With a constant volume of the surrounding solution V and intensive stirring, external diffusion restrictions can be neglected. In this case, for a fibrillary film, by analogy with a monolithic nonporous film [75], one can write a differential equation showing the simultaneous contribution of drug diffusion and PHB hydrolysis by the zero-pillage mechanism:


where Сd and Gd are the total concentration of the drug that entered the external volume V by the time t and the concentration of the mobile drug fraction capable of diffusing in the polymer sample with the corresponding effective diffusion coefficient Deff, independent of the coordinate and time, and kh is the fiber mass loss constant, mainly due to the hydrolysis of the ester groups of PHB.

In a fibrillar film formed by a random interlacing of ultrathin fibers, Deff is determined by two coupled processes: the diffusion mobility of drug molecules in the fiber volume (Df) and its diffusion transport in the interfibrillar space filled with a solvent (Dw). Consequently, the total transfer can be approximated by the model of a pseudo-two-layer medium. Following the models of Crank [76] and Mackey-Mears [77], the relationship between the effective diffusion coefficient (Deff), individual diffusion constants (Df, Dw) and the geometric characteristics of the system (LM, Rf and Lw) has the form:


here Rf and Lw are the average characteristic dimensions of the length of the diffusion path of the drug in the fiber and interfiber space, respectively, and LM is the thickness of the fibrillar film.

Taking into account the symmetry of the film during double-sided desorption, its size and fiber diameter are divided by 2. For a cylindrical fiber, the length of the diffusion path is determined by its radius, while for Lw, we used a correction for the increase in the diffusion path due to the bending of the drug molecule around randomly arranged cylindrical fibers. The correction for impenetrable obstacles was first introduced in the classical work of Mackey and Mears [77] and was recently used to describe transport in a magnetic composite based on chitosan and PHB [78]:


where φf is the volume fraction of polymer fibers in the fibrillar film, the values of which, as well as the average radii, are given earlier in Table 1.

[Zn-ТFP] %τ 1010 sτ 1010 s, annealed at140 0Сα12,*%∆НM, J/gαс, %
Dense areaAmorphous areaDense areaAmorphous area

Table 1.

Effective correlation time (τ) and volume ratio of paracrystalline (dense) (α1) and amorphous (α2) regions in an ultrathin PHB fiber containing Zn-ТFP.

Subscripts 1 and 2 refer, respectively, to more dense (paracrystalline) and less dense (amorphous) regions of the intercrystalline space. ∆НM—specific heat of fusion and αс—degree of crystallinity of PHB.

Taking into account the introduced correlation of the diffusion path length in the aqueous phase, and also taking into account the correction for the degree of crystallinity of the polymer Fc = (1–αc), proposed in monograph [79], Eq. (4) takes a more detailed form:


allowing you to go to the comparison of the contribution of two processes that determine the total transfer process, namely, drug diffusion in the fiber and diffusion in the aqueous interfibrillar space of the PHB film. A preliminary estimate of the values of the diffusion coefficients of the drug in PHB (Df) showed that they are several orders of magnitude lower than the corresponding coefficient in the aqueous interfibrillar space, the last term on the right-hand side of Eq. (5) becomes much less than the first, which leads to a simplification of the sought expression for the diffusion coefficient of the drug in the hospital:


where the geometric (φf, LM) and structural (Fc = 1 – αс) characteristics were determined using the SEM and differential scanning calorimetry (DSC) data, respectively.

Simplification of Eq. (5) to expression (6) is possible only if the limiting stage of drug desorption from a fibrillar film of thickness LM is its diffusion in the polymer medium of the fiber. This is not universal, since with an increase in the volume of the interfibrillar space of the system, i.e. with a decrease in the proportion of fiber and its less dense packing, as well as with an increase in the concentration of hydrophilic groups in the polymer, i.e. with an increase in their affinity for a polar solvent (for example, for water) and a corresponding increase in Df, it is possible to take into account both terms, the values of which can be quite comparable.

