Design of Biomedical Polymers

  • Matthew ParrottEmail author
  • Stuart DunnEmail author
Living reference work entry
Part of the Polymers and Polymeric Composites: A Reference Series book series (POPOC)


The utilization of polymers for biomedical applications (“biomedical polymers”) has led to significant advancements in medicine. Biomedical polymers have made a profound impact on human health and improved the quality of life for many patients. Current and evolving biomedical challenges posed by disease, environmental triggers, and physiological processes demand the development of biomedical polymers with specific properties and function. To address these challenges, the design of biomedical polymers has become of paramount importance. Designing polymers with specific structures opens the door to tailored properties and function. In this chapter, we cover the design of biomedical polymers for a variety of applications. We show that key polymer structures and properties are crucial to desired functionality for a given application. The biomedical applications we cover include (1) drug delivery, (2) imaging and tracking biomedical polymers in vivo, (3) scaffolds for tissue engineering, (4) medical devices, (5) surgery and wound repair, and (6) biosensors. By looking at the polymer structure-property-function relationships provided herein, we hope that this will enable improved designs of biomedical polymers to realize enhanced performance and efficacy in transforming human health.





3,9-bis (ethylidene 2,4,8,10-tetraoxaspiro [5,5] undecane) and 1,6-hexanediol


Dynamic polyconjugates


Enhanced permeability and retention




Hyaluronic acid


Human mesenchymal stem cells




Half maximal inhibitory concentration




Matrix metalloproteinase 2








Poly(aspartic acid)


Poly(beta-amino ester)


Poly(butyl and amino vinyl ether)s


Phosphate buffered saline






Poly(ethylene glycol)




Poly(ethylene oxide)


Poly(glycolic acid)


Poly(glutamic acid)


Poly(lactic acid)




Poly(lactic-co-glycolic acid)






Particle Replication In Non-wetting Templates


Poly(sebacic anhydride)


Small interfering RNA


Ultrahigh molecular weight poly(ethylene)

1 Biomedical Polymers

Biomedical polymers embrace a wide range of different classes with unique structures and characteristics. Most notably, there are natural and synthetic biomedical polymers that may be utilized directly, modified, or synthesized from starting monomers. The molecular weight of polymers greatly impacts their physical properties. In a typical polymerization, as the polymer chains grow, the critical molecular weight for chain entanglements is reached. At this point, many of the physical properties become notable such as the glass transition temperature (Tg) and melting temperature (Tm). Tg refers to long-range segmental motion of the polymer chains, which are in a vitrified state above the Tg with limited mobility. Other polymers are able to pack into domains that afford a semicrystalline morphology with a melting temperature (Tm). The morphology of the polymer greatly affects the physical properties. For a given polymer, there may be dispersity in the distribution in the molecular weight of the chains, which is measured as the polydispersity index (PDI). A completely uniform molecular weight of a polymer has monodispersity (PDI = 1); however the majority of polymers are polydisperse (PDI >1). The PDI can influence polymer processing and the uniformity of properties. In this chapter, we highlight the physical properties of polymers and how these properties can be utilized for biomedical applications.

To be classified as a biomedical polymer, biocompatibility must be satisfied. That is, the biomedical polymer must be tolerated by the body and not elicit side effects, allergies, or toxicity. Biodegradable polymers break down into monomers and constituents that may be absorbed in the body. These breakdown products must be safe and biocompatible as well. Polymers and their components clear from the body via different means depending on size and molecular weight. For example, polymers smaller than 5 nm are rapidly excreted in the kidney, while 10–100 nm polymers can circulate systemically for hours [1]. The mechanism of absorption or clearance of the polymer may depend on the biomedical application. The target application must be matched to the function of the polymer. For instance, in gene therapy, polymers must be able to effectively deliver nucleic acids without cytotoxicity. In this book chapter, the goal is to provide a foundation of structure-property-function relationships within different classes of biomedical polymers and applications. As a result, we want this information to elucidate the design of specific polymers and provide a platform to improve contemporary biomedical polymers.

2 Biodegradable Polymers

The biologically induced breakdown of polymers into their monomers and constituent parts provides a means for absorption and clearance from the body in biomedical applications. In many applications, it is crucial that polymers break down into their monomers. For example, in drug delivery, the degradation of polymers may facilitate the release of drug. For implants, the degradation of biomedical polymers evades the necessity for extra surgery to remove implants. In tissue engineering, the breakdown products of polymers can facilitate natural tissue growth.

Polymers may be biologically degraded through hydrolytic, enzymatic, acidic, and reducing mechanisms. Here, we cover the different classes of biodegradable polymers followed by their physical properties. Important physical properties relevant to synthetic biodegradable polymers include their morphology, thermal transition temperatures, and modulus. Next, we discuss the polymer’s mechanism of degradation and rate of degradation. This information should serve as a toolbox to select a biodegradable polymer with specific physical and degradation properties for a particular biomedical application. The chemical structures of biodegradable polymers can be found in Fig. 1 for synthetic biodegradable polymers and Fig. 2 for natural biodegradable polymers.
Fig. 1

Chemical structures of synthetic biodegradable polymers

Fig. 2

Chemical structures of polysaccharides and polypeptides

2.1 Synthetic Biodegradable Polymers

  1. (a)


    Within the polyester family, notable materials include aliphatic poly(ester)s, poly(ortho ester)s, and poly(phosphoester)s. These polymers can embrace a wide range of morphologies from completely amorphous rubbery or glassy materials to semicrystalline rigid polymers. The main mechanisms of degradation are by hydrolysis and enzymatic cleavage. Enzymes involved in the hydrolysis of esters include esterases and lipases. The rate of degradation can vary on time scales from hours to years based on the chemical structure and composition. Furthermore, acidic environments increase the rate of hydrolysis.
    1. (i)

      Poly(lactic acid) (PLA), Poly(glycolic acid) (PGA), and Poly(lactic-co-glycolic acid) (PLGA)

      PLA is a glassy semicrystalline polymer with a Tg of 60–65 °C and Tm of 150–160 °C. The degradation times of PLA can last for months [2, 3].

      PGA is a glassy semicrystalline polymer with a Tg of 35–40 °C and Tm of 225–230 °C with degradation times on the order of months. By creating of copolymers of PLA and PGA, namely, poly(lactic-co-glycolic acid) (PLGA), an amorphous material is created with a Tg ranging from 40 to 60 °C. Depending on the molecular weight and ratio of glycolic acid to lactic acid units, this copolymer can provide degradation times on a much shorter time scale from hours to days [4, 5]. PLA, PGA, and PLGA are FDA-approved and offer high mechanical strength with moduli between 1 and 10 gigapascals (GPa). These polymers can be degraded hydrolytically and through enzymatic cleavage by esterases and lipases.

    2. (ii)

      Poly(caprolactone) (PCL)

      PCL has a long aliphatic unit that provides flexibility and semicrystallinity, resulting in a Tg of −60 °C and a Tm of 60 °C. PCL has a modulus between 0.1 and 1 GPa. Furthermore, the increased chain length of methylene groups provides a high degree of hydrophobicity [6]. Consequently, the hydrophobicity minimizes the ability of water to hydrolyze the ester bonds. PCL has notably slow degradation that is influenced by crystalline microstructure [7]. The degradation of PCL can last for months to years, which occurs through hydrolysis and hydrolase-type enzymes, i.e., lipases, esterases, and proteases.

    3. (iii)

      Poly(ortho ester)s

      Polymer surface erosion is another mechanism by which degradation occurs and can be noted for poly(ortho ester)s. In surface erosion, the exterior of the polymer surface degrades first before the inside bulk of the material. This may be exemplified most notably for poly(ortho ester)s, such as those based on 3,9-bis (ethylidene 2,4,8,10-tetraoxaspiro [5,5] undecane) and 1,6-hexanediol (DETOSU-HD), which has a Tg of 22 °C. The degradation of DETOSU-HD proceeds over years [8]. Tunable degradation kinetics from days to months can be realized based on the amount of comonomer incorporated. DETOSU-based copolymers can have a modulus between 0.5 and 1 GPa. With hydrolysis being the primary route of degradation in surface erosion, the backbone of poly(ortho ester)s may further be designed to degrade rapidly under acidic conditions.

    4. (iv)


      With a backbone similar to nucleic acids, poly(phosphoester)s are a versatile class of poly(ester)s that are amenable to designed functionality. Examples of poly(phosphoester)s include poly(phosphate)s, poly(phosphite)s, poly(phosphonate)s, and poly(phosphoramidate)s. These structures allow for chemical modification in the backbone or side chain emanating from the phosphorous group. Moreover, these groups attached to the phosphorous heavily influence the physical properties and rate of degradation. For example, poly(bisphenol A-ethyl phosphate) (BPA-EOP) has a Tg of 103 °C, breakdown temperature of 310 °C, elastic modulus of 508 MPa, and shear modulus of 183 MPa [9]. By incorporating a phenyl group through a phosphonate linkage, poly(bisphenol A-phenylphosphonate) (BPA-PP) has Tg of 115 °C, breakdown temperature of 539 °C, elastic modulus of 627 MPa, and shear modulus of 230 MPa. The in vivo degradation of these materials showed that greater than 80% weight loss of BPA-EOP was encountered after 70 weeks, while less than 20% weight loss was found for BPA-PP. Therefore, by controlling the side chain for poly(phosphoester)s, the rate of degradation could be tuned dramatically. Poly(phosphoester)s can be degraded under physiological conditions by hydrolysis as well as by enzymes such as phosphatases and phosphodiesterases.

  2. (b)


    Similar to poly(ester)s, the degradation rate and drug release properties of poly(anhydride)s can be controlled by chemical structure and composition. Poly(anhydride)s provide a surface erosion degradation mechanism through hydrolysis of the anhydride bond. Two common poly(anhydride)s are poly(sebacic anhydride) (PSA) and poly[bis(p-carboxyphenoxy)propane] (PCPP). PSA has a Tg and Tm of 62 °C and 79 °C, while PCPP has a Tg and Tm of 92 °C and 230 °C. PCPP has a flexural strength of 2.76 MPa and Young’s modulus of 8.84 MPa [10]. PCPP alone showed constant erosion degradation kinetics over several months. By creating copolymers of PCPP and PSA, the rate of degradation has been increased by increasing the amount of PSA. Furthermore, a simultaneous decrease in Tg and Tm is observed when increasing the amount of PSA in the copolymer. With PCPP-SA (21:79) copolymer, complete degradation was reached in less than 2 weeks.


