Functionalized silk fibroin nanofibers as drug carriers: Advantages and challenges

Journal Pre-proof Functionalized silk fibroin nanofiber as drug carriers: Advantages and challenges

Mehdi Farokhi, Fatemeh Mottaghitalab, Rui L. Reis, Seeram Ramakrishna, Subhas C. Kundu PII:

S0168-3659(20)30112-7

DOI:

https://doi.org/10.1016/j.jconrel.2020.02.022

Reference:

COREL 10175

To appear in:

Journal of Controlled Release

Received date:

2 December 2019

Revised date:

11 February 2020

Accepted date:

11 February 2020

Please cite this article as: M. Farokhi, F. Mottaghitalab, R.L. Reis, et al., Functionalized silk fibroin nanofiber as drug carriers: Advantages and challenges, Journal of Controlled Release (2020), https://doi.org/10.1016/j.jconrel.2020.02.022

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© 2020 Published by Elsevier.

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Functionalized Silk Fibroin Nanofiber as Drug Carriers: Advantages and Challenges Mehdi Farokhia*, Fatemeh Mottaghitalabb*, Rui L. Reisc, Seeram Ramakrishnad, Subhas C

Kunduc a b

National Cell Bank of Iran, Pasteur Institute of Iran, Tehran, Iran

Nanotechnology Research Centre, Faculty of Pharmacy, Tehran University of Medical Sciences, Tehran, Iran c

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Centre for Nanofibers and Nanotechnology, National University of Singapore, Singapore,

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117576, Singapore

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*Corresponding authors:

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Dr. Mehdi Farokhi: Email: [email protected], Tel/Fax:+9821 64112358

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Dr. Fatemeh Mottaghitalab:Email: [email protected], Tel/Fax:+9821 64121530

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3Bs Research Group, I3Bs - Research Institute on Biomaterials, Biodegradable and Biomimetics, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, University of Minho, Guimaraes, Portugal

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Journal Pre-proof Abstract Electrospinning is well thought of as the most potent nanofiber producing technique, which is applicable in biomedical fields. The generation of electrospun nanofibers as drug carriers has widely been shown much interest over the past years. Electrospun nanofibers meet various advantages as drug delivery platforms including high surface area, acceptable mechanical properties based on the choice of the polymer, high processability, and the possibilities for

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surface modifications. Silk fibroin protein has gained a great attention as a drug delivery carriers

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due to ease of purification, sterilization, processability without using chemical crosslinkers, good

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biocompatibility, tailorable biodegradability, low immunogenicity, and high capacity to stabilize

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the loaded drugs. These characteristics along with advantageous benefits of electrospinning

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provide opportunities for producing suitable nanofibers based on silk fibroin for drug delivery purposes. It is also possible to incorporate various functional moieties to the electrospun silk

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fibroin nanofibers to enhance its biological activities. This review covers the progress in

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electrospinning of silk fibroin as a drug carrier in recent years.

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Keywords: Electrospinning; Silk fibroin; Nanofibers; Drug delivery systems.

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1. Introduction The main purpose of advanced drug carriers is to deliver the drug into the desired place and thereby lower the systemic dose to decrease the side effects. For this, it is essential to develop optimized systems in terms of structure and function. Recently, the nanofibrous drug carriers gained much attention due to the tailorable spatiotemporal structure. These carriers can easily be

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implanted on the targeted sites and release the drugs locally. Delivering the drug at the local area

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represents many advantages such as fine-tuning of the drug release rate at the site of disease,

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which is highly preferable in comparison to intravenous or oral delivery routes. Moreover, off-

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target side effects can potentially be avoided due to the delivery of drugs locally [1]. Besides, the bioactive molecules e.g., nucleic acids, proteins, and peptides generally cause systemic toxicity

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in high doses. They are metabolized rapidly in the body; thus, use of nanofibers as a local drug

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delivery carrier acts as a protective agent to preserve drug bioactivity. The release kinetics of drug from nanofibrous mats can also be sustained by tuning the binding efficiency between

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pharmaceutical agents and materials, nanofiber porosity, and diameter [2]. However, some

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challenges for local drug delivery are remained e.g., administration—achieving suitable surface coverage, fixation of device at the injured site, and diffusion of the high amount of drug within the tissue. Moreover, it is necessary to use flexible materials with good loading efficiency for implantation because the irregular shape of the surgical area limits the application of rigid biomaterials. For this, using nanofibrous mats or hydrogel containing an appropriate amount of drugs for covering the tissue are considered to be good strategies for implantation. It is also essential to investigate the inflammatory responses and healing mechanisms after local administration of drugs post-surgery. For instance, acute inflammation after surgery causes

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Journal Pre-proof edema which promotes more drug release into the implanted site. Moreover, the foreign body reactions often form a dense layer of collagenous fibers around the implant which make a barrier for normal diffusion of drug [3]. The degradation rate of the polymeric matrix is another problem for local drug delivery because the fast degradation rate of some polymers may lead to the burst release of the drug at the site of injury. To address these issues, designing an appropriate local drug delivery systems for maximum healing is essential.

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Different methods are developed for fabricating nanofibrous carriers; however, electrospinning is

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mostly preferred because of simplicity, low cost, and ease of processability from different

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polymeric sources. Furthermore, it is possible to produce nanofibers on a large scale applicable

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for industrial purposes [4]. The electrospinning setup comprises a syringe pump, a

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negative/positive high electrical field, and a grounded collector (Figure 1a). The electrospinning is based on applying high voltage to induce electrostatic repulsion to the highly charged

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polymeric solution to collect random, aligned, and oriented nanofibers on the surface of the

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collector. The repulsion forces overcome the tensions at the surface of the polymeric solution upon applying the high voltage. Finally, the charged jet of the polymer is expelled from the

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Taylor cone, the unstable jet moves rapidly between the collector and capillary tip and the solvent is evaporated. Then, the polymer nanofibers are placed on the collector plate (Figure 1b). It is worthy to note that simple or non-woven electrospun nanofibers are mostly used for biomedical and industrial applications; however, it is attempted to prepare more complicated membranes. Considering the effect of the electric field on the features of polymeric jet might help to better fabricate sheets with ordered nanofibers. To prepare nanofibers with desired orientations, the rotating collectors and drums can be used. The electric field can be manipulated

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Journal Pre-proof as well. Recent attempts of this technology lead to the greater understanding and emphasis on the

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production of commercially and more ordered sheets [5, 6].

Figure 1. a) Schematic illustration of electrospinning setup comprising of a syringe filled with the polymer solution, a syringe pump with a constant and controlled speed, and a high voltage applying current to the needle for charging the fluid. By applying the suitable voltage, a Taylor cone is formed and the jet ejects to the collector plate. b) The influence of the increasing voltage on formation of Taylor cone (dark-colored tip). At the low voltage, pendant drop is firstly formed and secondly the cone is formed on the tip. By gradual increased in the voltage, the drop 5

Journal Pre-proof size is decreased and the cone is formed at the tip of the needle. More increase in the voltages triggers the formations of fibers without Taylor cone formation on the blunt tip of the capillary. Reprinted with permission from [7]. The capability of electrospinning in producing nanoscale sheets provides opportunities for developing the versatile nanofibrous scaffolds for different biomedical applications [8]. Many nano/microfibrous meshes are fabricated from natural and synthetic sources such as chitosan

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(CS), hyaluronic acid (HA), collagen, gelatin, dextran (Dex), poly (lactic acid) (PLA),

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poly(lactic-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL) and their combinations [9, 10]. Among the natural biopolymers, silk protein fibroin (SF) is taken into consideration as one

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of the most useful polymers as a drug carrier. This is due to its high biocompatibility, tailorable

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biodegradability, low bacterial attachment, high structural integrity, and the breadth of

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applications, possibilities of fabrication in customizable sizes, good mechanical characteristics, simpler to re-engineer, interactions with bio-molecular components, and cost-effectiveness.

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Recently, nanofibrous SF based scaffolds have been used as potential carriers for drug delivery

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[11, 12]. Moreover, the in vitro and in vivo biodegradation of SF is tailorable during processing

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[13]. Manipulating the crystallinity and biodegradation rate of SF matrix can also control the release kinetics of drugs [14]. Besides, the post-modification of nanofibrous SF mats containing drugs enhances the cell behaviors in terms of proliferation, differentiation, and attachment [15]. It has been reported that the electrospinning process has not been reduced the activity of the incorporated bioactive molecules e.g., growth factors into the electrospun solution [16]. We and others have reviewed some of the biomedical aspects of SF-based scaffolds for tissue engineering applications [17, 18] and drug delivery [19, 20]. Here, we review the studies on the use of electrospun SF nanofibers for delivering different types of biomolecules. Based on our knowledge, there is no comprehensive review paper on this topic. The first part of the paper 6

Journal Pre-proof emphasizes on the role of electrospun nanofibers as drug delivery systems and the second part deals with some important characteristics of SF polymer especially electrospun SF nanofibers in biomedical applications. Finally, the capabilities of various electrospun SF nanofibers functionalized with growth factors, hormones, antibiotics, antioxidants, vitamins, and ions are discussed in detail. 2. Recent progress and trending in electrospun nanofibers

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Micro- and nanoscale fibers have been fabricated through various methods including wet

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spinning, rotary spinning, self-assembly, microfluidics, and electrospinning during the past few

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decades [21]. Among them, electrospinning has been extensively used due to its high feasibility

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and simplicity. In 1902, electrospinning has been first used by Morton [22] and Cooley [23] for dispersing fluids by electrostatic force. In 1934, polymeric filaments have been fabricated

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through electrospinning by Formhals [23]. Since then, various synthetic and natural polymers

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(>75 different types) have been electrospun to nanofibers [9]. In 2002, Kenawy et al. have used electrospinning to produce nanofibers applicable to drug delivery for the first time [24]. From

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then till now, different types of electrospun nanofibers have been used for drug delivery due to

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acceptable properties including simple processability, high mechanical strength, the capacity for different functionalization, and the ease of producing nanofibers with the high surface area. The structural properties of polymeric solution dictates the characteristics of electrospun membranes in terms of morphology, porosity, and possibilities for surface modification, which consequently influence the release kinetics of drugs from the polymeric matrices [25]. Various therapeutic components e.g., siRNA, DNA, oligonucleotides, peptides, proteins, and small molecules can be incorporated into nanofibrous sheets/mats fabricated by electrospinning. To enhance the capabilities of nanofibers for delivery of therapeutic agents, their surfaces can be modified for

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Journal Pre-proof better conjugation with bioactive components via physical or chemical modifications. Nanofibers modification facilitates the immobilization of specific ligands that promotes the cellular activities in terms of proliferation and differentiation and mimic the structure of extracellular matrix (ECM) [26]. 3. Advantages of electrospun nanofibers for drug delivery Controlling the release rate of drugs has many advantages such as adjusting the kinetics of drug

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release at the optimum dose that increases the bioavailability and solubility of drugs. Oral

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administration of drugs in the forms of tablets, granules, and capsules is the common route for

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drug delivery. Moreover, parenteral administrations of drugs e.g., subcutaneous, intravenous,

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intramuscular, and intra-arterial are also used [27]. Moreover, transdermal drug delivery is another effective method to avoid some restrictions of such delivery system e.g., low

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bioavailability of some oral drugs [28]. Nanofibrous mats as transdermal drug delivery produced

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by electrospinning possess many beneficial characteristics as drug vehicles. By using electrospinning, it is possible to generate nanofibers from different natural and synthetic

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polymers [29]. Based on clinical application, the release kinetics of drugs from nanofibrous mats

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can be tailored by optimizing the porosity, morphology, and diameter of nanofibers and also by adjusting the drug/polymer ratio [30]. The frequency of topical drug administration can be reduced by using electrospun nanofiber mats with the ability to sustain the release rate of drugs. Besides, the nanofibers can be used as mass transporters due to high porosity and good interconnectivity between the pores [31, 32]. For desired biomedical applications, it is possible to incorporate both hydrophilic and/or hydrophobic drugs in the core or shell of the nano- or microfibers during the electrospinning process [33]. In some cases, electrospun nanofibers have been considered more beneficial in drug delivery approaches due to the ability to control the

