Introduction

Extensive studies have shown the importance of using biocompatible and biodegradable polymers for drug delivery applications (Hughes 2017). Broadly speaking, many natural, viz., polysaccharides and proteins, and synthetic, viz. poly(α-hydroxy esters) polymers play a pivotal role in this respect, by virtue of their biocompatibility and biodegradability. However, the use of natural polymers in certain delivery applications can be disadvantageous from degradation perspective, exhibiting heightened, uncontrolled and inconsistent degradation (Abedalwafa et al. 2013). In contrast, uniqueness of synthetic polymers lies in the feasibility to tailor their chemical, physical and mechanical features, controlling their degradation rate (Dhandayuthapani et al. 2011; Gentile et al. 2014).

Specifically, synthetic polyesters show great promises, including polymers such as Poly-epsiloncaprolactone (PCL), poly(glycolic acid) (PGA),poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA). By comparison with other polyesters, PCL degradation is exclusively very slow, limiting its utilization in some specific biomedical fields and necessitating its modification such as surface decoration or forming composite systems with other natural and synthetic polymers (Abedalwafa et al. 2013). PGA, on the other hand, exhibited faster degradation rate despite its structural similarity with PLA, and hence, this polymer cannot meet up with the requirements of advanced delivery systems (Gentile et al. 2014; Vroman and Tighzert 2009).

In the family of polyesters, PLA and PLGA, also known as polylactide and polylactide-co-glycolide respectively, have been the most used for drug delivery applications, owing to their attractive mechanical and processing characteristics (Martins et al. 2018; Reis et al. 2017). In addition, contrary to others biodegradable polyesters which are synthesized from petrochemical sources such as polyhydroxyalkanoates (PHA) and poly(butylene succinate) (PBS), PLA and PLGA polymers are produced by lactic and glycolic acid fermentation from sugars which make them more green and eco-friendly (Dischert and Soucaille 2015; Juturu and Wu 2016; Koivistoinen et al. 2013).

Several systems based on PLA or PLGA have been prepared to deliver different payloads ranging from small drug molecules to large proteins and nucleic acids in a sustained release manner (Sun et al. 2014). The rate of drug release and cellular uptake from such systems could be easily tunable by changing several factors such as particle size, polymer crystallinity, polymer molecular weight (Mwt), polymer chemical composition, drug loading and drug-polymer interaction (Kamaly et al. 2016).

Being biocompatible, PLA and PLGA produce safe and non-toxic degradation products which made them good candidates for many medical and pharmaceutical applications (Alsaheb et al. 2015). The biodegradation of PLA and PLGA-based homo and copolymers involve their hydrolytic degradation to lactic and glycolic acids which are finally eliminated from the body as CO2 and water (Tačić et al. 2017). The biodegradation of PLA and PLGA based drug delivery systems (DDS) is very important with regard to controlling the release of loaded cargos (Kamaly et al. 2016). As the polymer degrades and erodes, the drug release kinetics is a function of polymer degradation as well as drug diffusion through polymer matrix (Xu et al. 2017).

Various microparticles (MPs) and nanoparticles (NPs) can be designed using FDA-approved biodegradable and biocompatible PLA and PLGA. These particulate systems possess excellent biocompatibility, controllable biodegradability and erosion and high safety profiles (Agrahari et al. 2016). Therapeutic agents and biologics can be encapsulated within such polymeric particulates for localized and systemic delivery (Yan et al. 2012). Many multifaceted merits are realized in PLA and PLGA based DDS such as sustainment of drug release, targeted delivery, improving the bioavailability, and enhancement of the stability of encapsulated biopharmaceuticals towards enzymatic degradation (Moghimi et al. 2001; Mundargi et al. 2008). With great effort, in the last years, different PLA and PLGA based drugs delivery systems have been developed, in particular for the delivery of anticancer drugs (Haggag et al. 2019; Jain et al. 2011, 2016; Shao et al. 2018) and for a specific organ targeting such as brain or liver targeting (Bao et al. 2015; Gao et al. 2015; Jose et al. 2014; Patel et al. 2018; Xia et al. 2012; Zhu et al. 2016). Moreover, these delivery systems have been tested with great results for pulmonary administration of different drugs (Feng et al. 2014; Gaspar et al. 2019; Takami and Murakami 2011). Furthermore, therapeutic gene delivery based on these systems has been explored especially loading pDNA, mRNA, siRNA and microRNA (Gomes dos Reis et al. 2019; Matta and Maalouf 2019; Nishio et al. 2018; Terry et al. 2019; Xu et al. 2018; Zhao et al. 2018).