2.2 Natural Biodegradable Polymers

  1. (a)


    Composed of sugar monomers, polysaccharides are mostly water-soluble polymers that are inherent in nature, biocompatible, and biodegradable. Each polysaccharide has a unique structure and particular functional groups. They are mainly degraded by enzymes specific to the sugar and polymer. For example, dextran may be degraded by dextranases such as dextranhydrolase, glucodextranase, and dextran glucosidase, in addition to others. Their natural origin and biodegradability make polysaccharides appealing for biomedical applications. Here, we list common polysaccharides found in biomedical applications.
    1. (i)


      Dextran consists of alpha-1,6 glycosidic linkages between glucose units. Branches are often found in dextran and occur via alpha-1,3 linkages. Dextran is a water-soluble, neutral polysaccharide. Dextran can easily be modified and rendered insoluble for therapeutic applications. For example, dextran has been derivatized with acetal groups by modification of the hydroxyl groups with 2-methoxypropene [11]. This derivatization produces a pH-sensitive acetalated dextran wherein acetal groups hydrolyze under mild acid conditions. For vaccine application, the protein ovalbumin was encapsulated in nanoparticles formed by a double emulsion. Ovalbumin particles increased major histocompatibility complex class I presentation when compared to free protein. Given the promise of these particles for vaccines against tumors and viruses, the acetalated dextran may be useful in sutures and scaffolds.

    2. (ii)


      Alginate is a carboxylic acid-containing, negatively charged polysaccharide composed of mannuronate and guluronate units linked by glycosidic bonds. The carboxylic acid moiety allows for facile and specific chemical modification. Furthermore, alginate may be chemically crosslinked by calcium to form gels, which find great biomedical use. For example, alginate has been crosslinked ionically and photochemically to form macroscopic gels incorporating small interfering RNA and cells for gene knockdown [12]. Alginate was ionically crosslinked using calcium and photochemically crosslinked using an initiator combined with UV irradiation. The photocrosslinked alginate provided controlled degradation through hydrolysis of ester linkages, sustained release of small interfering RNA, and gene knockdown.

    3. (iii)


      Hyaluronan is a negatively charged polysaccharide containing carboxylic acid and amide groups. Specifically, hyaluronan is composed of N-acetylglucosamine and glucuronic acid units. It is crucial in tumor development and can be harnessed by nanomaterials as a targeting ligand [13]. For example, lipid nanoparticles have been functionalized with hyaluronan to target epithelial cancer cells that overexpress hyaluronan receptors (CD44 and CD168). The effect of hyaluronan molecular weight on biological responses has been evaluated using nanoparticles [14]. Low molecular weight hyaluronan-decorated nanoparticles have low binding to CD44 receptor, while high molecular weight shows high binding and affinity for CD44 receptor. Furthermore, low molecular weight hyaluronan may serve as a substitute for poly(ethylene glycol) in stealthing nanoparticles for passive delivery.

    4. (iv)

      Chitin and Chitosan

      Chitin is mostly composed of N-acetylglucosamine units connected by beta-1,4 linkages. Deacetylated glucosamine units are found, and their percentage depends on the source for production. Chitin is insoluble in water due to extensive hydrogen bonding. Chitosan exists largely as the deacetylated derivative of chitin with some acetylated glucosamine units present. Chitosan can be made soluble under acidic conditions and has many advantageous properties for biomedical applications [15]. Specifically, it is renewable, nontoxic, and hydrolyzable by lysozyme. In addition, chitosan has film-forming, hydrating, and wound healing properties wherein it can be used for surgical sutures, dental implants, artificial skin, and rebuilding of the bone. The primary amine on chitosan allows for facile functionalization with chemical moieties that enhance interaction with cells or delivery of hydrophobic drugs. For instance, N-trimethyl chitosan can provide interaction with negatively charged cell membranes, and n-lauryl-carboxymethylchitosan can form micelles that encapsulate paclitaxel [15].

    5. (v)

      Chondroitin Sulfate

      Chondroitin sulfate is a negatively charged polysaccharide containing sulfate, carboxylic acid, and amide functional groups. It is composed of glucuronic acid and N-acetylgalactosamine units and can form proteoglycan aggregates through binding to protein. The aggregate can interact with tissue via electrostatics. Chondroitin sulfate has anti-inflammatory properties, prevents production of cartilage cytokines, and elicits apoptosis of articular chondrocytes [16]. The mineralization process and repair in the bone has been accelerated by using chondroitin sulfate. This anionic polysaccharide has been used to treat osteoarthritis [17]. Chondroitin sulfate was found to be more effective than glucosamine in reducing knee pain.

    6. (vi)


      Fucoidan is an anionic polysaccharide containing sulfate groups and fucose units. Fucoidan encompasses a family of fucoidans that can differ in chemical composition. For example, some fucoidans are composed of monosaccharides like glucose or galactose, acetyl groups, or proteins. Fucoidan has a number of desirable pharmacological characteristics such as anti-inflammatory, anticoagulant, antithrombotic, and anti-oxidative properties [18]. Fucoidan has been used for drug delivery by crosslinking with chitosan to form a “fucosphere” [19]. Bovine serum albumin was encapsulated in fucospheres, and its release was controlled by varying concentration of polymers, albumin, and preparation method. In addition the cytokine, granulocyte-macrophage colony-stimulating factor (GM-CSF), has been encapsulated in fucospheres [20]. Fucospheres demonstrated slow and sustained release from 90 to 140 days with retention of released cargo. Neutropenia and aplastic anemia can be treated with delivery of GM-CSF.

  2. (b)


    Polypeptides are polymers composed of amino acid monomer units connected via amide bonds in the backbone. They are natural, water-soluble polymers with a variety of functional groups and properties. Notable polypeptides in biomedical applications contain functional groups such as carboxylic acids and amines. Polypeptides are degraded by proteases that cleave the amide backbone of the polymer.
    1. (i)

      Poly(l-lysine) (PLL)

      PLL is a cationic polypeptide composed of lysine units, which bear a free primary amine. This primary amine allows for facile functionalization chemistry. PLL exists in two primary forms, most commonly as α-PLL where the long aliphatic amine group exists as the side chain. In ε-PLL, the long aliphatic amine resides in the backbone of the polymer. The inherent cationic nature of PLL enables strong complexation with polyanions. This property can find great use in complexing with nucleic acids for gene delivery or silencing. For example, PLL has been complexed with DNA and exhibited effective transfection [21].

    2. (ii)

      Poly(aspartic acid) (PAsp)

      PAsp is an anionic polypeptide composed of aspartic acid residues. It exists in two main polymeric forms: alpha and beta. In α-PAsp, there is a methylene group between the carboxylic acid and the backbone, while for β-PAsp, this methylene group is found in the backbone. The degradability, biocompatibility, low cytotoxicity, and functional carboxylic acid group of PAsp make it a good candidate for biomedical applications. For instance, PAsp-based supramolecular assemblies have been prepared for gene delivery [22]. Benzene-functionalized PAsp was combined with cyclodextrin-core PAsp polycations through host-guest interactions to form assemblies that were then mixed with DNA to form complexes that could effectively transfect cells.

    3. (iii)

      Poly(glutamic acid) (PGlu)

      PGlu is another anionic polypeptide that is composed of glutamic acid units. It mainly exists in two forms: alpha and gamma. In the alpha form, PGlu has an ethylene group between the carboxylic acid and the backbone, while in the gamma form, this ethylene group is in the backbone; therefore, there is no spacer between the backbone and carboxylic acid group. The anionic polyglutamic acid can be tethered to a functional group and can bind electrostatically to cationic nanoparticles. For example, PGlu-based ligands have been used as nanoparticle coatings to enable ligand-specific gene delivery in vitro [23] and alter biodistribution and gene delivery efficacy in vivo [24].

  3. (c)


    Proteins are large assemblies of polypeptides. They may consist of a variety of polypeptides connected by linkages such as disulfide bonds. Proteins can adopt unique tertiary structures such as beta sheets and alpha helices. Proteins serve many important roles biochemically that facilitate physiological processes. Similar to polypeptides, proteins are degraded enzymatically by proteases.
    1. (i)


      Albumin is a water-soluble protein that represents a group of proteins that make up the largest fraction of proteins in the blood. Human and bovine serum albumin are the most common albumin proteins. Human serum albumin is a non-glycosylated, anionic protein with enzymatic and antioxidant properties. It contains numerous disulfide bridges and at least one free thiol. Human serum albumin adopts an ellipsoidal structure and is largely composed of alpha helices. Bovine serum albumin (BSA) PRINT® particles encapsulating self-replicating RNA have been prepared toward vaccine use. To avoid dissolution, BSA particles were rendered transiently insoluble by using a reducible disulfide crosslinker [25]. These particles were complexed with cationic and transfecting lipid and effectively delivered RNA replication to Vero cells [25].

    2. (ii)


      Gelatin is a derivative of collagen that has been partially and irreversibly hydrolyzed. Collagen serves as the major protein in connective tissue while being the most abundant protein in mammals. Collagen contains a large number of glycine residues, followed by proline. The linkages between collagen strands are broken in gelatin, which can assume a variety of supramolecular structures ranging from globular to fibrillar with triple-stranded helices. Acid-solubilized collagen is a liquid at 4 °C and a hydrogel at 37 °C and neutral pH. This thermogelling property enables the encapsulation of cargos for biomedical applications. For example, small interfering RNA and cells have been encapsulated in macroscopic collagen hydrogels for gene silencing [12]. This may serve as an injectable biodegradable polymer for sustained inhibition of gene expression.


3 Nondegradable Biocompatible Polymers

A polymer is considered biocompatible when it does not pose side effects, allergic reactions, and toxicity while allowing the body to function normally. This opens the area of broad classes of polymers to biomedical applications. Here we will cover different nondegradable, biocompatible polymers that have found use in biomedical applications. Then, their physicochemical, thermal, and mechanical properties may be covered. The intent in this section is to provide a repertoire of biocompatible polymers that may be used in biomedical applications. The chemical structures of these biocompatible polymers are illustrated in Fig. 3.
  1. (a)

    Poly(ethylene glycol) (PEG)

    PEG is a water-soluble, neutral polyether composed of ethylene oxide repeat units. PEG also has broad solubility in organic solvents. PEG can be identified as polyoxyethylene or poly(ethylene oxide) (PEO). PEG is generally used for polymers with a molecular weight less than 20 kDa, PEO for polymers greater than 20 kDa, and polyoxyethylene is used for any molecular weight. It has a Tg of −66 °C and Tm of 66 °C [26]. PEG is utilized in a wide range of molecular weights down to oligomers. For general biomedical applications, the molecular weight of PEG is commonly between 2 and 10 kDa. PEG has notably been used on the surfaces of particles to provide a hydrophilic stealth layer that minimizes interactions with serum proteins and the mononuclear phagocytic system [27]. For instance, it was found that high density of PEG in brush conformation provided the longest circulation half-life, lowest protein binding, and lowest macrophage association when compared to low-density PEG and unPEGylated nanoparticles [28].

  2. (b)


    Poloxamers (or commercially “Pluronics®”) are ABA triblock copolymers of PEO-poly(propylene oxide)-PEO. A wide range of Pluronics are commercial with different molecular weights and percentage of PEO [29]. These copolymers have a distinct critical micelle concentration and hydrophilic-lipophilic balance. Examples of two Pluronics are F127 and L61. Pluronic F127 has 200 ethylene oxide units, 65 propylene oxide units, and a molecular weight of 12,600 Da. Pluronic L61 has 4.6 ethylene oxide units, 31 propylene oxide units, and a molecular weight of 2 kDa. The water solubility of Pluronics decreases as the relative fraction of propylene oxide units increases. Given the large amount and percentage of ethylene oxide units in Pluronic F127, it has water solubility greater than 10%. Conversely, the greater relative number of propylene oxide units in Pluronic L61 renders it insoluble water. Poloxamers exhibit a reversible solgel transition when going from low temperature to high temperature. To improve stability of the hydrogel, poloxamer 407 was derivatized with thiol and acrylate groups [30]. The solgel transition was achieved at body temperature, and the crosslinking was realized under physiological-mimicking conditions. The resulting stability was enhanced, and release of drug was prolonged four times relative to the normal hydrogel. This study demonstrated the potential of poloxamer hydrogels for controlled drug release, cell encapsulation, and tissue engineering.