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Journal Pre-proof initial burst release of drugs with controlled zero-order drug release profile [34, 35]. Overall, electrospinning has shown a great potential in producing ultrafine nanofibers; however, some drawbacks such as inability to fully tune the shape, size and process parameters, and choosing biocompatible solvent may hinder the applications of this method for specific aim. 4. Common electrospinning setups for fabricating drug loaded nanofibers

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4.1. Blend electrospinning Blend electrospinning is a low cost and simple technique in which the drug is blended with the

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solution of polymer before performing the electrospinning parameters. High drug loading with

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homogenous drug spreading with risk of burst release of drug obtain with this method. The

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physico-chemical properties of the polymers define the quality of interactions between drugs and

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polymeric solutions that consequently influence the drug dispersion within the nanofibers, drug encapsulation efficiency, and release kinetics of drugs [36]. It is reported that by choosing an

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appropriate solvent and adjusting the processing conditions, many biomolecules such as

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dexamethasone [37], retinoic acid [38], proteins (e.g. bovine serum albumin), and peptides can be incorporated into the electrospun mats [39]. Two common mechanisms are considered for the

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release of drugs from nanofibers including dissolution/erosion or desorption/diffusion from the sheets. The drug release mechanism from nonbiodegradable polymers is mainly based on diffusion [40, 41]. However, the blend electrospinning for protein delivery faces some challenges due to applying some harsh conditions during protein encapsulation. As many synthetic polymers like PLGA, PCL, polyurethane (PU) are soluble in organic solvents; however, they have also negative effects on the bioactivity of proteins by changing their conformations [42]. 4.2. Coaxial electrospinning

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Journal Pre-proof To increase the functionality of electrospun nanofibers, recently some modifications are applied during the electrospinning process. One of the approaches is using co-axial electrospinning for producing nanofibers from two polymers with a core-shell structure to benefit from the properties of both polymers [43, 44]. It is advantageous in drug delivery applications because it can control the release rate of drugs from the polymeric matrices by coating the inner polymer with a suitable polymeric shell [45]. By using core-shell nanofibers, it is possible to load the

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bioactive molecules in the core to maintain the bioactivity and functionality of the drugs during

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the electrospinning process. Moreover, the coaxial electrospinning provides opportunities to

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prepare core-shell nanofibrous mats from both miscible and immiscible polymers with increased

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drug loading efficiency. It can also control the release rate of drugs from the matrices while applying less harsh conditions during nanofibers processing. Moreover, it is possible to blend

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non-spinnable and spinnable polymers within the single polymeric mat in which the non-

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spinnable polymer forms core and the shell is covered by the spinnable polymer. The release rate

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of drugs can be tuned by the degradation rate of the core-shell nanofibers [41].

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4.3. Emulsion electrospinning

Emulsion electrospinning is a novel type of electrospinning that produces a core-shell structure from water in oil (W/O) emulsions for encapsulating hydrophilic drugs. The oil in water (O/W) emulsion is also used for encapsulating hydrophobic compounds [46]. The concept of this technique is related to the solution using for fabricating electrospun fibers. In an emulsion, the droplets of one polymer are insoluble in the other polymeric solution, which usually happens between a polar and non-polar solution. By using this technique, it is possible to load drug that is soluble in an organic solvent into a hydrophilic media or vice versa. Drugs with low molecular weight can easily be distributed into the nanofiber or core-shell fibrous mats. Compared with 10

Journal Pre-proof blending electrospinning, there is no need to use a common solvent in emulsion electrospinning because the polymer and drug are dissolved in applicable solvents. Consequently, it is possible to use many hydrophilic pharmaceutical agents with hydrophobic polymers with minimum negative effect of organic solvents on pharmaceutical agents through the electrospinning process [47, 48]. There are many polymeric phases such as natural polymers (e.g., proteins, polysaccharides) and synthetic polymers (e.g., polyvinyl alcohol (PVA), polyethylene oxide (PEO), and PCL) that are

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applicable for emulsion electrospinning [46].

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4.4. Surface modification electrospinning

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By using surface modification electrospinning, the surface of the conductive collector is coated

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with a particular chemical moiety by addition of certain functional groups that can cover the

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surface by offering a similar environment than the tissue. This strategy is used to prevent the burst release and control the release rate of the loaded bioactive agents on a specific surface [48].

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In some cases, the applied materials are extremely hydrophobic that limit their applications in

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clinical use. This situation plus insufficient functional groups on the surface of hydrophobic

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materials restrict this technique. Alternatively, various strategies are reported to modify the surface properties of ultrafine fibers to alleviate the non-specific absorption of plasma proteins and foreign body reactions [49]. The most common electrospinning methods are presented in Figure 2.

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Figure 2. Schematic illustration of different types of electrospinning. a) Blend electrospinning, b) Coaxial electrospinning, c) Emulsion electrospinning, and d) Surface modification electrospinning. Reprinted with permission from [48]. 5. Drug release kinetics of electrospun nanofibers Three different mechanisms are considered for drug release including diffusion from the bulk polymer or water-filled pores, osmotic pumping, and degradation/erosion. In diffusion and osmotic pumping, the drug molecules are transported through the scaffolds and then release. However, in degradation/erosion, the drug releases through the degradation of the polymers. In a

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Journal Pre-proof diffusion mechanism, the drug molecules gradually move out from the carriers via various chemicals (as a driving force) like concentration gradient. The drug diffusion takes place through a polymeric matrix (called monolithic matrix-based drug delivery system) or across a polymeric shell that covers a drug reservoir core (called reservoir-based drug delivery system). Conversely, drug reservoirs are surrounded via dissolvable/degradable polymeric shells or drug is dispersed in dissolvable/degradable polymeric solution (monolithic matrix drug delivery system) and the

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drug can release from the polymeric matrix core via degradation or dissolution of the polymer

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[50].

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The drug release from nanofibrous matrix also occurs via a combination of four mechanisms

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including diffusion, degradation of the polymeric matrix, dissolution of the drug, and partitioning

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of the drug within the polymer. The drug release from non-swellable degradable and nonbiodegradable polymers happens via the solid polymer matrix before diffusion into the bulk. The

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rate of drug release from polymer matrix is related to the ratio of water penetration into the

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polymer matrix, drug and polymer solubility, partitioning of the drug between the bulk and polymer, and diffusion rate of the drug within the polymer. The average distance of drug

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diffusion from the non-biodegradable polymer is constant due to their fixed geometry. In the case of degradable polymers, the average distance of drug diffusion is not fixed and changes over time. It is dependent on the rate of polymeric degradation [51]. Glassy or semi-crystalline polymers are used to fabricate drug delivery carriers with sustained release behavior because water is slowly diffused into these substrates. Crosslinking the polymeric chains is also useful for hindering the water diffusion and may be a good strategy for preparing sustained release systems [52]. The low solubility of hydrophobic drugs in aqueous conditions plus low hydrophilicity of hydrophobic polymer matrix also contributes to the slow 13

Journal Pre-proof release rate. Increasing surface charge and polarizable molecules at the surface of nanofibers during the electrospinning process may lead to fast drug release as a result of short diffusion distances and high concentration gradients of drugs [53, 54]. The release kinetic of hydrophobic agents from hydrophobic electrospun mats due to high surface area to volume ratio is fast during the first days. This converts to the sustained pattern of release for longer periods [38]. However, the hydrophilic drugs generally have low compatibility with hydrophobic polymers and low

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solubility in nonpolar solvents. This probably results in partition to the surface of nanofibers and

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subsequent fast release [55].

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6. Post-spinning modifications of the electrospun nanofibers

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To date, different nanofibrous scaffolds based on synthetic polymers are prepared by

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electrospinning. Some of them possess undesired surface characteristics, which limit their usage in some biomedical fields. To enhance the performance of these materials, their surfaces are

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modified by immobilizations of various drugs, growth factors, and other bioactive molecules via

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covalent and non-covalent interactions. Nevertheless, it is challenging to select the appropriate

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surface modification strategy to preserve the structure of the bulk nanofibrous mats. There are some techniques such as gamma radiation, ion beam, X-ray, and ozone oxidation that generate functional groups on the surface of the polymeric sheets. However, these methods may degrade and damage the nanofibers and change their mechanical properties [56-59]. Wet chemical treatment is the conventional and simple method for introducing high density of functional groups on the surfaces of materials. For example, by hydrolyzing the material with NaOH, the carboxyl group can be immobilized on the surface of polymers, while the amine group can be added via aminolysis [58, 60]. These strategies are not appropriate for modifying the surface of the nanofibers because they may cause irregular and unwanted etching and highly affects the 14

Journal Pre-proof surface properties of the mats [61]. Moreover, the harsh conditions of these processes enhance the degradability of the nanofibers and decrease their mechanical strengths [58]. Plasma treatment can be used as a powerful alternative method. This technique is able to introduce desired functional groups on the nanofiber surface through a simple and tunable process without changing the bulk characteristics of the material. In plasma treatment, different functional motifs can be added on the surface of nanofibers by using various gas sources e.g., oxygen, air, and

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methane. It is considered as a single step procedure and alternative green method over traditional

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wet modification method [62, 63]. The surface graft polymerization is another useful technique

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for improving the surface hydrophilicity of nanofibers and adding multifunctional chemical

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groups to enhance cellular activities in terms of growth, attachment, and differentiation. The surface graft polymerization is commonly triggered by plasma treatment and UV irradiation to

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activate the free radicals for further polymerization [64]. Another surface modification technique

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is chemical species comprises of pair of electrons that can induce a covalent bond on such electrophile materials. Nucleophiles like amines, hydroxyl, carboxyl, thiols, and alkoxides are

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able to undergo adding or substitution reaction with the chemical groups of the biomolecules.

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This facilitates to add desired groups on nanofibrous substrates [59]. Esterification is another surface modification strategy with the ability to make ester groups through reactions with organic or inorganic acidic solutions. By this technique, biomolecules containing carboxyl or hydroxyl groups can be immobilized on the surface of different nanofibers [59, 65]. 7. Physiochemical properties of silk fibroin Silk is a natural polymer with a fibrous structure that is produced by some arthropods [66]. However, silk is mostly produced by spiders and silkworms. The common form of silk for textile uses is Bombyx mori (B. mori) from mulberry silk, while the silk produced from spiders can be 15

Journal Pre-proof applied in drug delivery applications and other biomedical fields [67, 68]. The SF contains light and heavy chains with 25 and 325 kDa molecular weight, respectively. The heavy chain consists of a repeated sequence comprising alanine (43%), glycine (30%), and serine (12%) in the core. SF is covered by a hydrophilic protein known as a sericin (20–310 kDa), which must be removed during the degumming process due to its potential to stimulate the immune response in the host [69]. SF protein contains crystalline and amorphous conformations in both light and heavy

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chains in which the crystalline regions are highly repeated within the amorphous regions [70,

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71]. The crystalline sequences are mostly comprised of β-sheet secondary structure, while in the

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amorphous regions, other secondary structures of proteins are located. In the crystalline parts, the highly repetitive (GA)n sequences in the B. mori SF and the long poly(A)n stretches exist in the

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non-mulberry silk types and the spider silk mostly adopt an antiparallel β-sheet conformation

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[71]. The interactions and organization of the β-sheets in the crystalline sites define the

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nanocrystallite size (aspect ratio), intercrystallite distances, and nano- arrangement of crystallite, which consequently regulates the organization of crystalline domains [71, 72]. The crystalline

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structure of SF also affects the mechanical strength and the rate of flexibility in SF fibers [73].

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We previously reported that SF could degrade in phosphate buffer saline (PBS) upon transition of β-sheet into random coil conformation. This decreases the adhesion force and Young's modulus of SF fibers [74]. SF fibers generally have the tensile strength of 0.5 GPa, breaking elongation at 15–62 % , and toughness of 104 J kg−1 (Figure 3) [75]. Moreover, the existence of different chemical groups like amines, phenol, alcohol, thiol, and carboxyl in the structure of SF facilitates the interaction with biomolecules and antibodies for specific cell types. By activating the different SF functional groups, it is possible to incorporate various drugs within or on the

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Journal Pre-proof surface of SF fibers. Thus, SF is introduced as a powerful drug carrier that can be applied in

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different biomedical fields (Figure 4) [20].

