Medical, Dental, and Pharmaceutical Applications

7  Medical, Dental, and Pharmaceutical Applications 7.1 Biopolymers The biomedical use of synthetic biodegradable polymers is believed to have begun in the late 1960s with the approval of the first bioabsorbable sutures. Since that time, numerous applications in the biomedical field have been made including uses such as wound enclosures, body implants, tissue engineering materials, drug delivery materials, and in vivo sensing materials. The term biodegradable polymers in medicine refers to those polymers, which are slowly converted to nontoxic degradation products in the body, without specifying the elimination route from the body. There are also a number of alternative terms known as “bioresorbable,” “bioabsorbable,” and “bioerodible” polymers, which fall under the general definition of biodegradable polymers. There is considerable confusion in the prior art as to the exact meaning of each of these terms. In this chapter the following definitions are used:

or hydrolytically unstable backbone. The very instability of these polymers, which leads to biodegradation, has proven to be important in medical applications. The most widely used biodegradable polymers in medical devices are: • Biodegradable organic polymers

• poly(lactic acid) or polylactide (PLA), and its various stereoisomer forms, especially:



- poly(l-lactic acid) or poly(l-lactide) (PLLA)



- poly(d-lactic acid) or poly(d-lactide) (PDLA)



- poly(d,l-lactic acid) or poly(d,l-lactide) (PDLLA)

- poly(l-lactide-co-d,l-lactide) (PLDLLA)

- stereocomplex PLA (scPLA) (blend of PLLA and PDLA)



•  poly(glycolic acid) or polyglycolide (PGA) and its copolymers, especially:



• “Bioresorbable” refers to biodegradable polymers that undergo bulk erosion and converted under physiological conditions to degradation products, which are resorbed within the body. • “Bioabsorbable” refers to biodegradable polymers that are dissolved in bodily fluids. A bioabsorbable polymer can be bioresorbable if the dissolved polymer is metabolized and/or excreted by a living organism. The two terms are often used interchangeably in the literature.



•  polyhydroxyalkanoates (PHAs), homopolymers and copolymers, especially:



- poly(3-hydroxybutyrate) (P3HB)

- poly(4-hydroxybutyrate) (P4HB)



- poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBHV)



- poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx)

• poly(ε-caprolactone) (PCL), and its copolymers, especially:

• “Bioerodible” refers to water-insoluble biodegradable polymers that undergo surface erosion and converted under physiological conditions to water-soluble products, by means of physical (e.g., dissolution) and chemical processes, which are resorbed within the body. Biodegradation may occur by enzymatic mediation, degradation in the presence of water (hydrolysis) and/ or other chemical species in the body, or both. The most prevailing type of degradation of biodegradable polymers is hydrolysis of ester linkages or other labile bonds

- poly(l-lactic acid-co-glycolic acid) or poly(llactide acid-co-glycolide) (PLGA)



- poly(lactide-co-ε-caprolactone) (PLCL)

- poly(glycolide-co-ε-caprolactone) (PGCL) • poly(propylene fumarate) (PPF)



• poly(p-dioxanone) (PDO or PDS)

• poly(trimethylene carbonate) (PTMC) and its copolymers, especially:





- poly(lactide-co-trimethylene (PLTMC)

Biopolymers: Applications and Trends. http://dx.doi.org/10.1016/B978-0-323-35399-1.00007-7 © 2015 Published by Elsevier Inc. All rights reserved.

carbonate)

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- poly(glycolide-co-trimethylene carbonate) (PGTMC)

• polyanhydrides

• poly(ortho esters) (POEs)

• Biodegradable inorganic polymers

• polyphosphazenes (PPHOs)

• calcium polyphosphates (CPPs)

The main chemical, physical, and structural characteristics of these polymers and methods of their production are commented in Chapters 1 and 2 of this book. There are also numerous publications in the literature on the medical applications of biodegradable polymers [1–5]. An overview of the properties of the main biodegradable polymers used in medicine is presented in Table 7.1. Some of these polymers and their copolymers have been studied extensively for biomedical applications such as sutures, staples and mesh for wound closure, fracture fixation, bone augmentation and ligament reconstruction in orthopedics, ligation clips and vascular grafts in cardiovascular surgery, and dental repairs. They have also been used to prepare biodegradable drug delivery systems capable of releasing the drug or a biologically active substance over the desired length of time. The main biomedical applications of biodegradable polymers are summarized in the following sections.

A relative advantage of these materials is that of eliminating the need for a second surgical intervention, sometimes necessary for removing the implant, thanks to the gradual biodegradability, which favors regeneration of the preexisting tissues. On the other hand, these materials have the drawback, as far as biomedical use is concerned, of having a low hydrophilicity, which reduces their wettability and the adherence of the organism cells, namely the endothelial cells. Low hydrophilicity is the main drawback of known aliphatic polyesters (with the exception of PGA) in medical applications. This is of crucial importance for the swelling behavior in physiological systems. Surface energies of most polyesters are rather low, indicative of their nonhydrophilic character, resulting in low wettability and too low degradation rates. The low wettability induces low biocompatibility and insufficient interaction between implant material and living tissue. Another drawback in the use of known bioresorbable polyesters is the relatively low controllability of the resorption kinetics leading to a negative impact on medical applications. Due to the bulk erosion of known materials, resorption kinetics are hard to control, and can lead to unfavorable and undesirable low or high degradation rates. Low degradation rates are known for PHAs and PCL (2012, WO2012174580 A1, UNIV GRAZ TECH; MEDICAL UNIVERSITY OF GRAZ MUG; AUSTRIA TECH & SYSTEM TECH; HERAEUS MEDICAL GMBH).

Table 7.1  Comparison of Properties of Biodegradable Polymers (Compiled by US2013085563 A1, ABBOTT CARDIOVASCULAR SYSTEMS, 2013)a

Biopolymer

Tm (°C)

Tg (°C)

Tensile strength (MPa)

Tensile modulus (MPa)

Elongation at break (%)

Absorption rate

PLLA

175

65

28–50

1200–2700

6

1.5–5 years

P4HB

60

−51

50

70

1000

8–52 weeks

PCL

57

−62

16

400

80

2 years

PDO

110a

−10a

1.5a,b

30b

35c

6–12 weeksa 6 weeksb

PGA

225

35

70

6900

<3

6 weeks

PDLLA

Amorphous

50–53

16

400

80

2 years

P3HB

180

1

36

2500

3

2 years

P3HB, Poly(3-hydroxybutyrate); P4HB, Poly(4-hydroxybutyrate); PCL, Poly(ε-caprolactone); PDLLA, Poly(d,l-lactic acid), pol(d,l-lactide); PDO (or PDS), Poly(p-dioxanone); PGA, Poly(glycolic acid), polyglycolide; PLLA, Poly(l-lactic acid), poly(l-lactide). aAll except PDO from Martin DP, Williams SF. Biochemical Engineering Journal 2003;16(2):97–105. bMedical Device Manufacturing & Technology, 2005. cBronzino JD, editor. The biomedical engineering handbook, Press in Cooperation with IEEE Press Boca Raton, USA; 1995.

7: Medical, Dental, and Pharmaceutical Applications

Poly(l-Lactide) Poly(l-lactide) (PLLA) is a biodegradable and biocompatible, thermoplastic, aliphatic polyester resulting from polymerization of l, l-lactide (also known as l-lactide). PLLA has a glass transition temperature (Tg) of 53–63 °C, a melting temperature (Tm) of 170–185 °C and a crystallinity of about 37%. PLLA is currently used in a number of biomedical applications, such as sutures, stents, dialysis media, and drug delivery devices. It is also being evaluated as a material for tissue engineering. PLLA is preferred in cases where high mechanical strength and toughness are required, for example, in orthopedic devices.

Poly(d,l-Lactide) Poly(d,l-lactides) (PDLLAs) is a group of a biodegradable and biocompatible, thermoplastic, aliphatic polyesters resulting from polymerization of d,l-lactide using stannous octoate as the catalyst. There are three types of d,l-lactide, meso-d,llactide, racemic d,l-lactide (or rac-lactide, which is an equimolar mixture of l-lactide and d-lactide), and blend-d,l-lactide. PDLLAs are used for the preparation of bioabsorbable sutures, controlled drug release systems, stents, stent coatings, and tissue engineering scaffolds. PDLLAs are glassy, amorphous polymers that degrade by bulk hydrolysis in vivo. Upon irradiation, a reduction in molecular weight of PDLLA is observed. A well-known copolymer of PDLLA is poly(l-lactide-co-d,l-lactide) (PLDLLA), which is also amorphous and has mechanical features similar to those of PLLA without the inconvenience of long degradation times and high crystallinity. PLDLLA mainly is used as scaffold for bone engineering.

Polyglycolide Polyglycolide or poly(glycolic acid) (PGA) was one of the first developed synthetic biodegradable polymers for medical use (Dexon™ sutures). PGA has high crystallinity (40–55%), a Tg of 35–40 °C, a high Tm (about 225 °C), low solubility in organic solvents, and a very high tensile strength (12.5 GPa) [6]. Owing to its hydrophilic nature, a PGA suture tends to lose its mechanical strength rapidly (50%) over a period of 2 weeks and is absorbed in about 4 weeks after implantation [7].

Poly(l-Lactide-co-Glycolide) Poly(l-lactide-co-glycolide) (PLGA) is the most investigated biodegradable polymer for medical and pharmaceutical applications. It is obtained by the

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random copolymerization of l-lactide and glycolide. Since PLLA and PGA have significantly different properties, the proper selection of the ratio of the two monomers allows for optimization of the PLGA properties for the intended application. PLGA has lower crystallinity and Tm than PGA and PLLA. For example, while PGA and PLLA are partially crystalline, a 50/50 PLGA is entirely amorphous (see Table 7.2). These morphological changes result in an increase in the rates of hydration and hydrolysis. Thus, PLGA copolymers tend to degrade more rapidly than PGA and PLLA. PLGA has been used mainly in the manufacture of sutures (e.g., Vicryl®, Ethicon Inc.), as a vehicle in controlled release medicines, and as scaffold in tissue engineering because of its good adhesion and proliferation properties [8].