  3. (c)

    Poly(2-oxazoline)s (POx)

    As PEG is utilized in several biomedical formulations, a functional analogue class of polymers to poly(ethylene glycol) has been pursued: poly(2-oxazoline)s (POx). POx are a class of polymers that have a polypeptide-like structure and tunable water solubility. POx consists of an ethyleneimine backbone where the nitrogen atom is substituted as an amide with a variety of groups and solubilities. For example, poly(2-ethyl-2-oxazoline) is water soluble, while poly(2-nonyl-2-oxazoline) is water insoluble. POx fit into a variety of biomedical applications ranging from drug and gene delivery to membranes and stimuli-responsive materials [30, 31]. POx have been conjugated to biomolecules such as bovine serum albumin and insulin to tailor properties [32]. For insulin POx conjugates, blood glucose levels were lowered for four times longer than insulin alone. Similar to PEG, albumin POx conjugates mitigated immunogenic properties.

  4. (d)

    Poly(N-(2-hydroxypropyl)methacrylamide) (PHPMA)

    PHPMA is a water-soluble, neutral polymer that is nontoxic and nonimmunogenic. PHPMA has an alcohol-containing side chain that allows for functionalization in biomedical applications. PHPMA is commonly used as a drug carrier by conjugating drugs through peptide linkers that can be cleaved by intracellular enzymes. For instance, the protein synthesis inhibitor puromycin and the antibiotic daunomycin have been bound to PHPMA copolymer via different peptide linkers [33]. When incubated with lysosomal cathepsin proteases, the drugs were released over 20 h. The conjugates exhibited differential toxicity in a mouse leukemia model depending on the identity of the peptide linker.

  5. (e)

    Poly(ethyleneimine) (PEI)

    PEI is a water-soluble cationic polymer with an ethyleneimine backbone, which provides a secondary amine. PEI’s Tg is −28 °C and Tm is between 40 and 47 °C [34]. One of the most prominent uses of PEI in biomedical applications involves non-viral gene delivery. The amino groups and cationic nature of PEI facilitate polyplex formation with DNA through electrostatics. The molecular weight and structure of PEI have a notable effect on the physicochemical properties and activity of the polyplex, specifically, condensation, size, charge, biodistribution, and transfection efficiency. High molecular weight PEI has shown greater DNA condensation and cytotoxicity, while low molecular weight has exhibited higher transfection efficiency [35].

  6. (f)

    Poly(ethylene) (PE)

    PE is a simple poly(olefin) that is composed of two methylene groups in the backbone per repeat unit. PE molecular weights typically range from 1.4 kDa to 3.5 MDa for solid material. Waxy solids are found for lower molecular weights with degree of polymerization between 8 and 100. At degrees of polymerization less than 8, the alkanes exist as liquids or gases. Ultrahigh molecular weight (UHMW) PE is most commonly used in biomedical applications such as a bearing surface in joint replacements. UHMW PE has a molecular weight range from 2 to 6 Mg/mol, a melting point of 125–135 °C, and a tensile modulus of 0.8–1.5 GPa [36]. UHMW PE joint replacement devices have lifetimes greater than 15 years if implanted correctly, and current research into the design and properties of the material may further improve performance.

  7. (g)

    Poly(styrene) (PS)

    PS is a glassy, hydrophobic thermoplastic polymer with a Tg of 100 °C [26]. It contains an aromatic phenyl group side chain and may be produced in several forms. Polystyrene has an ultimate tensile strength of 40 MPa, elongation of 7%, and tensile modulus of 3 GPa. PS has found use in biomedical applications for the fabrication of shape- and size-specific particles for biological studies. Specifically, the phagocytosis of PS particles by alveolar macrophages has been investigated [37, 38]. The shape of PS particles was found to play a dominant role in phagocytosis initiation, while size was found to dictate completion of phagocytosis depending on volume compared to the size of the cell. Wormlike PS particles, with aspect ratio greater than 20, showed minimal phagocytosis compared to equal volume spherical particles.

  8. (h)

    Poly(methyl methacrylate) (PMMA)

    PMMA is glassy, hydrophobic thermoplastic polymer with a Tg of 115 °C for its syndiotactic form [26]. PMMA has a Young’s modulus ranging from 1.8 to 3.1 GPa. The tensile strength of PMMA ranges from 48 to 76 MPa. PMMA may be used in amphiphilic block copolymers to form nanoparticles for biomedical applications. For example, PMMA-core/PEG-shell block copolymer nanoparticles have been prepared and functionalized with a ligand that can chelate radioactive copper-64 to study structural effects on biodistribution [39]. In another study, the effect of plasma protein nanoparticle coating on biodistribution and phagocytosis by the reticuloendothelial system was studied for radiolabeled PMMA nanoparticles [40].

  9. (i)

    Poly(acrylic acid) (PAA)

    PAA is a glassy, water-soluble polymer with a Tg of 105 °C [26]. PAA is an anionic polymer bearing a carboxylic acid residue directly off the vinyl backbone. This acid group allows for facile functionalization and also provides additional properties. For example, the acid group allows PAA to swell several times its volume and finds use as a super-absorber. For functionalization of PAA, the carboxylic acid groups have been modified with PEG to serve as stealth coating. Specifically, cationic hydrogel nanoparticles were complexed with PAA-PEG through electrostatics to neutralize surface charge and alter biodistribution in vivo, facilitating accumulation in hepatocytes [41]. Furthermore, PAA-coated nanoparticles encapsulating small interfering RNA enabled gene knockdown in vivo.

  10. (j)

    Poly(vinyl ether)s

    Poly(vinyl ether)s consist of a vinyl backbone with an ether side chain. A variety of chemical groups may be selected as the side chain, which influences properties. For example, poly(methyl vinyl ether) has a Tg of −31 °C and Tm of 144 °C, while poly(butyl vinyl ether) has a Tg of −55 °C and Tm of 64 °C [26]. For biomedical applications, additional groups such as primary amines may be chosen as the side chain functionality. One particularly interesting example is that of an amphipathic poly(vinyl ether) containing butyl and amino side chains, which is termed PBAVE. Due to its amphipathic nature, PBAVE has membrane lytic activity and was able to transfect DNA efficiently in vitro [42]. When PBAVE was conjugated with small interfering RNA and targeting ligands, in vivo gene knockdown was achieved [43].

Fig. 3

Chemical structure of nondegradable, biocompatible polymers

4 Polymer Architectures and Structures

Polymers may adopt a wide variety of architectures and structures. Here we will cover the different polymer architectures and structures along with some of their properties. The architecture of a polymer chain refers to the branching from a linear chain, hence the term branched polymer. Dendrimers represent a particular treelike architecture. Crosslinked polymeric networks are chemically, physically, or ionically connected completely to other polymer chains. Polymer complexes are formed form complementary polyelectrolytes that self-assemble into unique structures. Micelles are self-assembled particles from amphiphilic block copolymers. Nanoparticles are a broad class of self-assembled or top-down fabricated polymeric particles with sizes between 1 and 100 nm. Covering a few of the different polymer architectures and structures will supply a foundation for understanding the properties of biomedical polymers. Figure 4 shows representative cartoons of linear polymers, block copolymers, polymeric networks, and branched polymers.
Fig. 4

Illustration of linear polymers, block copolymers, polymeric networks, and branched polymers

  1. (a)


    Linear polymers consist of linear chains of monomeric units. They may adopt conformations ranging from random coils to globules. Linear polymers can have different tacticities, that is, the order of adjacent chiral center’s stereochemistry. These include isotactic (all substituent on the same side), syndiotactic (substituent alternate on the chain), and atactic (substituents are random along the chain). These conformations can influence the packing of chains and corresponding physical properties. Nonetheless, the properties of linear polymers depend largely on their molecular weight, which influences the chain entanglement and packing.

  2. (b)


    Branched polymers contain side chains emanating from the backbone based on the original monomer. An example of branched PEI shows the architecture in Fig. 5. The degree and type of branching influence polymer properties such as rheology, melting, and degradation in addition to application efficacy [44]. For example, the intrinsic viscosity of a branched polymer increases with molecular weight less than that for a linear polymer. Furthermore, the thermal transition temperatures (Tg, Tm) of poly(ethylene) are influenced by branching. Linear high-density poly(ethylene) has solid crystallization with a Tm of 135 °C [44]. The addition of a few branched units (known as low-density poly(ethylene)) depresses the melting temperature to 115 °C [44]. Introducing many branches in poly(ethylene) renders the polymer amorphous without a Tm. Additionally, for poly(ester)s, the rate hydrolysis is slower for branched polymers versus linear polymers.
    Fig. 5

    Chemical structure of branched PEI

  3. (c)

    Polymer Networks

    Polymeric networks are a special type of polymer structure in which all chains are connected via crosslinks. The crosslinks may take form through chemical linkages, physical domains (such as glassy nodules in rubbery matrices), or ionic bonds (using complementary ions to connect polyelectrolytes). For chemical means, this may be accomplished by using monomers and crosslinkers with appropriate functionality. For example, PEG diacrylate can be polymerized to yield a crosslinked network (Fig. 6). Similarly, gelatin can be functionalized with acrylamide groups that can be crosslinked through free radical polymerization. For ionic crosslinking, the carboxylic acid groups of alginate may be crosslinked using calcium(2+).
    Fig. 6

    Chemical structures of crosslinked PEG and gelatin

  4. (d)


    Dendrimers constitute a unique synthetic three-dimensional class of polymers with monodispersity, nanometer sizes, and treelike structures. The structures are composed of a core from which branched units emanate. Each subsequent branching point is referred to as generation X (number). The degree of branching, or number of generations, contributes to the size, molecular weight, and number of surface groups. Dendrimers terminated with reactive end groups allow for further functionalization with different moieties such as other polymers (e.g., PEG). There are two main synthetic approaches for dendrimers: (1) divergent growth, development of a dendrimer originates from a core, and (2) convergent growth – the dendrimer surface is reacted together inward to a focal core. The flexible synthetic approaches to dendrimers allow access to tunable cores, branches, and surface groups with defined physical properties.

    One of the most well-known dendrimers is poly(amidoamine) (PAMAM), which is illustrated in Fig. 7. PAMAM dendrimers are synthesized largely from ethylenediamine and methyl acrylate where the surface end groups can be controlled based on the final reaction.
    Fig. 7

    Chemical structure of PAMAM generation 1

  5. (e)


    Polymeric micelles are generally composed of amphiphilic block copolymers that self-assemble into core-shell structures where the core is hydrophobic and the shell is hydrophilic. The size and physical properties of the polymeric micelles can be tuned by the identity and molecular weight of the blocks of the amphiphilic copolymer. One polymer class that has already been covered, Pluronics, readily forms micelle architectures. In aqueous solution, the hydrophobic segments of the block copolymer self-assemble into a core with the hydrophilic blocks on the exterior. These micelles can find use as surfactants in a variety of technologies such as drug delivery, lubricants, detergents, and soaps.