We aim in this manuscript to briefly review the most popular biodegradable polymers, PLA and PLGA, as well as micro- and nanoparticle systems with regard to their biocompatibility and biodegradation and their impact as delivery systems for vaccines, drugs and genes with optimal efficacy.

Poly (lactic acid) (PLA) and poly (lactic-co-glycolic acid) (PLGA)

PLA and PLGA are among the most highly studied polymers used in medical applications for drug delivery (Brown et al. 2015; Ramot et al. 2016). PLA present asymmetric centers in its backbone resulting in the formation of either d or l forms (PDLA or PLLA), while PLGA is the copolymer of d,l-lactic acid with glycolic acid (Silva et al. 2015). Figure 1 depicts the chemical structures of PLA and PLGA polymers. It is worth to say that PLLA acquires crystalline form while PDLA with disordered polymer chains is amorphous (Müller et al. 2013). Moreover, the absence of side chain in case of glycolic acid in PLGA enhance chain packing and hence polymer crystallinity when compared to PLLA of similar molecular weight (Tesfamariam 2016).

Fig. 1
figure 1

Chemical structure of A PLA and B PLGA

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PLA and PLGA are easily shaped to DDS of various architectures, being soluble in many solvents including chloroform, dichloromethane, tetrahydrofuran (THF), acetone and ethyl acetate (Mohammadi-Samani and Taghipour 2015). They are insoluble in water; however, they absorb water and degrade by hydrolysis of their ester linkage. The hydrophobic side chain (methyl groups) of PLA renders the polymer with lower water uptake properties and slower degradation compared to PLGA (Fig. 1). Therefore, the drug release from PLGA would be much faster than that from PLA (Makadia and Siegel 2011). Many polymeric DDS, including MPs and NPs, were developed using PLA, PLGA and their copolymers for delivering various drugs (Kapoor et al. 2015; Mohammadi-Samani and Taghipour 2015; Tyler et al. 2016).

Biocompatibility of PLA and PLGA

The biocompatibility of PLA and PLGA was tested and proofed both in vitro and in vivo (Mir et al. 2017). The ultimate biocompatibility of polymers is influenced, not only by polymers but also by their nature of degradation products and degradation rate (Engineer et al. 2011). As mentioned above, PLA and PLGA polymers degrade by esterase enzyme, forming lactic and glycolic acids that enter Kreb`s cycle to be eliminated as CO2 and water through respiration, feces, and urine (Silva et al. 2015). Accumulation of these polymers, when used in low concentration, in the human body organs has never been reported. Nevertheless, in some cases due to failure on the elimination of degradation by-products, the in situ accumulation of acidic (lactic and glycolic acids) degradation products become dangerous, altering biological response when in contact with tissues at high local concentrations (Ramot et al. 2016).

Also, the release of small non-degraded fragments and particulates during polymer erosion process could activate different unexpected immune responses by macrophages (Lyu and Untereker 2009). Moreover, particulates type of PLA and PLGA based systems is of prime importance, when considering toxicological issues. Due to their enhanced cellular uptake and biodistribution, NPs might have affected cell growth, viability and tissue responses (Dailey et al. 2006; Makadia and Siegel 2011). Thus, investigation of particles safety profile, on a cellular or tissue level is highly warranted.