The diffusion equation for an infinite cylinder which satisfactorily approximates the shape of the fiber and provided that the drug at the initial moment is uniformly distributed over the volume of the polymer was presented by Crank in his classic work [76]:


where r is the radial coordinate of diffusion, and the symbol Gd, as in Eq. (2), denotes the concentration of the mobile fraction of the drug in a cylindrical fiber with a corresponding constant diffusion coefficient Df and Rf as before denotes the average radius of the fiber. The initial and boundary conditions corresponding to drug desorption from a cylindrical fiber are as follows: Gd = G0d at t = 0 (at the initial moment of time) and under the condition 0 < x < Rf:

Gd = 0 at r = R (at the fiber-solution interface) and under the condition t > 0.

The second boundary condition is written from considerations of symmetry and indicates the absence of flux at the center of a single fiber ∂Gd/ ∂r = 0 at r = 0.

It was shown in [80] that the solution of Eq. (7) has the form of a power function and makes it possible to obtain the dependence of the cumulative amount of pharmaceuticals coming from the fiber into the environment by the diffusion mechanism ΔMD(t) on its contact time with this medium t:


where ΔMD∞ is the limiting value of ΔMD under the condition t → ∞.

Note that Eqs. (7) and (8), describing the diffusion of drug through the side walls of an infinite cylinder, are valid provided that the length of the fiber exceeds its radius by at least five times [81], and this requirement is strictly fulfilled for fibers of practically infinite length, obtained by the technology of electrospinning. In addition, the last equation is valid provided that the inequality ΔMD/ΔMD∞ < 0.4. The combination of Eqs. (2) and (8) gives the final expression for the release of the drug from the cylindrical fibers, both taking into account the weight loss during hydrolysis and due to the diffusion of the drug:


where kc = kh - [Df/Rf2]. The positive sign in the equation shows that under the given conditions for PHB, the inequality kh > [Df/Rf2] is fulfilled.

The experimental curves shown in Figure 4 correspond to Eq. (9), which allows the calculation of the fiber diffusion coefficients Df. For calculations, these drug release curves were presented in diffusion coordinates ΔMD/ΔMD∞ ∼ t1/2, as shown in Figure 5. Comparison of the experimental results presented in Figure 3 with the corresponding symbols, and calculations according to Eqs. (7) and (9), shown by solid lines, indicates their good agreement and demonstrates the possibility of using this model to assess the diffusion characteristics of ultrathin cylindrical fibers. Table 2 exhibits the concentration of mobile СD and immobilized Сh drug fractions.

Figure 5.

Kinetic curves of DPD release from PHB fibers, presented in diffusion coordinates. The concentrations of DPD are 1 (1), 3 (2), and 5 wt.% (3); LM is the total thickness of the fibrillar film.

DPD, wt %Rf × 104, cm2Df × 1012, cm2/sСD × 102, g/gСh × 102, g/gkc × 105, s−1kh × 105, s−1 (53%)1.65 (34%)4.41.4 (56%)0.89 (32%)6.11.1
19.8*6.9*0.78 (78%)0.22 (22%)9.48.5*

Table 2.

Sorption, kinetic, and diffusion characteristics of the PHB-DPD fiber system. Concentration of mobile СD and fixed Сh fractions of DPD in fiber; СExt—concentration of DPD in the interfiber space. The percentages of each fraction are indicated in parentheses.

The conditional calculations of the corresponding characteristics, calculated approximating the ellipsoids of revolution with cylinders of larger diameter.

The effect of a sharp drug release at the initial portion of the kinetic curves is called the “burst effect” [82]. In our case, it is mainly associated with the residual drug concentration displaced from the fiber volume during the electrospinning process. The appearance of this fraction does not exceed 10% and is observed only for ultrafine fibers with a high drug loading (> 3%). When the sample is immersed in an aqueous medium, it is removed from the film due to rapid diffusion mobility in the hydrated interfibrillar space. It should also be noted that the concentration of mobile drug molecules decreases with an increase in crystallinity and therefore reaches a maximum value for fibers containing 1% DPD, having the lowest crystallinity of 38%. The last line of Table 2 shows the constants of the drug yield due to the loss of fiber mass during hydrolysis (kh). It can also be noted here that the rate of hydrolysis decreases with an increase in the crystallinity of the fiber and the lower, the higher the concentration of the drug in the fiber.

