**5. Applications of PLA in the biomedical field**

Due to its bioresorbability and biocompatibility in the human body, PLA has been employed to manufacture tissue engineering scaffolds, delivery system materials, or covering membranes, different bioabsorbable medical implants, as well as in dermatology and cosmetics.

### **5.1 Tissue engineering**

262 Biomedical Science, Engineering and Technology

While plasma treatment has been successfully applied to improving PLA wettability and cell affinity, its main disadvantage is that surface rearrangements caused by thermally activated macromolecular motions, to minimize its interfacial energy, can also influence the surface modification. Moreover, the potential influence of plasma on the degradation of PLA

PLA does not carry reactive side-chain functional groups. Consequently, the first step of chemical modification is typically a simple surface hydrolysis (with an alkali) or an aminolysis treatment. The hydrophilic –COOH and –OH or reactive –NH2 groups introduced by cleavage of the ester bonds can be used to bind bioactive molecules such as arginine-glycine-aspartic acid (RGD)-containing peptides, chitosan (CS), arginine and lysine,

The synthetic RGD-containing peptides could be immobilized on PLA after treatment by hydrolysis or aminolysis (Stupack et al., 2001). Materials prepared by this method provide suitable recognition sites for cell adhesion receptors and biodegradation rates, making them suitable for various applications in fields such as tissue engineering and implant technology. It has also been determined that RGD-conjugated poly(lactic acid-*co*-lysine)(arginineglycine-aspartic acid) nanoparticles (PLA-PLL-RGD NP) are non-toxic and bind more efficiently to human umbilical vein endothelial cells (HUVECs) as compared to bare PLA-PLL NP in vitro. Targeted imaging results obtained in vivo showed that PLA-PLL-RGD can selectively bind to BACP-37 breast cancer cells. Lieb et al. also demonstrated largely increased cell densities and cell proliferation on surfaces modified with RGD-anchored monoaminated poly(ethylene glycol)-*block*-poly(*D*,*L*-lactic acid) (H2N-PEG-PLA), mediated

Chitosan (CS) is a biopolymer displaying good biocompatibility, non-toxicity and biodegradability, produced by the alkaline *N*-deacetylation of chitin. The immobilization of this polymer on PLA has been accomplished by coating the surface with chitosan, modified with the photosensitive hetero-bifunctional cross-linking reagent 4-azidobenzoic acid, and irradiation with ultraviolet light to photolyze the azide groups and covalently link the two polymers (Zhu et al., 2002). The -OH and -NH2 groups of chitosan provide further opportunities to introduce a wide range of functional groups on the surface. Thus CS molecules immobilized on PLA were modified with a heparin (Hp) solution to form a polyelectrolyte complex on the surface, which inhibited platelet adhesion and activation,

The surface attributes of PLA can be tailored to enhance its hydrophilicity and biocompatiblity through various methods. Unfortunately, all these established methods for surface modification are inherently flawed to some extent. For example, a single plasma treatment can merely improve cell adhesion but cannot accelerate cell growth; non-covalent attachment of a functional material onto a PLA surface is not stable and permanent. An excellent method suggested to solve the second issue is the use of 1,6 hexanediamine for surface aminolysis, followed by conjugation with biocompatible macromolecules such as gelatin, chitosan, or collagen (Zhu et al., 2004). Hong et al. have

PEG, collagen, and so on to regulate cell adhesion or protein adsorption.

through RGD–integrin interactions (Lieb et al., 2005).

cannot be ignored.

**4.2.2 Chemical modification** 

and enhanced cell adhesion.