Although larger polymeric microparticles cannot be taken up by macrophages (Champion et al. 2008; Gustafson et al. 2015), macrophages in proximity to polymer surface secretes polymer damaging agents that help MPs erosion process and help starting phagocytosis process (Ronneberger et al. 1996). Furthermore, the interior of macrophage lysosomal vesicles has highly acidic pH (~ 3) (Galloway et al. 1983), compared to the extracellular environment, that can accelerate the process of PLA and PLGA degradation (Zhou and Deng 2002).

It is worthy to mention that biocompatibility is not only the polymer’s intrinsic property-dependent or particulate type-dependent, but also, biological environment-dependent and hence, intensity and length of specific polymer-tissue interactions can be varied greatly in different organs, tissues and species (Makadia and Siegel 2011; Ramot et al. 2016, 2015). Thus, site of administration of PLA and PLGA based DDS greatly impacted foreign body responses. The prolonged tissues/particulates interface and their continual presence inside the tissues can elicit intolerable inflammation, injury and immunological rejection (Ramot et al. 2016). For example, implantation of PLA and PLGA based DDS, with a high surface area and low injection volume via subcutaneous (SC) or intramuscular (IM) routes may increase the risk of a persistent local foreign body reaction over the duration time of polymer degradation.

PLA and PLGA local tissue responses involve three steps: (i) Organization of the inflammatory responses; (ii) Monocytes migration to DDS injection site which differentiate into macrophages and fibrous capsule development; (iii) Rapid degradation of the polymer and enhanced formation of fibrous tissues generated in second step (KyungáKim 2016). A diagrammatic illustration of tissue response steps to PLA and PLGA DDS is presented in Fig. 2.

Fig. 2
figure 2

Tissue response steps to PLA and PLGA drug delivery systems

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Specifically, recognition of the intravenous (IV) NPs by the reticuloendothelial system (RES) is heightened, resulting in their phagocytosis. However, the rationale design of functional long-circulating PLGA NPs via surface modification with hydrophilic polymers such as polyethylene glycol (PEG), poloxamers and polysaccharides greatly stabilizes NPs against opsonization (Gref et al. 2000; Mahapatro and Singh 2011; Yoo et al. 2010). Interestingly, the alleviating influence of incorporated therapeutic agents insides PLA and PLGA DDS must be considered as such tissue responses can be greatly diminished due to drug encapsulation (Ramot et al. 2016). For instance, loading PLGA MPs with anti-inflammatory drugs such as dexamethasone has been demonstrated to provide long term control over inflammation and fibrosis, improving their performance following SC implantation (Bhardwaj et al. 2010; Dang et al. 2011).

Concerning utilization of PLA and PLGA based DDS to other administration sites such as intravitreal, intraparenchymal “for brain delivery”, intratracheal and intracochlear implantations, previous studies evidenced that they meet the biocompatibility requirements. Mild tissue responses have been shown, if any with the absence of histological alterations or cell ultrastructure modification (Ensari et al. 2015; Kim et al. 2011; Lewitus et al. 2011; Rong et al. 2014).

Biodegradation of PLA and PLGA

PLA and PLGA belong to the large family of biodegradable aliphatic polyesters (Vroman and Tighzert 2009). PLA and PLGA polymers, prepared with different isomer composition and polymer molecular weight, have variable melting points and crystallinity (Ahmed and Varshney 2011). Moreover, the biodegradation of these polymers in the human body can be controlled (i.e. retarded) by the incorporation of d-lactide isomer in their backbones. The fact that our bodies do not produce enzymes that can utilize d-lactic acid results in reduced degradability of polymers containing d-lactide monomer (Gorrasi and Pantani 2013).