Figure 3. Schematic representation of hierarchical structure of spider silk and attributed

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mechanical strength and toughness. Spider silk is comprised of fibrils constituting of nanocrystallites in an amorphous structure. The oriented nanocrystals are comprised of β-sheets that

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are tightly packed, while in the amorphous region, there are ransom coils, helices, β-turns, and βspirals conformations. The amino acids sequence of silk defines the type of conformation. The amino acid distribution and ordered remarkably distinct the mechanical strength of SF fibers in terms of flexibility and elasticity (amorphous matrix), and strength (crystallites). Based on the stress-strain analysis, the toughness of SF fibers is achieved. This can be observed by the area underneath a stress-strain curve. Reprinted with permission from [76].

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Figure 4. Different types of modification strategies for changing the surface properties of SF. In the heavy chain of SF, GAGAGS is the basic repeat. The molar density of reactive groups in the

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heavy chain of SF is estimated as percentage values. BMP-2: bone morphogenetic protein-2,

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HRP: horseradish peroxidase, and RGD: arginyl glycyl aspartic acid. Reprinted with permission from [77].

growth factors

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8. Functionalized electrospun silk fibroin nanofibers for loading and controlled release of

8.1. Functionalized electrospun silk fibroin nanofibers for bone tissue engineering SF based drug carriers have been produced by different methods. Various parameters define the choice of fabrication methods including the application of drug carriers, drug stability, and the desired release rate of the drug. Among various fabrication methods, electrospinning is reputed as a suitable technique to develop silk based drug delivery systems. The proper properties of electrospun mats can be achieved by controlling the solution viscosity, type of polymer, voltages, 18

Journal Pre-proof and feeding rate. Recently, SF nanofibers produced by electrospinning have been used as bone tissue constructs [20]. However, SF is not inherently osteoconductive; thus, it is necessary to incorporate a suitable osteogenic factor to induce the activities of osteoblast cells [78]. In a study,

nanofibrous

SF/chitosan/nanohydroxyapatite/bone

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(SF/CS/nHAp/BMP-2, SCHB2) scaffolds were prepared by electrospinning. BMP-2 was placed within the core of the nanofibers, while the shell was comprised of SF/CS/nHAp with two

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thicknesses (SCHB2-thick and SCHB-thin) (Figure 5) [79]. During 14 days, BMP-2 had a

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controlled release profile from both structures. However, the concentration of BMP-2 in the

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release media was significantly higher for SCHB2-thin nanofibers than SCHB2-thick nanofibers

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during all time points. The cumulative release study showed that about 20% of BMP-2 was released from both nanofibers on day one; however, SCHB-thin nanofibers showed faster release

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kinetics. The BMP-2 incorporated SF/CS/nHAP nanofibers showed better performance in

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provoking the differentiation of human-bone-marrow-derived mesenchymal stem cells (hBMSCs) into an osteogenic lineage compared with neat SF/CS and SF/CS/nHAp nanofibrous

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scaffolds. SCHB-thin scaffolds also showed better in vitro results than other groups. During 28

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days, the cells exhibited the higher expression levels of alkaline phosphatase (ALP) as an early osteogenic marker, osteopontin (OPN) and collagen I (COL I) as middle osteogenic markers, and osteocalcin (OCN) as a late osteogenic marker. The amount of OCN in SCHB2-thin nanofibers was higher than in other groups indicating the highest potential in inducing osteogenesis. Moreover, the capability of SCHB2-thin nanofibers in regenerating ectopic bones was also confirmed in animal model [79]. Taken together, nanofibrous SCHB2-thin mats were proposed as potential bone-tissue constructs.

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Figure 5. Schematic representation of coaxial electrospinning setup for preparing SCHB2-thick and SCHB2-thin nanofibrous mats/scaffolds. The effect of the produced scaffolds on

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differentiating hBMSCs is also depicted. Reprinted with permission from [79].

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In another study, a composite based on SF nanofibers and HAp comprising BMP-2 showed a great potency in preserving the bioactivity of the released BMP-2, while controlling the sustained release of the growth factor by tuning the amount of BMP-2 loaded on HAp and SF. BMP-2 had a controlled release rate from all the nanofibrous SF/HAp mats within 21 days. It is worthy to note that the BMP-2 release was highly related to the ratio of drugs loaded on SF and HAp, respectively. It is assumed that BMP-2 release rates can be tailored by changing the distribution of drugs on SF nanofibers and HAp to develop optimal microenvironment for bone repair [80]. Despite the beneficial aspects of BMP-2 in inducing bone metabolism and regeneration, high

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Journal Pre-proof concentrations might cause immunogenicity, heterotopic bone repair, and edema [81]. Therefore, finding the new substitutes can be helpful. For example, a new 24-amino acid polypeptide known as P24 (S[PO4]KIPKASSVPTELSAISTLYLDDD) has been used as a substitute for BMP-2. The P24 is achieved from the knuckle epitope of BMP-2, which is comparable to BMP2 in terms of osteogenic functions [82]. BMP-2 polypeptide was used to modify the surface of graphene oxide (GO) to produce GO-P24 [83]. The BMP-2 modified GO with the negative surface charge was

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assembled on electrospun chitosan-coated SF nanofibers with positive charge via electrostatic

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interaction. The authors suggested that by the addition of BMP-2 modified GO in the SF matrix,

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some cellular behaviors e.g., cell adhesion, growth, differentiation, osteogenesis, and

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mineralization of BMSCs may improve. The results showed that about 3.3% of P24 was loaded on the nanofibers. The physical adsorption of P24 on the surface of the nanofibers facilitated its

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fast release kinetics within 3 days. After that, slow and sustained release was observed. About

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49.43%±2.93% of P24 was release during 21 days. This data indicated that GO–P24 was more appropriate to support P24 constantly and steadily act with the cells adhered to the nanofibrous

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mat as well as prevented the local fast diffusion of P24 [83]. Recently, SF/PCL nanofibrous

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sheets modified by polyglutamate conjugated BMP-2 peptide (E7-BMP-2) was prepared by electrospinning [84]. In comparison to covalent binding or non-specific physical adsorption of peptides and biopolymers, adding of 7-amino acid glutamate sequence may enhance the coupling of peptide to the mineralized HAp, thereby inducing its sustained release [85]. E7-BMP-2 peptide showed 66.9%, 80.4%, and 85.0% binding during 1, 2, and 24 hours, respectively. About 6.2% of the peptide was released initially from the mineralized nanofibers. This observation was similar to the trend obtained for E7-conjugated peptide binding, and the release from the synthetic HAp. This result indicated that the E7 domain induced higher loading of BMP-2 on the

21

Journal Pre-proof mineralized HAp on the fiber surface than BMP-2 peptide. Also, BMP-2 peptide had higher release kinetics than E7-BMP-2 peptide. Moreover, adipose tissue-derived stem cells (ADSCs) showed good differentiation into osteogenic lineages on E7-BMP-2 peptide-modified scaffold as confirmed by the overexpression of osteogenic markers in vitro. It can be concluded that electrospun SF nanofibers provided the good regions for mineralization, which consequently increased the compressive modulus of the scaffold [84].

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Small molecules can also be used as potent osteoinductive factors that can alleviate the

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challenges of using proteins and peptides because they have low molecular mass (<1000 Da) and

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do not provoke the host’s immune response. Unlike growth factors, the functionality of small

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molecules is not related to their structural integrities. Over the past decade, a wide variety of

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osteoinductive small molecules have been discovered [86]. For instance, Rosuvastatin (RSV), a weak hydrophilic small molecule, showed a great potential in osteoinduction than other statins

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like atorvastatin and simvastatin. However, using high doses of RSV might cause cytotoxicity

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because it interferes with the synthesis of cholesterol, which is one the main components for preserving the integrity of the cell membrane [87, 88]. It has been reported that PVA/SF

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nanofibers in the shape of the core-shell substrate could control the release rate of RSV in vitro [89]. A biphasic release profile was observed for the RSV. RSV had an initial fast release from the core-sheet nanofibers which was sustained over time. It seems that the high surface area of the electrospun nanofibers was responsible for the un-controllable initial burst release of RSV. The sustained release was attributed to the lower hydrophilicity of SF in a shell than PVA in the core of coaxial fibers. SF provided a thick layer over PVA nanofibers, which limited the release of RSV and sustained its release rate during 336h. Moreover, SF/PVA nanofibers created a potential delivery system for RSV with improving the growth and osteogenic differentiation of

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Journal Pre-proof hADSCs as confirmed by Alizarin red staining and overexpression of osteogenic markers like ALP, OCN, COLI, and Runt-related transcription factor 2 (RUNX2) [89]. In another study, Urtica dioica L. (nettle) as osteogenic molecule was incorporated into electrospun SF nanofibers by using co-electrospinning process [90]. By addition of nettle, the diameter of SF nanofibers was increased. The results showed that the release of nettle was based on fickian diffusion. The data also confirmed that the released nettle improved the differentiation of hADSCs by the

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expression of osteogenic differentiation-related genes e.g., ALP, Runx2, Col I, OCN in a dose-

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dependent manner [90].

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Taken together, the application of SF to produce nanofiber mats provide the functional sits for

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biomineralization. SF polymers content different amino acids like glutamic acid and aspartic acid

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providing carboxyl groups that make sits for mineralization process. Moreover, these regions facilitate the addition of various bioactive molecules on the nanofiber surface to achieve a drug

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delivery system for various growth factors delivery [91, 92].

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8.2. Functionalized electrospun silk fibroin nanofibers for nerve tissue engineering

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Functionalized nanofibers are good strategies for bridging and recovering the injured axons. It is possible to produce parallel arrays of nanofibers by electrospinning technique applicable for nerve regeneration. SF has good size and shape stabilities in the physiological microenvironment due to its hydrophobic mature. During electrospinning of SF, it is possible to incorporate functional biomolecules to the nanofibers. This helps to control the release rate of the bioactive molecule and preserve its functionality and stability [93]. For nerve injuries with long gaps, nerve guide conduits (NGCs) can be combined with various biomolecules e.g., brain-derived neurotrophic factor (BDNF), nerve growth factor (NGF), ciliary neurotrophic factor (CNTF), and glial cell line derived neurotrophic factor (GDNF) to reach maximum regeneration outcome. 23

Journal Pre-proof Neurotrophic factors stimulate the growth, adhesion, migration, and differentiation of nerve cells and play a main role in regenerating nerve injuries [12, 94]. Thus, neurotrophic factors loaded NGCs are potential drug delivery carriers that not only increase the axonal growth in vitro but also provoke the regeneration of long gaps in the peripheral nerve in vivo. For instance, aligned SF/P (LLACL) nanofibers were formed by co-axial electrospinning that NGF was encapsulated in the core of SF. These aligned nanofibers were used for regenerating peripheral nervous system

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(PNS) [95]. NGF revealed no initial burst release from the nanofibers and had a steady release

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profile with zero-order kinetics. NGF was slowly released from the nanofibers during 60 days

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and a negligible amount of the growth factor was released after this period. The bioactivity of the

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released NGF was preserved for 60 days. This prolonged-release time may be related to the strong electrostatic interactions between negatively charged SF (pI 4.3) and positively charged

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NGF (pI 9.3). Another possible mechanism is attributed to the hydrogen bonding interaction

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between SF and NGF due to they were both proteins which have carboxyl groups and amino groups. After 12 weeks post-implantation, the aligned nanofibrous NGCs containing NGF

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exhibited a high rate of repair compared with those nanofibers without NGF exhibiting the high

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capacity of growth factor to promote peripheral nerve regeneration [95]. Similarly, co-axial electrospinning was used to prepare PLA/SF nanofibers containing NGF. The SF was placed on the core of the scaffold and NGF was encapsulated in the nanofibers [96]. To enhance the surface hydrophilicity, the air plasma treatment was used that had no negative effect on the structural properties of nanofibers. After 9 days, the released NGF from PLA/SF and plasmatreated PLA/SF were 0.13% and 3.84%, respectively. However, the release rates of NGF were 6.74% and 6.90% for PLA/SF and plasma-treated PLA/SF on day 14, respectively. Furthermore, neurite outgrowth of PC12 cells was observed on day 11 [96].