Poly(ε-caprolactone) Poly(ε-caprolactone) (PCL) is a biodegradable polyester derived from fossil fuel resources with a Tm = 57–60 °C and a Tg = −(60–62) °C. This polymer is often used as an additive for resins to improve their processing characteristics and their end use properties (e.g., impact resistance).

Polyhydroxyalkanoates Polyhydroxyalkanoates (PHAs) are produced by soil bacteria and are degraded upon subsequent exposure to these same bacteria. The PHA polymers may constitute up to 90% of the dry cell weight of bacteria and are found as discrete granules inside the bacterial cells. These PHA granules accumulate in response to nutrient limitation and serve as carbon and energy reserve materials. Table 7.2  Crystallinity and Thermal Properties of PGA, PLGA, PLLA, and PDLLA (2005, US2005079470 A1, DENTIGENIX INC.) Biopolymer

Crystallinity (%)

Tm

Tg

PGA

46–52

225

36

PLGA (90/10)

40

210

37

PLGA (50/50)

0

–

55

PLLA

37

185

57

PDLLA

0

–

n/a

PDLLA, Poly(d,l-lactide); PGA, Polyglycolide or poly(glycolic acid); PLA, Poly(lactic acid), polylactide; PLGA, Poly(l-lactic acid-co-glycolic acid), poly(l-lactide-co-glycolide). Adapted from Wong ad Mooney, 1997 [9].

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There are currently only few commercially available PHAs, namely P3HB, P4HB, PHBHV, and PHBHHx out of a great variety of PHAs (more than 100 different monomers have been incorporated into polymers under controlled fermentation conditions). A preferred PHA for medical applications is P4HB. P4HB is a biocompatible, bioresorbable, strong, and ductile polymer. P4HB degrades primarily by bulk hydrolysis. Surface erosion also plays a role in the degradation process, and is believed to be a hydrolytic process mediated by enzymes. Its degradation products are much less inflammatory, making scaffolds of P4HB well suited to applications where tissues and cells are sensitive to acidic environments [10]. The degradation rate of P4HB in vivo is fast relative to other PHAs, however, its bioresorption rate is slower than many of the materials used as bioresorbable sutures such as PGA. Additionally, P4HB implants maintain their molecular weight during the process of bioresorption. Because of its excellent mechanical properties, maintenance of high molecular weight, processability, biocompatibility, and bioresorbability, P4HB is useful as bioresorbable wound closure material for suturing and stapling devices (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.). It can be used also in many types of implant applications including orthopedic, craniomaxillofacial, dental, and cardiovascular, as well as in cardiology, plastic and reconstructive surgery, general surgery, ear, nose, and throat surgery, and oral surgery. While the PHAs offer a wide range of mechanical properties which are potentially useful in medical applications, their use particularly in vivo as bioresorbable polymers has been limited by their slow hydrolysis.

Poly(Propylene Fumarate) Poly(propylene fumarate) (PPF) is a biodegradable polymer that has been used as a biomaterial. PPF is an unsaturated linear polyester that degrades in the presence of water into propylene glycol and fumaric acid, degradation products that are easily cleared from the human body by normal metabolic processes. Because the fumarate double bonds in PPF are reactive and crosslink at low temperatures, PPF can be an effective in situ polymerizable biomaterial. The high mechanical strength of cured PPF matrices and their ability to be crosslinked in situ make them especially suitable for orthopedic applications, such as bone cement, orthopedic scaffold for bone tissue regeneration, and

Biopolymers: Applications and Trends

in drug delivery systems (1995 WO9529710 A1; 2000, WO0062630 A1, UNIV WM MARSH RICE).

Poly(p-dioxanone) Poly(p-dioxanone) (PDO or PDS) homopolymer has a Tm of 105–110 °C and a Tg of −20 to −10 °C, when the crystallization has not yet progressed. PDO will crystallize when stored at room temperature (1996, EP0691359 A2; 2012, WO2012040316 A1, ETHICON INC.). In addition, PDO is hydrolytically unstable, which renders it biodegradable, and is unable to tolerate standard environmental conditions. Methods have been proposed to control the instability of PDO, for example, by the addition of hydrolytic stabilizers, such as calcium salts, or by carefully controlled packaging in order to isolate the PDO products from the environment. As a result, the commercial use of PDO has been generally limited to highly specialized applications, wherein the exposure to the environment can be controlled, for example, biodegradable sutures and surgical devices, such as, for example, surgical clips (1997, EP0763559 A2, UNION CARBIDE CHEM PLASTIC).

Poly(Trimethylene Carbonate) Poly(trimethylene carbonate) (PTMC) is a biodegradable and biocompatible polymer, which has a Tg of approximately −17 °C. This amorphous and flexible polymer can be crosslinked by irradiation in an inert atmosphere to form a creep resistant and formstable network. Such elastomeric polymers are especially used as scaffolding materials for the engineering of soft tissues, or as depots for controlled release systems, or in the design of implants, such as antiadhesion membranes or vascular prostheses. In vivo, PTMC degrades relatively rapid by surface erosion without the release of acidic degradation products. PTMC has a very low tensile modulus and tensile strength, and it is mainly used as a component of copolymers and blends. Copolymers of TMC and d,l-lactide present intermediate physical properties between those of PTMC and PDLLA, and depending on the composition of copolymers, the values of Tg, tensile modulus, tensile strength, and degradation rate can be tuned. This makes these polymers suitable for numerous biomedical applications. However, amorphous non-crosslinked copolymers with Tgs below physiological temperatures have poor form stability at body temperature. Therefore, after crosslinking, autoclaving can be used as a method of sterilization of medical devices prepared from these copolymers.

7: Medical, Dental, and Pharmaceutical Applications

Polyanhydrides Polyanydrides are surface-eroding polymers, which maintain their mechanical integrity by preserving the molecular weight of the polymer and exhibit a gradual loss in size. Polyanydrides have a hydrophobic backbone and a water-labile linkage, which allows heterogeneous degradation in vivo from the surface of the polymer, leading to near zero-order release kinetics. The water-labile anhydride linkage provides the basis for the use of a variety of backbones, each having a unique degradation rate. Thus, the rate of degradation in vivo can be controlled by controlling the length and composition of the polymer backbone. Examples of polyanhydrides include poly[bis(p-carboxyphenoxy) propane anhydride] and poly[bis(p-carboxy) methane anhydride]. Polyanhydrides find applications in orthopedic, dental, and controlled release applications (1990, WO9009783 A1; 1994, US5356630 A, MASSACHUSETTS INST TECHNOLOGY).

Poly(Ortho Esters) Poly(ortho esters) (POEs) are surface-eroding polymers. The hydrophobic character of the polymer limits water penetration and hydrolysis to only the exterior surface of the polymer matrix [11]. Thus the surface erosion is much faster than that of the bulk. The chemical and physical properties of POEs depend on the chemical structures of the constituent monomers [12]; for example, reaction of bis(ketene acetal) with rigid trans-cyclohexane dimethanol produce a rigid polymer with a Tg = 110 °C, whereas that of the flexible diol 1,6-hexanediol produces a soft material having a Tg = 20 °C. Mixture of the two diol results in polymers having intermediate Tgs. The surface-eroding property of POEs is attractive for a variety of tissue engineering applications. The monomer release would be steady over the lifetime of the matrices, in contrast to PLA and PGA. In addition, the gradual loss of polymer from the surface of the scaffold may allow the surrounding tissue to serially fill the space vacated by the polymer (1997, WO9745533 A1, UNIV MICHIGAN).

Polyphosphazenes Polyphosphazenes (PPHOs) are high molecular weight polymers with an inorganic backbone of alternating phosphorus and nitrogen atoms with two side groups on each phosphorus atom. Polymers with a wide range of properties can be synthesized from this polymer backbone by incorporating different side groups, by varying the side group in single-substituent

295

PPHOs, and/or by using two or more co-substituents. Individual PPHOs may be hydrophobic, amphiphilic, or hydrophilic; water-stable or water-erodible; crystalline or amorphous; bioinert or bioactive. The hydrophobicity of the PPHO is increased by adding hydrophobic side groups, such as aromatic groups, to the backbone. The degradation rate can be regulated by increasing the hydrophobicity of the PPHO (2000, US6077916 A, PENN STATE RES FOUND).

Calcium Polyphosphates Ceramics based on calcium polyphosphates (CPPs) are used in biomedical or dental applications that require biodegradable structures for implants and the like. These materials are nontoxic, nonimmunogenic, and are composed of calcium and phosphate ions, the main constituents of bone. CPPs are inorganic polymers (Ca(PO3)2)n consisting of networks of oxygen-bridged (PO4)3− tetrahedra and shared Ca2+ ions (one per pair of phosphate tetrahedra). Porous structures made of CPP are biodegradable and, as such, offer potential for a number of novel biomaterial applications including use as substrates for forming tissue-engineered implants for the repair and augmentation of degraded soft and hard tissues and, in particular, for anchoring soft connective tissues to bone (e.g., cartilage or ligament to bone) [13,14]. Porous CPP substrates of desired structure can be formed by sintering CPP powders of appropriate size.

7.2  Wound Enclosures, Body Implants, and Tissue Engineering Materials Sutures are the major area of application for biodegradable polymers in the medical devices market. Other applications include staples, screws, and rods for pinning and repairing ligaments, orthopedic moldings, cardiovascular and intestinal supports [15].