**4.3 Outlook of PLA surface modification** 

Since the introduction of the concept in 1988, tissue engineering, a technique invented to reconstruct living tissues by associating the cells with biomaterials that provide a scaffold on which they can proliferate three-dimensionally and under physiological conditions, has emerged as a potential alternative to tissue or organ transplantation and has thus attracted great attention in science, engineering, and medicine. To meet the diverse needs of tissue engineering, scaffolds made from various materials have been tested in this field. Although certain metals are somewhat good choices for medical implants due to their superior mechanical properties, their lack of degradability in a biological environment makes them disadvantageous for scaffold applications (Liu & Ma, 2004). Inorganic/ceramic materials such as HAP or calcium phosphates, being studied for mineralized tissue engineering with good osteoconductivity, are also limited due to poor processability into porous structures (Ilan & Ladd, 2002). In contrast, polymers have great design flexibility because their composition and structure can be tailored to meet specific needs (Huang et al., 2007). Degradable polymers frequently used for tissue engineering applications are linear aliphatic polyesters such as PGA, PLA, and their copolymers (PLGA), which are fabricated into scaffolds. These polymers are among the few synthetic polymers approved by the FDA for human clinical applications. The drawbacks of these polyesters include their hydrophobicity and lack of functional groups, which limits cell adhesion, an important factor when constructing polymeric scaffolds. Another drawback is their slow hydrolytic degradation (Iwata & Doi, 1998).

An ideal scaffold used for tissue engineering should possess the following properties: 1) Be biocompatible, so that the scaffold can be well integrated into host tissues without resulting in any immune response; 2) It should be porous with appropriate pore size, size distribution and mechanical function, to allow cell or tissue growth and the removal of metabolic waste; 3) It must be mechanically able to withstand local stress and maintain the pore structure for tissue regeneration; 4) Very importantly, the scaffold should be biodegradable (Ma, 2004). Synthetic scaffolds are considered important components of a successful tissue engineering strategy (Wang, 2007). Hybrid three dimentional porous scaffolds of synthetic and naturally derived biodegradable polymers are particularly promising because they combine the advantages of the two types of materials. They should maintain sufficient mechanical strength while providing specific cell-surface receptors during the tissue remodeling process that stimulate both in vitro and in vivo cell growth (Chen et al., 2002). PLA-based hybrid materials have been successfully tested clinically for that purpose so far, and tests on other tissues including bladder (Engelhardt et al., 2011), cartilage (John et al., 2003), liver (Lv et al., 2007), adipose (Mauney et al., 2007), and bone tissues (Mathieu et al., 2006) have also been reported.

Jiang et al. (Jiang et al., 2010) functionalized chitosan/PLGA by heparin immobilization with controlled loading efficiency. One of the main benefits of introducing chitosan into PLGA microspheres is that chitosan imparts functionality due to its reactive amino groups, so that biomolecules such as heparin could be attached (Jiang et al., 2006). The compressive strength and modulus remained in the range of human trabecular bone after the heparinization process. More importantly, heparinized chitosan/PLGA scaffolds with a low heparin loading (1.71 g/scaffold) showed a stimulatory effect on cell differentiation, as indicated by enhanced osteocalcin expression as compared with a non-heparinized chitosan/PLGA scaffold. Based on these results, Jiang et al. (Jiang et al., 2006) continued to evaluate the novel scaffolds for bone regeneration in vivo. In the rabbit ulnar critical-sized-defect model created, successful bridging of the critical-sized defect on the sides both adjacent to and away from the radius occurred using chitosan/PLGA-based scaffolds. However, the addition of chitosan to PLGA led to somewhat higher inflammation and lower mineralization than for the PLGA counterpart, which is a major problem that remains to be solved.

Three-dimensional (3D) electrospun fibrous scaffolds have been suggested as a potential tissue engineering tool for bone regeneration. Shim et al. (Shim et al., 2010) thus reported a 3D microfibrous PLLA scaffold fabricated using electrospinning techniques with a subsequent mechanical expansion process. The use of these 3D scaffolds for the proliferation of osteoblasts was examined. The 3D scaffolds led to a 1.8-fold higher level of osteoblast proliferation than generally achieved for electrospun 2D nanofibrous membranes. In vivo results further showed that 3D electrospun microfibrous matrices provided a favorable substrate for cell infiltration and bone formation after 2 and 4 weeks when using a rabbit calvarial defect model.

3D printing technology has rapidly expanded in the tissue engineering field since it was first developed at the Massachusetts Institute of Technology. Ge et al. (Ge et al., 2009) developed 3D-printed poly(lactic acid-co-glycolic acid) (PLGA) scaffolds which could support the proliferation and osteogenic differentiation of osteoblasts. Based on their in vitro study, they also evaluated PLGA scaffolds for bone regeneration within a rabbit model (Ge et al., 2009). In both the intra-periosteum and the iliac bone defect models, the implanted scaffolds facilitated new bone tissue formation and maturation over a time period of 24 weeks.