The totality of the results known in the literature and presented in this chapter allows us to draw two conclusions. First, the intercrystalline regions of ultrathin PHB fibers have a close packing of chains, which is significantly higher than the density in the amorphous region of the film, which is confirmed by EPR measuring. Second, under the same conditions, the values of drug diffusion coefficients in ultrathin and highly crystalline PHB fibers, as well as in its spherical microparticles containing DPD [83], are several orders of magnitude lower than similar characteristics measured for PHB films.


6. Conclusion

Over the past 15 years, there has been not only significant progress at the level of theory and modeling of the processes of directed transport of biologically active compounds, but at the same time a large number of polymer and hybrid materials were created for modern therapy, providing an innovative component for drug delivery. Today, the main trend is the accumulation of experimental data showing how nanotechnological processes, including electrospinning, affect the characteristics of the developed means of targeted drug delivery.


  1. 1. Ariga K, Vinu A, Miyahara M. Recent progresses in bio-inorganic nanohybrids. Current Nanoscience. 2006;2(3):197
  2. 2. Colilla M, Manzano M, Vallet-Ragí M. Recent advances in ceramic implants as drug delivery systems for biomedical applications. International Journal of Nanomedicine. 2013;3(4):403
  3. 3. Dubey KA, Chaudhari CV, Bhardwaj YK, Varshney L. Polymers, blends and nanocomposites for implants, scaffolds and controlled drug release applications. Advanced Structured Materials. 2017;66:1
  4. 4. Hernández-Gómez LH, Beltrán-Fernández JA, Ramírez-Jarquín M, Pava-Chipol N, Pérez-Montejo S. Characterization of scaffold structures for the development of prostheses and biocompatible materials. Advanced Structured Materials. 2019;92:471
  5. 5. Mann JL, Yu AC, Agmon G, Appel EA. Supramolecular polymeric biomaterials. Biomaterials Science. 2018;6(1):10
  6. 6. LaVan DA, McGuire T, Langer R. Small-scale systems for in vivo drug delivery. Nature Biotechnology. 2003;21(10):1184
  7. 7. Ishihara T, Mizushima T. Techniques for efficient entrapment of pharmaceuticals in biodegradable solid micro/nanoparticles. Expert Opinion on Drug Delivery. 2010;7(5):565
  8. 8. Haidar MK, Eroğlu H. Nanofibers: New insights for drug delivery and tissue Engineering. Current Topics in Medicinal Chemistry. 2017;17(13):1564
  9. 9. Bhardwaj N, Kundu SC. Electrospinning: A fascinating fiber fabrication technique. Biotechnology Advances. 2010;28(3):325
  10. 10. Zhang Y, Chen Q, Ge J, Liu Z. Controlled display of enzyme activity with a stretchable hydrogel. Chemical Communications. 2013;49:84
  11. 11. Park S, Lee Y, Kim DN, Park S, Jang E, Koh WG. Entrapment of enzyme-linked magnetic nanoparticles within poly(ethylene glycol) hydrogel microparticles prepared by photopatterning. Reactive and Functional Polymers. 2009;69(5):293. DOI: 10.1016/j.reactfunctpolym.2009.02.001
  12. 12. Thornton PD, Mart RJ, Ulijn RV. Enzyme-responsive polymer hydrogen hydrogel particles for controlled release. Advanced Materials. 2007;19(9):1252
  13. 13. Gupta B, Revagade N, Hilborn J. Poly(lactic acid) fiber: An overview. Progress in Polymer Science. 2007;32(4):455
  14. 14. Zhang B, Yan X, He H-W, Ning X, Long Y-Z. Solvent-free electro-spinning: Opportunities and challenges. Polymer Chemistry. 2017;8(2):334
  15. 15. Streicher RM, Schmidt M, Fiorito S. Nanosurfaces and nanostructures for artificial orthopedic implants. Nanomedicine. 2007;2(6):861