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Journal Pre-proof It is reported that gradients patterns of biological and/or chemicals agents on the surface of nanofibers could affect the cell migration and alignment [97, 98]. The guidance of growth cones might be happened by various growth factors. In the other word, the concentration gradient of released bioactive molecules can be sensed by growth cones [99]. This process is mediated by cell signaling pathways that are provoked by the gradient of concentrations of biological cues [100]. In this regard, dual-gradient and aligned electrospun silk nanofibers loaded with NGF

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were prepared [101]. Incorporating NGF in silk nanofibers did not change β‐ sheet conformation

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of silk. After 5 days, no release of NGF was detected from the nanofibrous electrospun silk,

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indicating the strong adsorption of NGF on nanofibers. During the release times, the bioactivity

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of NGF was maintained and the functionalized silk nanofibers facilitated the neuronal outgrowth in longer intervals. It was also observed that silk nanofibers with gradient concentration of NGF

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had more effect on stimulating the neuronal growth compared with a single concentration of

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NGF. As depicted in Figure 6A, B, less alignment and growth of axons were observed on the nanofibers containing lower concentrations of NGF. On the contrary, the axons were highly

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elongated through two directions guided by silk fibers in the sample containing higher amount of

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NGF (Figure 6C). In the samples with gradient concentrations of NGF, the axons were aligned in even lower concentrations, which was in contrast with uniform concentrations group (Figure 6D, E). Better axonal growth was observed at higher concentrations of NGF (500 NG/mL) that was happened in one direction compared with the single concentration of NGF (Figure 6F) [101]. In general, the topography of ECM has a vital role in mediating the cellular activities in terms of growth, adhesion, migration, recruitment, and organization. The biological functionality of PNS is related to the highly ordered axon bundles. To better mimic this situation, it is necessary to guide the alignment of cells on the scaffolds. Accordingly, it is possible to develop highly

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Journal Pre-proof aligned nanofibers with electrospinning for stimulating the guided growth of axons and promoting its elongation [102]. Based on the literature, electrospun SF nanofibers comprising of

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biological moieties are respectable choices for nerve repair.

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Figure 6. Morphology of the neurons on NGF loaded electrospun silk nanofibers and NGF

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gradient electrospun silk nanofibers. The neurons derived from DRGs were seeded on electrospun nanofibers with 5, 50, and 500 ng/mL−1 (A, B, and C, respectively) NGF, or

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gradient electrospun nanofibers containing 5, 50, and 500 ng/mL−1 (D, E, and F, respectively)

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NGF. The neurons were stained with antiBIII tubulin antibody and evaluated under a fluorescence microscopy after 2 days. Scale bars: 50 μm. Reprinted with permission from [101]. 8.3. Functionalized electrospun silk fibroin nanofibers for skin tissue engineering Nanofiber dressings are one the best scaffold for fast rapid and the best aesthetic regeneration of the wound. Electrospun nanofibers prevent the bacterial infection and wound dry up, permit the gas and nutrient transition, and absorb the wound exudates. Moreover, it is possible to incorporate various biomolecules and antibiotics into electrospun nanofibers to endow extra functions such as anti-inflammatory properties and improve tissue growth [103]. For chronic

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Journal Pre-proof wounds, sustained and prolonged release of growth factors can be achieved by using electrospun nanofibers to provide better wound healing [104]. SF dressings are able to induce the in vitro growth, attachment, and spreading of fibroblasts and keratinocytes [105, 106]. Moreover, SF dressing stimulated the less inflammatory response and infiltration of immune cells compared with commercial dressings [107]. To improve the bio-functionality of SF dressings, they were loaded with epidermal growth factor (EGF) to enhance their regeneration capacity. EGF

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enhances the growth and migration of keratinocytes and has specific receptors on cell

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membranes of both keratinocytes and fibroblasts; thus, it is capable to induce in vivo wound

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repair [108]. In a study, EGF was incorporated into SF nanofibers by electrospinning. The

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release profile of EGF showed a burst release in the first day followed by sustained and stable kinetics for 6 days. After 7 days of exposure to the wound, about 24.7 ± 0.8% of EGF was

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released from silk mats. Upon placing silk mats on the top of the wound, only the nanofibers that

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were directly in contact with wound wet microenvironment could release the EGF in a diffusion manner [109]. Besides, biofunctionalized SF nanofibers provoked the re-epithelialization of the

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wound site by keratinocyte activation. Electrospun SF mats not only provided suitable

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mechanical support for the wounds but also induced the wound repair by controlled and direct release of EGF. Grafting the SF nanofibrous mats on the wound enhanced the rate of wound closure during 24 and 48 hours that confirmed the suitable mechanical strength of electrospun nanofibers [109]. Recently, non-mulberry SF (NMSF)/PVA based electrospun sheets biofunctionalized with EGF and ciprofloxacin HCl had been fabricated to regenerate skin defects [110]. NMSFs e.g., Antheraea assama silk fibroin (AASF) and Philosamia ricini silk fibroin (PRSF) have Arginine, Glycine, and Aspartate (RGD) motif. These motifs are suitable for binding with integrin

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Journal Pre-proof receptors of cells and consequently the cell-materials interactions. This probably facilitated wound regeneration [111]. In this study, the aqueous solution of PVA was blended with aqueous solutions of three types of SF including AASF, PRSF, and Bombyx mori silk fibroin (BMSF). Electrospun nanofibers possess good biocompatibility, high water vapor transmission rate (~2330 g m-2 day-1), acceptable water retention capacity (440%), high elasticity (~2.6 MPa), and suitable antimicrobial activity. EGF had an initial burst release from all types of mats within the

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first hour. The initial release rate of EGF from PVA was 12.04 ± 1.45%; while it was 15.83 ±

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1.04% for PVA-BMSF. The higher initial burst release of EGF from PVA-AASF (20.24 ±

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1.50%) and PVA-PRSF (19.99 ± 1.75%) was observed than other mats. The release of growth

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factor was continued during 24h and reached the plateau after 72h (Figure 7A). The cell proliferation rate was higher on the EGF impregnated mats, indicating that the bioactivity of the

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released EGF was preserved (Figure 7B and C). The in vitro release pattern of EGF may not be

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exactly similar to the in vivo release profile due to various situations in the wound microenvironment. Furthermore, EGF functionalized NMSF nanofibers improved the growth

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rate of HaCaT cells and human dermal fibroblasts in comparison to non-functionalized mats

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exhibiting the beneficial effect of EGF delivery. The in vivo results also showed a high rate of wound regeneration, increased re-epithelialization, formation of vascularized granulation tissue, and high maturation of wound were occurred by using NMSF mats [110].

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Figure 7. (A) Release kinetics of EGF after immersing the prepared mats in PBS at 37 οC, (B

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and C). Rate of the proliferation of HaCaT and HDF cells on bulk nanofibers and EGF-

[110].

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functionalized nanofibers for 7 days. (*p ≤ 0.05, **p ≤ 0.01). Reprinted with permission from

Pignatelli et al. prepared SF nanofibers containing human platelet lysate (hPL) by electrospinning for wound regeneration [14]. The release profile of the SF fibers with various crystallinities was considered by tracking the release of albumin (alb). SF-alb and SF-alb-hPL mats with low crystallinity were rapidly dissolved during 1h and the total amount of alb was completely released. About 41 ± 2%–48 ± 2% of alb was released from the mats with 30% crystallinity. However, the mats with 45% crystallinity had the release of 6 ± 0.2%–10 ± 0.7%. The release of alb was slowly continued after the initial release. This study showed that the

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Journal Pre-proof release profile of BSA can be controlled by optimizing the crystallinity of SF using simple and mild vapor methods [14]. The effect of electrospun SF/gelatin (SF/GT) nanofibers containing astragaloside IV (AS) on acute trauma was also investigated [112]. It had been reported that AS, as the main component of Chinese medicine Astragali Radix, was able to enhance wound repair and neovascularization, improving the growth and migration of keratinocytes and preventing scar formation by

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regulating collagen deposition [113]. The cumulative release study showed that more than 80%

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of the drug was released during 12h and reached 95% after 36h. The authors claimed that the

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AS-loaded SF/GT nanofiber with rapid drug release in 12h, and slowly release in 12-48h were

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appropriate for wound regeneration. After 3, 6, and 9 days post-treatment, AS-loaded SF/GT

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nanofiber dressing exhibited a remarkable increased in the rate of wound closure than control group. Moreover, the content of vascular endothelial growth factor (VEGF) and the number of

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macrophages were significantly increased in the wound area after 7 days by using AS-loaded SF/GT nanofiber dressing. AS-loaded SF/GT nanofiber dressing also suppressed the expression

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of alpha smooth muscle actin (α-SMA) in comparison to the control. The histological evidence

nanofibers [112].

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and scanning electron microscopy (SEM) also confirmed the anti-scar properties of the prepared

In conclusion, nanofibrous matrices have beneficial features with extensive applications in skin regeneration. The highly porous structure of electrospun nanofibers facilitates the suitable respiration of cells and makes an ideal microenvironment for wound repair. For selecting biomaterials as wound dressing, it is necessary to consider the rate of water transmission, suitable air ventilation, inhibiting the entrance of microbes into the wound area, suitable coverage and optimal attachment to the wound site [28]. The long exposure of SF mats with the 30

Journal Pre-proof wound microenvironment due to its slow degradation rate does not induce inflammation and does not interrupt the process of wound repair. 8.4. Functionalized electrospun silk fibroin nanofibers to promote angiogenesis One of the obstacles of tissue constructs is an ineffective ability to induce angiogenesis. Therefore, designing and developing a delivery system is essential to sustain the release of

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angiogenic factors for preserving suitable local therapeutic concentrations in the defected tissue,

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while diminishing potential unwanted systemic side effects. Some studies reported the effect of angiogenic factors in inducing the recruitment of endothelial progenitor cells by activating the

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endogenous stem cells to provoke neo-vascularization. Recently, various controlled delivery

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systems have been designed based on SF nanofibrous mats for delivering angiogenic factors

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[114]. VEGF, a key angiogenic growth factor, plays a vital role in neovascularization [115] by provoking the recruitment, growth, spreading, and differentiation of endothelial cells. Upon

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binding VEGF to specific cell receptors, the VEGF signalizing pathways trigger the formation of

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new vessels [116]. In a study, we prepared a nanocomposite containing freeze-dried calcium

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phosphate and VEGF loaded electrospun SF nanofibers. To control the release rate of VEGF, an external layer of electrospun PLGA nanofibers was further electrospun on the surface of calcium phosphate/SF substrate [117]. We found that the low viscosity of the SF solution limits its capability for producing nanofibers using the electrospinning method. To increase the viscosity of SF solution, it can be blended with high molecular weight polymers such as polyethylene oxide (PEO). However, PEO should be removed after nanofiber formation, which can decrease the amount of growth factor encapsulated in the scaffold [118, 119]. Alternatively, we concentrated the SF solution by applying 41–55 °C temperature to reach enough viscosity and form stable and continuous nanofibers. Afterward, VEGF was incorporated into the concentrated 31

Journal Pre-proof SF solution with suitable viscosity for further electrospinning. VEGF had fast release kinetics without PLGA layer during 7 days; while, it had much slower release rates from those scaffolds containing the electrospun PLGA layer. PLGA nanofibers played a crucial role in controlling the release rates of VEGF from the bio-hybrid nanocomposite. Moreover, the bioactivity of the released VEGF was preserved over 83%, suggesting the usefulness of the developed system for stimulating the angiogenesis within the defected area. Furthermore, the electrospun SF

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nanofibers containing VEGF induced the blood vessels formation and supported the regeneration

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of the defected bone in vivo [117].