7.2.1 Sutures Sutures are fibrous devices used in surgical procedures for holding cut tissue surfaces in apposition for a period of time sufficient for healing. Sutures are used in a variety of applications including wound fixation, wound closure, general tissue approximation,

296

Biopolymers: Applications and Trends

and attachment of implants. Sutures are classified mainly into three categories based on:

• Physical structure

• Monofilament

• Multifilament/braid

• Material composition

• Natural

• Synthetic • Behavior in tissue

• Absorbable

• Non-absorbable

Monofilament sutures are made of a single strand of material. Monofilament sutures have less tissue drag than multifilament sutures. Multifilament sutures are made of several strands of suture that are twisted or braided together. Braided multifilaments are characterized by a higher flexibility but have a higher roughness than monofilaments. To get smooth multifilaments often a coating is applied on the suture. Sutures are often made of nonbioabsorbable polymers such as polyolefins, nylon, cotton, and the like, and are generally removed after a period of time. Bioabsorbable sutures are made of biodegradable polymers such as collagen, silk, PLGA, PDO, and the like, and do not require subsequent removal. Bioabsorbable sutures are used generally for soft tissue approximation where only short-term wound support is required. Catgut was the first bioabsorbable suture and was described already in AD 175 by Galen [16]. Catgut sutures are made from the submucosa layer of the intestines of animals (e.g., sheep, beef), which consist mainly of collagen. Animal intestinal tubes are split longitudinally, cleaned and spun or twisted to form strands. Chromic gut suture is a variant, which permits the suture to retain its strength for a longer period of time to allow a wound or incision to heal properly. Such a suture is tanned by immersion in a solution of tanning agent such as a chromium salt, which increases the resistance of collagen to hydrolytic attack (1999, US5954748 A, UNITED STATES SURGICAL CORP.). Catgut sutures are stored in a tubing solution, from which they are removed by the surgeon just prior to use, otherwise, they will dry out and become too hard for use. Such storage solutions are well known in the art, and typically include water

(10 wt%) and alcohol (89 wt%), for example, ethanol and/or isopropanol, and optionally triethanolamine (1 wt%). US1254031 A (1918, DAVIS & GECK INC.) discloses a suture comprising a plurality of animal, preferably beef, tendon fibers braided together to constitute a single strand, and a coating of collagen coalesced with the braided fibers causing adhesion of the fibers and filling up the interstices of the strand. US5891167 A (1999, UNITED STATES SURGICAL CORP.) discloses a catgut suture including a dehydrothermal crosslinked collagen coating which is insoluble in tubing solution. The collagen coating is formed by immersing the catgut suture in a solution of collagen in acidified water. Optionally, a plasticizer may be included in the coating solution. The suture is then heated in a vacuum to effect crosslinking of the collagen. In US5954748 A (1999, UNITED STATES SURGICAL CORP.) the coating is a crosslinked gelatin formed by immersing the gut suture in aqueous gelatin solution, fixing the gelatin in a solution of crosslinking agent such as glutaraldehyde, buffering the coating, and drying the suture at a temperature of at least about 50 °C. In many instances, however, catgut may cause adverse tissue reaction in the sutured flesh. This, together with the fact that it requires storage under moist conditions, makes it less than an ideal suture material. In Europe and Japan, catgut sutures have been banned due to concerns that they could transmit bovine spongiform encephalopathy (BSE) or mad cow disease, although the herds from which catgut is harvested are certified BSE-free [17]. GB1048088 A (1962, ETHICON INC./DU PONT) discloses a suture comprising oriented monofilaments of PLLA. BE654236 A (1965, AMERICAN CYANAMID CO) discloses a suture of at least one filament of stretched and oriented PGA. US3867190 A (1975, AMERICAN CYANAMID CO) discloses surgical elements including sutures consisting of PLGA containing about 15–85 mol% glycolic acid and 85–15 mol% lactic acid, which is claimed to have enhanced tissue absorption as compared with PLA, and enhanced solubility in organic solvents as compared with PGA and, hence can be cast into sheets during preparation and implantation. Such surgical element may be used alone or in combination with PGA or a coating for PGA surgical elements. Unfortunately, the copolymers formed by co-reacting increasing amounts of glycolide with the

7: Medical, Dental, and Pharmaceutical Applications

lactide were said to have the disadvantage of forming surgical articles which lacked dimensional stability in vivo (1981, US4243775 A, AMERICAN CYANAMID CO). US4243775 A (1981, AMERICAN CYANAMID CO) discloses surgical products manufactured preferably of PGTMC. In a preferred embodiment, the sterile surgical article is in the form of a suture or a ligature. In a most preferred embodiment, the sterile surgical article is in the form of a needle and suture combination. US4052988 A (1977, ETHICON INC.) discloses absorbable sutures and other surgical devices manufactured of PDO. EP0908142 A2 (1999, ETHICON INC.) discloses a coated braided suture with improved knot strength comprising a bioabsorbable suture (e.g., made of PLGA, namely Polyglactin 095, l-lactide/glycolide = 95/5 mol%) having a low molecular weight bioabsorbable polymer (e.g., PGCL, glycolide/εcaprolactone  =  10/90) dispersed throughout the braided suture. EP0668083 A1 (1995, UNITED STATES SURGICAL CORP.) discloses bioabsorbable sutures and other surgical devices manufactured of a bioabsorbable filament having a generally cylindrical core portion made of PDO with a diameter in the range of 0.01–0.8 mm, and a shell portion made of PCL having a thickness of 0.001–0.5 mm. EP2425865 A1 (2012, AESCULAP AG) discloses a surgical thread comprising a thread body and a coating made essentially of PHA. The PHA is selected from the group consisting of P3HB, P4HB, poly(3hydroxyvalerate) (P3HV), poly(4-hydroxyvalerate), poly(5-hydroxyvalerate), poly(3-hydroxyhexanoate) (P3HHx or PHHx), poly(4-hydroxyhexanoate), poly(5hydroxyhexanoate), poly(6-hydroxyhexanoate), copolymers thereof, optical isomers thereof, and combinations, in particular blends, thereof. In one of its embodiments the thread body is made of PLA, PGA, PCL, PTC, PDO, copolymers thereof and combinations, in particular blends, thereof. A preferred copolymer is PLGA. The thread is claimed to have improved knot characteristics, in particular in view of knot fitting, knot retention, and knot run-down behavior, due to a particular PHA coating. In this manner, the security of knots may be improved and a risk of wound dehiscence may be minimized significantly. Another advantage is related to an improved degradation characteristic of the thread. Thus, the hydroxyalkanoic acids produced during degradation of PHA, in

297

particular 3-hydroxybutyric acid and/or 4-hydroxybutyric acid as released from P3HB and/or P4HB, have a minor acidity, whereby postoperative hyperacidity of the wound environment, and thus undesired enhancement of inflammation responses may be prevented, in particular during the initial wound-healing period. Yet another advantage is related to the excellent biological anchoring of the thread. Thus, there is fibrous connective tissue encapsulation in the vicinity of the incision site even at an early time postsurgery, whereby tissue integration of the thread, and thus security of wound sealing and wound closure are improved. The various commercially available bioabsorbable sutures are listed in Table 7.3.

7.2.2 Fasteners Surgical fasteners include linearly insertable (i.e., push-in type) fasteners and rotationally insertable (i.e., screw-in type) fasteners. Linearly insertable surgical fasteners offer an alternative to rotationally insertable surgical fasteners and as a means for attaching or reattaching soft tissue to bone. Tacks, rivets, staples, suture anchors, plugs, and soft tissue anchors are among the most common forms of linearly insertable surgical fasteners. Examples of commercially available bioabsorbable fastener devices are: •  ArthroRivet™ implant (Arthrotek Inc.) made of Lactosorb® (Biomet, Inc.) copolymer (82% PLLA, 18% PGA);

• Bard® Permasorb® Disposable Fixation Device with PDLLA fasteners designed for mesh fixation and for approximation of soft tissue on open and laparoscopic surgery;

• Sorbafix® fastener (Bard Davol Inc.) designed for prosthetic mesh fixation and for approximation of soft tissue. Absorption of the PDLLA fasteners is nearly complete at approximately 12-month postimplantation, leaving less foreign material behind.

Suture fastener devices, including suture anchors, buttons, or pledgets, are typically used to reattach tissue to bone. Often the procedures involve the attachment of tendon, ligament, or other soft tissue to bones in the shoulder, knee, elbow, wrist, hand, and ankle. In one approach, bone anchors are inserted into the bone, and then soft tissue such as ligament or tendon

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Biopolymers: Applications and Trends

Table 7.3 Commercially Available Bioabsorbable Sutures Commercial name

Biopolymer

Form

Function

Company

Surgical gut suture—plain

Beef serosa or sheep submucosa

Monofilament (virtual)

In general soft tissue approximation and/or ligation, including use in ophthalmic procedures, except for cardiovascular and neurological tissues

Ethicon Inc.

Vicryl rapide™

PLGA (polyglactin 910)

Braided

In superficial soft tissue approximation of the skin and mucosa, where only shortterm wound support (7–10 days) is required

Ethicon Inc.

Coated Vicryl™

PLGA (polyglactin 910)

Braided/ monofilament

In general soft tissue approximation and/or ligation, including use in ophthalmic procedures, except for cardiovascular and neurological tissues

Ethicon Inc.

Coated Vicryl™ plus

PLGA (polyglactin 910) and triclosan (Irgacare® MP)

Braided/ monofilament

In general soft tissue approximation and/or ligation, except for ophthalmic, cardiovascular and neurological tissues

Ethicon Inc.

Polysorb™

PLGA (Lactomer™) coated with a mixture of PGLC and calcium stearoyl lactylate

Braided multifilament

In soft tissue approximation or ligation and ophthalmic surgery, except for cardiovascular or neural tissue

Covidiena

Monocryl™

PGCL (Poliglecaprone 25)

Monofilament

In general soft tissue approximation and/ or ligation, but not for use in cardiovascular or neurological tissues, microsurgery or ophthalmic surgery

Ethicon Inc.