  16. 16. Li D, Frey MW, Baeumner AJ. Electrospun polylactic acid nanofiber membranes as substrates for biosensor assemblies. Journal of Membrane Science. 2006;279(1-2):354
  17. 17. Mikheev AY, Shlyapnikov YM, Kanev IL, Avseenko AV, Morozov VN. Filtering and optical properties of free standing electrospun nanomats from nylon-4,6. European Polymer Journal. 2016;75:317
  18. 18. Miguel SP, Figueira DR, Simões D, Ferreira P, Correia IJ. Electrospun polymeric nanofibres as wound dressings: A review. Colloids and Surfaces B: Biointerfaces. 2018;169:60
  19. 19. Zhou B, Li Y, Deng H, Hu Y, Li B. Antibacterial multilayer films fabricated by layer-by-layer immobilizing lysozyme and gold nanoparticles on nanofibers. Colloids and Surfaces B: Biointerfaces. 2014;116:432
  20. 20. Cheng H, Yang X, Che X, Yang M, Zhai G. Biomedical application and controlled drug release of electrospun fibrous materials. Materials Science and Engineering. 2018;90:750
  21. 21. Rebia RA, Rozet S, Tamada Y, Tanaka T. Biodegradable PHBH/PVA blend nanofibers: Fabrication, characterization, in vitro degradation, and in vitro biocompatibility. Polymer Degradation and Stability. 2018;154:124
  22. 22. Malafeev KV, Moskalyuk OA, Yudin VE, Popova EN, Ivan’kova EM. Synthesis and properties of fibers prepared from lactic acid–glycolic acid copolymer. Polymer Science, Series A. 2017;59(1):53
  23. 23. Cao K, Liu Y, Olkhov AA, Siracusan V, Iordanskii AL. PLLA-PHB fiber membranes obtained by solvent-free electrospinning for short-time drug delivery. Drug Delivery and Translational Research. 2018;8:291
  24. 24. Bose S, Tarafder S. Calcium phosphate ceramic systems in growth factor and drug delivery for bone tissue engineering: A review. Acta Biomaterialia. 2012;8(4):1401
  25. 25. Sokolova VV, Radtke I, Heumann R, Epple M. Effective transfection of cells with multi-shell calcium phosphate-DNA nanoparticles. Biomaterials. 2006;27(16):3147
  26. 26. Dorati R, DeTrizio A, Modena T, et al. Biodegradable scaffolds for bone regeneration combined with drug-delivery systems in osteomyelitis therapy. Pharmaceuticals. 2017;10(4):96. DOI: 10.3390/ph10040096
  27. 27. Nandi SK, Bandyopadhyay S, Das P, Samanta I, Mukherjee P, Roy S, et al. Understanding osteomyelitis and its treatment through local drug delivery system. Biotechnology Advances. 2016;34(8):1305
  28. 28. Das S, Baker AB. Biomaterials and nanotherapeutics for enhancing skin wound healing. Frontiers in Bioengineering and Biotechnology. 2016;4:82
  29. 29. Tran CD, Duri S, Harkins AL. Recyclable synthesis, characterization and antimicrobial activity of chitozan-based polysaccharide composite materials. Journal of Biomedical Materials Research Part A. 2013;0(8):2248
  30. 30. Tan H, Chu CR, Payne K, Marra KG. Injectable in situ forming biodegradable chitosan-hyaluronic acid based hydrogels for cartilage tissue engineering. Biomaterials. 2009;30(13):2499. DOI: 10.1016/j.biomaterials.2008.12.080
  31. 31. Song Y, Nagai N, Saijo S, Kaji H, Nishizawa M, Abe T. In situ formation of injectable chitosan-gelatin hydrogels through double crosslinking for sustained intraocular drug delivery. Materials Science & Engineering. C, Materials for Biological Applications. 2018;1(88):1
  32. 32. Vikhoreva G, Bannikova G, Stolbushkina P, Panov A, Drozd N, Makarov V, et al. Preparation and anticoagulant activity of a low-molecular-weight sulfated chitosan. Carbohydrate Polymers. 2005;62:327
  33. 33. Xue J, Zhao W, Nie S, Sun S, Zhao C. Blood compatibility of polyethersulfone membrane by blending a sulfated derivative of chitosan. Carbohydrate Polymers. 2013;95:64. DOI: 10.1016/j.carbpol.2013.02.033