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In another study, nanofibrous SF/PVA scaffolds were developed by electrospinning that

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comprised of collagen peptide (CP) and S-Nitrosoglutathione (GSNO) for inducing wound repair

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[120]. It seems that CP has better bioavailability to the wound microenvironment than parent collagen due to its smaller size; thus, it can be more useful for repairing the lost ECM. GSNO, an

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endogenous S-nitrosothiol, is the main biomolecule triggering the nitric oxide (NO) signaling

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pathway that promotes neovascularization both in vitro and in vivo studies [121]. The electrospun nanofibers (SF-PVA, CP-SF-PVA, and CP-GSNO-SF-PVA) exhibited highly

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porous, random, bead-less, and continuous nanofibers. More than 70% of NO (4.02 mmol/L) was released from the nanofibers within the first four hours which was gradually increased during the time. The initial burst release of NO might be related to the hydrophilicity of nanofibers. About 50% and 80% of NO were released after 2.24 h and 7.14 h, respectively. NO had a burst release during the first 4h and reached a plateau after 12 hours. Controlling the release rate of NO can mediate the repair of ischemic wounds. Moreover, the release of CP was significantly increased during time and reached the maximum value after 4h. The burst release of CP might be attributed to the hydrophilicity of the scaffold and high solubility of CP in medium that could provide good

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Journal Pre-proof nutritional support for growth and proliferation of fibroblast cells. Combining the use of CP as a regenerative biomolecule and GSNO as a NO donor in the structure of the nanofibrous SF mats provided a promising construct for repairing ischemic non-healing wounds [120]. Recently, nanofibrous SF hydrogels were mixed with desferrioxamine (DFO) to develop a substrate with sustained release for stimulating the formation of new blood vessel[122]. Small molecule drugs e.g., DFO had been applied to stabilize HIF-1α (hypoxia-inducible factor-1

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alpha) to form hypoxia condition for inducing neovascularization in bone and wound

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regeneration [123]. DFO had a zero-order release profile for more than 40 days without using

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chemical crosslinker. In comparison to chemical crosslinked carrier systems, longer sustained

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release kinetic inside SF nanofiber hydrogels indicated robust interactions between SF

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nanofibers and DFO. The DFO blended nanofibers improved the formation of vascular tubes by human umbilical vein endothelial cells (HUVECs) compared with DFO free hydrogels that only

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formed fragile and thin tubes after 12 and 24 h. By addition and increasing the DFO content,

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thicker tubes were formed by the cells (Figure 8). Based on the in vivo evidences, more angiogenesis, higher functionality, and migration of cells within the hydrogel were seen by using

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DFO blended nanofibrous SF hydrogels than DFO free hydrogels [122].

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Figure 8. The capability of the nanofibrous DFO-loaded silk hydrogels in enhancing

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angiogenesis: (a, b and c) ELIZA assay indicating the expression of HIF-1α, VEGF, and SDF-1α in fibroblast after 3, 7, and 15 days; (d) Fluorescent images representing the tube formation on

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SD groups with various DFO content after 4h, 12h, and 24h cell culture. Scale bars: 100 μm; (e) Quantification of tube length in different segments of the fluorescent images. SF: silk nanofiber hydrogels, FD: DFO, and SD: DFO-loaded silk nanofiber hydrogels. Statistically significant *P≤0.05, **P≤0.01 and ***P≤0.001. Reprinted with permission from [122]. In general, it is preferred to locally deliver angiogenic factors to the site of implantation for improved vascularization. However, denaturation of angiogenic factors during nanofibrous fabrication and long storage time and high cost for purification of growth factors limit their usage as tissue constructs. Moreover, it is necessary to maintain the bioactivity and functionality of growth factors during the scaffold fabrication process. Substitution of traditional angiogenic 34

Journal Pre-proof factors with less sensitive bioactive molecules with the ability to promote angiogenic signaling pathways makes possible to add them during the electrospinning process without affecting their bioactivity. 9. Dual growth factors delivery using electrospun silk fibroin nanofibers Despite the advantages of nanocarriers as a drug/growth factor delivery system, they do not fully

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provide all the requirements for optimal tissue repair. Most of the delivery systems comprise

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only one growth factor with sustained release behavior. However, it seems that these systems cannot provide the optimal condition for tissue regeneration; thus, incorporating multiple growth

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factors into one system might better mimic the tissue’s microenvironment. In recent studies,

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attempts were made to use multiple growth factors for their synergistic effects on the

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regeneration of the defected tissues. By using electrospinning procedure, it is possible to incorporate multiple agents in a single system by applying co-axial or multi-jet methods. In a

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study, co-axial electrospinning was performed to prepare nanofibrous SF/PLGA scaffolds for co-

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delivery of rhBMP2 and dexamethasone (DXM) for bone tissue engineering application [124].

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DXM had a burst release in the first 12h from SF/PLGA nanofibers followed by slow sustained release rates for 21 days. The release behavior of DXM was correlated with the presence of rhBMP2. For example, higher DXM was released at higher concentrations of rhBMP2. DXM was able to interact with amino groups of SF via its fluorine residues which facilitated the sustained release of the drug from SF nanofiber due to the formation of the hydrogen bond. However, the rhBMP2 was loaded in the core and thus had slower release kinetics than DXM, which was loaded on the shell of the nanocomposite. PLGA formed the hydrophobic core and prevented the entrance of the water molecule; thus, rhBMP2 was released first from the core to the shell of the nanofibrous scaffold. Then, it was diffused from the shell to the outer 35

Journal Pre-proof microenvironment. The strong Van der Waals bonds between rhBMP-2 and PLGA restricted the fast release of the growth factor from the core of the carrier into the shell. With the increase of the rhBMP2 content, the release kinetic of rhBMP2 was decreased from the nanofibers with similar diameter. Furthermore, DXM provoked the growth and differentiation of rBMSCs upon fast release with high quantity during the first 7 days. However, rhBMP2 accumulated up to 7 days that was necessary for osteogenic differentiation of rBMSCs [124]. Previously, we

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developed a nanostructure based on SF/calcium phosphate/PLGA for dual delivery of VEGF and

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platelet-derived growth factor (PDGF) [125]. It is noteworthy that new blood vessels formed by

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VEGF are immature and need PDGF for further maturation to be more stable with less

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permeability [126]. The data confirmed the potential of SF nanofibrous to support tissue regeneration and to form a suitable reservoir for the incorporated growth factors. The release

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results also showed that PDGF had a slower release rate than VEGF during 28 days and the

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bioactivities of both growth factors were preserved. This indicated that the electrospinning parameters did not affect the functionality of growth factors. After 10 weeks post-implantation of

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the scaffold in calvarial defect of rabbit, the biofunctionalized scaffolds promoted new blood

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vessel formation and new bone formation [125]. In another experiment, it was reported that core-shell silk fibroin/poly(L-lactic acid-co-εcaprolactone)-polyethylene oxide (SF/PLCL-PEO) containing connective tissue growth factor (CTGF) and fibroblast growth factor 2 (FGF-2) improved the differentiation of MSCs into fibrogenic lineage

[127]. FGF-2 was released through the dissolution controlled fashion.

Moreover, the CTGF encapsulated in the SF core, which was further surrounded by PLCL-PEO shell that acted as a matrix/reservoir combinatory system. FGF had burst release within the first 8h; however, only 5–8% of the growth factor was released per day in the first week. The

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Journal Pre-proof cumulative release profile of FGF was about 81.7 ± 2.6% and 91.6 ± 1.8% at 7 and 14 days, respectively. Furthermore, the release of CTGF exhibited no burst release compared with FGF-2 with cumulative release rate of 10.1 ± 0.9% in the first 3 days. The release of CTGF was also much slower and sustained (23.1 ± 0.6%) during 14 days [127]. In many investigations, the advantages of using dual delivery systems than single drug delivery for PNS regeneration were also confirmed. For instance, electrospinning was used to prepare

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multichannel SF for co-delivery of NGF and ciliary neurotrophic factor (CNTF) to induce PNS

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repair [128]. NGF and CTNF had a synergistic role in regenerating the defected nerve by

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provoking sensitive fibers and motor neurons, respectively. The growth factors were

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incorporated into the scaffold via electrostatic bonds. NGF was released from the nanofibrous

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silk conduit after 1h. After one week, the release of NGF was almost stopped. In addition. CTNF had a rapid release rate during one day which subsequently reached zero-order release. The slope

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of the release curve was almost constant for 6 days, suggesting the sustained release of CTNF

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from silk nanofibers within one week. Besides, the release profiles revealed that only 1% of NGF and 12% of CNTF were released from the scaffold. At pH 7.4, the negatively charged SF (pI =

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4.3) made an electrostatic interaction with positively charged NGF (pI = 9.3) and CNTF (pI = 5). Moreover, the high molecular weight of CNTF (23kDa) and NGF (27 kDa) prevented their diffusion from nanofibers. These two parameters might be responsible for the slow release rate of growth factors over one week [128]. Similarly, NGF and CNTF had a synergistic effect in multi-channel electrospun SF conduits in stimulating nerve repair. The primary neurons from rat dorsal root ganglia (DRGs) were cultured and grown on the surface of nanofibers and adhered to the fibers directionally based on the nanofiber orientation (Figure 9) [129]. The presence of NGF in biofunctionalized nanofibers induced a 3-fold improved in neurite length and outgrowth.

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Journal Pre-proof CNTF also promoted growth, alignment, and migration of glial cells in the functionalized electrospun nanofibers. During the first 4 days, both growth factors were not released because they might have entrapped and strongly absorbed within the conduit or they had a release rate

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below ELIZA kit detection [129].

Figure 9. The outgrowth of DRGs cells confirming the preservation of growth factors bioactivity. Florescent images from DRGs cells cultured on (A and e) unfunctionalized nanofibers, (B and f) NGF-functionalized nanofibers, (C and g) CNTF loaded nanofibers, (D and h) NGF and CNTF-functionalized nanofibers after 5 days. The anti BIII tubulin antibody was used for neurons staining. Higher magnification of neurons (e, f, g, and h) was taken after 3 days post-seeding to avoid overlapping of neurons that were observed after a long time in culture. Scale bars: A, B, C, D: 100 mm; e, f, g, h: 20 mm Reprinted with permission from [129].

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Journal Pre-proof In another study, electrospun SF nerve conduits containing GDNF and NGF exhibited sustained release with preserving bioactive GDNF and NGF more than 4 weeks as confirmed by neuronal differentiation of Neuro-2A and PC12 cells, respectively [130]. The orientation and topography of the drug-loaded electrospun SF conduit dictated the direction and extent of axonal growth from DRG sensory neurons. The conduit also induced the functional recovery of spinal cord (SpC) motor neurons, indicating its potential for promoting the regeneration of both motor and

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sensory neurons [130]. Incorporating CNTF, BDNF or both into the nanofibrous SF mats could

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enhance the functionality of defected nerve without causing negative effects on the morphology

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and chemical structure [93]. During 4 weeks, NGF and GDNF were released simultaneously

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with no significant differences. Both growth factors showed a constant 6-7 pg/day release rates during this period. However, only 0.5% of the total growth factors was released after 4 weeks.

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The growth factors had a very slow release rate, allowing the nanofibers to continue to supply

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CNTF and BDNF to cells for a long period. The growth factors also remained bioactive and available to the cells, as rat retinal ganglion cells (RGCs) showed longer axonal outgrowth when

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in close contact with the biofunctionalized nanofibers. In comparison to unfunctionalized

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nanofibers, the growth of neurites enhanced 2-fold on nanofibers containing BDNF, 2.5-fold on nanofibers containing CNTF and nearly 3-fold on nanofibers containing both growth factors [93]. In another study, VEGF and BDNF were concomitantly incorporated into a conduit for the simultaneous induction of nerve repair. The mechanism of action of these growth factors was different because VEGF was used to promote angiogenesis and vascular permeability in the defected area, while BDNF was used to improve nerve regeneration and neuronal survival. Coaxial electrospinning was used to develop the conduit by switching the position of BDNF or

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Journal Pre-proof VEGF in the core or shell of sheets [131]. The scaffolds containing SF-VEGF in the core and SF-BDNF in the shell had a better influence on the proliferation of Schwann cells than those containing SF-BDNF in the core and SF-VEGF in the shell. However, a higher rates of neovascularization and nerve repair were found in rats treated with conduits containing SFBDNF in the core and SF-VEGF in the shell. These findings were attributed to the release kinetics of both growth factors, indicating that in the presence of sufficient blood vessels formed

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by VEGF, nerve repair could be promoted as a function of BDNF. The growth factors had burst

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release on the first 4 days that reached constant and stable rates during 16 days. Based on growth

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factors location, the release profiles were different. For example, the growth factor located in

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shell had faster release rate than those located in the core. In addition, the release of VEGF was faster than BDNF even in the same location, indicating the different degradability of growth

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factors or stronger ionic interactions between BDNF and SF than VEGF. Overall, core-shell SF

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nanofibers with controlled release behaviors that contained BDNF in the core and VEGF in the shell were able to improve the regeneration of cavernous nerve regeneration [131]. Liu et al. also

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used electrospinning to fabricate SF nanofibers for dual delivery of BDNF and VEGF [132].