  6. (f)


    Polymeric nanoparticles can be formed through a variety of processes such as nanoemulsification, complexation, nanoprecipitation, stretching, and Particle Replication In Non-wetting Templates (PRINT®). Through these processes, nanoparticles are obtained with sizes based on the material and methodology employed.
    1. (i)


      Nanoemulsions consist of a lipid oil core stabilized by surfactants to yield stable dispersions. To prepare nanoemulsions, conventional approaches utilize high-shear or high-pressure homogenization and ultrasonication. Exemplar surfactants include poly(ethylene glycol) with aliphatic tails and triglycerides. For instance, a formulation was developed from d-alpha-tocopheryl polyethylene glycol 1000 succinate (TPGS), polyoxyethylene 20-stearyl ether (Brij 78), and miglyol 812 to yield 200 nm “BTM” nanoparticles [45, 46]. A microwave synthesis approach was employed for the BTM composition to obtain sub-30 nm nanoparticle sizes by accessing the phase inversion temperature [47]. Similar to nanoemulsions, solid lipid nanoparticles have been extensively studied [48] and consist of a solid lipid core instead of an oil core. These nanoparticles provide physical and chemical stability of the dispersion.

    2. (ii)


      When mixing oppositely charged polyelectrolytes together such as a polycation with a polyanion, nanoparticle formation may take place. An example of this type of nanoparticle would be observed from the combination of a negatively charged nucleic acid like DNA with a positively charged amine-containing polymer such as PEI [49].

    3. (iii)


      Hydrophobic polymeric nanoparticles can be synthesized using nanoprecipitation. In this approach, the polymer is dissolved in a water-miscible organic solvent, which is dispensed generally dropwise into an aqueous solution containing surfactant. By rapid solvent diffusion, nanoparticles are formed and can reach sub-100 nm sizes. Notable examples involve the synthesis of PLGA nanoparticles stabilized by poly(vinyl alcohol) or Pluronics [50, 51].

    4. (iv)


      A unique approach was taken to fabricate elliptical disk PLGA nanoparticles starting from spherical particles [52]. Initial spherical particles were embedded into a polymer film that was stretched to yield non-spherical particles. These particles were capable of shape-switching modulated by temperature, pH, or chemical means from minutes to days.

    5. (v)


      Particle Replication In Non-wetting Templates (PRINT®) technology is a specialized particle molding process spun-off from soft lithography to fabricate nanoparticles with control over size, shape, composition, and surface properties [53, 54]. Perfluorinated elastomeric molds with shape- and size-specific cavities are filled with liquid monomer precursors, polymers, or a variety of materials through capillary or melt-fill processes. After curing or solidification, nanoparticles are extracted from the cavities generally using an adhesive polymer, which then may be dissolved to harvest the particles. PRINT technology has been utilized to synthesize biocompatible hydrogel PEG-based particles to biodegradable PLGA particles with a variety of shapes and sizes such as cylinders, cubes, and hex nuts [55, 56].


5 Biomedical Applications

After covering the different types and identities of polymers as well as their corresponding physicochemical properties and structures, it is now possible to delve into their applications. We will present the use of biomedical polymers in the following applications: (1) drug delivery, (2) imaging and tracking biomedical polymers in vivo, (3) scaffolds for tissue engineering, (4) medical devices, (5) surgery and wound repair, and (6) biosensors. Each application and sub-focus therein embraces key design criteria of biomedical polymers to realize desirable performance. Here, we review the literature on efficacious approaches and polymer systems that provide target behavior for given applications. Further, we highlight the crucial features of polymers that achieved success in their respective application. This may serve as a reference to understand the design of polymers for a variety of biomedical applications. In turn, we hope that this will lead to further iterations, designs, and development of polymeric systems for a range of applications with improved performance.

5.1 Drug Delivery

In the battle against human disease, drug delivery has come to the forefront of intensive research. Treatment of diseases using therapeutic interventions can be accomplished by delivering drugs that can remedy malfunctioning and restore normal physiology. Delivering treatments can be enhanced with the use of polymeric materials. Herein, we will cover the literature of biomedical polymers used for drug delivery, which will include different polymer architectures and cargos such as chemotherapeutics and nucleic acids. This literature may serve as a repertoire of biomedical polymers for researchers to select or adopt for their particular disease treatment goal.

5.1.1 Mechanism of Drug Release

The therapeutic drug may be released from the biomedical polymer through chemical and/or diffusive processes. In chemical processes, the drug may be bound to the polymer through a covalent linkage as a conjugate. There are a variety of linkers used for polymer-drug conjugates with specific degradation mechanisms. In diffusive processes, the drug may be encapsulated within a nanoparticle or polymer matrix and diffuse out. The diffusion of drug from the polymer can depend on the pore size, degree of complexation, crystallinity, or hydrophobicity.

Conjugation of therapeutic drugs to polymers can enhance their efficacy in vivo. For example, polymer conjugates can provide prolonged circulation in the blood, increased tissue accumulation and cell uptake, improved stability, and triggered release in specific biological locations. Polymer-drug conjugation is notably different than other self-assembled and top-down systems in that polymer-drug conjugates use a covalent linkage for drug delivery. Hydrophilic polymers, such as PEG, are often conjugated to drugs to enhance their solubility and provide in vivo stealth-like behavior.

Between the polymer and drug, the linker structure is of crucial importance as this determines how and when the drug is released. Extensive efforts into designing linkers for polymer-drug conjugates have enabled triggered release in specific tissues and cellular compartments for effective drug delivery. Hydrolyzable and acid-sensitive linkers include esters, hydrazones, beta-thiopropionate, and maleamate bonds. Reductively cleavable linkers include disulfides, while enzymatically cleavable linkers involve the use of peptide sequences. Examples of these linkers are provided with literature references under the different architectures that will be covered.

For diffusion of the drug from a polymeric material, there are several factors that govern the rate of release. For example, with a crosslinked nanoparticle, the pore size and crosslink density influence the rate of release: higher crosslink density corresponds with a slower rate of drug release. For micelles, the degree of hydrophobicity will dictate the noncovalent interaction strength and tendency of the drug to stay in the core of the micelle. In polyelectrolyte complexes, the binding between polyanion and polycation will influence the rate of release based on the ionic strength required to break the ionic interactions. For solid polymers, the rate of dissolution may be influenced by the crystallinity of the material and may determine the diffusion of drug from the polymer matrix. Examples of these drug diffusion systems will be covered for the different polymeric architectures.

5.1.2 Therapeutic Cargos Chemotherapeutics

Cancer is a devastating health problem that is expected to be the leading cause of death over the next few years. In 2015, the United States is expected to have 549,430 cancer deaths, which accounts for nearly 1 of every 4 deaths [57]. Direct medical costs for cancer in the United States were estimated at $88.7 billion. Given the huge economic and healthcare burden created by cancer, there is a critical need for effective cancer treatments. Chemotherapeutic drugs are effective in eradicating cancer cells; however, they inflict damage to healthy cells as well [58]. Biomedical polymers have enabled improved drug delivery to tumor tissues [54, 59, 60, 61]. DNA

DNA used for gene therapy can adopt multiple conformations and has a size ranging from 1 to 200 kb pairs. After entry into the cytoplasm, DNA must cross the nuclear pore membrane to reach the nucleus for its expression. For gene therapy, DNA must place the important genes at particular locations in the chromosome for the production of target proteins. For efficient delivery to target cells and the nucleus, the identity and structure of biomedical polymer surrogates make a tremendous impact on translation for expression of desired proteins. For example, amine-containing cationic polymers are often utilized to provide electrostatic complexation with nucleic acids, cellular internalization, and endosomal escape into the cytoplasm. Small Interfering RNA (siRNA)

siRNAs are short (20–25 nucleotides), double-stranded RNA that serve an important role in RNA interference. The antisense strand of siRNA targets the complementary strand of a certain gene. After entry into the cytoplasm, siRNA may be recognized and incorporated into the RNA-induced silencing complex where siRNA is unwound and binds the complementary messenger RNA sequence. This binding results in the cleavage of the duplex, which prevents the expression of the target protein. To efficiently deliver siRNA into the cell and cytoplasm using biomedical polymers, the structure and physicochemical properties play a huge role in achieving effective gene silencing. Similar to the delivery of DNA, amine-containing cationic polymers are often utilized in transfection with siRNA.

5.1.3 Polymer Architectures for Drug Delivery

To deliver a drug to a particular tissue and cell, certain polymer architectures may provide distinct advantages. A variety of polymer architectures have been developed with specific physicochemical properties that enable delivery of select drugs. The sizes observed with the different architectures range from sub-5 nm to greater than 100 nm. Chemical compositions of the biomedical polymers vary from homopolymers to block copolymers and may include hydrophobic, hydrophilic, ionic, and targeting moieties. Here we cover the different architectures found for biomedical polymers used in drug delivery. These architectures include polymer conjugates, dendrimers, micelles, and nanoparticles. Polymeric Conjugates
Delivery of chemotherapeutic drugs has been realized using β-cyclodextrin (CD)-based polymers. β-CD linear polymers were synthesized and conjugated with camptothecin. When incorporated into the polymer, camptothecin had more than three orders of magnitude greater aqueous solubility [62]. Camptothecin conjugates showed cell-line- and structure-dependent half maximal inhibitory concentrations (IC50s) in vitro where the parent polymer had low cytotoxicity. The different structures investigated included different peptide linkers (Fig. 8). The rate of drug release from the conjugates was significantly faster in plasma compared to phosphate buffered saline (PBS), which may be attributed to cleavage by enzymes found in the bloodstream. The peptide linker and molecular weight of the CD conjugates were then evaluated in vivo [63]. High molar mass (97 kDa) polymer conjugates showed greater antitumor efficacy than the native drug. Conjugates with the shorter peptide linker (glycine) were found to be less toxic than the longer peptide linker (triglycine). The CD conjugates afforded long-term control over tumor growth, which may be attributed to sustained release of the drug in acidic, intracellular compartments.
Fig. 8

Chemical structure of cyclodextrin polymer-drug conjugates

A hydrophilic polymer that has been used in polymeric drug conjugates derives from N-(2-hydroxypropyl)methacrylamide (HPMA). In one study, a conjugate between poly(HPMA) and an angiogenesis inhibitor (TNP-470) was established using a peptide linker [64]. This peptide linker (Gly-Phe-Leu-Gly) was found to be enzymatically degradable and demonstrated selective release from the polymer by the lysosomal enzyme cathepsin B. The conjugate accumulated selectively in tumor tissue, minimized accumulation in normal organs, and enhanced the activity of the drug in vivo.