  34. 34. Huang X, Wang R, Lu T, Zhou D, Zhao W, Sun S, et al. Heparin-like chitosan hydrogels with tunable swelling behavior, prolonged clotting times, and prevented contact activation and complement activation. Biomacromolecules. 2016;17:4011. DOI: 10.1021/acs.biomac.6b01386
  35. 35. Zhang L, Ma Y, Pan X, Chen S, Zhuang H, Wang S. A composite hydrogel of chitosan/heparin/poly (γ-glutamic acid) loaded with superoxide dismutase for wound healing. Carbohydrate Polymers. 2018;180:168
  36. 36. Qiao X, Peng X, Qiao J, Jiang Z, Han B, Yang C, et al. Evaluation of a photocrosslinkable hydroxyethyl chitosan hydrogel as a potential drug release system for glaucoma surgery. Journal of Materials Science. Materials in Medicine. 2017;28:149
  37. 37. Nguyen HX, O’Leary EA. An in vitro thrombolysis study using a mixture of fast-acting and slower release microspheres. Pharmaceutical Research. 2016;33:1552. DOI: 10.1007/s11095-016-1897-1
  38. 38. Dobiasch S, Szanyi S, Kjaev A, Werner J, Strauss A, Weis C, et al. Synthesis and functionalization of protease-activated nanoparticles with tissue plasminogen activator peptides as targeting moiety and diagnostic tool for pancreatic cancer. Journal of Nanobiotechnology. 2016;14:81. DOI: 10.1186/s12951-016-0236-3
  39. 39. Zhang LL, Li P, Li YM, Wang AQ. Preparation and characterization of magnetic alginate-chitosan hydrogel beads loaded matrine. Drug Development and Industrial Pharmacy. 2012;38(7):872. DOI: 10.3109/03639045.2011.630397
  40. 40. Zhang Y, Sun Y, Yang X, Hilborn J, Heerschap A, Ossipov DA. Injectable in situ forming hybrid iron oxide-hyaluronic acid hydrogel for magnetic resonance imaging and drug delivery. Macromolecular Bioscience. 2014;14(9):1249
  41. 41. Iordanskii AL, Rogovina SZ, Berlin AA. Current state and developmental prospects for nanopatterned implants containing drugs. Review Journal of Chemistry. 2013;3(2):117-132
  42. 42. Leoni L, Desai TA. Nanoporous biocapsules for the encapsulation of insulinoma cells: Biotransport and biocompatibility considerations. IEEE Transactions on Biomedical Engineering. 2001;48(11):1335
  43. 43. Mandracchia D, Tripodo G. Chapter 1: Micro and nano-drug delivery systems. In: Silk-based Drug Delivery Systems. 2020. pp. 1-24. DOI: 10.1039/9781839162664-00001. eISBN: 978-1-83916-266-4
  44. 44. Rajangam T, An SS. Fibrinogen and fibrin based micro and nano scaffolds incorporated with drugs, proteins, cells and genes for therapeutic biomedical applications. International Journal of Nanomedicine. 2013;8:3641
  45. 45. Ziv-Polat O, Skaat H, Shahar A, Margel S. Novel magnetic fibrin hydrogel scaffolds containing thrombin and growth factors conjugated iron oxide nanoparticles for tissue engineering. International Journal of Nanomedicine. 2012;7:1259
  46. 46. Agapova OI. Silk fibroin and spidroin bioengineering constructions for regenerative medicine and tissue engineering (review). Sovremennye Tehnologii v Medicine (Modern Technologies in Medicine). 2017;9(2):190-206
  47. 47. Gomzyak VI, Demina VA, Razuvaeva EV, Sedush NG, Chvalun SN. Biodegradable polymer materials for medical applications: From implants to organs. Fine Chemical Technologies. 2017;12(5):5-20. (In Russ.)
  48. 48. Kinard LA, Kasper FK, Mikos AG. Synthesis of oligo(poly(ethylene glycol) fumarate). Nature Protocols. 2012;7(6):1219
  49. 49. Santoro M, Tatara AM, Mikos AG. Gelatin carriers for drug and cell delivery in tissue engineering. Journal of Controlled Release. 2014;0:210. DOI: 10.1016/j.jconrel.2014.04.014
  50. 50. Modaresi SM, Mehr SE, Faramarzi MA, Gharehdaghi EE, Azarnia M, Modarressi MH, et al. Preparation and characterization of self-assembled chitosan nanoparticles for the sustained delivery of streptokinase: An in vivo study. Pharmaceutical Development and Technology. 2014;19:593. DOI: 10.3109/10837450.2013.813542