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Both growth factors had a high release rate within the first 4 days that was decreased during 16 days. BDNF and VEGF had cumulative release rates of 9.2×103 and 6.9×103 pg, respectively. After 16 days, the cumulative release profile of NGF and BDNF showed 11.9% and 15.9%, respectively. VEGF had higher release kinetics than BDNF that may be attributed to the strong interactions between BDNF and SF compared with VEGF. New blood vessel formation induced by VEGF was also confirmed by the expression of endothelial markers without chronic inflammatory responses. The dual growth factors loaded SF nanofibers exhibited an outstanding

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Journal Pre-proof effect on improving the nerve repair after 54 and 8 weeks post-implantation than nonfunctionalized scaffolds [132]. Taken together, tissue regeneration is a complicated process that needs multiple factors for stimulating tissue healing. Designing and developing multiple biomolecule delivery systems are essential to reach maximum therapeutic outcomes. Till now, various multifunctional drug delivery systems are prepared for optimal tissue regenerations. The key for developing proper

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platforms to mimic the natural tissue regeneration process is to integrate biomolecules in an

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appropriate polymeric carrier. However, dual or multiple growth factors delivery do not always

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have a complementary effect on tissue healing as a result of the dose-dependent activity of both

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bioactive molecules. Consequently, it is crucial to use growth factors that have a synergistic

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effect on tissue regeneration. For this, further investigations are crucial to explore a suitable combination of growth factors with optimum doses and maximum effects.

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10. Electrospun silk fibroin nanofibers for anticancer delivery

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Despite the usefulness of the current methods for diagnosis and treatment of cancers, their low efficacy and side effects on normal cells encourage researchers to find new approaches with

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higher sensitivity, accuracy, and efficacy. For this, nanotechnology offered new routes for basic and clinical research on cancer [133, 134]. Furthermore, it is found that cancer cell monolayers on the 2D tissue culture plates could not fully mimic the conditions on tumor microenvironment in the body; thus, developing 3D culture model with multicellular tumor spheroids is essential. To date, various 3D constructs were prepared for in vitro study of cancer. Among these systems, electrospun fibers with nano/micro dimensions were known as suitable scaffolds to imitate the interactions between cells and cell-matrix because they are highly similar to collagen fibers in the ECM of tumor cells in terms of structure and dimension [135]. In addition, the extraordinary 41

Journal Pre-proof properties of electrospun nanofiber enable their application for cancer detection and therapy, targeted cancer treatment, and preparing smart drug carriers for anticancer delivery [136]. Accordingly, electrospun SF/PCL nanofibers were prepared for the controlled release of titanocene as a breast anticancer drug. The systems exhibited good potential in inducing apoptosis of MCF-7 breast cancer cells [137]. The rate of apoptosis induction was affected by interactions between titanocene and amino residues in the structure of SF, degree of SF

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crystallinity, and the percentage of drug loading. In the beginning first hour, titanocene had a

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burst release (10-12%) from SF nanofibers. However, during 8 to 48 hours, the release of the

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drug was gradually sustained. Only a slight amount of drug was detectable after 6 days when

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using low concentrations of the drug. At this time, significantly higher amounts of drug (>85%) were released for higher drug concentrations. On the sixth day, no drug release was observed and

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the curve was saturated at 85%. Some drugs could strongly bond with the SF matrix; thus, they

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could only release from the scaffold upon polymer degradation. However, the release of drugs from semi-crystalline polymer might be sustained as a result of limited water uptake in the

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crystalline regions that restrict the fast release of drugs [138]. The highly aligned β-sheet

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conformation of SF might affect the formation of the crystalline microstructure of PCL fibers like aligned lamellae and fibrillary structures [139] that consequently decreased the release rate of the titanocene [137]. In another study, injectable pH-sensitive SF nanofiber hydrogel was prepared to sustain the release of doxorubicin (DOX) for 8 weeks. In this situation, the anticancer drug could have long-term cytotoxicity against cancer cells both in vitro and in vivo [140]. DOX exhibited a pH-responsive release behavior from silk hydrogel; the release at pH 4.5 was faster than pH 7.0. The silk content within the hydrogel also affected the release behavior of DOX. After 8 weeks, DOX had a cumulative release of 19.7% and 35.5% at pH 7 and pH 4.5,

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Journal Pre-proof respectively, from 2 wt% silk nanofiber hydrogels. In the case of 1 and 0.5 wt% SF nanofibers, the cumulative release reach 25.7% and 36.1% at pH 7 to 60.2% and 70.1% at pH 4.5, respectively. The data showed that silk nanofiber hydrogels exhibited preferred injectability without compromising the main properties of a silk-based drug release system e.g., sustained release capacity, and pH-responsive characteristics, indicating a significant future for breast cancer treatment [140]. Similarly, SF nanospheres and silk sericin (SS) nanofibers were used as

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biotemplate to control the self-assembly and nucleation of silica to develop an optimized system

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for anticancer delivery [141]. SF and SS facilitated the assembly and nucleation of silica into

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mono-dispersed Si/SS nanofibers and Si/SF nanospheres (Figure 10). Both substrates exhibited

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acceptable drug loading efficiency with the ability to control the in vitro release rate of DOX. By increasing the drug/carrier ratio, the loading efficiency of the drug was increased which was

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inconsistent with the other studies. After mixing DOX with the nanocomposite for 24h, 29% and

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33% of the drug was loaded into Si/SF and Si/SS, respectively. DOX had a burst release in the initial phase at pH 5.4 from both Si/SF and Si/SS substrates. At pH 7.4, DOX had a sustained

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release rate for 7 days. Approximately 75% of the drug was released from nanofibers and

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nanospheres at pH 5.4 after one week, while only 20% of the drug was released during this time at pH 7.4 from both substrates. It seems that the acidic condition interrupted the electrostatic interactions between the carrier and DOX, and induced the burst release of the drug. Furthermore, Si/SF nanospheres were successfully uptake by human cervical carcinoma (HeLa) cells and were accumulated around the nuclei of cells. The Si/SS nanofibers were only able to attach to the surface of cancer cells. The data confirmed the higher efficacy of DOX-loaded Si/SF nanospheres and Si/SS in treating cancer than free DOX [141].

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Figure 10. Schematic representation of the self-assembly and nucleation of silica into

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nanocomposites. They are mediated by silk protein sericin (SS) nanofibers and self-assembled

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silk protein fibroin (SF) nanospheres. (A) Addition of APTES into SF (top) and SS (bottom) solutions for the formation of an array of silica nuclei on the surface of SF nanospheres (top) and

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SS nanofibers (bottom) collected in the solution, and (B) subsequent addition of tetraethylorthosilicate (TEOS) into the first solution to form silica on the nuclei of silica to

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promote the growth of silica on SF nanospheres (top) and SS nanofibers (bottom). (C) More addition of TEOS to induce continuous growth of silica on SF nanospheres (top) and SS

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nanofibers (bottom) to form spherical (top) and fibrous (bottom) silica nanocomposites. Reprinted with permission from [141].

In another study, smooth SF fiber bead-on-string fiber and coaxial bead-on-string fiber were developed and the release rate of drugs was studied at different pH values [142]. The cumulative release profile of DOX exhibited that the release of drug was correlated to the morphological aspects of the electrospun nanofibers. The release of drug from bead-on-string construct was slower and less than smooth fibers. DOX was less released from coaxial bead-on-string fibers than bead-on-string fiber. Moreover, the drug release data revealed that the electrospun materials had pH-sensitive releasing properties. In an acidic condition, higher levels of the drug were 44

Journal Pre-proof released with faster kinetics. Generally, coaxial bead-on-string fiber had a better efficacy to be used as a sustained delivery platform for anticancer drugs [142]. Li et al. also fabricated electrospun SF nanofibers containing curcumin loaded-SF nanospheres for co-delivery of drugs [143]. The hydrophobic curcumin was incorporated into SF nanospheres by ethanol precipitation of self-assembled SF. Afterward, the curcumin-loaded SF nanospheres and hydrophilic DOX.HCL were entrapped into SF nanofibers by applying the colloid electrospinning method.

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The nanofibrous mats showed dual drug release profiles in a sustained manner. The release rate

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of drugs from the core and shell of SF nanofibers was controlled by the content of loaded drugs

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and the degree of crystallinity that were tuned via a water-annealing process at various

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temperatures. Fickian model was a possible mechanism for drug release. The current drug delivery system may be potentially used as local multi-drug delivery carriers for the treatment of

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different medical conditions, like breast cancer or skin cancer [143].

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In summary, progress in developing in vitro 3D cancer models enables better diagnosis and

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treatment of cancer. The high capacities of electrospun SF nanofibers in terms of direction, alignment, modification, and encapsulation of drugs provided the opportunity to develop 3D

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cancer models to mimic the malignant tumors microenvironment. Moreover, different implantable and injectable SF based drug delivery carriers allow the delivery of chemotherapeutics without negative side effects [144]. 11. Electrospun silk fibroin nanofibers for antibiotic delivery Despite the advances in developing surgical methods and producing various antimicrobial agents, one of the main consequences of surgeries are still infectious diseases. To avoid infection after surgery, most clinicians prescribe systemic administration of high doses of antibiotics for a long time. However, the systemic toxicity and short half-life of antibiotics limit their usage. 45

Journal Pre-proof Therefore, local administration of antibiotics was proposed as a new strategy to decrease systemic toxicity and increase the therapeutic outcomes [68, 145]. The SF suffers from lacking antibacterial property. For this reason, it is usually loaded and modified by various functional nanomaterials to enhance its antibacterial properties. As an example, SF/GO (graphene oxide)blended nanofibers fabrication by electrospinning method revealed that the addition of GO could increase the inhibitory effect against Gram-positive Staphylococcus aureus (S. aureus and Gram-

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negative Escherichia coli (E. coli). The bacterial membrane destroys by GO, leading to the

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efflux of intracellular organelle and killing bacteria [146]. In another study, polyethylenimine

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(PEI) with suitable antibacterial characteristics at the various concentrations (10, 20, and 30%

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(w/w)) was blended with SF through electrospinning. The bare SF nanofibers had no antibacterial effect against P. aeruginosa and S. aureus, while those containing PEI had high

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antibacterial effect against P. aeruginosa and S. aureus [147]. The antimicrobial effect of PEI

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raise from its polycationic nature and the ability to interact with the membrane of bacteria and disturbed the bacterial cell [148]. The addition of Manuka honey (MH) in electrospun SF

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nanofibers is another useful method to improve the antimicrobial activity of fibrous SF matrices

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[149]. The antibacterial effect of MH/SF nanofibers was dose-dependent; the increase in the content of MH resulted in higher antibacterial effect. Furthermore, the nanofibrous MH/SF highly suppressed the growth of E. coli and P. aeruginosa than S. aureus and MRSA [149]. The content of methylglyoxal is responsible for antibacterial activity of MH against S. aureus. Unique Manuka Factor (UMF™) is the medical grade of MH, which is mainly used against S. aureus [150]. In another experiment, an antibacterial peptide motif (Cys-KR12) derived from human cathelicidin peptide (LL37) was immobilized on electrospun SF constructs for wound care applications [151]. The nanofibrous composite had an inhibitory effect against S.

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Journal Pre-proof epidermidis, S. aureus, E. coli, and P. aeruginosa without the formation of biofilm on the surface of the membrane. The mats also increased the growth rates of fibroblasts and keratinocytes and induced the differentiation of keratinocytes by promoting cell-cell interactions (Figure 11). Thus, Cys-KR12-immobilized SF nanofiber membranes were suggested as suitable substrates for

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wound healing purposes [151].