The biodegradation of polyesters including PLA and PLGA mainly occurs by hydrolytic degradation which may be enzyme catalyzed (Elsawy et al. 2017). Several factors can affect hydrolytic degradation process including (i) water permeability and solubility, (ii) chemical composition and molecular weight and molecular weight distribution, (iii) the mechanism of hydrolysis (noncatalytic, autocatalytic, enzymatic), (iv) additives (acidic, basic, monomers, solvents, drugs), (v) polymer crystallinity, (vi) glass transition temperature (glassy, rubbery) and other physicochemical factors (e.g. ionic strength, pH) (Antheunis et al. 2009; Frank et al. 2005; Makadia and Siegel 2011; Xu et al. 2017; Zolnik and Burgess 2007).

The most important factor affecting the degradation of PLA and PLGA and particles prepared by them is polymer molecular weight (Anderson and Shive 1997) and degree of crystallization (Mitchell and Hirt 2015). Crystalline parts of the polymer show more resistance to degradation when compared to amorphous regions (Göpferich 1996). The crystallinity of the polymer depends on its composition i.e. increasing the percentage of glycolide monomer in PLGA backbones decreases polymer crystallinity as it decreases chain rearrangement to produce crystalline structures (Keles et al. 2015; Xu et al. 2017).Moreover, it was found that increasing PLA and PLGA molecular weights, their degradation rate decreases because of the high chain–chain crossing between long polymer backbones which lead to the resistance of chain cleavage (Makadia and Siegel 2011).

The extent of water uptake affects the hydrolytic degradation of PLA and PLGA as water reacts with polyesters, resulting in reverse polycondensation (Alexis 2005). Various factors affect water uptake of these polymers including; molecular weight, purity, morphology, and the polymer processing techniques (Antheunis et al. 2010; Dorati et al. 2007; Keles et al. 2015). The abundance of water is critical to start ester bond degradation to oligomers and monomers, especially for high molecular weight polymers. Furthermore, polymer porosity helps in increasing polymer biodegradation by increasing water uptake as well as cellular migration through polymer macropores (Rodriguez et al. 2016).

PLA and PLGA degradation was found to increase in acidic (Zolnik and Burgess 2007) and basic (Wang et al. 1998) conditions, hence, the incorporation of large amounts of acid and basic drugs (D’Souza et al. 2015) and/or additives (Yoo et al. 2005) to them enhances degradation kinetics. It is also worth to note that phagolysosomal vacuoles of macrophages have acidic content (pH ~ 3) which, when secreted on the interface with the polymers, facilitate their biodegradation (Anderson and Shive 1997).

Methods used for the characterization of polymer degradation

Many techniques have been used to investigate the process of PLA and PLGA degradation through tracking the physical or chemical changes that happen during polymer breakdown to oligomers and monomers (Engineer et al. 2011; Keles et al. 2015). Figure 3 shows different techniques used to probe polymers degradation. The different techniques usually involve either direct investigation of the variations in polymer molecular weight with time (e.g. gel permeation chromatography (GPC) (Bawa et al. 2018) and time-of-Flight secondary ion mass spectrometry (TOF–SIMS) (Marchany et al. 2015)) or testing the changes in different physical properties which are correlated to polymer size (e.g. crystallinity and thermal properties i.e. differential scanning calorimetry (DSC) and thermo gravimetric analysis (TGA)) (Turek et al. 2015). Moreover, degradation can be assessed by tracing properties of the polymer such as loss of molecular mass which can be demonstrated using GPC and changing of polymer crystallinity which can be assessed by studying changes in its thermal properties.

Fig. 3
figure 3

Different techniques used to probe degradation of polymers

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Other techniques involve following the buildup of monomer molecules during degradation and this could be followed using pH changes for acidic and basic monomers, high performance liquid chromatography (HPLC) (Giunchedi et al. 1998), Fourier transform infrared spectroscopy (FTIR) (Leroy et al. 2017), nuclear magnetic resonance (NMR) and Raman spectroscopy (Kasyapi et al. 2015). Qualitative methods by imaging polymer using scanning electron microscopy (SEM) (Liu et al. 2006) and atomic force microscopy (AFM) (Giannouli et al. 2018) have been used to check polymer degradation by probing morphological changes with time.