The current clinical use of PLA-based scaffolds nevertheless remains very limited (Iwasa et al., 2009), mainly because of the risk of disease transmission and immune response. This can be illustrated by taking cartilage tissue engineering as an example. Traditional autologous chondrocyte implantation (ACI), first introduced by Brittberg et al. in 1994 (Brittberg et al., 1994), has yielded good clinical results (Bentley et al., 2003). To date, none of the short- or mid-term clinical and histological results using scaffolds were reported to be better than ACI. As for the scaffolds, collagen and hyaluronan-based matrices are among the most popular scaffolds in clinical use nowadays, since they offer substrates which are normally essential elements in native articular cartilage (Iwasa et al., 2009). Among the very few cases of scaffolds in clinical use is the copolymer of PGA/PLA (polyglactin, vicryl) and polydioxanone, which is used for cartilage repair under the trade name of BioSeed®-B and BioSeed®-C (Biotissue Technologies AG, Freiburg, Germany).

In summary, tissue engineering is one of the most exciting interdisciplinary fields today and is growing rapidly with time. The inclusive criteria for studies on scaffolds capable of clinical application were in vivo or clinical studies and thus certain artificially designed scaffold features (such as pore size, interpore connectivity, etc.) are necessary for optimal tissue engineering applications (accelerated tissue regeneration). Suggestions for future directions include the use of designer scaffolds with in vivo experimentation, and coupling scaffold design with cell printing to create designer material/biofactor hybrids to optimize tissue engineering treatments (Hollister, 2005).

### **5.2 Delivery systems**

264 Biomedical Science, Engineering and Technology

scaffolds. These polymers are among the few synthetic polymers approved by the FDA for human clinical applications. The drawbacks of these polyesters include their hydrophobicity and lack of functional groups, which limits cell adhesion, an important factor when constructing polymeric scaffolds. Another drawback is their slow hydrolytic degradation

An ideal scaffold used for tissue engineering should possess the following properties: 1) Be biocompatible, so that the scaffold can be well integrated into host tissues without resulting in any immune response; 2) It should be porous with appropriate pore size, size distribution and mechanical function, to allow cell or tissue growth and the removal of metabolic waste; 3) It must be mechanically able to withstand local stress and maintain the pore structure for tissue regeneration; 4) Very importantly, the scaffold should be biodegradable (Ma, 2004). Synthetic scaffolds are considered important components of a successful tissue engineering strategy (Wang, 2007). Hybrid three dimentional porous scaffolds of synthetic and naturally derived biodegradable polymers are particularly promising because they combine the advantages of the two types of materials. They should maintain sufficient mechanical strength while providing specific cell-surface receptors during the tissue remodeling process that stimulate both in vitro and in vivo cell growth (Chen et al., 2002). PLA-based hybrid materials have been successfully tested clinically for that purpose so far, and tests on other tissues including bladder (Engelhardt et al., 2011), cartilage (John et al., 2003), liver (Lv et al., 2007), adipose (Mauney et al., 2007), and bone tissues (Mathieu et al., 2006) have also been

Jiang et al. (Jiang et al., 2010) functionalized chitosan/PLGA by heparin immobilization with controlled loading efficiency. One of the main benefits of introducing chitosan into PLGA microspheres is that chitosan imparts functionality due to its reactive amino groups, so that biomolecules such as heparin could be attached (Jiang et al., 2006). The compressive strength and modulus remained in the range of human trabecular bone after the heparinization process. More importantly, heparinized chitosan/PLGA scaffolds with a low heparin loading (1.71 g/scaffold) showed a stimulatory effect on cell differentiation, as indicated by enhanced osteocalcin expression as compared with a non-heparinized chitosan/PLGA scaffold. Based on these results, Jiang et al. (Jiang et al., 2006) continued to evaluate the novel scaffolds for bone regeneration in vivo. In the rabbit ulnar critical-sized-defect model created, successful bridging of the critical-sized defect on the sides both adjacent to and away from the radius occurred using chitosan/PLGA-based scaffolds. However, the addition of chitosan to PLGA led to somewhat higher inflammation and lower mineralization than for the PLGA counterpart, which is a major problem that remains to be

Three-dimensional (3D) electrospun fibrous scaffolds have been suggested as a potential tissue engineering tool for bone regeneration. Shim et al. (Shim et al., 2010) thus reported a 3D microfibrous PLLA scaffold fabricated using electrospinning techniques with a subsequent mechanical expansion process. The use of these 3D scaffolds for the proliferation of osteoblasts was examined. The 3D scaffolds led to a 1.8-fold higher level of osteoblast proliferation than generally achieved for electrospun 2D nanofibrous membranes. In vivo results further showed that 3D electrospun microfibrous matrices provided a favorable substrate for cell infiltration and bone formation after 2 and 4 weeks when using a rabbit

(Iwata & Doi, 1998).

reported.

solved.

calvarial defect model.