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Figure 11. (A) Western blotting of involucrin (differentiation marker for keratinocyte). (B) The expression rate of involucrin based on band intensity analysis (n = 3, mean ± SD). Asterisks 48

Journal Pre-proof stand for significant differences in comparison to pristine SF. (C) Confocal immunofluorescence images of human keratinocytes (HaCaT) cultured on the tissue culture plate, pristine SF, K200, and K500 (red, F-actin; green, involucrin; blue, DAPI). Reprinted with permission from [151]. Ojah et al. used co-axial electrospinning to prepare core-shell nanocomposite containing amoxicillin trihydrate-loaded SF (AMOXBMSF) in the core and PVA in the shell. The effect of surface functionalization of the core-shell nanofibers by oxygen (O2) dielectric barrier discharge

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(DBD) plasma treatment in release of drug was further assessed [152]. AMOX had steady and

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continuous release kinetics (>60%) during 200h. AMOX had a faster release rate from AMOXBMSF/PVA/O2 than AMOX-BMSF/PVA nanofibers during 55h. After that, the release rate of

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AMOX was sustained and reached the plateau after 120h. Moreover, upon surface modification,

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the hydrophilicity of MOXBMSF/ PVA/O2 was increased, which induced the faster release rate

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of AMOX from the nanofibers. The released AMOX had also an inhibitory effect against E. coli and S. aureus bacteria [152]. In another study, doxycycline monohydrate (DCMH) and thyme

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essential oil (TEO) as antibacterial agents were incorporated into electrospun SF/gelatin

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nanofibers. TEO showed a burst release from the nanofibers during the first 3h, while the release

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of DCMH was almost sustained during 48 h. The release of TEO reached the plateau after 21-48 hours, indicating that the drug concentration was not remarkably increased in the release medium. Moreover, TEO-loaded nanofibers exhibited more antibacterial activity against S. aureus and Klebsiella pneumonia bacteria than DCMH-loaded ones [153]. In another study, Ceftazidime (CTZ) was loaded on silk fibroin/gelatin (SF/GT) nanofibers. CTZ had initial burst release that might be related to the high surface area to volume ratio of the nanofibers. Moreover, the release kinetics of CTZ was attributed to the drug solubility and ionic/Van der Waals bond interactions between the SF/GT nanofibers and the drug. The SF/GT nanofibers containing CTZ had a proper antibacterial effect against Pseudomonas aeruginosa [154]. 49

Journal Pre-proof Electrospun SF nanofibers are suitable substrates against bacterial growth or infection. Some issues are still remained to be further addressed in antibacterial nanofibers. First, the synthesis of SF nanofibers containing antibacterial agents, which till now is largely restricted to the laboratory using simple setups (e.g., syringe needles) to synthesis small-sized nano-membranes. This needs to be confirmed on large scales. Second, it is also needed to produce the biocide-SF nanofibers containing antibiotic agents that show stronger antibacterial activity than one side-

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loaded nanofibers. Fabrication of electrospun SF nanofibers containing antibiotic agents with

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ability to control both Gram-positive and Gram-negative bacteria is more applicable system

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[155].

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12. Electrospun silk fibroin nanofibers for antioxidant delivery

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Antioxidants are critical factors for controlling the repair of damaged tissues. Antioxidants control the production of reactive oxygen species (ROS) upon the inflammation phase; thus,

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decrease the side effects of wounds [156, 157]. These components are highly available in various

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plants and are able to induce the process of wound regeneration. Green electrospinning was used

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to fabricate nanofibrous SF/PEO mats containing grape seed extract (GSE) [158]. GSE is comprised of biological flavonoids e.g., oligomeric procyanidolic complexes; thus, it is considered as a promising antioxidant agent [159]. The release rate of GSE had three separate parts. A burst release occurred at the first 50h, which related to the GSE attached on the surface of the nanofibrous mats. After that, a continuous release rate from 50 to 100 hours was observed. Finally, a constant release rate happened until the end of the experiment. The release profile of GSE might also be affected by the nature of GSE itself and the porosity and pore size of the nanofibrous mats. GSE loaded SF/PEO nanofibers significantly increased the proliferation of skin fibroblasts compared with bare SF/PEO nanofibers and highly protected the cells from tert50

Journal Pre-proof butyl hydroperoxide (t-BHP)-induced oxidative stress. The t-BHP is commonly used as an oxidative damage model due to the ability to trigger cell apoptosis and injuries [160]. The GSE loaded nanofibers also exhibited outstanding properties as skin substitutes to improve wound healing [158]. Another antioxidant agent is melanin with the ability to breakdown the free radical chains [161]. Melanin incorporation into random and aligned electrospun SF nanofibers showed a high free radical scavenging properties [162]. The melanin loaded nanofibers showed high

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capacity to break the free radical chain compared with bare SF nanofibers. However, the

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scavenging activity of random fibers was more than aligned nanofibers, which might be

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attributed to the higher thickness of aligned nanofibers than the other ones [162]. In another

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study, Fenugreek as a natural antioxidant agent was loaded into SF nanofibers by using coelectrospinning [163]. The release rate of fenugreek form nanofibers with a higher amount of

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drug was fast within 5h followed by a gradual release till 24h. This fast release was attributed to

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the release of fenugreek that was deposited within the nanofiber pores. However, the release of fenugreek from those nanofibers containing lower amounts of this molecule was almost slow and

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constant during 24h. It might be related to the hydrophobic nature of SF nanofibers that restrict

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water penetration which inhibits drug diffusion into the release medium. By increasing the Fenugreek content, the scavenging activity was increased as confirmed by 1,1-diphenyl-2picrylhydrazyl (DPPH) scavenging assay. Moreover, the proliferation rate of fibroblasts was increased on nanofibers (Figure 12). The presence of phenolics and flavonoids in Fenugreek increased the antioxidant activity of Fenugreek loaded SF nanofibers (49.3%) compared with bare SF nanofibers (5.6%). The antioxidant activity of bare SF nanofibers was correlated to the presence of tyrosine and tryptophan amino acids in its structures that comprise phenolic side chains [164]. Adding curcumin into electrospun SF/poly(L-lactic acid-co-ecaprolactone)(P(LLA-

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Journal Pre-proof CL) nanofibers also increased the antioxidant activity of the nanofibers [165]. At the first 12h, curcumin had a slight fast release followed by more constant release during 72h. In general, higher concentrations of loaded drug provided higher diffusion driving force that induced the release of drug. Bare SF/P(LLA-CL) nanofiber had a scavenging activity of 32.2%. Curcuminloaded SF/P(LLA-CL) nanofibers exhibited a remarkable scavenging effect on the DPPH radical, which was gradually improved by increasing the content of curcumin from 2.0% to 6.0% (w/w).

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Thus, curcumin increased the antimicrobial effect of the nanofibers that broadened their usage in

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biomedical fields [165].

Figure 12. FDA staining showing the growth rate of Swiss albino 3T6 fibroblast cells on different types of nanofibers after (a−e) 48 h, (f−j) 72 h: (a and f) SF nanofibers, (b,g) SF−Fenugreek nanofibers (1:0.1), (c and h) SF−Fenugreek nanofibers (1:0.2), (d and i) SF−Fenugreek nanofibers (1:0.5), (e and j) SF−Fenugreek nanofibers (1:1). Reprinted with permission from [163]. 13. Electrospun silk fibroin nanofibers for vitamin delivery

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Journal Pre-proof Vitamins, as essential components in the body, play vital role as antioxidants and hormone, control the cell behavior in terms of growth and differentiation, and induce cell signaling pathways. In recent years, electrospun SF nanofibers are used for vitamin delivery that has gained good attention for tissue engineering purposes. For example, electrospun SF mats loaded with L-ascorbic acid 2-phosphate [vitamin c (VC-2-p)] were prepared by the electrospinning method. The data revealed the beneficial aspects of both bare and VC-2-p-loaded SF nanofibers

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for stimulating the growth, adhesion, and spreading of L929 fibroblast cells [166]. The

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expression of collagen type I was significantly increased by VC-2-p-loaded SF nanofibers

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compared with neat SF nanofibers. VC-2-p had a burst release from SF nanofibers during the

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first 20 min, which reached a plateau over time. It was suggested that the loading of VC-2-p close to the surface of electrospun nanofibers plus easy diffusion of water-soluble of VC-2-p into

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the water might be responsible for the fast release of vitamin C from mats. The increased

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concentration of VC-2-p in nanofibers could promote the driving force for drug diffusion which

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facilitated VC-2-p to easily diffuse from the SF nanofibers [166]. D-α-tocopherolpolyethylene glycol 1000 succinate (VE-TPGS) is also reputed as a water-soluble

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derivative of vitamin E (VE) that exhibits high capacity in inducing skin repair due to antiinflammatory effect, high anti-oxidant activity, anti-scar properties, cost-effectiveness, and good availability [167]. Core-shell SF/PVA/aloe vera (AV) produced via electrospinning was used for VE delivery [168]. Concurrently, about 1 and 5 mg VE were loaded in starch nanoparticles and then were incorporated into the SF-PVA-AV nanofibers. VE showed a fast release rate from the nanofiber via Fickian diffusion. After 4h, about 15.93% and 64.20% of VE were released from the samples containing 1 mg and 5 mg VE, respectively. The release rate of VE from those nanofibers containing 5 mg of the drug was remarkably higher than the mats comprising 1 mg

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Journal Pre-proof VE. This phenomenon might be attributed to the high water solubility of VE. The higher hydrophilicity induced more water penetration into the nanofibrous mats that can dissolve the high amount of VE. Therefore, higher free spaces were formed in the nanofibrous matrix that induced the diffusion of drug from the nanofibers to the release media. The addition of AV and VE into the nanofibers improved the growth, attachment, and collagen production of fibroblast cells. Moreover, VE incorporation enhanced the antioxidant activity of the scaffold to preserve

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the cells from the oxidative products. This was helpful for wound healing purposes [168].

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Similarly, incorporation of VE TPGS increased the scavenging activity of nanofibrous SF mats

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against oxidative stress induced by t-BHP [169]. VE TPGS had a slight burst release (⁓7%) at the initial 30 min that was constantly continued over 72h. A possible mechanism for VE TPGS

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release was drug diffusion. Furthermore, the presence of hydrophobic α-tocopherol succinate and

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hydrophilic PEG in the structure of VE TPGS increased the intermicellar interactions. PEG

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group was also able to make hydrogen bonds with amide groups of SF. The aforementioned parameters were the main reasons for sustained release of VE TPGS from SF nanofibers [169].

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Bhutto et al. blended vitamin B5 with P(LLA-CL) and P(LLA-CL)/SF solutions to prepare

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aligned nanofibers to promote the growth of Schwann cells [170]. The drug showed a release rate of 9% during 1h, which was increased to 75% within 10h and consequently reached 80% in a plateau manner after 24h. More than 85% of the hydrophilic vitamin B5 was released for 3 days. The localization of vitamin within the nanofibers was the cause of initial burst release because the vitamin was located on the surface of the nanofiber that was simply dissolved in PBS solution and release from the scaffold during the first hours [170]. In another study, the activated form of vitamin B12, mecobalamin, was incorporated into aligned SF nanofibers to stimulate the neurite outgrowth and survival of neurons by activation of Erk1/2 and Akt signaling pathway.

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Journal Pre-proof The vitamin loaded nanofibers exhibited acceptable biocompatibility and low inflammatory responses. They also enhanced the migration of Schwann cells and DRG neurons in alignment

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with scaffold direction (Figure 13) [171].

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Figure 13. Immunofluorescence staining of DRG neurons cultured on flat silk fibroin (SF) film, aligned SF scaffold (ASF), and mecobalamin loaded aligned SF scaffold (MASF). Scale bar: 200

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μm. Reprinted with permission from [171].

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14. Electrospun silk fibroin nanofibers for ion delivery Metal oxides are attractive materials for various applications because their surface defects make extraordinary optical, electrical, sensing, and chemical catalyzing characteristics [172]. Nevertheless, hard processing of metal oxide into nanoparticles and nanowire plus insufficient biocompatibility, and biodegradability restrict their application in biomedical fields. To improve the biological properties, they can be combined with organic materials, which form a flexible composite with higher biocompatibility [173], and extend their application as flexible biosensors and antibacterial dressings [174]. Among metal oxide nanoparticles, silver nanoparticles

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Journal Pre-proof (AgNPs) are of interest due to their high chemical stability, good electrical conductivity, and acceptable antibacterial and catalytic activities. The biological and chemical behavior of AgNPs can be enhanced by incorporation into a polymeric matrix. Colloidal AgNPs are tended to aggregate to reduce their surface energy like other metallic nanoparticles [175]. The antimicrobial activity of AgNPs is also reduced upon aggregation. One of the strategies that alleviate the aggregation of AgNPs is the incorporation of the metal oxide into a polymer

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solution and then electrospun into nanofibers. Accordingly, an ultrafine nanofibrous composite

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based on SF/Ag was fabricated that contained both silk I (random coil) and silk II (β sheet) mats

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[176]. The Ag ions showed a first-order release rate from both types of nanofibers for 8h.