siRNA polymer-drug conjugates have been pursued by many groups. siRNA is a prime candidate for conjugation due to its moderate size, negative charge, and high susceptibility to degradation by RNAses. siRNA conjugates have been synthesized with acid and reductively labile linkers. For example, siRNA was conjugated to PEG via acid-labile beta-thiopropionate bond [65] (Fig. 9). The PEG was end-functionalized with lactose to target hepatoma cells. When this conjugate was combined with PLL, targeted polyion complexes were formed, which were capable of significant gene silencing in hepatocarcinoma cells. Another example of an siRNA conjugate with a reductively labile linker involves siRNA conjugated to PEG via a disulfide bond [66]. This siRNA conjugate was mixed with poly(ethyleneimine) (PEI) to form polyelectrolyte complex, which effectively silenced target gene expression in prostate cancer cells.
Fig. 9

Chemical structure of siRNA conjugated to PEG-ligand using acid-labile beta-thiopropionate bond

Amphipathic cationic polyvinyl ethers were synthesized with different alkyl group sizes [42] to evaluate potential for transfection. Transfection of DNA increased with alkyl group size where the butyl group showed the greatest transfection. The butyl group may provide the best transfection due to its hydrophobicity and ability to lyse membranes. This polymer, termed PBAVE, which is composed of poly(butyl and amino vinyl ether)s, was translated into dynamic polyconjugates (DPCs) for the delivery of siRNA [43]. This was accomplished by reversibly masking the amines with acid-labile PEGs and targeting ligands via maleamate bonds (Fig. 10). Furthermore, siRNA was conjugated to the polymer through a reductively labile disulfide bond. After endocytosis of the polyconjugate by the target cell, the endosomal vesicle gradually becomes more acidic. At this point, the acid-labile groups hydrolyze from the conjugate and reveal the PBAVE, which is able to lyse the endosome for entry into the cytoplasm. Herein, the intracellular environment is reducing and can cleave the disulfide bond between siRNA and the polymer to facilitate delivery of the cargo to the RNA-induced silencing complex. This dynamic polyconjugate system was able to target and effectively deliver siRNA to liver cells in vivo for gene knockdown.
Fig. 10

Chemical structure of maleamate-masked PBAVE and subsequent unmasking process Polymeric Micelles

For chemotherapeutic drug delivery, biocompatible polymeric micelles have shown high drug loading, encapsulation of poorly water-soluble drugs, prolonged circulation in blood, enhanced tumor accumulation, and therapeutic efficacy. For example, filamentous micelles (filomicelles) with high physical stability were prepared from block copolymers of poly(ethylene oxide) (PEO) and PCL or poly(ethylethylene) as degradable or inert materials, respectively [67]. In one particular study, the hydrodynamic diameter of filomicelles ranged from 25 to 60 nm, and the persistence length ranged from 0.5 to 5 μm. The size of the filomicelles was tuned by varying the block length of the copolymer with higher molecular weight yielding longer filomicelles. Compared to their spherical counterparts, the filomicelles persisted in circulation ten times longer, which was about 1 week after intravenous injection. The non-spherical shape allows for evasion of macrophage phagocytosis, and circulation time increased with persistence length of the filomicelle. Paclitaxel-loaded filomicelles showed efficacious drug delivery and improved therapeutic effect compared to spherical counterparts. Further investigation of paclitaxel-loaded micelles showed that filamentous micelles nearly double the maximum tolerated dose compared to spherical micelles [68]. Incorporation of fluorescent dye into the filomicelles showed that they migrate into a tumor and the micellar fragments penetrate into the tumor stroma.

Another micellar example with block copolymers involves poly(glutamic acid) (PGlu) and PEG in the delivery of cisplatin [69]. The formed micelles adopted a diameter of 28 nm, and the incorporation of cisplatin occurred via ligand exchange of Pt(II) from the chloride to the carboxylate of glutamic acid residues. Steady release of cisplatin occurred under physiological conditions. Drug-loaded micelles showed prolonged circulation and high tumor accumulation with complete tumor regression without toxicity. Combining PGlu-PEG block copolymer with a platinum anticancer drug and PGlu homopolymer at different concentrations allowed access to different sizes of micelles below 100 nm for evaluation in tumor models [70]. In highly permeable colon tumors, all particles accumulated and showed similar anticancer efficacy. However, in poorly permeable pancreatic tumors, the accumulation of drug increased with decreasing micelle size, and anticancer efficacy was greatest for the smallest, 30 nm micelles. Further derivatization of the PEG-PGlu polymeric micelle system with cyclic RGD integrin-targeting peptide and a platinum anticancer drug showed rapid tumor accumulation and penetration into tumor parenchyma [71]. Targeted micelles showed enhanced tumor delivery and anticancer efficacy in vivo, which may be attributed to a proposed selective active transcytosis internalization pathway via integrin receptors.

Given the anticancer efficacy observed with polymeric micelles, they have been utilized in clinical trials mainly due to their size, improved pharmacokinetics, and enhanced tumor accumulation of drug. A block copolymer composed of PEG and poly(d,l-lactide) was combined with paclitaxel to form 20 to 50 nm polymeric micelles, referred to as Genexol-PM® [72]. This formulation has been studied in clinical trials and is approved in South Korea for the treatment of breast cancer and non-small cell lung cancer. This formulation exhibited a higher maximum tolerated dose, better tumor accumulation, and improved efficacy compared to free drug and Taxol®, a commercially developed form of paclitaxel formulated with Cremophor EL®.

For the clinical trials using Pluronics, the formulation consisted of doxorubicin and Pluronics F127 and L61 (named SP1049C), which yielded ~25 nm micelles. Phase II clinical trials have been carried out using SP1049 C for advanced adenocarcinoma of the esophagus and gastroesophageal junction [73]. This study showed SP1049C provided substantial antitumor activity and a positive safety profile. In a murine leukemia model, SP1049C was found to decrease tumorigenicity and aggressiveness of cancer cells in vivo with enhanced activity against cancer stem cells [74].

Anticancer activity has also been seen in vivo for a polymeric micelle NK911, which has gone into clinical trials. NK911 is a block copolymer of PEG and poly(aspartic acid) (PAsp) and is partially conjugated with doxorubicin [75]. This copolymer is formulated with free doxorubicin to produce ~40 nm micelles that gradually releases doxorubicin over 8–24 h. The micelle showed prolonged blood circulation and high tumor accumulation. Furthermore, NK911 exhibited higher anticancer activity than free doxorubicin in multiple tumor models. In clinical trials, NK911 was found to be well tolerated [76].

A variety of poly(2-oxazoline)s (POx) block copolymers have been synthesized and studied for drug delivery [77]. One interesting example involves the first demonstration of formulating third-generation taxoids with POx micelles [78]. These POx were composed of methyl and butyl POx with a molecular weight of 10 kg/mol and PDI of 1.14. All taxoids could be loaded at nearly 1:1 ratio (50% drug) with resulting sizes around or below 100 nm. One of the taxoids showed enhanced efficacy in a multidrug resistant cancer cell line compared to paclitaxel in vitro. Furthermore, this taxoid demonstrated enhanced tumor inhibition in two orthotopic multidrug resistant tumor models in vivo.

The stability of polymeric micelles can be improved through crosslinking approaches. One approach harnesses block copolymers with ionic and nonionic hydrophilic segments. These block ionomers were combined with oppositely charged polymers to form block ionomer complexes. This approach has been utilized to deliver small molecule drugs, proteins, and nucleic acids [79]. In addition to electrostatically crosslinked micelles, covalently crosslinked micelles have been explored. For example, a biodegradable PAsp-based, PEG-terminated triblock copolymer was synthesized with crosslinkable thiol and pH-responsive tertiary amino groups [80] (Fig. 11). This system was termed interlayer-crosslinked micelle, which was capable of loading doxorubicin in the partially hydrated core and had a disulfide-crosslinked interlayer to prevent core expansion at neutral pH. After internalization by cells, triggered release of drug occurred in the acidic, reducing environment of the lysosome. In vitro and in vivo anticancer efficacy was observed for these micelles with improved performance compared to free drug and PEG-PCL micelles.
Fig. 11

Chemical structure of thiol and tertiary amine poly(aspartic acid) in interlayer-crosslinked micelles

Polymeric micelles have been utilized for the delivery of nucleic acids. For example, siRNA was linked to phosphothiol ethanol through a disulfide bond, and this conjugate was mixed with phosphthiolethanol-PEG to form reducible micelles [81]. This system showed protection of siRNA from nuclease degradation, cytocompatibility, and gene knockdown 50-fold more effective than free siRNA. Another example of micellar delivery of siRNA in combination with paclitaxel used matrix metalloproteinase 2 (MMP2)-sensitive copolymers [82]. Specifically, PEG was linked through a cleavable peptide linkage to 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine. This system provided high stability, siRNA condensation, drug encapsulation, MMP2-triggered tumor targeting, and increased cell uptake after MMP2 exposure. Dendrimers
Efficacious chemotherapeutic delivery via dendrimer conjugates has been demonstrated. For example, camptothecin was conjugated to PEGylated PLL dendrimer through aspartic acid [83]. The dendrimer conjugate had a long blood circulation time and ca. 14 times tumor uptake of the drug relative to free drug. With a single injection of the dendrimer conjugate in tumor-bearing mice, there was prolonged survival, while in another model all mice survived to the end of the study when injected with multiple doses. Another example involved the derivatization of a bow-tie dendrimer (Fig. 12) with doxorubicin through an acid-labile hydrazone bond, which provided prolonged blood circulation of the drug and nine times higher tumor accumulation relative to free drug [84]. A single initial intravenous injection resulted in complete tumor regression and survival of 60 days with efficacy similar to that of commercially available Doxil (liposomal formulation of doxorubicin).
Fig. 12

Chemical structure of poly(ester) dendrimer-PEG bow-tie hybrid

The incorporation of drugs with dendrimers through noncovalent approaches has been pursued, harnessing a variety of interactions. For example, high-generation poly(propylene imine) dendrimers were utilized for the encapsulation of Rose Bengal within their cavity via host-guest-like interactions [85]. However, for drug release, specific hydrolysis conditions were required to remove the dendritic shell. Additional noncovalent approaches include multiple hydrogen bonding and hydrophobic interactions. Poly(aryl ether)s have been utilized to enhance the water solubility of hydrophobic compounds [86]. The hydrophobic nature of certain dendrimers such as poly(arylene ether)-based polymers can be exploited to form micelles by functionalizing the outer surface with hydrophilic groups such as PEG. Micellar dendrimers have an intriguing advantage over conventional polymer micelles composed of amphiphilic block copolymers. Specifically, the micellar dendrimers remain stable under dilute conditions while the stability of block copolymers depends on the critical micelle concentration.