Degradation pathways and drug release models

The degradation of PLA and its copolymers occur by three hydrolytic pathways: (i) surface degradation, (ii) bulk degradation and (iii) bulk degradation with autocatalysis (Gajjar and King 2014) (Fig. 4). For the surface degradation or erosion, the cleavage of ester bonds of polymers occurs mainly on the surface, resulting in the formation of monomers and oligomers that diffuse to the degradation medium faster than those produced in the polymer bulk. Bulk erosion happens when degradation medium penetrates polymer matrix and random hydrolysis occurs throughout the polymer bulk, reducing its backbone molecular weight. During the gradual bulk erosion process, as long as the degradation medium diffuses throughout the hydrolyzing matrix, monomers and oligomers diffuse to the degradation medium and the polymer density decrease (von Burkersroda et al. 2002). When the bulk degradation results in the formation of a higher concentration of acidic degradation products in the polymer interior compared to the polymer surface, autocatalysis may happen, accelerating the internal degradation. As the polymer interior degradation increases, monomers and oligomers diffuse rapidly, producing polymer matrix with hollowed core (Antheunis et al. 2010).

Fig. 4
figure 4

Degradation pathways for PLA and its copolymers: A surface degradation, B bulk degradation, C bulk degradation with autocatalysis

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PLA and PLGA degrade with second order (2nd) order kinetics depending on both ester bond and water concentration (Lyu and Untereker 2009). As hydrolysis or autocatalysis proceeds, the number average molecular weight (Mn) of polymers decreases with the increase in chain end concentration (CE) and this could be quantified by the following equation:

$$CE = \frac{X}{V} = \frac{\rho }{Mn} = \frac{\rho }{{N\!\cdot \!M_{0} }}$$
(1)

where X, V, and ρ: total number of chains, volume, and density of the samples; N: degree of polymerization; M0: monomer molecular weight.2nd order hydrolysis reaction can be described as follows:

$$\frac{d CE}{dt} = k_{2} C_{e} C_{w}$$
(2)

where k2: rate constant; Ce: ester bond concentration; Cw: water concentration.

By combining Eqs. (1, 2), the number average molecular weight (Mn) can be calculated from 2nd order hydrolysis equation:

$$\frac{1}{Mn} = \frac{1}{{Mn_{0} }} + \frac{1}{{M_{0} }}k_{2} C_{e} t$$
(3)

where Mn: number average molecular weight at time t; Mn0: initial number average molecular weight before starting hydrolysis reaction. In the case of auto-accelerated degradation or autocatalysis, the degradation rate is proportional to the chain bond concentration as well as water and hydrolytic bond concentrations as follows:

$$\log Mn = \log Mn_{0} - k_{3} C_{e} C_{w} t$$
(4)

where k3: autocatalysis rate constant.

The release of different drugs from PLA or PLGA matrix is the sum of surface and bulk diffusion as well as matrix erosion mechanisms (Makadia and Siegel 2011). The drug release occurs mainly by diffusion when the polymer weight loss is minimal. This could be described by the following equation:

$$\frac{\partial C(r,t)}{\partial t} = \frac{1}{{r^{2} }}\frac{\partial }{\partial r}\left( {r^{2} D(M_{w} )\frac{\partial C(r,t)}{\partial r}} \right)$$
(5)

where C: drug concentration; r: radial position; t: time; D(Mw): polymer molecular weight dependent diffusivity of the drug.