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Moreover, the cumulative release rate of Ag ion from silk I nanofibers was remarkably higher than silk II ones due to the amorphous random coil conformation. So, it is possible to control the

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release rate of ions by controlling nanofibers morphology. The nanofibrous SF/Ag scaffolds with

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low Ag contents (0.1% Ag, w/v) also had acceptable antibacterial activity against P. aeruginosa. It had also lower antibacterial effect against S. epidermidis and S. aureus at 1% Ag (w/v)

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concentration with slight toxicity to L929 cells [176]. It was also demonstrated that various

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compounds of Ag ions in the forms of silver phosphate, silver tetrafluoroboratem, silver acetate, and silver sulfadiazine can be incorporated into the polymeric solution of SF matrix and form a composite nanofiber [177]. The composite SF/Ag nanofibers were able to control the release rate of Ag in distilled water. The amount of Ag ions released from the SF nanofibers comprising silver phosphate was higher than other samples. A similar release profile of Ag ions was detected for nanofibrous SF mats containing silver acetate and silver tetrafluoroborate. Compared with other groups, SF nanofibers containing silver sulphadiazine had a lower release rate. The highest cumulative release of Ag ions was from SF nanofibers containing silver tetrafluoroborate due to

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Journal Pre-proof the high solubility of this type of Ag compound. All Ag compounds incorporated in SF nanofibers showed toxicity on fibroblasts and normal human keratinocytes. Therefore, it is essential to optimize the structures to minimize the toxicity to epidermal cells [177]. In another investigation, it was reported that increasing the concentration of silver sulfadiazine within SF nanofibers could decrease the number of fibroblast cells and human keratinocytes [178]. SF nanofibers containing 1 wt% silver sulfadiazine had faster and better wound repair than

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commercial Acticoat dressing; however, it hindered the adhesion of epidermal cells on

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nanofibrous SF mats [178].

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SF nanofibers containing gold nanoparticle (AuNPs) are also useful for tissue regeneration.

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Accordingly, a tubular SF nanocomposite containing AuNPs was fabricated by electrospinning

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as nerve conduits [179]. The prepared electrospun conduit improved the growth and adhesion of Schwann cells without considerable immunogenicity and toxicity in vivo. The nerve myelination

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was increased after implantation of conduits pre-seeded with Schwann cell (Figure 14). The

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existence of AuNPs in the electrospun conduits did not provoke the inflammatory responses in the host. The strong interaction between nanoparticles and SF nanofibers preserved the AuNPs

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within the nanofibers and restricted their migration to the surrounding tissue [179].

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Figure 14. Detail ultrastructure of the sciatic nerve regeneration observed under TEM using silk

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nanocomposite. Each axon shows the myelin sheath encapsulation stained black. The myelin sheath cover in a lamellar shape over the regenerated axon. In the group of AuNP-SF and AuNP-

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SF cells, more axons with thick deposition of myelin are seen in groups after 18 months’ posttreatment than only SF and SF with cell groups. Reprinted with permission from [179].

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In another study, electrospun SF nanofibers doped with AuNPs were prepared [15]. The SFAuNPs composite nanofiber had 70% higher Young’s modulus than bulk SF nanofibers. The proliferation of hMSCs was two-fold higher on surface-modified SFAuNPs containing RGD peptide than bulk SF nanofibers [15]. Similarly, it was shown that incorporating AuNPs in SF matrix enhanced the mechanical properties and decreased the degradation rate of the SF scaffold [180]. The SF matrices with or without AuNPs provoked medium to high level of inflammatory responses and granulation tissue formation, wound closure, and neovascularization. However,

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Journal Pre-proof the addition of AuNPs into SF matrices did not have a remarkable effect on wound healing stages as a result of the low and passive loading of AuNPs inside the SF matrix [180]. 15. Conclusion and future perspective The controlled delivery systems are essential for regenerative medicine and treating the chronic diseases. They are necessary to avoid the burst release of drugs, which decrease the systemic

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toxicity and inevitable side effects. This also increases the therapeutic outcomes in the targeted

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site. It is possible to fabricate different delivery carriers using different polymers. Among them, the electrospun silk fibroin mats are of interest for the drug delivery applications.

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Electrospinning enables the fabrication of silk nanofibers with tailorable diameter and thickness

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to control the release kinetics of the drug in a specific therapeutic window. Despite the

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advantageous properties of electrospun nanofibers as drug delivery platforms, they still face some drawbacks. One of the main challenges is the burst release of the loaded drug which might

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be due to the accumulation of the drug near the surface of the nanofibers [181]. This

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phenomenon probably occurred when the drug is mixed with or encapsulated into the polymeric

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solution before performing the electrospinning process. Moreover, preserving the bioactivity of the drug during the electrospinning process is another important challenge. Applying ultrasonication or high-voltage during the electrospinning may change the structure and biofunctionality of the biomolecule [182]. Among different electrospinning processes, emulsion and co-axial electrospinning have been more interested due to the ability to shield pharmaceutical agents and decrease the complications of initial burst release. The core-shell structures produce by co-axial electrospinning protect the drug from the electric field because the drug is loaded within the core of the nanofibers and coated by the polymeric shell. In addition, the drug-

59

Journal Pre-proof polymer interaction only occurs during the fabrication process of nanofibers, which remarkably reduce the unwanted reaction between the polymeric solutions and the loaded biomolecules [36]. Among various polymeric structures for electrospinning, silk fibroin has gain more attention in recent years. There are many reports confirming the in vitro capabilities of the electrospun silk fibroin mats for drug delivery. There are still limited in vivo investigations that show the potential of electrospun silk fibroin mats as controlled drug delivery systems. Despite the good

of

amount of information regarding silk structure, the conditions for loading, binding, and release

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mechanisms of different drugs from silk fibroin nanofibers are not fully understood.

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Furthermore, the commercially and FDA approved delivery systems are mostly based on

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synthetic polymers e.g., PCL, and PLGA; while the FDA approved natural polymers such as

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albumin, collagen, gelatin, and silk fibroin are still under investigations [183-185]. The first FDA approved of silk material has been the biomedical suture and has also approved for regenerating

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soft tissues (Seri-scaffold -2008, Serica Technologies). The 510k clearance (Silk Voice -

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K180631 - 2018, Sofragen) as an injectable filler applicable for vocal fold insufficiency has also received approval [185]. Nevertheless, silk sutures have been replaced with sutures made from

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synthetic polymeric because of some degrees of immunogenicity and differences in degradation rates after implantation at various sites [73, 186]. Native silk fiber has been considered as an allergenic agent [187] that induce Type I allergy e.g., asthma and increase the level of Immunoglobulin E (IgE) [188]. It is assumed that the sericin glycoprotein coated on the outer layer of silk fibers is responsible for stimulating the allergic reaction and pro-inflammatory responses of silk [189]. On the contrary to many studies that reported minimal immune stimulation by fibroin, the role of fibroin in stimulating the type I allergic responses was also reported [188]. Another study also showed that the higher β-sheet and silk II conformation in 3D

60

Journal Pre-proof fibroin structure induced more immunogenic responses than 2D fibroin film with silk I conformation [190]. Another main challenge of using electrospun silk nanofibers for drug delivery is using organic or acidic solutions such as hexafluoro-2-propanol and formic acid to prepare electrospinnable silk solution. Moreover, the silk nanofibers produced by electrospinning usually have silk I or random coil conformations with low stability. These nanofibrous mats need to immerse into

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organic solvents such as methanol to induce the formation of more stable antiparallel β-sheet

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conformation (silk II) in the nanofibers. However, using organic and caustic solvents might

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damage the structure and bioactivity of the biomolecules, especially sensitive drugs [191].

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Therefore, it is preferred to apply more mild conditions and avoid using organic solvents to

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minimize the toxic effects of the residual solvents. As reported in the text; however, silk fibroin plays a role in stabilizing the pharmaceutical agents and/or small molecule as a function of

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adsorption, entrapment, covalent binding, and/or encapsulation. Preserving the stability of drugs

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is one of the important aspects that must be considered when designing a sustained drug delivery

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system during the time of release.

The physico-chemical characteristics of silk fibroin affect the capacity of drug binding and may be rather limited. For example, silk fibroin has a PI 4.2 due to the presence of negatively charged side chains in its hydrophilic heavy chain. At pH values more than 4.2, they can interact with positively charged drugs compared with negatively charged ones. By using SF modifications, it is possible to add some proper functional moieties on silk fibroin for interaction with negatively charged drugs [20]. In addition, it is necessary to adjust the degradability of the delivery systems to optimize the release kinetics of the loaded drug according to the healing time rate. Most of the polymeric drug 61

Journal Pre-proof carriers have a fast degradation rate during weeks or months. The high degradability restricts their applications, especially in load-bearing purposes. Moreover, some mostly used polymers e.g. PLGA produce acid residues upon hydrolytic degradation that stimulate pro-inflammatory responses in the implanted site. The local decrease in pH, also induces the degradation of proteins through acid catalysis. Silk polymer has a slow degradation rate in vitro, thus the main mechanism for the release of drugs from silk is diffusion. Additionally, silk is a protein that

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commonly degrade by proteolysis and produce non-toxic byproducts that safely metabolize in

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vivo [192, 193]. By manipulation of silk crystallinity, it is possible to both control its

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degradability and drug release kinetics. The crystalline domains of silk fibroin are responsible for

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fibroin self-assembly and strong physical interactions with pharmaceutical agents. It is feasible to control the rate of silk biodegradation by manipulating the degree of crystallinity through

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methanol or water vapor annealing method [185]. Extensive efforts have been performed to

na

investigate the relationship between the properties of silk solution and the resulting electrospun silk fibroin properties. However, predicting the characteristics of the electrospun silk fibroin

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nanofibers is still challenging because it is related to several parameters. For example, one of the

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main factors that affect the spinnability of silk fibroin solution is viscosity which is associated with the concentration of silk fibroin solution and molecular weight. To develop electrospun silk fibroin as drug delivery systems, it is possible to adjust electrospinning parameters, the concentration of silk fibroin solution, and blending silk fibroin polymer with hydrophilic polymers. By adjusting these parameters, it is possible to prepare a silk fibroin nanofibers with optimal properties for controlling the drug release. Another feature that makes silk fibroin as a potential drug delivery carrier is the versatility of methods for sterilization. Most of the polymers used as drug delivery systems, whether natural or

62

Journal Pre-proof synthetic, face limitations for sterilization due to sensitivity to thermal treatments. Autoclaving, exposure to ethylene oxide vapor, and gamma irradiation are the common sterilization methods for silk fibroin polymer [194]. However, the high heat produced by autoclave can destruct the structure of proteins and drugs loaded in silk fibroin matrices. In this case, it is necessary to use alternative sterilization methods or incorporate the drugs after sterilization. Batch-to batch variety of silk is another challenge because it can affect its properties. This limitation can be

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overcome by using genetically engineered silk to reduce these inconsistencies. Totally, through

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many potential modification methods applicable for silk, a silk-based nanofibrous mat is a very

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interesting substrate in the field of drug delivery systems.

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Acknowledgments

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This work was supported by Pasteur Institute of Iran. SCK presently holds European Research

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Area Chair position supported by the European Union Framework Programme for Research and Innovation HORIZON 2020 under grant agreement nº 668983 - FoReCaST

References

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None

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Declaration of Competing Interest

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Journal Pre-proof Highlights 

Delivering the drug at the local area represents many advantages such as fine-tuning of the drug release rate at the site of disease.

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Different types of electrospun nanofibers have been used for drug delivery due to acceptable properties including simple processability, the capacity for different

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functionalization, and the ease of producing nanofibers with high surface area. Some advantages of using silk fibroin as a drug carrier are performing mild aqueous

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processes for loading sensitive drugs e.g., protein, peptides to provide proper resistance to

The existence of different chemical groups such as amines, phenol, alcohol, thiol, and

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enzymatic, dissolution, and thermal degradation.

pharmaceutical agents.

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Electrospinning enables the fabrication of silk nanofibers with tailorable diameter and

ur

thickness to control the release kinetics of the drug in a specific therapeutic window.

Jo



lP

carboxyl in the backbone of silk fibroin also facilitates the interaction with different

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