PLL has been used for dendrimer synthesis and DNA transfection. Dendritic PLL was synthesized with different generations [21], and the fifth- and sixth-generation PLL dendrimers were able to effectively transfect cells in vitro. These fifth- and sixth-generation dendrimers have 64 and 128 surface amine groups and were subsequently functionalized with arginine and histidine to increase the positive charge for complexation [87]. Arginine-modified dendritic PLL enhanced DNA transfection 3- to 12-fold in multiple cell lines, while the histidine derivative did not. For the histidine derivative, acidic conditions were required to effectively complex DNA. Using this effective complexation, subsequent transfection showed enhanced DNA delivery compared to the native polymer. Nanoparticles
Prodrugs have been designed for controlled release of chemotherapeutics conjugated to polymeric nanoparticles [88]. These prodrugs were a follow-up iteration from the development of acid-sensitive silyl ether crosslinkers with tunable degradation rates and cargo release kinetics [89]. For the prodrugs, asymmetric bifunctional silyl ether conjugates of camptothecin, dasatinib, and gemcitabine were synthesized. The silyl ether bond between the drug and polymerizable unit allowed for release of the native, active therapeutic after hydrolysis (Fig. 13). These prodrugs were synthesized in one step and were polymerized into PEG-based nanoparticles. By controlling the steric bulk of the substituent on the silicon atom, the release of drug could be tuned and was accelerated under acidic conditions. Corresponding in vitro studies showed that drug release could be tuned to elicit cytotoxicity similar to free drug in addition to eliciting minimal toxicity without regard to drug loading.
Fig. 13

Structure and mechanism of degradation for silyl ether prodrugs

Docetaxel was loaded into 200 × 200 nm cylindrical PLGA nanoparticles with encapsulation efficiencies greater than 90%. PRINT allows for high drug encapsulation regardless of loading while maintaining the physical properties of nanoparticles. By controlling the loading of docetaxel, the IC50 was tuned and reached higher potency than the commercial standard Taxotere® in vitro. PRINT has also been used to generate better medicines and specific nanoparticle theranostics, which have been reviewed [53, 90].

Paclitaxel has been encapsulated in BTM nanoemulsions with ca. 200 nm size and approximately 6% drug loading (drug/oil). These lipid-based, paclitaxel-loaded BTM nanoparticles showed long-term stability at 4 °C and 37 °C in PBS with slow release of cargo, lack of initial burst release, and potential for lyophilization without the need for cryoprotectants [45]. These particles were targeted to A431 tumors in vivo using epidermal growth factor receptor-binding Z-domain [46]. Delivery of paclitaxel-loaded nanoparticles in vivo provided anticancer efficacy in nude mice bearing resistant NCI/ADR-RES tumors. Furthermore, BTM nanoparticles overcame P-glycoprotein-mediated multidrug resistance by depleting ATP and inhibiting P-glycoprotein [91]. The delivery and efficacy of taxanes using lipid-based nanoparticles have been extensively reviewed elsewhere [92].

siRNA has been encapsulated in PRINT hydrogel nanoparticles [93]. This was accomplished through noncovalent and covalent means. Noncovalent complexation and physical entrapment enabled gene silencing; however, to retain siRNA during systemic administration or particle modification, a covalent approach was pursued. For reversible covalent incorporation, a polymerizable siRNA macromonomer with a degradable disulfide linkage was polymerized into PEG-based hydrogel nanoparticles (Fig. 14). A functional amine monomer was utilized for cellular internalization and endosomal escape. The content of this amine was screened to optimize cytocompatibility and gene silencing efficiency. This system allowed for stable incorporation in serum, triggered release of siRNA under intracellular reducing conditions, and effective gene silencing in vitro.
Fig. 14

Chemical structure and mechanism of release for reductively responsive siRNA-conjugated hydrogel nanoparticles

A double emulsion solvent evaporation technique has been utilized for the preparation of PLGA nanoparticles in the delivery of siRNA [94]. In this study, PLGA nanoparticles loaded with siRNA were administered in a single dose to the mouse reproductive track for effective and continual gene silencing in the vaginal mucosa. Another example of PLGA nanoparticle delivery of siRNA utilized a film-stretching approach to fabricate nanoneedles [95]. Gene knockdown was enhanced by nanoneedle-shaped PLGA nanoparticles compared to their spherical counterparts, attributed to cell permeabilization and induced cytoplasmic delivery. Furthermore, toxicity was minimized by using shape-shifting nanoparticles that lose the sharp, cell-penetrating nanoneedle edges. siRNA has also been encapsulated in PLGA PRINT nanoparticles [96]. In this study, 80 × 320 nm nanoparticles were synthesized with high encapsulation of siRNA and then surface-modified with cationic, transfecting lipids. These nanoparticles were internalized by multiple cancer cell lines and elicited knockdown of therapeutically relevant gene expression for the treatment of prostate cancer.

Nanoprecipitation has been used for the preparation of cisplatin prodrug PLGA-PEG nanoparticles [97]. These nanoparticles were decorated with prostate-specific membrane antigen (PSMA) targeting aptamer to deliver cisplatin to prostate cancer cells. The cisplatin prodrug was prepared by installing hexanoic acid groups in the axial positions to afford hydrophobicity and affinity for PLGA. The nanoparticles had a drug loading of approximately 6% and size ca. 140 nm. The targeting ligand was conjugated to nanoparticles post-precipitation through amine-carboxylic acid coupling using 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide. Targeting ligand-bearing nanoparticles was internalized through endocytosis by PSMA-positive LNCaP prostate cancer cells. Furthermore, the aptamer-functionalized cisplatin prodrug nanoparticles were more potent than cisplatin or nontargeted nanoparticles in LNCaP cells. This work established a potential method for systemic delivery of platinum drugs targeted to prostate cancer cells.

Poly(beta-amino ester)s (PBAEs) are a member of the poly(ester) class of materials for nanoparticle DNA delivery. A combinatorial library of PBAEs was synthesized, to understand the structure-property relationships for polymeric gene delivery [98]. These polymers condensed DNA to form nanoparticles that effectively delivered genes to cells with low cytotoxicity. Effective delivery of DNA was achieved with linear polymers of ca. 10 kDa, hydroxyl side chains, primary amine end groups, and tertiary amines. Therefore, the structure, functional groups, and molecular weight of the polymer influenced the gene delivery efficiency. These large-scale studies may be utilized in the design of biomedical polymers for drug delivery.

Similar to the combinatorial library approach for PBAEs, another class of polymers was synthesized through a high-throughput approach. Core-shell nanoparticles were synthesized via a combinatorial method for intracellular delivery [99]. Block copolymers of PEG and epoxides were crosslinked with a library of different amines (Fig. 15). These core-shell nanoparticles were able to effectively deliver plasmid DNA and small interfering RNA to cells in vitro. Some highlighted properties of efficacious particles include tertiary amines with one to two reactive sites, amines with buffering capacity, and thin hydrophilic shells. Covalent attachment of cholesterol allowed for gene silencing in the liver in vivo. Again, the polymer structure and functional groups played a huge role in determining drug delivery efficacy.
Fig. 15

Chemical structure of epoxide-containing block copolymers and their crosslinking with diamines

Delivery of DNA and siRNA has been accomplished using CD-based polycations. The beta-CD polymers contain amidine groups in the backbone as the cation. When the polycation was mixed with nucleic acid, nanoparticle formation occurred and gene delivery was achieved. To avoid undesirable interactions after intravenous administration, functionalization with PEG was pursued. The optimal approach involved use of adamantane-terminated PEG, which formed an inclusion complex with the hydrophobic pocket of beta-CD [100]. Further functionalization of these polycations with imidazole enhanced gene delivery efficiency [101] (Fig. 16). The pH-buffering properties provided by imidazole facilitated endosomal escape and gene delivery. Furthermore, the imidazole polycation was shown to have weaker binding with DNA and faster release under intracellular conditions. Here, the introduction of imidazole groups at the ends of the polymer notably altered the structure, properties, and efficacy of the polyplex. Further development of this work led to the first targeted delivery of siRNA to elicit gene knockdown in humans with melanoma [102].
Fig. 16

Chemical structure of cyclodextrin-based polycations for nucleic acid delivery

5.2 Imaging and Tracking Biomedical Polymers In Vivo

To understand the behavior of biomedical polymers in vivo and improve the efficacy of drug delivery, imaging modalities are of crucial importance. Imaging polymers can elucidate the role of polymer size, composition, and architecture on in vivo behavior. To analyze the in vivo behavior of polymers, conventional approaches use fluorescent dyes, multiple blood draws, and sacrifice of numerous mice with extensive analysis of collected tissue. Noninvasive imaging modalities include gamma camera imaging (scintigraphy), magnetic resonance imaging (MRI), single-photon emission computed tomography (SPECT), and positron-emission tomography (PET). These modalities are powerful and may provide sensitive, quantitative, and noninvasive analysis of tracers in real time. Here we will cover the use of select modalities to image or track biomedical polymers with linear, dendritic, and nanoparticle architectures.

5.2.1 Linear Polymers

HPMA copolymer-doxorubicin was the first linear polymeric conjugate to be evaluated clinically in 1994 [103]. The copolymer was composed of HPMA and a peptide linker side chain conjugated with doxorubicin. The peptide linker was Gly(D,L)Phe-Leu-Gly, which is cleavable by the lysosomal enzyme cathepsin B. This copolymer had a molecular weight of ca. 30 kDa and a doxorubicin content of 8.5 wt%. Since the copolymer contained doxorubicin, it formed small micelles in solution of approximately 6 nm in diameter. In Phase I trials, the copolymer was labeled with 131I through simple mixing in an iodogen reaction vial and purified by G-25 Sephadex column for imaging by gamma camera (scintigraphy) and quantitative analysis [104]. Pharmacokinetic analysis showed the labeled copolymer had a distribution half-life of 1.8 h and an elimination half-life of 93 h. Moreover, imaging studies showed that the copolymer is taken up by some tumors.

In Phase II studies, the HPMA copolymer-doxorubicin conjugate was evaluated in breast, lung, and colorectal cancers [105]. Here, an imaging analogue that contained 1 mol% methacryloyltyrosinamide was utilized for radioiodination using 123I for scintigraphy and SPECT analysis. The polymer showed bi-exponential kinetics with a distribution half-life of 3 h and elimination half-life of 41.2 h and area under the curve (AUC) of 5 mM·h. Radioimaging in one of the patients with breast cancer showed 5.9% of the injected dose after 24 h. The tumor of this patient was one of the major sites of the body with radioactivity present after 48 h. These results illustrated that therapeutics conjugated to polymers have altered and enhanced pharmacokinetics and anticancer activity.

5.2.2 Dendrimers

Nano-sized macromolecules with specific sizes can be useful as contrast agents. Given that the body may process macromolecules distinctly based on nanometer differences, dendrimers find great utility for elucidating the effect of nanometer sizes on biodistribution. These can be utilized with Gd(III), which enabled sufficient contrast by MRI with sub-millimolar concentrations at 1% of the concentration of iodine required as a contrast agent. Two common chelators for Gd(III) are diethylenetriamine pentaacetic acid (DPTA) and 1,4,7,10-tetraazacyclododecane tetraacetic acid (DOTA). PAMAM and polypropyleneimine diaminobutane (DAB) dendrimers are water-soluble and provide functional handles for conjugation with chelators. A difference in a few nanometers and internal core (PAMAM vs. DAB) was shown to significantly alter the pharmacokinetics [106, 107]. For example, DAB dendrimers showed notably more liver accumulation and were more rapidly excreted than PAMAM dendrimers. Dendrimers of 10–14 nm showed minimal renal excretion, while 3–6 nm dendrimers were rapidly excreted by the kidney. Furthermore, 6 nm dendrimers were found to leak out of tumor vessels much quicker than 10 or 14 nm dendrimers. This study demonstrated the potential for additional applications of dendrimers with particular physicochemical properties as contrast agents in imaging functional anatomy such as tumor blood vessels and the lymphatic system.