There has long been a desire to achieve the targeted delivery of bioactive compounds to areas in the body to maximize therapeutic potential and minimize side-effects. Many types of particles have been tested as delivery tools for biomedical applications such as liposomes, solid lipid nanoparticles, and biodegradable polyesters like PLA and PLGA (Torchilin, 2006). With its excellent biocompatibility, biodegradability, mechanical strength, heat processability, and solubility in organic solvents, PLA can be used to produce dosage forms such as pellets, microcapsules, microparticles (MP), nanoparticles (NP), etc. MP and NP of PLA, modified or unmodified, are increasingly investigated for sustained release and targeted drug, peptide/protein, and RNA/DNA delivery applications because of their small size enabling their permeation through biological barriers such as the blood-brain barrier (Roney et al., 2007).

Although PLA-based materials such as PLGA have been FDA-approved and are clinically available, they lack chemical functionalities to facilitate specific cell interactions. Furthermore, their potential for the sustained release of hydrophilic molecules (e.g. proteins) is often limited (Fahmy et al., 2005). Frequent undesired effects include low encapsulation efficiency and high burst release of the encapsulated biomolecule within the first few hours or days, which is mainly due to the desorption of surface-associated hydrophilic molecules having weak interactions with the polymer (Fahmy et al., 2005). To circumvent these limitations and establish therapeutic efficacy, large doses or site-specific administration are often required for devices comprised of polyester biomaterials. In an attempt to address these problems, numerous groups have introduced functional groups (such as amine functionalities) on these materials, either through direct conjugation or device fabrication with additives (Betram et al., 2009).

As for drug release from MPs or NPs, it is generally controlled by both drug diffusion and polymer degradation. To ensure the efficacy of drug delivery, control over the particle size and particle size distribution is critical, since smaller particles and narrower size distributions facilitate the design of targeted drug delivery systems. These involve binding fragments specific to a tumor-associated surface antigen, with a ligand binding to its corresponding receptor on the tumor cell surface, which can be attached on the surface of the PLA-based materials. Furthermore, polymers that display a physicochemical response to changes in their environment are being intensively explored as potential drug and gene delivery systems. The use of stimuli-responsive nanocarriers offers an attractive opportunity for targeted delivery, in which the delivery system becomes an active participant rather than a passive vehicle. The advantage of stimuli-responsive nanocarriers becomes obvious when the stimuli are unique to disease pathology, allowing the nanocarrier to respond specifically to the pathological characteristics. For instance, in solid tumors, the extracellular pH decreases significantly from 7.4 (the pH value under normal physiological conditions) to about 6.5 (Vaupel et al., 1989; Haag, 2004). In addition, the pH ranges from 4.5 in lysosomes to about 8.0 in mitochondria. Given these pH shifts, therapeutic compounds with a *p*Ka between 5.0 and 8.0 are able to exhibit dramatic changes in physicochemical properties (Ganta et al., 2008). Another option is thermo-sensitive polymeric micelles, containing a hydrophobic core and a thermosensitive shell, the later changing from a hydrophilic nature at body temperature to a collapsed hydrophobic polymer at a hyperthermic condition of 42°C (Na et al., 2006). Investigations concerned with this theme include responses induced by chemical substances, changes in temperature (Tyagi et al., 2004) or pH (Sethuraman & Bae, 2007), electric signals (Sawahata et al., 1990) or other environmental conditions (Qiu & Park, 2001).