Monocryl™ plus

PGCL (Poliglecaprone 25) and triclosan (Irgacare® MP)

Monofilament

In general soft tissue approximation and/ or ligation, except for cardiovascular or neurological tissues, microsurgery or ophthalmic surgery

Ethicon Inc.

7: Medical, Dental, and Pharmaceutical Applications

299

Table 7.3  Commercially Available Bioabsorbable Sutures—cont’d Commercial name

Biopolymer

Form

Function

Company

PDS™ Plus41

PDO and triclosan (Irgacare® MP)

Monofilament

In soft tissue approximation, including use in pediatric cardiovascular tissue where growth is expected to occur and ophthalmic surgery (other than contact with cornea and sclera); not indicated in adult cardiovascular tissue, microsurgery, and neural tissue; particularly useful where the combination of an absorbable suture and extended wound support (up to 6 weeks) is desirable

Ethicon Inc.

Dexon™ S

PGA

Monofilament/ braided

In soft tissue approximation and/or ligation, including use in ophthalmic procedures, except for cardiovascular tissue or in neural tissue

Covidien

Bondek® plus

PGA with polyglyd™ coating

Braided multifilament

In soft tissue approximation and/or ligation, including use in ophthalmic procedures, except for cardiovascular or neurological procedures.

Teleflexb

PDS™ II

PDO

Monofilament

In soft tissue approximation, including use in pediatric cardiovascular tissue where growth is expected to occur and ophthalmic surgery. PDS suture is not indicated in adult cardiovascular tissue, microsurgery and neural tissue. These sutures are particularly useful where the combination of an absorbable suture and extended wound support (up to 6 weeks) is desirable Continued

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Table 7.3  Commercially Available Bioabsorbable Sutures—cont’d Commercial name

Biopolymer

Form

Function

Company

Caprosyn™

Glycolide-co-εcaprolactoneco-trimethylene carbonate-lactide block copolymer (Polyglytone™ 6211)

Monofilament

Frequent uses: plastic surgery, general subcuticular closure, OB/GYN, urological, and ENT (ear, nose, and throat)

Covidien

Biosyn™

Glycolide(60%)-cop-dioxanone (14%)-cotrimethylene carbonate(26%) block copolymer (Glycomer™ 631)

Monofilament

In soft tissue approximation, subcuticular skin closure and/or ligation, including use in ophthalmic surgery, except for use in cardiovascular or neurological surgery

Covidien

Maxon™, Maxon™ CV

PGTMC (polyglyconate)

Monofilament/ braided

In soft tissue approximation and/or ligation, including use in pediatric cardiovascular tissue, where growth is expected to occur, and in peripheral vascular surgery; not indicated for use in adult cardiovascular tissue, ophthalmic surgery, microsurgery and neural tissues

Covidien

Phantom Fiber™

P4HB

Braided

Multiple applications: elbow, shoulder, knee, hip, foot, and ankle

Tornier

MonoMax®

P4HB

Monofilament

Soft tissue approximation

Braun Surgical

P4HB, Poly(4-hydroxybutyrate); PDO (or PDS), Poly(p-dioxanone); PGA, Poly(glycolic acid), polyglycolide; PGCL, Poly(glycolide-co-εcaprolactone); PGTMC, Poly(glycolide-co-trimethylene carbonate); PLGA, Poly(lactide-co-glycolide). aFormerly United States Surgical Corp. bDeknatel acquired by Teleflex Inc. in 2000.

may be sutured to the anchor point. The procedure may be performed in an open manner or preferably using a minimally invasive technique, whereby the device is deployed by a suitable delivery device. Early suture anchors were made of metal. However, the use of metallic implants has been associated with a number of problems: loosening, intra-articular migration, and breakage of implant [18]. Metallic implants can also impair adequate visualization with either radiographs or magnetic resonance imaging (MRI).

Bioabsorbable suture anchors provide a viable alternative to metallic implants, overcoming the problems with retained hardware, radiographic visualization, and revision surgery. Early bioabsorbable suture anchors were made of PGA that hydrolyzed quickly, which resulted in fixation failure and recurrence of shoulder instability if used for shoulder reconstructions. Next generation of bioabsorbable anchors were a blend of PLA and PGA, resulting in slower hydrolyzation and added strength. New bioabsorbable suture

7: Medical, Dental, and Pharmaceutical Applications

anchors made of PLLA has a lengthened half-life compared to sutures made of PLGA copolymers [19]. Examples of bioabsorbable suture fastener devices currently in use, which are representative of the state of the art include:

• PushLock® Suture Anchor (Arthrex) made of PLDLLA, or biocomposite of PLDLLA and β-tricalcium phosphate (β-TCP) (see also Section 7.2.6);

• SutureTak® (Arthrex) made of PLDLLA or BioComposite SutureTak® made of PLDLLA and β-TCP is a family of bioabsorbable suture anchors indicated for use in arthroscopic stability procedures of the femoroacetabular join (see also Section 7.2.6); • Bionx Biodegradable Anchor (Bionx Implants); •  Bioroc EZT™ Suture Bone Fastener/screw (Depuy Mitek, a Johnson & Johnson Co.; ex Innovasive Devices1); • Bio-Statak® Resorbable Soft Tissue Attachment Device (Zimmer) made of PLLA to anchor sutures;



• Acufex TAG Bioabsorbable Anchors (Smith & Nephew Endoscopy).

7.2.3 Staples While traditional suturing remains a popular method for closing skin openings, the use of fasteners, in particular staples as a skin closure technique has become increasingly popular, especially in surgical settings where the opening is created through a purposeful incision. In these settings, the incision tends to make a clean, straight cut with the opposing sides of the incision having consistent and nonjagged surfaces. Typically, stapling of a skin opening, for example, is accomplished by manually approximating the opposing sides of the skin opening and then positioning the stapler so that a staple will span the opening. Surgical staples initially have a U-shaped configuration, and are employed by using a tool to insert the legs of the staples into the skin and cause the legs to bend inwardly to hold the staple in place, while the cross-member of the staple traversing the skin opening (see Figure 7.1). Generally, the staple is made of a deformable material

1. Johnson & Johnson Co. acquired in 1999 Innovasive Devices.

301

Figure 7.1  Staples (1989, EP0326426 A2, NIPPON MEDICAL SUPPLY).

such as surgical stainless steel and the legs of the staple are driven into an anvil causing the staple to deform so as to retain the skin tissue in a compressed manner within the staple. This process can be repeated along the length of the opening such that the entire incision is held closed during the healing process. After a sufficient amount of healing has occurred, the staples are withdrawn using a removal tool which causes the legs to unbend sufficiently to withdraw the staple from the tissue. Although effective, metallic staples suffer from the drawback of requiring postoperative removal. The use of biodegradable polymers eliminated the need for postoperative removal of the staples. Examples of biodegradable polymers used in the manufacture of bioabsorbable staples and clips are disclosed in the following patents: GB2069650 A (1981, ETHICON INC.) uses (co) polymers of lactide, glycolide, and p-dioxane; US4532927 A (1985, ETHICON INC.) uses PLGA; EP0185453 A2 (1986, ETHICON INC.) uses a blend of PLGA and PDO (>25 wt%); WO2012040316 A1 (2012, ETHICON INC.) uses a blend of PLGA and PDO (16–24 wt%); EP0209371 A1 (1987) and US4889119 A (1989) of ETHICON INC. use a glycolide-based blend of polymers comprising PGA and PLA; US4839130 A (1989, UNITED STATES SURGICAL CORP.) uses a multiphase, polymeric composition derived from lactide-based and glycolide-based polymers; and EP0326426 A2 (1989, NIPPON MEDICAL SUPPLY) uses PLA, PGA, or PLGA. Due to the nature of bioabsorbable polymers, however, bioabsorbable staples could not be inserted with the same deformation approach used by metallic staples. In fact, bioabsorbable staples were purposefully designed to avoid any deformation requirement, as deformation was seen as a potential failure mechanism. An example of such a design is illustrated by the inwardly biased skin staple of US5089009 (1992, UNITED STATES SURGICAL CORP.), wherein the bioabsorbable polymer is PLGA. Thus, as the physical and chemical properties of bioabsorbable surgical staples evolved, the development of designs and insertion methods associated with

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bioabsorbable staples has focused on avoiding deformation of the bioabsorbable fastener. Examples of commercially available bioabsorbable staples are:

• FasT-Fix™ AB Meniscal Repair System (Smith & Nephew), a meniscal implant made of PLLA;

• Gore® Seamguard® Bioabsorbable Staple Line Reinforcement (W. L. Gore & Associates Inc.) made of PGTMC; specifically engineered to reduce the incidence of perioperative leaks and bleeding in a variety of open and minimally invasive surgeries. It adds a consistent, strong, yet thin 0.50 mm reinforcement to the staple line that mechanically increases the strength of the resection line by evenly distributing the staple pressure exerted on the tissue over a wider surface area.

Representative patents of bioabsorbable meniscus repair devices (tacks) are shown in Table 7.4. The most commonly used device was the Meniscal Arrow (Linvatec) because of its ease of insertion and early success rates. The current version of the Meniscal Arrow (Contour™ Meniscus Arrow™) has a lowprofile head and is barbed along the entire length of the implant shaft to improve fixation strength. It is composed of a faster-resorbing self-reinforced copolymer PLLA (L/D = 80/20), which retains its strength for up to 24 weeks and then gradually resorbs [20,21]. US4884572 (1989) and US4895148 A (1990) of CONCEPT disclose a repair tack (10) that is primarily intended for use in arthroscopic surgery for the repair of torn meniscus tissue (see drawing in Figure 7.2). The repair tack is formed from a bioabsorbable polymer or copolymer, preferably derived from glycolic and lactic acids. The bioabsorbable repair tack is chosen to have a degradation time in excess of the required healing time for the tissue.