When polymers are formulated into MPs or NPs, the effect of population size distribution on drug release can be described by the following equation:

$$\frac{1}{{M_{\infty } }}\frac{{\partial M_{drug} }}{\partial r} = f\left( {r;\lambda ,k} \right) = \frac{k}{\lambda }\left( {\frac{r}{\lambda }} \right)^{k - 1} \exp \left( { - \left( {\frac{r}{\lambda }} \right)^{k} } \right)$$
(6)

where M∞: drug total mass in the population; Mdrug: drug mass in one particle and r is the particle radius.

Biomedical applications of PLA and PLGA nano- and microsystems

During the last decade, several research groups have prepared nano- and microsystems using PLA and PLGA for many therapeutic applications. They aimed to use these biodegradable and biocompatible polyesters to modify drug release and enhance its activity and optimal therapeutic effectiveness (Lee et al. 2016; Sharma et al. 2016). In the following sections, we will discuss part of extensive applications of the currently developed PLA and PLGA based nano and microsystems with emphasis on their beneficial effect in the delivery of biopharmaceuticals, viz., drugs, vaccines and gene delivery (Tables 1, 2 and 3). Examples of recent research studies in the last ten years which deal with biodegradation and biocompatibility of PLA and PLGA nano- and microparticulate delivery systems and their main achievements are covered in this review.

PLA and PLGA based MPs and NPs delivery systems for vaccine delivery

The potential of polyesters to produce delivery platforms adequate for vaccination for preventive and therapeutic purposes has been increasingly explored. Their attractiveness in protein and peptide delivery stemmed from their ability to act as safe and efficacious vaccines adjuvants and to co-encapsulate antigens and/or immunomodulators and other vaccine adjuvants (Ahmed and Varshney 2011; Demento et al. 2012; Kasturi et al. 2011; Mundargi et al. 2008; Schlosser et al. 2008; Wischke et al. 2009). Antigen-containing polymeric particulates get due recognition of strong and long lasting T-lymphocytes responses as compared to the soluble antigens (Cruz et al. 2017). Other striking features include protection of their payload from proteolytic degradation, sustained antigen availability, peptide dosage reduction, minimized immunization times and reduced toxicity (Demento et al. 2012; Getts et al. 2014; Tam et al. 2016).

As a consequence, the coordination of more robust optimal humoral and cellular arms of the immune response, mimicking the natural infection influence, could be achieved (Demento et al. 2012; Kasturi et al. 2011). The mechanism behind such stimulatory capacity is the phagocytosis of the particles by dendritic cells (DCs), the most professional cellular target, leading to the maturation markers up-regulation (such as CD40 and CD80) and antibody and T-lymphocytes activation and proliferative response to different antigens (Anderson et al. 2008; Clawson et al. 2010; Danhier et al. 2012; Luzardo-Alvarez et al. 2005; Rosas et al. 2001; Waeckerle-Men and Groettrup 2005). Interestingly, anti-cancer immunotherapy could be realized by such immune system alerts, through their capability to encapsulate ligands for Toll-like receptors (TLRs) (Hamdy et al. 2011; Kranz et al. 2016; Pinzon-Charry et al. 2007; Seth et al. 2017; Zhang et al. 2011).

The utilization of particulate formulated vaccines with controlled size (preferentially in the micron range), charge, polymer composition, molecular weight, degradation rate and, hydrophobicity could guarantee non-specific long lasting DCs targeting (Feng et al. 2006; Kanchan and Panda 2007; Kazzaz et al. 2006). In particular, degradability governed the second phase of the mechanism behind the release of peptides/proteins, following the initial phase of diffusion, and hence, their immunization response and the proposed regimen (Taha et al. 2012). In addition, nature of co-adjuvants and delivery route play a critical role in the efficacy of these protective immunity generators (Ashhurst et al. 2018; Newman et al. 2002; Renukuntla et al. 2013). A selection of some examples on the application of PLA and PLGA based MPs and NPs delivery systems for vaccine delivery is reviewed in Table 1, with a focus on the biocompatibility of the prepared systems and their therapeutic outcomes.