The use of nucleic acids as therapeutic agents for genetic diseases has been extensively studied (Torchilin, 2008). However, a major limitation of this technique lies in the low delivery efficiency of the therapeutic DNA to the diseased site. To address this issue, various strategies have been explored including vectors engineered from viruses (Brun et al., 2008) and PLGA in NP formulations. PLGA NP have shown particular promise in improving the delivery efficacy (Kocbek et al., 2007). Besides, the physical characteristics of the nanoparticles can be manipulated to escape the degradative endosomal lumen, resulting in cytosolic localization. To develop novel administration paths, hybrid versions of research have been conducted on this subject, yet the results are mostly based on animal models or in vitro results, making it difficult to draw final conclusions. From clinical trials, substantial obstacles to their use, such as immunogenicity and inflammatory potential, have also been demonstrated (Nafee et al., 2007). Therefore, there is still a long way to go before real clinical applications come through.

encapsulation efficiency and high burst release of the encapsulated biomolecule within the first few hours or days, which is mainly due to the desorption of surface-associated hydrophilic molecules having weak interactions with the polymer (Fahmy et al., 2005). To circumvent these limitations and establish therapeutic efficacy, large doses or site-specific administration are often required for devices comprised of polyester biomaterials. In an attempt to address these problems, numerous groups have introduced functional groups (such as amine functionalities) on these materials, either through direct conjugation or

As for drug release from MPs or NPs, it is generally controlled by both drug diffusion and polymer degradation. To ensure the efficacy of drug delivery, control over the particle size and particle size distribution is critical, since smaller particles and narrower size distributions facilitate the design of targeted drug delivery systems. These involve binding fragments specific to a tumor-associated surface antigen, with a ligand binding to its corresponding receptor on the tumor cell surface, which can be attached on the surface of the PLA-based materials. Furthermore, polymers that display a physicochemical response to changes in their environment are being intensively explored as potential drug and gene delivery systems. The use of stimuli-responsive nanocarriers offers an attractive opportunity for targeted delivery, in which the delivery system becomes an active participant rather than a passive vehicle. The advantage of stimuli-responsive nanocarriers becomes obvious when the stimuli are unique to disease pathology, allowing the nanocarrier to respond specifically to the pathological characteristics. For instance, in solid tumors, the extracellular pH decreases significantly from 7.4 (the pH value under normal physiological conditions) to about 6.5 (Vaupel et al., 1989; Haag, 2004). In addition, the pH ranges from 4.5 in lysosomes to about 8.0 in mitochondria. Given these pH shifts, therapeutic compounds with a *p*Ka between 5.0 and 8.0 are able to exhibit dramatic changes in physicochemical properties (Ganta et al., 2008). Another option is thermo-sensitive polymeric micelles, containing a hydrophobic core and a thermosensitive shell, the later changing from a hydrophilic nature at body temperature to a collapsed hydrophobic polymer at a hyperthermic condition of 42°C (Na et al., 2006). Investigations concerned with this theme include responses induced by chemical substances, changes in temperature (Tyagi et al., 2004) or pH (Sethuraman & Bae, 2007), electric signals (Sawahata et al., 1990) or other environmental conditions (Qiu & Park,

The use of nucleic acids as therapeutic agents for genetic diseases has been extensively studied (Torchilin, 2008). However, a major limitation of this technique lies in the low delivery efficiency of the therapeutic DNA to the diseased site. To address this issue, various strategies have been explored including vectors engineered from viruses (Brun et al., 2008) and PLGA in NP formulations. PLGA NP have shown particular promise in improving the delivery efficacy (Kocbek et al., 2007). Besides, the physical characteristics of the nanoparticles can be manipulated to escape the degradative endosomal lumen, resulting in cytosolic localization. To develop novel administration paths, hybrid versions of research have been conducted on this subject, yet the results are mostly based on animal models or in vitro results, making it difficult to draw final conclusions. From clinical trials, substantial obstacles to their use, such as immunogenicity and inflammatory potential, have also been demonstrated (Nafee et al., 2007). Therefore, there is still a long way to go before real clinical

device fabrication with additives (Betram et al., 2009).

2001).

applications come through.