•  SubQ It!™ bioabsorbable skin closure system (Opus KSD, Inc.) comprising a disposable stapler preloaded with bioabsorbable fasteners made of PLGA (l-lactide/glycolide = 82/12), which are inserted subcutaneously. Although it has many other applications, the SubQ It!™ system was specifically designed for closing small incisions used for 5 and 10 mm trocars in laparoscopic procedures. Table 7.4 displays a list of representative patents disclosing bioabsorbable surgical fasteners.

7.2.4  Meniscus Repair Devices (Tacks) Meniscus repair devices (tacks) are orthopedic fixation devices that are used to secure fixation of longitudinal vertical meniscus lesions (bucket-handle lesions) located in the vascularized area of the meniscus in combination with suitable immobilization. The devices are often used in minimally invasive surgery. Bioabsorbable meniscal repair devices, including arrows, screws, darts, and staples, belong to the socalled third generation of meniscal repair devices. Most of these devices were made of rigid PLLA, which retains its strength for up to 12 months and requires 2–3 years or more to completely resorb. Examples of bioabsorbable meniscus repair devices, which are representative of the state of the art are: • BioStinger® Meniscal Fixation System (Linvatec), a meniscal implant made of PLLA;



• Biofix™ Meniscal Arrow (Bioscience Ltd.), a meniscal implant made of PLLA; • Meniscal Arrow (Linvatec); • Contour™ Meniscus Arrow™ (Conmed-Linvatec);

• RapidLoc™ Meniscal Repair System (Mitek), meniscal implant made of PLLA or PDO.

7.2.5  Meniscus Regeneration An unmet need in meniscus repair relates to defects located in the avascular region of the menisci where no blood vessels are present. Tears or defects in this region are not expected to heal. The only available treatment for avascular tears is a meniscectomy, where the portion of the meniscus surrounding the tear is removed. This procedure is unsatisfactory, as the meniscectomy disturbs the ability of the meniscus to function properly. Therefore, a meniscus regeneration product able to facilitate repair of avascular tears is highly desirable. PLLA [22], PGA [23,24], PLGA [25], PLCL [26], PHAs (e.g., PHBHV) [27] (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.), and aliphatic polyurethanes (PUs) [28] have been considered for use as meniscus regeneration scaffolds. These meniscus regeneration devices can be fabricated using a variety of different processing techniques, including the use of salt leaching, melt, solvent, fiber, and foam processing techniques. Devices may be formed, for example, as sponges, foams, nonwovens, or woven materials [29]. US2005232967 A1 (2005) and US2007031470 A1 (2007) of DEPUY MITEK INC. disclose a biocompatible meniscus repair device comprising a

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303

Table 7.4  Bioabsorbable Surgical Fasteners (Rivets, Tacks, Staples, and Tissue Screws) Used in Soft Tissue Reattachment Function

Biopolymer

Patent

Tack for the repair of bodily tissue (meniscal cartilage, bone, skin, and ligaments) in an in vivo surgical procedure

PCL, PLLA, PDLLA, PGA, PDLLGA (d,l-lactide/glycolide = 95/5, 90/10, 85/15, 75/25,50/50), PDLLCL (d,llactide/ε-caprolactone = 90/10, 75/25, 50/50), poly(ester amides), copolyoxalates, aliphatic polycarbonates, poly(glutamic-co-leucine)

US4884572 A (1989); US4895148 A (1990, CONCEPT)

Tack for the repair of meniscal cartilage

PLA, PGA, PCL, PTMC

WO9718761 A1 (1997); WO9718762 A1 (1997, INNOVASIVE DEVICES INC.)

Tack for the repair of bodily tissue (reattaching soft tissue to bone)

PCL, PLLA, PDLLA, PGA, PDLLGA (d,l-lactide/glycolide = 95/5, 90/10, 85/15, 75/25,50/50), PDLLCL (d,l, lactide/ε-caprolactone = 90/10, 75/25, 50/50), poly(ester amides), copolyoxalates, aliphatic polycarbonates, poly(glutamic-co-leucine)

US5129906 A (1992, LINVATEC CORP.)

Staples

PLA/PGA blend

US4889119 A (1989, ETHICON INC.)

Staples

PLGA, PDO, pol(amino acid)s (e.g., casein, albumin, etc.), PVOH, catgut

WO8301193 A (1983, UNITED STATES SURGICAL CORP.)

Staples, clips

PLGA

WO8401508 A1 (1984, UNITED STATES SURGICAL CORP.)

Staples

PLA, PGA, pol(amino acid)s (e.g., casein, albumin, etc.), PVOH, catgut

FR2322578 A (1977, UNITED STATES SURGICAL CORP.)

Staples

PLA, PGA, PDO

GB2075144 A (1981); US4428376 A (1984; ETHICON INC.)

Fasteners (staples)

PLGA

US2004059378 A1 (2004, INCISIVE SURGICAL IN)

Staples, clips

PLA/PGA blend, PDO

EP0185453 A2 (1986, ETHICON INC.)

Screw fasteners for the repair of meniscal cartilage

PCL/PGA blend, PLA/PGA blend, P3HB, PDO, POE, crosslinked collagen

US5730744 A (1998, JUSTIN DANIEL F; WINTERS JR THOMAS F)

Catgut, Collagen derived from sheep or beef intestinal submucosa; P3HB, Poly(3-hydroxybutyrate); PCL, Poly(ε-caprolactone); PDLLA, Poly(d,llactic acid), poly(d,l-lactide); PDLLCL, Poly(d,l-lactide-co-ε-caprolactone); PDLLGA, Poly(d,l-lactide-co-glycolide); PDO or PDS, Poly(pdioxanone); PGA, Poly(glycolic acid), polyglycolide; PLA, Poly(lactic acid), polylactide; PLGA, poly(lactic acid-co-glycolic acid); PLLA, Poly(l-lactic acid), poly(l-lactide); POE, Poly(ortho ester); PTMC, Poly(trimethylene carbonate); PVOH, Poly(vinyl alcohol).

biocompatible tissue repair scaffold adapted to be placed in contact with a defect in a meniscus, wherein the scaffold comprises a high-density, dry laid nonwoven polymeric material and a biocompatible foam, wherein the scaffold has a tensile modulus greater than about 1.5 MPa and a suture pull-out strength

greater than about 6 N. The nonwoven polymeric material may be formed from polymer(s) derived from, for example, glycolide, lactide, ε-caprolactone, trimethylene carbonate (TMC), or dioxanone monomer; preferably from PDO, or PLGA. The nonwoven polymeric material is heat-set.

304

Figure 7.2 Repair of torn meniscus tissue (1989, US4884572 A; 1990, US4895148 A, CONCEPT). 10, Repair tack; 15, Crossbar grip portion; 20, Applicator; 30, Needle; 31, Needle’s sharp end; 34, 36, Meniscus portions.

Actifit® (Orteq® Ltd.) is a commercial bioabsorbable meniscal porous scaffold. It is a segmented aliphatic PU consisted of PCL soft segments (80%) and urethane hard segments (20%) made of 1,4-butanediisocyanate and 1,4-butanediol [30].

7.2.6  Orthopedic Fixation Devices Orthopedic fixation devices, including plates, screws, pins, rods, anchors, and staples, are used in orthopedic surgery procedures, such as bone fracture fixation, autograft ankle stabilization, reconstruction surgery of the anterior cruciate ligament (ACL) and the posterior cruciate ligament (“PCL”), replacement of the intervertebral discs, and posterior spinal fixation (2011, WO2011041714, UNIV DREXEL). Commonly used osteosynthesis devices are produced from metals, mainly from medical grade stainless steel, pure titanium, titanium alloys, vitallium, chrome cobalt, and suitable biocompatible nonbiodegradable polymers. Metallic implants used successfully for fracture reduction have certain disadvantages. They show large differences in tensile modulus as compared with moduli of bones, frequently undergo in vivo corrosion, are not degradable, and have to be removed after the bone fracture is healed leaving cavitations in the bone.

Biopolymers: Applications and Trends

Metallic implants can also cause several other problems if left intact as the bone reacts to the metallic implants over time forming deposits on them. Metallic implants provide also potential sites for infection. Migration of the metallic implant presents another problem if the implant is left in the patient’s body. This problem is of particular importance with patients of a young age, where the skull growth can cause significant migration of the metallic implant, namely a plate in the brain. Furthermore, when orthopedic fixation devices made of metals are performed on patients with osteoporosis, because the strength of the bone is less than the strength of the metal, various problems such as loosening and redislocation can occur. To avoid these problems use is made of orthopedic fixation devices made of biodegradable polymers. These internal fixation devices do not have to be removed from the implantation site, do not corrode, have tensile moduli close to those of the bone, transfer increasing load to the bone, and yet assure adequate fixation of bony fragments (1993, EP0530585 A2, SYNTHES AG). Compared with conventional metal fixation devices, bioabsorbable polymer-based fixation devices have the advantages of causing no long-term implant palpability, no long-term temperature sensitivity, no stress shielding, and no interference with postoperative diagnostic imaging. These advantages may lead to better bone healing, reduced cost, elimination of need for a subsequent surgery for implant removal, and fewer complications from infections. Biopolymers such as PGA, PLA, PLGA, and P4HB have been used in the production of internal fixation devices, such as plates, screws, pins, and rods to hold bone together following surgery, or to repair broken bones. Other polymers, such as PDO, have also been considered for use in the manufacture of surgical internal fixation devices. The use of bioabsorbable polymers for construction of internal fixation devices is the subject of numerous publications [31–36]. EP0108635 A2 (1984), US4550449 A (1985), and US4905680 A (1990) of JOHNSON & JOHNSON ORTHOPAEDICS disclose bone fixation devices (e.g., plates, wires, rods, pins, staples, cable ties, and clips) made of PLLA. The polymer has a very high molecular weight and is strong enough to be fabricated into bone plates, screws, and other internal fixation devices. The polymer will maintain its strength for a long enough period of time for the bone, onto which it is placed, to heal, and it will be absorbed by the body over an extended period of time. As the polymer is absorbed, the bone plate will lose its