Table 1 Applications of PLA and PLGA based MPs and NPs delivery systems for vaccine delivery
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PLA and PLGA based MPs and NPs systems for drug delivery

Past efforts have realized great achievements in therapeutic strategies using drugs, both small molecules and macromolecules, in combination with PLA and PLGA biomaterials (Makadia and Siegel 2011). In particular, polyesters based therapeutics gain success in improving chemotherapeutic efficacy, both in vitro and in vivo, and anti-inflammatory response for treatment of cancer (Hu and Zhang 2012) and inflammatory diseases (Danhier et al. 2012) respectively in various administration routes.

Bioinert and biodegradable PLA and PLGA drug carrier materials hold outstanding advantages, including (i) sustainable delivery (Mansor et al. 2018), (ii) proper control of drug release kinetics (iii) diminished fluctuations of blood drug concentrations (Makadia and Siegel 2011), (iv) augmented cellular uptake of NPs via endocytosis (Bi et al. 2016; Priemel et al. 2018), (v) enhanced stability (Xiong et al. 2016), (vi) optimal clinical utility (Lü et al. 2009) and (vii) improved medication adherence (Liu et al. 2006).

In recent years, researchers have become more focused on designing surface modified PLA and PLGA particulates, possessing prolonged systemic circulation time (Meng et al. 2018). The most relevant way to avoid RES recognition is the PEGylation, via masking hydrophobicity of the particles and imparting stealth characteristics by hydrophilic PEG brush (Mansor et al. 2018). Experimental research has also proved other favorable properties of PEGylated particles such as reducing aggregation while enhancing stability and biological potency (Goudarzi et al. 2018). Moreover, considerable interest has been conducted in the development of ligand coupled NPs like peptide, antibodies, and nucleotides ligands for enhanced localization and active targeting capabilities to specific organs or cells (Dinarvand et al. 2011). Examples of recent publications that have shown promising results towards effective sustained PLA and PLGA based MPs and NPs for drug delivery are listed in Table 2. An emphasis is placed on findings of investigations done to assess degradation and/or toxicity issues of the prepared DDS.

Table 2 Application of PLA and PLGA based MPs and NPs systems for drug delivery
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PLA and PLGA based MPs and NPs systems for gene delivery

Gene therapy interests the researchers of the scientific community by its potential in fighting diseases from their genetic source providing personalized medicine to the patients (Naldini 2015). This specific therapy is therefore promising for the treatment of complex diseases such as chronic infections, cancer and inherited diseases, especially in substitution of standard treatments when they fail to cure the disease.

Many achievements have been reached in the field of gene delivery of a variety of nucleic acids such as pDNA (Baoum et al. 2010), mRNA (Yasar et al. 2018), siRNA, and microRNA. These nucleic acids can both introduce genes that encode a functional protein that is vital in preventing disease progression or block the translation of specific mRNAs to prevent a toxic effect. Multiple obstacles are present after the administration of nucleic acids such as the presence of nuclease, the nucleic acid instability, biological membrane and endocytosis. The overcome of these obstacles can be achieved by the use of nanotechnology and the incorporation of nucleic acid in different nano and microparticles (Wong et al. 2017; Xiao et al. 2019).