The biodistribution profiles of aliphatic polyester dendrimers have been determined through radiolabeling with Tc(I) and Re(I) [108] using SPECT. Specifically the core of poly(2,2-bis(hydroxymethyl)propanoic acid) dendrimers with different generations was modified with a chelator. Radiolabeling with 99mTc allowed for real-time in vivo monitoring biodistribution; these dendrimers were rapidly cleared from blood by the kidneys and not retained in any organs after injection in rats. Another set of poly(ester)-based dendrimers functionalized with PEG chains was synthesized in a “bow-tie” structure for elucidating biodistribution and tumor uptake [109]. The effect of molecular weight and number of PEG chains for these bow-tie dendrimers was explored. Due to the polyester composition, these dendrimers were demonstrated to be degradable. Dendrimers were labeled with 125I and evaluated for biodistribution. Higher molecular weight bow-tie dendrimers with molecular weight greater than 40 kDa demonstrate prolonged circulation half-life and high tumor accumulation with 109 kDa provided the longest half-life and 52 kDa providing the highest tumor accumulation.

5.2.3 Nanoparticles

An interesting imaging approach involved an amphiphilic core-shell nanoparticles that were prepared from diblock copolymers of poly(acrylic acid) and poly(styrene) or poly(methyl acrylate) [110]. Particles were crosslinked with a diamine to yield shell-crosslinked nanoparticles, and acid groups were functionalized with a macrocyclic chelator (TETA derivative) for chelation of copper-64 (Fig. 17) and imaging by PET. Nanoparticles with different sizes (18–37 nm), core composition, and surface PEGylation were studied for their biodistribution. It was found that 18 nm NPs with a rigid polystyrene core showed the longest blood retention time and lowest liver accumulation. This study showed the importance of polymer and particle design parameters (size, flexibility, and surface chemistry) on biological behavior in vivo.
Fig. 17

Functionalization of poly(acrylic acid)-containing shell-crosslinked nanoparticles with macrocyclic chelator TETA

To optimize the radiolabeling of shell-crosslinked nanoparticles, the spacer length, crosslinking extent, and charge density were investigated [111]. The block lengths of PAA-block-PS were varied from 30 to 136 repeat units. A nine-atom spacer (DOTAamine) with three carboxylates was compared to a five-atom spacer (DOTAlysine) with four carboxylates. The longer spacer DOTAamine provided enhanced coupling due to steric hindrance factors and electrostatics based on the negatively charged nanoparticle. Twenty percent crosslinked nanoparticles enabled significantly increased amount of 64 Cu-accessible DOTAs compared to 50% crosslinked nanoparticles. This can be attributed to more functional acrylic acid units present on the surface for the lower degree of crosslinking. The specific activity was greatest for the block copolymer with the highest PAA content, which provided small particle size and minimal aggregation. Here, we see that in preparing radiolabeled nanoparticles, several factors must be considered to maximize specific activity for imaging.

5.3 Scaffolds for Tissue Engineering

Polymeric materials are widely utilized as scaffolds in tissue engineering to restore lost or damaged tissue. This goal is often accomplished with a combination of polymeric scaffold, cells, and biological factors. The scaffold serves as an artificial extracellular matrix designed to mimic the environment and physical characteristics of the specific tissue. The scaffold materials are designed to elicit desirable biological responses. Some of the most important properties of the scaffold include high porosity, appropriate pore size, biocompatibility and biodegradability, mechanical strength, and surface chemistry. All of these properties are influenced by the type and preparation of polymeric material. Further investigation into scaffold designs has revealed that micro-/nanostructure, geometry, and topology can make a profound impact on function and physiological responses. Controlling the properties and design features of scaffolds for a particular tissue engineering application depends on the identity and physicochemical characteristics of the polymeric material.
  1. (i)


    Polyesters represent a large class of polymers that are used as tissue regenerative scaffolds. Their mechanical strength, biocompatibility, and tunable biodegradation times make them ideal candidates as scaffolds in tissue engineering. Electrospinning is a common method to fabricate fibers from polymers for biomedical applications. Using electrospinning, a three-dimensional collecting method was employed to fabricate PLGA/PCL fibrous tubes with different interconnected structures, patterned architectures, and macroscopic configurations for potential tissue engineering applications [112]. Electrospun PLGA/PCL micro- and nanofiber scaffolds were synthesized with diameters ranging from 280 nm to 8 μm with small (700 nm) and large (20 μm) pores [113]. These scaffolds were utilized for nerve regeneration in lesioned rats. There were three groups of rats studied: (1) sciatic nerves were transected; (2) a 10 mm gap was left after removal of a 5–7 mm segment of the sciatic nerve; (3) electrospun tubes were implanted post neurotmesis. The multi-scaled structure of these electrospun tubes enabled the repair of a nerve gap of 10 mm in vivo in a rat sciatic nerve, while the sciatic nerves failed to reconnect in control groups.

    Polymer hybrid materials generally show enhanced attachment and proliferation of cells. For instance, electrospun PCL was compared to PCL coated with gelatin and calcium phosphate (apatite) [114]. This coated porous scaffold mimicked the bone extracellular matrix and provided higher rate of proliferation of preosteoblastic cells after 1 week culture relative to PCL alone. Another hybrid material that has been studied for bioactivity is PLGA-collagen-hydroxyapatite [115]. Biomineralization was observed to occur preferentially on the hybrid membrane side compared to PLGA alone using human mesenchymal stem cells (hMSCs). This study demonstrated the development of a hybrid material system as a potential new class of biomimetic scaffolds that facilitate bone tissue engineering.

  2. (ii)


    To maintain signaling in the nervous system, muscle contraction, and wound healing, electrical stimulation can play a crucial role. Nerve cell stimulation can be provided by using the conducting polymer, poly(aniline) (PANI) [116]. PANI is biocompatible and has been copolymerized or blended with other biodegradable polymers to provide a range of mechanoelectric properties. For example, blends of poly(lactide-co-caprolactone) (PLCL) and PANI electrospun into fibers provided unique cellular behavior [117]. Specifically, fibroblasts and myoblasts exhibited higher levels of adhesion to PLCL-PANI compared to PLCL alone. Furthermore, the growth of fibroblasts was enhanced by stimulation under direct current flows. Another study developed a mesh of PLGA and PANI as an electrically active scaffold for coordinating the beating of cardiomyocytes [118]. Abnormalities in the functioning of cardiomyocytes are relevant to myocardial infarction. Synchronous beating of all cardiomyocytes was realized via electrical stimulation of the PLGA-PANI fibers to mimic the heart.

  3. (iii)

    Hyaluronic Acid

    Combination of hyaluronic acid (HA) with other biopolymers can enhance biological functions. HA is a key scaffold for tissue engineering. In particular, HA is FDA-approved, biocompatible, and is a native extracellular component. Covalent incorporation of fibronectin through photocrosslinking with HA provided increased viability of cultured endothelial cells compared to fibrinogen adsorbed to HA [119]. Gelatin-containing HA hydrogel particles embedded into a HA network have controlled the adhesion and differentiation of mesenchymal stem cells [120]. HMSCs are readily attached, migrated deeply, and formed an interconnected population in this network, while isolated spheroids were encountered with gelatin-free networks. Furthermore, production of collagen and mineral deposition were noted in gelatin-containing networks, suggesting osteogenic differentiation, while markers for adipogenesis were found in gelatin-free networks.

    Another example of a hybrid material for cartilage repair involved hydrogels composed of methacrylated glycol chitosan with HA prepared by photocrosslinking [121]. These polymers were photocrosslinked with a riboflavin initiator using visible light. High cell viability of encapsulated chondrocytes (~80 to 87%) was observed over 21 days. Furthermore, it was found that the incorporation of HA increased cell proliferation and cartilaginous extracellular matrix depositing. From these studies, we see that HA has provided enhanced biological compatibility and function for tissue engineering.

  4. (iv)

    Gelatin and PEG

    Three-dimensional structures, such as those microengineered through projection stereolithography, have been utilized for eliciting specific biological responses. For example, microengineered structures were prepared starting from gelatin functionalized with methacrylic anhydride [122]. The mechanical properties of this crosslinked gelatin were controlled by varying the porosity and concentration of the prepolymer. Cell growth was supported on defined geometries with uniform cell distribution, high cell density, and homogeneity. To maximize live cell fabrication and minimize damage to cellular DNA through UV irradiation, visible light-based projection stereolithography was developed [123]. Here, hydrogel scaffolds were prepared from PEG diacrylate with different shapes and architectures. Adipose stem cells maintained high viability in the scaffolds, and porous scaffold architectures provided higher cell viability and activity. The three-dimensional structure of the scaffold was shown to influence the biological responses.


5.4 Medical Devices: Replacements for Heart Valves, Arteries, and Joints

Advances in tissue engineering enable the production of replacement organs in great demand. This notable shortage of organs is being addressed by developing new methods to provide biological replacements such as heart valves, arteries, and joints. Conventional approaches to produce biological replacement have substantial drawbacks. These include the potential for thromboembolism, infection, immune responses, and inability to adapt for growth. Conversely, tissue engineering approaches are aimed toward delivering and integrating living and biocompatible material that avoid the drawbacks.

The heart valve is critical in the cardiovascular network for appropriate one-directional blood flow. Heart valve disease poses a significant threat for death. One approach to create heart valves involves the transplant of autologous cells onto a scaffold shaped like a heart valve. Modern approaches are harnessing three-dimensional printing for the production of heart valves. For example, hybrid hydrogels composed of methacrylated HA and gelatin have been synthesized for the encapsulation of aortic valve cells [124]. With increased gelatin content, lower stiffness, higher viscosity, cell spreading, and enhanced maintenance of fibroblast phenotype were realized. High cell viability was observed in addition to the deposition of collagen and glycosaminoglycans. Another three-dimensional printing technique provided living alginate/gelatin hydrogels with region-specific incorporation of two cell types [125]. Viability was maintained for aortic smooth muscle cells and aortic valve leaflet cells, which expressed muscle actin and elevated vimentin, respectively.

Connections among tissues and organs to the heart are provided by blood vessels. Blood vessel diseases represent a significant percentage of deaths due to complications with atherosclerosis and subsequent myocardial infarctions. Engineering of blood vessels may allow for facile surgery and repaired function instead of challenging operations like cardiac bypass surgery. Poly(glactin)/PGA tubular scaffolds have been fabricated to construct tissue-engineered artery conduits [126]. Harvested artery or vein cells were seeded on scaffolds to create a vascular construct that was then used as arteries in lambs. These vascular grafts demonstrated growth and development of endothelial lining and extracellular matrix and behaved well in pulmonary circulation. Another example of a vascular graft involved PCL, which was utilized in a rat abdominal aorta replacement model [127]. PCL micro- and nanofibers were evaluated over 18 months. PCL maintained structural integrity and showed an absence of aneurysmal dilation, thrombosis, and minimal hyperplasia.