7: Medical, Dental, and Pharmaceutical Applications

strength. At the same time, the bone will be healing and be capable of assuming its normal load. The main drawbacks of bioabsorbable polymerbased fixation devices, such as screws and plates are their relatively poor mechanical properties, and frequently, early loss of stability of fracture fixation. The latter is mainly observed with implants produced using methods which lead to at least partial orientation of a polymer. This seems to be due to loosening of the contact between the screw head and the plate hole, resulting from nonuniform swelling of the fixation elements in the body fluids, and the premature erosion of the screw thread, which in the majority of designs is a copy of the metallic screw thread (1993, EP0530585 A2, SYNTHES AG). Moreover, forming bioabsorbable bone screws for the fixation of plates to facial and cranial bones is even more difficult because of their reduced size. To achieve the desired strength, the bioabsorbable screws must be larger in size than a comparable metallic screw. Furthermore, current bioabsorbable polymer-based orthopedic fixation devices do not actively promote bone healing and regrowth, leaving voids in the tissue once the implanted device is fully degraded (2011, WO2011041714, UNIV DREXEL). It has also been observed that tissue response to bioabsorbable implants fabricated from hydroxyacid polyesters such as PGA, PLA, and the like is not uniformly acceptable [37]. These bulk-eroding materials break down over time due to hydrolysis to produce watersoluble, low molecular weight fragments. The fragments are then attacked by enzymes to produce lower molecular weight metabolites. Acid fragments that are produced during degradation of the polymer backbone have shown to cause local tissue inflammation [38]. The inflammation has been observed in vascular systems as well, and the extent of inflammation depends on the pH of the acid that in turn is dependent on the type and amount of acid produced during degradation. Other orthopedic studies have also documented an inflammatory reaction following implantation of PGA or PLGA orthopedic fixation devices [37,39]. This is particularly true when there is a large volume of material to be degraded. More particularly, a localized, sterile inflammatory response can initiate a cascade of biological events leading to osteolytic reactions, which are radiographically detectable, and which compromise local bone quality (2002, US2002123751 A1, MEDICINELODGE INC.). This inflammation is not typically observed in bioabsorbable polymers that degrade by

305

Figure 7.3  Side view of one of the embodiments of the bioabsorbable bone screw (1993, EP0530585 A2, SYNTHES AG). 1, Threaded shaft portion; 2, Pointed end; 3, Head portion; 4, Corrugated surface; 10, Bone screw.

surface erosion, such as POEs and polyanhydrides, as the amount of acid released at a given time is small to cause tissue inflammation (2005, EP1600182 A1, CORDIS CORP.). EP0530585 A2 (1993, SYNTHES AG) discloses a bioabsorbable bone screw (10) and a plate system with self-locking properties for fixing bone fractures. The bioabsorbable screw comprises a threaded shaft portion (1) for insertion into the bone and a head portion (3), or rigid connection in the screw hole of a bone plate; the diameter of said head portion (3) increases in the direction opposite to the shaft portion (1), and the outer surface of said head portion (3) being provided with a three-dimensional structure (4) in the form of corrugations (see Figure 7.3). Preferred bioabsorbable polymers are PLA, PGA, PDO, PLGA, PGTMC, and poly(ester amides). US5584836 A (1996, SMITH & NEPHEW RICHARDS INC.) discloses a bioabsorbable suture anchor of improved resistance to shear forces that arise during the torquing of the screw to affix a medical implant or graft to bone structure of a patient, the screw comprising: an elongate body (1), comprising a biodegradable polymer, the body having a longitudinal axis and an outer surface comprising means for engaging sides of a hole wherein the elongate body is adapted to fit; one end of the elongate body comprising means for securing a suture; and at least two cannulae (4) in the elongate body (1), said cannulae offset at a preselected distance from the longitudinal axis of the body, said cannulae sized for receiving means for applying torque to the body (see Figure 7.4).

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Figure 7.4  End (a) and side (b) views of an embodiment of a dual cannulated interference screw (1996, US5584836 A, SMITH & NEPHEW RICHARDS INC.). 1, Screw body; 2, Screw thread; 3, Top-end; 4, Two cannulae. Figure 7.6  Side view of an orthopedic fixation screw (2002, US2002123751 A1, MEDICINELODGE INC.). 5, Orthopedic fixation screw; 10, Body portion; 15, Head portion; 20, Shank; 25, Screw threads.

Figure 7.5 Biodegradable spacer for bone joining screw (1998, JPH1085232 A, TAKIRON CO). A, Spacer (washer); 1, Screw; 1a, Screw head; 2, Fractured bone; 3, Hole; 3a, Tapered seat.

The biodegradable polymer is selected from the group consisting of PLA, PGA, PLGA, PCL, PDO, POEs, poly(alkyl carbonates), and polyanhydrides. JPH1085232 A (1998, TAKIRON CO) discloses the use of a metallic screw with a biodegradable spacer (washer) (see Figure 7.5). The spacer is made preferably of a biodegradable polymer selected from PLA, PLGA, and PLCL. The screw and spacer are used to transfix a fracture; as the spacer resorbs, the compression is relieved so that the bone carries progressively more of the load. The spacer improves fixation or coupling strength and allows fractured bone to recover without causing pain to patient. US2002123751 A1 (2002, MEDICINELODGE INC.) discloses an orthopedic fixation screw comprising a body portion (10) adapted to extend below the surface of a bone and a head portion (15) adapted to extend above the surface of the bone (see Figure 7.6). The head portion is formed from a bioabsorbable polymer selected from PLA, PGA, PTMC, PDO, PCL,

and copolymers, preferably from POEs and polyanhydrides. The body portion is formed out of bone, allograft bone, or bioceramic (e.g., Bioglass™). The bioceramic is osteoconductive and osteoinductive. WO03020330 A2 (2003, GOGOLEWSKI SYLWESTER) discloses an implantable bioresorbable medical device, for example, a bone fixation plate, comprising a terpolymer having l-lactide (75–90%, more preferably 85%), d-lactide (4–11%, more preferably 5%), and glycolide (4–18%, more preferably 10%) repeat units. Biodegradable screws are becoming more prevalent in medical procedures because they can eliminate the need for a second removal operation after a first implantation operation, reduce stress shielding at the fixation site, reduce the opportunity for hardware migration, and also reduce or eliminate postoperative artifact imaging. Biodegradable screws, including interference screws, can be used in the fixation of soft tissue. Such screws can be used, for example, to fix soft tissue grafts to bone during cruciate ligament reconstruction surgeries of the knee. US3463158 A (1969), US3620218 A (1971), and US3739773 A (1973) of AMERICAN CYANAMID CO disclose fixation devices like rods, screws, plates, and cylinders made of PGA. However, PGA deteriorates quite rapidly after implantation into the organism. It has been observed that very clear signs of degradation have occurred within 2 weeks. Consequently, PGA cannot provide any absolute assurance of reliable support, i.e., maintain excellent mechanical properties, and in particular resistance to shock, for a long enough time to guarantee repair of the skeleton parts of heavy mammals.

7: Medical, Dental, and Pharmaceutical Applications

FR2439003 A1 (1980, ANVAR) discloses a fixation device constituted of a matrix of PLA or PLAbased copolymer containing reinforcing elements made of PGA or PGA-based copolymer. PLA is bioabsorbable, but much more slowly than PGA. In particular, it maintains a great part of its mechanical properties for at least 2 months. PLA is also less crystalline than PGA and has, therefore, better shock resistance. Therefore, by reinforcing a PLA matrix with PGA threads or other reinforcements, maximum benefit is derived from their respective properties, while avoiding the drawbacks. PLA contributes sufficient basic resilience and a good in vivo stability, while PGA reinforces the matrix without being impaired or damaged during casting, which is most important, and it is protected from the living milieu where it is easily attacked. Further it has been noticed that when during the degradation of PGA reinforcements initially embedded in the matrix was becoming apparent, the degradation rate of the remaining matrix was increased. WO2011041714 A (2011, UNIV DREXEL) discloses in one of its embodiments a surgical fixation screw for orthopedic surgery comprising an interconnective porous shell, wherein said shell comprises a nanocomposite material comprising surface-functionalized nanodiamonds and at least one biodegradable polymer, and a hollow core within said shell, wherein said hollow core is closed on the posterior end of said screw and open on the anterior end of said screw. The biodegradable polymer is selected from the group consisting of PLLA, PDLLA, PLDLLA, PGA, PLGA, PDLLA with bioactive glass, PLGA with bioactive glass, PLLA with β-TCP, PLLA with hydroxyapatite, PDO, PCL, PCL with alginate, P3HB, N-vinyl pyrrolidone copolymers, POE, chitosan, and hyaluronic acid. The biodegradable biocompatible polymer forms a covalent bond with the surface-functionalized nanodiamonds. The covalent bond is an amide bond formed between a carboxylate group of the biodegradable polymer and an amino group of the surface-functionalized nanodiamonds. The nanocomposite material is claimed to have improved mechanical property and stimulate regeneration of original biological material that is being replaced, thus avoiding cumbersome and problematic follow-up procedures. EP0052998 A1 (1982, HOWMEDICA) discloses a construction in which a layer of biodegradable polymer is disposed between a bone plate

307

Figure 7.7  Cross-sectional view of a fixation device (1987, WO8700419 A, MINNESOTA MINING & MFG). 1, Fixation device; 2, Steel bone plate; 3, Bone plate spacer; 4, Screws; 8, 9 bone portions; F, Fracture.