Some examples of delivery systems incorporating PLA are provided in Table 5 and in Fig. 5 (Chen et al., 2007; Sethuraman & Bae, 2007). Sethuraman et al. (Sethuraman & Bae, 2007) developed a novel drug delivery system for acidic tumors consisting of two components: 1) A polymeric micelle with a hydrophobic core of PLLA and a hydrophilic shell of PEG conjugated to TAT (a cell-penetrating peptide in HIV), and 2) a highly pH-sensitive copolymer of poly(methacryloyl sulfadimethoxine) (PSD) and PEG (PSD-b-PEG). The final carrier, which was able to shield the micelles and expose them at slightly more acidic tumor pH levels, was achieved by complexing PSD with the TAT of the micelles. The results obtained showed significantly higher uptake of TAT micelles at pH 6.6 in comparison with pH 7.4, and that TAT not only translocated into the cells but it was also traced to the surface of the nucleus [see Fig. 5].

Fig. 5. Test results for PLA-based drug delivery materials. Fluorescent microscopy images are shown on top for COS7 cells transfected by plasmid encoding enhanced green fluorescence protein (EGFP) with different carriers: (A) lipofectamine, (B) methoxypolyethyleneglycol-PLA-chitosan nanoparticles (MePEG-PLA-CS NP); the transfection efficiency, as detected by flow cytometry, is higher in (B) than in (A) (Reproduced with the permission from Chen, J. et al. (2007). Preparation, characterization and transfection efficiency of cationic PEGylated PLA nanoparticles as gene delivery systems, Journal of Biotechnology*,* Vol.130, No.2, pp.107. Copyright (2007) Elsevier) At the bottom are dual label confocal micrographs for MCF-7 cells incubated with TAT micelles: (a) Cells stained with fluorescein isothiocyanate (FITC) attached to TAT in the micelles; (b) the same nuclei as in (a) were stained with TOPRO-3; (c) superimposed images of (a) and (b); the yellow color shows the localization of TAT within the nuclei (Reproduced with the permission from Sethuraman, V. A. & Bae, Y. H. (2007). TAT peptide-based micelle system for potential active targeting of anti-cancer agents to acidic solid tumors. Journal of Controlled Release*,* Vol.118, No.2, pp.216. Copyright (2007) Elsevier).


Table 5. Investigations on PLA-based material as drug delivery systems. AP: peptide, CRKRLDRN; MPEG: methyl ether poly(ethylene glycol); PAE: poly(β-amino ester); PEI: polyethylenimine; cNGR: Cyclic Asn-Gly-Arg; DMAB: dimethyldidodecylammonium bromide.

In summary, some problems still remain to be tackled for this promising novela dministration method. A major issue is the presence of surfactants such as SDS or stabilizers such as PVA in the microparticles, necessary to achieve antigen binding and colloidal stability. Although present only at low concentrations, the acceptability of such components in human vaccines depends on the results of extensive and costly toxicological studies. Biodegradable polymers used for drug delivery to date have mostly been in the form of injectable microspheres or implant systems requiring complicated fabrication processes with organic solvents. In such systems, the organic solvents can denature components such as protein drugs being encapsulated. Besides, these delivery systems have relatively low transfection efficiencies in vitro as compared with reagents commercialized for cell transfection. The last problem concerns the lack of test results for these delivery systems using animal models or in clinical trials, which is of fundamental importance for real applications in biomedical therapy.

### **5.3 Other fields**

268 Biomedical Science, Engineering and Technology

Enhanced transport across the rat nasal mucosa

Increased uptake of the Lf-

Good control over blood glucose concentration

Induced seric and mucosal antibody production

Elicited strong CTL response

Complete drug release Zhang &

Good control over the release Venkatraman

Reduced inflammation of arthritis rabbit model

Initial burst followed by sustained release

Improved transfection

Presented high tumorspecific targeting ability

Effective silencing of the

nanoparticle internalization

Significantly improved transfection efficiencies

activity

gene in cells

Carrier for DNA Rapid and efficient

Table 5. Investigations on PLA-based material as drug delivery systems. AP: peptide, CRKRLDRN; MPEG: methyl ether poly(ethylene glycol); PAE: poly(β-amino ester); PEI: polyethylenimine; cNGR: Cyclic Asn-Gly-Arg; DMAB: dimethyldidodecylammonium