  51. 51. Jin HJ, Zhang H, Sun ML, Zhang BG, Zhang JW. Urokinase-coated chitosan nanoparticles for thrombolytic therapy: Preparation and pharmacodynamics in vivo. Journal of Thrombosis and Thrombolysis. 2013;36:458
  52. 52. Mourino V, Boccaccini AR. Bone tissue engineering therapeutics: Controlled drug delivery in three-dimensional scaffolds. Journal of the Royal Society Interface. 2010;7(43):209. DOI: 10.1098/rsif.2009.0379
  53. 53. Vatchha SP, Kohli A, Tripathi SK, Nanda SN, Pradhan P, Shiraz SM. Biodegradable implants in orthopaedics (review). Annals of International Medical and Dental Research. 2015;1(1):3-8
  54. 54. Zhao F, Yao D, Guo R, Deng L, Dong A, Zhang J. Composites of polymer hydrogels and nanoparticulate systems for biomedical and pharmaceutical applications. Nanomaterials. 2015;5(4):2054. DOI: 10.3390/nano5042054
  55. 55. Bukreeva TV, Sulyanov SN, Korotkov NY, Rumyantseva SS, Marchenko IV, Funtov KO, et al. In situ synthesis and characterization of magnetic nanoparticles in shells of biodegradable polyelectrolyte microcapsules. Materials Science and Engineering. 2014;45:225
  56. 56. Albuquerque PBS, Malafaia CB. Perspectives on the production, structural characteristics and potential applications of bioplastics derived from polyhydroxyalkanoates. International Journal of Biological Macromolecules. 2018;107:615
  57. 57. Sevastianov VI, Perova NV, Shishatskaya EI, Kalacheva GS, Volova TG. Production of purified polyhydroxyalkanoates (PHAs) for applications in contact with blood. Journal of Biomaterials Science, Polymer Edition. 2003;14:1029
  58. 58. Artsis MI, Bonartsev AP, Iordanskii AL, Bonartseva GA, Zaikov GE. Biodegradation and medical application of microbial poly(3-hydroxybutyrate). Molecular Crystals and Liquid Crystals. 2010;523:593. DOI: 10.1080/15421401003726519
  59. 59. Ivantsova EL, Kosenko RY, Iordanskii AL, Rogovina SZ, Prut EV, Filatova AG, et al. Structure and prolonged transport in a biodegradable poly(r-3-hydroxybutyrate)-drug system. Polymer Science, Series A. 2012;54(2):87-93
  60. 60. Karpova SG, Popov AA, Lomakin SM, Shilkina NG, Iordanskii AL, Klenina NS, et al. Changes in the structural parameters and molecular dynamics of polyhydroxybutyrate-chitosan mixed compositions under external influences. Russian Journal of Physical Chemistry B. 2013;7(3):225-231
  61. 61. Olkhov AA, Karpova SG, Staroverova OV, Kucherenko EL, Ishchenko AA, Iordanskii AL. Effect of external factors on the structure of ultrathin fibers of poly(3-Hydroxybutyrate) and dipyridamole. Fibre Chemistry. 2017;48(4):1
  62. 62. Karpova SG, Olkhov AA, Shilkina NG, Iordanskii AL. Investigation of biodegradable composites of poly(3-hydroxybutyrate) ultrathin fibers modified by a complex of iron(III) with tetraphenylporphyrin. Polymer Science, Series A. 2017;59(3):343. DOI: 10.1134/S0965545X17030075
  63. 63. Cho D, Zhmayev E, Joo YL. Structural studies of electrospun nylon 6 fibers from solution and melt. Polymer. 2011;52:4600
  64. 64. Vasserman AM, Buchachenko AL, Kovarskii AL, Neiman MB. Study of molecular motion in polymers by the paramagnetic probe method. Polymer Science U.S.S.R. 1968;10(8):2238. DOI: 10.1016/0032-3950(68)90317-1
  65. 65. Di Lorenzo ML, Gazzano M, Righetti MC. The role of the rigid amorphous fraction on cold crystallization of poly(3-hydroxybutyrate). Macromolecules. 2012;45(14):5684. DOI: 10.1021/ma3010907