and the bone surface. As this layer of biodegradable polymer degrades over time, the load carried by the bone increases and the load carried by the bone plate decreases. The biodegradable polymer is preferably selected from PLA, PGA, and PLGA. WO8700419 A (1987, MINNESOTA MINING & MFG) discloses a bone fracture fixation device in which a bone plate spacer (or washer), comprising a blend of a bioabsorbable polymer and a nonbioabsorbable polymer, is disposed between a bone plate and the bone surface (see Figure 7.7). The bioabsorbable polymer is preferably PGA, PLA, PDO, PTMC, or poly(ester amide (especially, poly[oxysuccinoyloxyhexane1,6-di(amidocarbonylmethylene)]). The nonbioabsorbable polymer is selected from polyolefins (e.g., polyethylene, polypropylene), polyamides (e.g., polyamide 6, polyamide 6.6, polyamide 12), and PUs (e.g., Biomer®, Ethicon). By proper selection of the polymers blended and the thickness of the spacer, it is possible to control the rate at which the mechanical properties of the spacer deteriorate to coincide with the rate of increase in bone strength due to healing and thereby avoid a catastrophic failure that may occur with a purely absorbable material. Examples of commercial bioabsorbable screws include: • RCI screw (Smith & Nephew); • Arthrex Bio-Interference Screw (Arthrex);

• PushLock® (Arthrex) made of PDLLA or biocomposite material consisting of PDLLA and β-TCP is a family of bioabsorbable knotless suture anchors indicated for use in arthroscopic stability procedures of the femoroacetabular joint;

308

• SutureTak® (Arthrex) made of PDLLA or BioComposite SutureTak® made of PDLLA and βTCP is a family of bioabsorbable suture anchors indicated for use in arthroscopic stability procedures of the femoroacetabular joint;

• RFS™ Screw System (Tornier) made of oriented PLGA (l-lactide/glycolide = 85/15). Examples of commercial bioabsorbable plates include: • Rapidsorb® plates (DePuy Synthes CMF) made of PLGA (l-lactide/glycolide = 85/15);

• Inion system (Stryker Leibinger GmbH & Co. KG.) composed of PLLA, PGA, PDLLA, and PTMC [5]; • Lactosorb® Trauma Plating System (Biomet, Inc.) made of PLGA (l-lactide/glycolide = 82/18);

•  Resorb X™ (KLS Martin) made of PDLLA (amorphous). Biodegradable pins, including bone-filling augmentation material, are used for bone and soft tissue fixation. Such devices have been used, for example, to stabilize wrist, foot, ankle, hand, elbow, shoulder, and knee fractures. Examples of commercial devices include: • Bioresorbable Pins® (Teknimed) made of PLDLLA (l-lactide/d,l-lactide = 70/30);

• OTPS Biodegradable Pins™ (Inion) made of PLTMC (l-lactide, d-lactide, and TMC); • Veofix resorbable pins (MedicMicro) made of PLLA; • RFS™ (Resorbable Fixation System) Pin System (Tornier) made of oriented PLGA (l-lactide/ glycolide = 85/15); •  Orthosorb™ pins (Biomet Orthopedics; formerly, Depuy Orthopaedics/Johnson & Johnson) made of PDO. Another surgical procedure for treating bones, which have been fractured through trauma or disease, is a method in which the bone marrow is scraped out, a structural framework furnished with a sheathing formed from a bioabsorbable polymer is inserted therein, and once inserted the structural framework is expanded and packed with bone cement (2005,

Biopolymers: Applications and Trends

WO2005112804 A1, MYERS SURGICAL SOLUTIONS LLC). However, it is argued that the sheath of the aforementioned patent application is manufactured of polymers such as collagen, and PLA fabric, and lacks pliability, making close adhesion to bone a problem (2013, WO2013146788 A1, UNIV NAGOYA NAT UNIV CORP.; NAT UNIV CORP. NAGOYA INST TECH). Furthermore, if the bone cement at a packing site leaks out from the bone and flows into blood vessels, there is a risk of causing a pulmonary embolism. To solve this problem, during packing of bone cement into a bone defect region such as a bone fracture site or the like, JP2000262609 A (2000, MITSUBISHI MATERIALS CORP.) discloses a biodegradable polymer encapsulating the bone cement paste in order to prevent the bone cement from leaking, and also to prevent delayed curing time. The biodegradable polymer is selected from fibrin sheets, collagen sheets, and homopolymers or copolymers of PLA and PGA. US2011245922 A1 (2011, NAGOYA INST TECHNOLOGY) discloses a material for filling bone defects having a flocculent three-dimensional structure comprising a fibrous substance containing a biodegradable polymer as a principal component, and further containing a siloxane, wherein the fibrous substance has an average diameter of 10 μm or more and 100 μm or less. The biodegradable polymer is selected from PLA, PLGA, PCL, or a natural polymer such as fibrin, collagen, alginic acid, hyaluronic acid, chitin, chitosan, or the like. However, it is argued that the biodegradable polymers used in the last two patent applications lack excellent elongation, and therefore cannot withstand the pressure produced during injection of the bone cement paste, thus presenting the problem of a risk of the bone cement leaking out from the three-dimensional structure and, as a result, flowing into blood vessels, as well as difficulty in packing the viscous bone cement paste into every corner of the bone (2013, WO2013146788 A1, UNIV NAGOYA NAT UNIV CORP.; NAT UNIV CORP. NAGOYA INST TECH). WO2009054851 A1 (2009, CELONOVA BIOSCIENCES INC.) discloses an expandable device comprising PPHO that may be used for the treatment of diseased or injured bone tissue, for example, osteoporosis. The device is inserted into the interior area of the bone such as a cavity in order to internally support the bone tissue during treatment.

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309

7.2.7  Bone Dowel Devices Bone dowels are rod-shaped fasteners that may be installed in a bone to hold fragments of fractured bones together during the healing process. Bone dowels are commonly used in spinal fusion procedures for a variety of reasons, for example, to treat patients with degenerative disc disease, deformities, as well as those involved in traumatic injuries. In posterior fusions, bone is typically removed from the hip area and placed in a traverse direction between adjacent vertebrae, often with the aid of spinal instruments. The instrument helps to hold the spine together so that a bone fusion can occur. Fusions in the lumbar area also can be done anteriorly, wherein the disc is removed and bone graft is placed between the two adjacent vertebral bodies. The bone dowel is then filled with a small amount of bone from the patient’s hip. A tapping device is then used to create screw threads in the vertebral bodies that will be fused by the bone dowel. The bone dowel is then screwed into place between the vertebrae (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.). Biodegradable polymers can be fashioned into dowels for spinal or other orthopedic repair. DE4106823 C1 (1992, LIEBSCHER KUNSTSTOFFTECHNIK; ARTHREX MED INSTR GMBH) discloses a bone dowel assembly for anchoring a suture. The dowel assembly is inserted into a hole drilled into the bone in order to affix a suture by means of which the tissue can be affixed to the bone. Preferably, the bone dowel assembly is made of a bioabsorbable polymer, such as PLA or PDO (see Figure 7.8). DE3445711 A1 (1986) and US5522894 A (1996) of DRAENERT KLAUS disclose a bone replacement material in the form of a bone dowel comprising a three-dimensional supporting structure having elementary bodies, which are substantially spherical, and held together so that each elementary body is held in rigid contact with at least three adjacent elementary bodies to define enclosed spaces which provide for bone in-growth and penetration. The elementary bodies are made preferably of a bioabsorbable polymer such as a polypeptide, PLA, PGA or one of their copolymers, gelatin, collagen or a calcium compound as the matrix, and preferably highly porous tricalcium phosphate, or hydroxyl apatite particles, or particles from a related calcium compound as fillers. Figure 7.9 shows an example of a bone dowel (10) to anchor bone screws made of the bone replacement material including substantially spherical elementary bodies (18) having a size

Figure 7.8  Bone dowel assembly (1992, DE4106823 C1, LIEBSCHER KUNSTSTOFFTECHNIK; ARTHREX MED INSTR GMBH). 1, Bone dowel assembly; 2, Dowel shank; 3, Borehole; 4, longitudinal slots; 5, Expansion cone; 6, End of expansion part; 7, End of dowel shank 2; 8, First aperture; 9, Ribs; 10, Free end of shank 2; 11, longitudinal axis; 12, Free end of expansion cone 5; 13, Second aperture; 14 and 15, Mutually opposite tines or reeds; 18, Suture.

Figure 7.9 Bone dowel (1986, DE3445711 A1; 1996, US5522894 A, DRAENERT KLAUS). 10, Bone dowel; 12, Rings or bulges; 14, Semispherical front end; 16, Cylindrical end section; 18, Spherical elementary bodies; 19, Network sheathing the matrix.

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of at least 200 μm and being together such that each elementary body is held in rigid contact with at least three adjacent elementary bodies to define enclosed spaces there between, which provide for bone in-growth and penetration, wherein the elementary bodies are provided and coated on the bone dowel. The bone dowel (10) has five rings or bulges (12), a nearly semispherical front end (14), and a cylindrical end section (16). Spherical elementary bodies (18) form the matrix structure consisting of spheres which is sheathed by network (19). After being incorporated into bone, the dowel (10) sheathes and/or reinforces the bone. The bone substitute is useful as a full prosthesis, as anchoring point of an articular prosthesis, or as coating on a bone or joint implant for treating bone defects or for replacing joints, or as bone pin for anchoring bone screws, or as seal for marrow cavities for closing the marrow cavity in amputation. Stable, strong implants are achieved and new bone formation is induced.