Joint replacements were first developed over a century ago and remain important materials for damaged joints. Ultrahigh molecular weight poly(ethylene) (UHMWPE) has been the main chosen material for joint replacement. This may be attributed to the appropriate properties and biocompatibility of UHMWPE. Specifically, UHMPWE is a semicrystalline material with a molecular weight of 3–6 megadaltons (MDa), yield/tensile strengths on the orders of megapascals (MPa), and elastic moduli around 1 GPa [128]. Although UHMWPE can function well for at least 15 years, there is growing concern about UHMWPE debris generated in vivo [36]. This has led to advances in processing, sterilization, and crosslinking of UHMWPE [129]. For example, vitamin E has been used to stabilize UHMWPE and showed good mechanical, wear, and oxidation properties and improved longevity of joint replacements [130]. These advancements and further iterations may enable increased longevity without the presence of debris for UHMWPE joint replacements in vivo.

5.5 Surgery and Wound Repair Materials

Polymeric materials have found extensive use in surgery and wound repair. Wounds can be defined by a disruption in the skin arising from a physiological nature or external damage. Three main phases constitute wound healing: inflammation, tissue formation (proliferation), and tissue remodeling. This healing process involves a complex interplay among cells, extracellular matrices, and signaling factors. To aid in the wound healing process, polymeric materials have found utility in promoting re-epithelization, accelerating healing, minimizing scar formation, and stimulating cellular activities. The properties of polymeric materials are important to the effect they exert in the wound healing process. Notable properties of surgery and wound repair materials include moisture permeability, tensile strength, elasticity, bioadhesion, biocompatibility, and bioactivity (increasing angiogenesis). Moreover, combining polymers can lead to synergistic and positive biological responses in wound repair.

Chitin and chitosan (deacetylated chitin) have been widely utilized in wound repair materials due to their inherent regenerative properties. For example, chitosan can activate macrophages to signal healing, encourage cell motility and angiogenesis, and exert antimicrobial activity [131]. Furthermore, chitosan is biocompatible and biodegradable, mucoadhesive, and amenable to chemical modification via free primary amine. A formulation involving chitin nanofibrils, chitosan glycolate, and chlorhexidine was prepared as a spray, gel, and gauze to evaluate efficacy in treating traumatic wounds [132]. In all clinical cases, the preparations demonstrated abundantly satisfactory results. Physiological repair was enhanced with the gel, scar formation was absent with the gauze, and bleeding abrasions were treated with the spray as a first-aid tool.

Blends of chitin and chitosan with additional polymeric materials have resulted in stimulated healing-impaired wound repair [133]. Specifically, a hydrogel was prepared from chitin/chitosan, alginate, and fucoidan as wound dressing material. This blended hydrogel elicited advanced granulation tissue and capillary formation in healing-impaired wounds compared to the commercial calcium alginate fiber material Kaltostat®. Another chitosan-based hydrogel was synthesized with layers of chitosan, alginate, and PGA for wound healing in diabetic rats [134]. This hydrogel provided accelerated wound healing compared to alginate hydrogels alone and increased collagen regeneration and epithelialization.

Materials for biologically interactive wound repair have been explored using crosslinked glycosaminoglycans. Specifically chondroitin sulfate, hyaluronan, and PEG were crosslinked into a polymeric network [135]. Increased re-epithelization was observed for a film of the polymeric network combined with the commercial film dressing Tegaderm™ compared to the commercial product alone. Furthermore, there was more collagen regeneration and organization for the polymer network, while the inflammatory response remained the same. Another example for wound healing was with chondroitin sulfate and involved the preparation of bilayer gelatin-chrondoitin-6-sulfate-HA biomatrices [136]. This material was applied for wound healing in severe combined immunodeficiency mice. The bilayer material served as a skin substitute, which promoted wound healing and had a high take rate of the skin graft.

5.6 Polymeric Biosensors

The detection of chemical and biochemical species is important in monitoring our health and the environment. Polymeric materials offer the capabilities for sensitive analyte detection with amplified signals. Recently, conjugated polymers have been the focus of developing biosensor materials. Conjugated polymers are composed of unsaturated backbones that can provide conductivity and energy transfer. The advantages of conjugated polymers include inherent material processing and mechanical properties combined with electrical and optical properties of semiconductors and metals. The sensing properties of conjugated polymers are influenced by chemical structure, conjugation length, conformation, and packing.

Conjugated polymers are often hydrophobic and water-insoluble. Common structures of conjugated polymers are provided in Fig. 18. For biological applications, water-soluble polymers are typically preferred. Therefore, there has been extensive effort into designing water-soluble conjugated polymers for biosensing through chemical modification and incorporation. The interactions between the conjugated polymer and biochemical agent of interest can be harnessed for increased and specific sensing.
Fig. 18

Chemical structures of biomedical polymers used as biosensors

For instance, a water-soluble cationic conjugated polymer, which can complex DNA, has been designed to change emission color based on conformational and aggregation changes [137]. A benzothiadiazole was incorporated into a quaternary amine-functionalized poly(fluorene-co-phenylene). Different colors were obtained for noncomplementary and complementary single-stranded DNA in addition to DNA-free solutions. This opened the door to multicolor biosensors based on tuning electrostatic and optical interactions. Another example utilized cationic poly(fluorene-alt-1,4-phenylene) derivatized with benzothiadiazole for detecting heparin [138]. This polymer allowed for multicolor detection and quantification of heparin by harnessing the electrostatic attraction between the polymer and heparin. Detection was observable by the naked eye from orange emission induced by fluorescence resonance energy transfer.

The detection of proteases has been realized with water-soluble poly(phenylene ethynylene)s functionalized with carboxylates through oligo(ethylene glycol) spacers [139]. Covalent incorporation of a fluorescence-quenching peptide yielded a nonemissive substrate. The polymer regained fluorescence after exposure to trypsin, which cleaved the peptide sequence incorporated into the polymer. A small molecule analogue of this system highlighted the effect of polymer signal amplification in biosensing. Another example for the detection of proteins involved a sulfonate-derivatized poly(phenylene vinylene) as a biosensor for cytochrome c [140]. At neutral pH, cytochrome c is positively charged and can form a complex with this conjugated polymer, resulting in a quenching of its fluorescence. The efficient quenching was attributed to the combined effect of electron transfer between cytochrome c combined with complex formation between cationic and anionic polyelectrolytes, which opened the door for improving poly(phenylene vinylene) biosensors.

Enzymatic cleavage and oxidative damage to DNA has been detected using a cationic conjugated polymer [141]. A quaternary ammonium-functionalized polythiophene allowed for real-time, label-free monitoring of DNA cleavage by harnessing the interpolyelectrolyte complex formed between the conjugated polymer and DNA. In complex formation, a planar conformation was assumed with an associated red shift in absorption. After DNA damage, the complex could not form and exhibited a distinct absorption. In addition, the DNA damage could be observed visually by color change. DNA has also been utilized indirectly for biosensing. For example, human thrombin was detected with high sensitivity using single-stranded DNA (aptamer)-conjugated polymer complexes [142]. A specific DNA aptamer changes conformation after binding human thrombin. This change in conformation was detected using a cationic poly(3-alkoxy-4-methylthiophene) derivative to provide a label-free optical signal. This approach provided a foundation for rapid detection and identification of proteins and high-throughput screening in drug discovery.

6 Perspectives and Future Trends

From the biomedical applications covered herein, we have seen unique relationships among polymer identity, structure, properties, and function. These relationships may provide blueprints for the design of a polymer to reach a specific biomedical goal. The approach to designing each biomedical polymer and its structure involves fine-tuning the chemical composition. The chemical composition may be selected based on the breadth of choices covered herein. Distinct structures and compositions may be accessed through novel or defined synthetic approaches. The range of accessible structures and forms of biomedical polymers spans from linear native polymers to self-assembled systems to synthetic hydrogels. Certain structures provide advantages for a given application. Moreover, we show there are key characteristics and advantages of polymer structures and composition that are similar within each application.

In drug delivery and imaging, key characteristics of biomedical polymers include biocompatibility, stability, chemical functionality, and unique biological behavior. Current and future directions include high-throughput synthesis and screening, enhanced stability of the material and cargo, and specific cargo release mechanisms and rates. For example, enhanced stability of the cargo may be realized by covalent conjugation and protection from degradation by polymer shielding. Cargo release may be tuned via labile linkages, disassembly mechanisms, or pore size of nanoparticles. The stability and biodegradation rate of the polymer can be tweaked based on the chemical composition, structure, and morphology such as increasing crystallinity. Additional focus directions involve controlling the structure, size, shape, and composition of the material. For instance, the size and shape of self-assembled systems can be determined by the molecular weight and hydrophobic-hydrophilic stoichiometry of polymers. Furthermore, top-down fabrication methods allow for versatile control over the composition, shape, size, and surface chemistry of nanoparticles. The repertoire of information presented on drug delivery and imaging may allow for better designs, development, and testing of biomedical polymers for therapeutic and diagnostic purposes.

In tissue engineering, key characteristics of biomedical polymers include mechanical strength, hydrophilicity/hydrophobicity, biodegradability, stability, conformation, and biofunctionality. Select biomedical polymers and combinations thereof have been shown to exert biological interactions and elicit particular biological responses. Furthermore, polymers functionalized with distinct chemical moieties have provided additional bio-interactions. The nano- and microscale features of polymeric materials substantially influence biological behavior. Forthcoming work for these applications involves the design and synthesis of novel polymers and combinations thereof, further chemical modification, and improved or new techniques for processing and production. For example, polymers may be produced by methods including electrospinning, lithographic processes, and three-dimensional printing. The future of biomedical polymers is open to creative polymer designs and material production. New designs and production methods may increase the breadth of unique structures and continually realize improved function and efficacy in biomedical applications.

7 Conclusions

In summary, we have covered different classes, structures, and properties of biomedical polymers with relevant biomedical applications. The structure-property-function relationships of biomedical polymers reveal design criteria that may be harnessed to realize their potential in the biomedical arena. Tailoring the structure and chemical composition of biomedical polymers to match the application may enable desirable performance. For example, specifying the method of drug incorporation and controlled release can allow for fine-tuned drug delivery. Similarly, controlling the physicochemical and mechanical properties of polymers can provide particular extracellular matrix-like environments for tissue engineering. The range of polymer structures covered herein provides a toolbox of crucial design features. Selecting and building upon these polymer designs may lead to the development of unique properties and improved performance. One of the forefront goals in designing biomedical polymers is ultimately aimed at transforming human health and improving the quality of life. We hope that this book chapter has provided the necessary information to better design and develop biomedical polymers that may advance medicine and healthcare.


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Copyright information

© Springer International Publishing AG, part of Springer Nature 2018

Authors and Affiliations

  1. 1.Department of Radiology and Biomedical Research Imaging CenterUniversity of North Carolina at Chapel HillChapel HillUSA
  2. 2.Lineberger Comprehensive Cancer CenterChapel HillUSA

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