Biodegradable polyesters including PLA and PLGA were highly investigated for preparing gene delivery vectors (Kumari et al. 2010). However, for delivery of therapeutic genes, there is a great challenge of formulating PLA and PLGA particles containing a high payload, owing to their negatively charged backbone. The loading of nucleic acids to PLA and PLGA NPs could be done by either direct encapsulation using double emulsion solvent evaporation technique (Mok and Park 2008) or by the pre cationization (adding cationic polymers) of polyester NPs to enhance the encapsulation efficiency of nucleic acids and accelerate their release (Luten et al. 2008). Cationization of polyester NPs can be done by incorporating cationic molecules such as polyethylenimine (PEI) (Frede et al. 2016; Kolte et al. 2017), chitosan (Tahara et al. 2011), polyarginine (Zhao et al. 2012), poly(2-dimethylaminoethyl methacrylate (pDMAEMA) (Qian et al. 2014), cationic lipids (Dioleyltrimethylammonium propane “DOTAP”–Dioctadecenyltrimethylammoniumpropane “DOTMA”) (Yasar et al. 2018; Zhu et al. 2015), Dimethylaminoethane carbamoyl cholesterol (DC-Chol) (Baoum et al. 2010; Gwak et al. 2016) and cationic cell-penetrating peptide (CCCP) (Jain et al. 2015).

It is well known that the smaller the particle size, the more efficient the gene delivery, however, PLGA particles are easily phagocytosed by immune cells when prepared in size less than 10 μm (Oyewumi et al. 2010). Several trials have been done to modify or sustain the release of nucleic acids from PLA and PLGA based systems. However, their hydrophobic nature renders extended nucleic acid release not feasible being rapidly removed from circulation after parenteral administration (Dinarvand et al. 2011). In order to solve the previously mentioned obstacle, many strategies have been explored including PEGylation, copolymerization with hydrophilic polymers, NPs coating with a surfactant, lipid and targeting ligands (Jain et al. 2015).

PLA and PLGA based NPs loaded with nucleic acids can be delivered to various animal models through different routes of administration such as IM, ID, SC and oral (Bala et al. 2004). Outstanding features of non-immunogenicity and non-toxicity of PLA and PLGA have been keyed the researchers to design PLA and PLGA particles loaded with nucleic acids cargo, such as oligonucleotides, plasmid DNA, and small interfering RNA (siRNA), as an alternative to immunogenic and toxic viral vectors (Xu and Zhang 2009). In addition to previously mentioned PLA and PLGA merits, such biomaterials could give favorable outcomes of potentiating poor intrinsic transfection efficiency of nucleic acids (Heo et al. 2014; Kolte et al. 2017; Xu and Zhang 2009), providing long term gene expression or gene silencing (Jain et al. 2015; Zhu et al. 2015), enhancing translocation inside the cell (Jain et al. 2015), maximizing protection from enzymatic degradation and enhancing gene delivery system cytocompatibility and tolerability (Baoum et al. 2010; Qian et al. 2014; Zhao et al. 2012).

In the future, nanotechnology and gene therapy will be inseparable because, the successful development of efficient non-toxic, non-immunogenic, and biodegradable NPs will improve the efficiency of gene therapy, resulting in breakthroughs for the cure of specific diseases (Wong et al. 2017). Indeed, the successful development and optimization of therapeutic nuclei acids and subsequent clinical gene therapy trials seem to pave the way for recent expansion of market authorization (Lächelt and Wagner 2015).

Special focuses on some research papers, covering the application of PLA and PLGA based MPs and NPs systems that are combined with different cationic polymers for gene delivery as well as addressing safety issue, are summarized in Table 3.

Table 3 Application of PLA and PLGA based MPs and NPs systems for gene delivery
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Conclusion

Development of new DDS with excellent safety profile, avoiding unwanted tissue responses and tunable degradation is highly warranted. Biocompatibility and biodegradability of PLA and PLGA polymers pose them as a promising choice in many local and systemic therapeutic applications. Their degradation rate can be traced and modeled using different techniques and was revealed to affect drug release kinetics from PLA and PLGA-based DDS. Insights into the described mechanisms of sustained drug release, PLA and PLGA MPs and NPs have been utilized for vaccine development as well as encapsulating various conventional drugs, proteins and nucleic acids for gene delivery. Based on the abovementioned data and therapeutic outcomes, the idea of the technology transfer of a number of new DDS could gain grounds in the near future. New therapies might be seen in clinics, specifically in the fields of vaccination and cancer immunotherapy.