  66. 66. Kamaev PP, Aliev II, Iordanskii AL, Wasserman AM. Molecular dynamics of the spin probes in dry and wet poly(3-hydroxybutyrate) films with different morphology. Polymer. 2001;42(2):515. DOI: 10.1016/S0032-3861(00)00339-6
  67. 67. Ozerin AN, Zurov YA, Selikhova VI, Chvalun SN, Bakeyev NF. Methods of investigation use of the method of measuring the absolute intensity of small angle x-ray scattering for the study of the structure of amorphous regions in oriented polyethylene films. Polymer Science U.S.S.R. 1976;18(9):2434-2442
  68. 68. Huynh CT, Kang SW, Li Y, Kim BS, Lee DS. Controlled release of human growth hormone from a biodegradable pH/temperature-sensitive hydrogel system. Soft Matter. 2011;7(19):8984
  69. 69. Hu C, Liu S, Zhang Y, Li B, Yang H, Fan C, et al. Long-term drug release from electrospun fibers for in vivo inflammation prevention in the prevention of peritendinous adhesions. Acta Biomaterialia. 2013;9:7381
  70. 70. Padmanabhan J, Kyriakides TR. Nanomaterials, inflammation, and tissue engineering. Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology;7(3):355
  71. 71. Keshavarz P, Ayatollahi S, Fathikalajahi J. Mathematical modeling of gas–liquid membrane contactors using random distribution of fibers. Journal of Membrane Science. 2008;325:98
  72. 72. Petlin DG, Amarah AA, Tverdokhlebov SI, Anissimov YG. A fiber distribution model for predicting drug release rates. Journal of Controlled Release. 2017;258:218
  73. 73. Nakielski P, Kowalczyk T, Zembrzycki K, Kowalewski TA. Experimental and numerical evaluation of drug release from nanofiber mats to brain tissue. Journal of Biomedical Materials Research—Part B: Applied Biomaterials. 2015;103(2):282
  74. 74. Iordanskii AL, Rogovina SZ, Kosenko RY, Ivantsova EL, Prut EV. Development of a biodegradable polyhydroxybutyrate-chitosan-rifampicin composition for controlled transport of biologically active compounds. Doklady Physical Chemistry. 2010;431(2):60-62. DOI: 10.1134/S0012501610040020
  75. 75. Sevim K, Pan J. A model for hydrolytic degradation and erosion of biodegradable polymers. Acta Biomaterialia. 2018;66:192-199. DOI: 10.1016/j.actbio.2017.11.023
  76. 76. Crank J. The Mathematics of Diffusion. Oxford: Clarendon Press, 1992.
  77. 77. Mackie JS, Meares P. The diffusion of electrolytes in a cation-exchange resin membrane I. Theoretical. Proceedings of the Royal Society of London. Series A. 1995;232(1191):498. DOI: 10.1098/rspa.1955.0234
  78. 78. Bychkova AV, Iordanskii AL, Kovarski AL, Sorokina ON, Kosenko RY, Markin VS, et al. Nanotechnologies in Russia. 2015;10(3–4):325
  79. 79. Mikhailov YM, Ganina LV, Shapaeva NV, Chalykh AE. Diffusion of Low-molecular mass substances in the region of polymer-glass transition. Polymer Science, Series B. 1999;41(5-6):120-123
  80. 80. Siepmann J, Siepmann F. Mathematical modeling of drug delivery. International Journal of Pharmaceutics. 2008;364:328
  81. 81. Jamstorr BЕ. Acta Universitatis Upsaliensis. Vol. 884. Uppsala: Digital Comprehensive Summaries of Uppsala Dissertations from the Faculty of Science and Technology. p. 79
  82. 82. Huang X, Brazel CS. On the importance and mechanisms of burst release in matrix-controlled drug delivery systems. Journal of Controlled Release. 2001;73(2–3):121. DOI: 10.1016/S0168-3659(01)00248-6
  83. 83. Bonartsev AP, Livshits VA, Makhina TA, Myshkina VL, Bonartseva GA, Iordanskii AL. Controlled release profiles of dipyridamole from biodegradable microspheres on the base of poly(3-hydroxybutyrate). Express Polymer Letters. 2007;1(12):797

Written By

Anatoly A. Olkhov, Svetlana G. Karpova, Anna V. Bychkova, Alexandre A. Vetcher and Alexey L. Iordanskii

Submitted: 30 November 2021 Reviewed: 09 June 2022 Published: 14 July 2022