In summary, some problems still remain to be tackled for this promising novela dministration method. A major issue is the presence of surfactants such as SDS or stabilizers such as PVA in the microparticles, necessary to achieve antigen binding and colloidal stability. Although present only at low concentrations, the acceptability of such components in human vaccines depends on the results of extensive and costly toxicological studies. Biodegradable polymers used for drug delivery to date have mostly been in the form of injectable microspheres or implant systems requiring complicated fabrication processes with organic solvents. In such systems, the organic solvents can denature components such as protein drugs being encapsulated. Besides, these delivery systems have relatively low transfection efficiencies in vitro as compared with reagents commercialized for cell transfection. The last problem concerns the lack of test results for these delivery systems

Initial burst followed by an exponential decrease

NP by bEnd.3 cells

and cytokine release

Vila et al., 2005

Hu et al., 2009

Xiong et al., 2007

Aline et al., 2009

Liggins et al., 2004

et al., 2005

Feng, 2006

Ataman-Önal et al., 2006

et al., 2005

Wu et al., 2010

Liu et al., 2011

Fay et al., 2010

Patil & Panyam, 2009

Freitas et al., 2005

Chen et al., 2007

Jie

Material Application Results Reference

Carrier for tetanus

lactoferrin (Lf)

Carrier for oral

p24 proteins

Carrier for paclitaxel

Carrier for paclitaxel

Carrier for nimesulide

Gene delivery systems

Carrier for 5-FU and paclitaxel

Drug carrier for cancer therapy

luciferase siRNA

Loaded with plasmid DNA

Carrier for HIV p24 proteins

Carrier for 5-FU and paclitaxel

toxoid

insulin

PEG-PLA NP Conjugated with

PLA NP Carrier for HIV

PLA-PEG particles

b-PLA

PLA

PLA

NP

NP

bromide.

PLA-b-Pluronic-

Surfactant-free PLA NP

microspheres

PEO-PLA copolymers

PLA-TPGS copolymers

microspheres

PEGylated PLA

PLA-PEG-PLA copolymer

PLA/MPEG-PAE

cNGR-PEG-PLA

DMAB coated PLGA NP

PLGA/PEI NP Carrier for

AP-PEG-

Due to its versatility, PLA has been investigated for membrane applications (e.g. wound covers), implants and medical devices (fixation rods, plates, pins, screws, sutures, etc.), and dermatological treatments (e.g. facial lipoatrophy and scar rejuvenation).

With respect to wound treatment, bacterial infections are one of the main factors impacting the healing process. One of the best approaches to treat wound infections is by the immobilization of drugs or antibacterial agents within the nanofibers by electrospinning, or the electrospinning of polymers with intrinsic antibacterial and wound-healing properties. Dozens of patents have been issued on that topic so far (Ghosh et al., 2007; Robinson et al., 2009). Silver nanoparticles (nAg) and the natural polysaccharide chitosan (as well as its quaternized derivatives) are most commonly used as antibacterial agents with a high intrinsic activity against a broad spectrum of bacteria (Rujitanaroj et al., 2008; Ignatova et al., 2009).

Metals are still the most popular materials for fracture fixation, but their disadvantages include stress shielding, accumulation in tissues, hypersensitivity, growth restriction, pain, corrosion, and interference with imaging techniques. Consequently, the focus of research is increasingly shifted to biomaterials like PLA, which offers satisfactory strength during the healing of bone tissue and then degrades over time (Mavrogenis et al., 2009). A commercial product with a proven track record in clinical applications is the VICRYLTM suture material, based on PGA/PLA copolymers (Mehta et al., 2005). The number of applications of PLA as fixation rods, plates, pins, screws, sutures, etc. in orthopaedics and dentistry is also increasing (Raghoebar et al., 2006).

Fig. 6. Applications of PLA in the biomedical field.

In 2004, Sculptra™ [poly(*L*-lactic acid)] was approved by the FDA as the first injectable facial "volumizer" in the treatment of lipoatrophy due to its significant therapeutic effectiveness (Burgess & Quiroga, 2005). The lipodystrophy syndrome is associated with the usage of highly active antiretroviral therapy (HAART) containing protease inhibitors or nucleoside reverse transcriptase inhibitors for HIV patients. The action mechanism of Sculptra™ is via stimulation of the fibroblastic activity with generation of collagen and other connective tissue fibers. In addition, it acts as dermal matrix adding support by thickening the dermis (Vleggaar & Bauer, 2004). Moreover, PLA can help improve the appearance of scars due to acne, surgery, trauma, or suture (Lowe & Beer, 2005, as cited in Beer & Rendon, 2006).