7.2.8  Ligaments and Tendons Ligaments and tendons are bands or sheets of fibrous connective tissue (“connective cords”), which provide support and stability to the musculoskeletal system. They have similar anatomical structures, but different biological functions. Both serve as loadbearing structures, but tendons attach muscle to bone, while ligaments attach bone to bone. Occasionally, usually due to excessive stress, unnatural movement or injury, a ligament or tendon might tear, either partially or fully. Examples of often injured cords are flexor tendons of the hand and the ACL of the knee. One generally accepted method of treatment for such a condition is replacing the torn ligament or tendon with an autograft. This involves harvesting another ligament or tendon from elsewhere in the body and transplanting it at the site of the torn ligament or tendon. Although this process is often successful, it can result in some loss of mobility at the donor location, as well as various other complications such as pain and local morbidity at the donor site. Another accepted method is using an allograft, which has its own drawbacks, such as risk of infectious diseases and high costs. EP0239775 A2 (1987, AMERICAN CYANAMID CO) discloses an implantable device composed of one or more bioabsorbable polymer(s) or combinations of bioabsorbable/nonabsorbable polymer(s) for the repair or augmentation of ligaments and tendons damaged by disease or injury. The device can be used as a flat braid for in-growth and orientation of new fibrous connective

Biopolymers: Applications and Trends

tissue (e.g., ligaments, tendons), in both intra-articular and extra-articular sites by maintaining structural stability during initial healing and then undergoing at least partial gradual absorption to prevent stress shielding and allow newly formed tissue to become correctly oriented and load bearing. The bioabsorbable filaments are made of biodegradable polymer including PGTMC, PGA, PDO, PLLA, PDLLA, and copolymers or blends of these polymers. Natural bioabsorbable polymers such as regenerated collagen or surgical gut may also be used. The biocompatible, nonabsorbable components include poly(ethylene terephthalate) (PET), poly(butylene terephthalate) (PBT), poly(ether ester) multiblock copolymers, polypropylene, high tensile strength/tensile modulus polyethylene, polyamide (including polyaramide), or polyether-type PUs. Once spun into filaments, the properties of the above materials may be improved for this application by various temperature/time/stress treatments. WO9616612 A1 (1996) and US6106556 A (2000) of OMEROS MED SYS INC. disclose a splice for repair of a severed connective cord comprising a rigid or semirigid reinforcement member inserted into or over the damaged portion of an injured tendon or ligament. The tendon or ligament is connected to the reinforcement member such that the cord–member combination can immediately withstand normal tensile forces. The interconnection can be mechanical, such as by pins extending through the sleeve reinforcement member and cord. The sleeve can be bioabsorbable over a sufficiently long period of time that the cord is healed by the time the sleeve is absorbed. The sleeve is made of a blend of bioabsorbing polymers, such as PDO, PGA, PLA, or PLGA, which is selected based on the healing characteristics of the particular connective cord repaired and the dimensional requirements for the splice in order to achieve the desired strength and bioabsorbing properties. In addition, the sleeve and/or the components securing it to the cord can be coated or impregnated with an agent or agents to enhance healing or decrease adhesion or scar formation such as hyaluronic acid, angiogenic factors, growth factors, and/or collagenase inhibitors. Such agents can immediately diffuse into the body directly adjacent to the repair and/or be released over time as the sleeve is absorbed. WO2010134943 A1 (2010, SOFT TISSUE REGENERATION INC.) discloses a device for ligament or tendon repair comprising biodegradable polymer fibers and biocompatible nondegradable polymer fibers, in a three-dimensional braided scaffold (see Figure 7.10). The two end sections (14a, b)

7: Medical, Dental, and Pharmaceutical Applications

of the braided scaffold are designed for attachment of the device at the site of implantation and are designed to allow bone cell in-growth, and one or more middle regions (12) are designed to allow ligament or tendon cell in-growth. The polymer of the middle section degrades more rapidly or the middle section is more porous than the end sections, wherein the device provides articular tissue function at the site of implantation as tissue integrates into the device. The

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biodegradable polymer fibers are made up of polymers selected from the group consisting of PLLA, PDLLA, PLGA, POEs, polyanhydrides, PPHO, PCL, PHAs, biodegradable PUs, polyanhydride-coimides, PPF, PDO, polysaccharides, collagen, silk, chitosan, and celluloses. Commercial bioabsorbable orthopedic devices for tendon repair are: • BioFiber™, an orthopedic scaffold made of a leno weave of P4HB monofilaments (Tornier) for tendon repair;

Figure 7.10  Perspective view of a multiregion ligament or tendon repair device (2010, WO2010134943 A1, SOFT TISSUE REGENERATION INC.). 10, Ligament or tendon repair device; 12, Middle region; 14a and 14b, End sections.

• BioFiber®-CM, an orthopedic scaffold made of P4HB coated with bovine collagen (Tornier) for tendon repair.

US2005149033 A1 (2005, MCGUIRE DAVID A) discloses a bioabsorbable implantable device for replacing sutures in construction of a composite graft in ligament replacement surgery. In one of its embodiments the components of the device are in the shape of a rivet or a staple (see Figure 7.11). The bioabsorbable device contains one of a group of biodegradable

Figure 7.11  Examples of composite graft fasteners (2005, US2005149033 A1, MCGUIRE DAVID A). 100, Rivetshaped graft fastener 100; 101, Female component; 102, Top (female component); 103, Top (male component); 104, Hollow sheath (female component); 105, Grooved (or toothed) shaft (male component); 110, Staple-shaped graft fastener; 111, Female component; 112, Top (female component); 114, Hollow sheaths (female component); 115, Grooved (or toothed) shafts (male component); 119, Male component; 120, Circle-shaped graft fastener; 122, Hollow female component; 123, Male component; 130, Semicircle or U-shaped graft fastener; 131, U-shaped graft fastener; 131′, U-shaped graft fastener; 132, Hollow female component; 132′, Hollow female component; 133, Male component; 133′, Male component.

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polymers comprising PDLA, PLLA, PGA, and PDO. The bioabsorbable implantable devices can be used for securing tendon grafts to bone blocks and for holding together the fibers of the tendon graft when the bone–tendon–bone graft is inserted into a patient during surgery.

7.2.9  Surgical Mesh A surgical mesh is a sterile woven material designed for permanently implantation within the body during open or laparoscopic procedures. A wide range of mesh implants is available with two main functions: to stabilize and strengthen soft tissue defects and to act as a sling to support prolapsed organs and viscera. In tissue repair applications, the mesh serves to mechanically strengthen the weakened area while simultaneously promoting long-term stability by acting as a scaffold for new tissue growth. It is most commonly used in pelvic organ prolapse, urinary stress incontinence, and hernia repair. Biodegradable surgical meshes are used in the treatment of hernias where the connective tissue has ruptured or as a sling material to support the repositioning and support of the bladder neck for female urinary incontinence. Such meshes (plugs) may also be used as soft tissue implants for reinforcement of soft tissue, for example, in the repair of abdominal aponeuroses and the abdominal wall, fascial and capsular defects, patellar and Achilles tendons, and replacement of infraspinatus tendons and cranial cruciate ligaments. Other uses include the bridging of fascial defects, as a trachea or other organ patch, organ salvage, slings (including an intestinal sling), dural grafting material, wound or burn dressing, and as a hemostatic tamponade. EP0159502 A2 (1985, AMERICAN CYANAMID CO) discloses a knitted surgical mesh-made fibers or PGA or PGTMC filaments. EP0714666 A1 (1996, ETHICON INC.) discloses in one of its embodiments a coated biocompatible substrate, such as a surgical mesh comprising a biocompatible, absorbable aliphatic polyester such as PGA, PLA, PCL, PTMC, PDO, and combinations thereof, coated with a resorbable hard tissue osteoconductive or osteoinductive calcium containing, nonfibrous, powdered compound, preferably a calcium phosphate such as hydroxyapatite, tri- or tetracalcium phosphate, or a bioactive glass, or mixtures thereof in a suitable carrier such as water, saline,

Biopolymers: Applications and Trends

water-soluble poly(ethylene glycol)s (PEGs), and combinations thereof. WO2012122215 A2 (2012, GALATEA CORP.) discloses in one of its embodiments a mastopexy implant for securing a breast of a patient in a target position, the breast having a nipple–areolar complex and an inframammary fold, the implant comprising a unitary flexible bioabsorbable mesh. An example of a suitable mesh is made of P4HB (TephaFlex®, Tepha Inc.). WO2013178229 A1 (2013, COLOPLAST AS) discloses a biocompatible nonwoven mesh comprising: (1) fibers of a biodegradable fibers material and (2) glue points in the form of domains of biodegradable polymer wherein the fibers are interconnected by the glue points. The biodegradable fiber material comprises homo- or copolymers of glycolide, l-lactide, d,l-lactide, d-lactide, meso-lactide, ε-caprolactone, 5-valerolactone, 1,4-dioxan-2-one, TMC, block copolymers of methoxy poly(ethylene glycol) or PEG and one or more of the monomers mentioned above, biodegradable PUs, biodegradable PU–urea, polyester diol-based PU, polypeptides, and PH3B. The biodegradable polymer of the glue points uses the same biodegradable polymers as above. WO2013155174 A1 (2013) and US2013317527 A1 (2013) of ETHICON INC. disclose tissue repair patches having a substantially flat or planar base member. The base member is preferably a mesh. The tissue repair patches can be made of biodegradable polymers such as PLA, PGA, PDO, PCL, PGTMC, and PLTMC. The tissue repair patch comprises additionally an adhesion barrier on at least one side of the base member, wherein the adhesion barrier comprises a biodegradable polymer selected from the group consisting of oxidized regenerated cellulose, PDO, PGCL (poliglecaprone 25), and combinations thereof. WO2006015276 A2 (2006, TEPHA INC.) discloses in one of its embodiments a surgical mesh knitted from noncurling P4HB fibers. The noncurling fibers are produced by extrusion, orientation, relaxation, and annealing the extruded filament. Table 7.5 lists various bioabsorbable meshes available in the market.

7.2.10 Slings A surgical sling is an implant that is intended to provide additional support to a particular tissue. It

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Table 7.5 Bioabsorbable Meshes Commercial name

Biopolymer

Other components

Form

Function

Company

Proceed™ Surgical Mesh

ORC

PP mesh Prolene™ Soft Mesh) encapsulated by PDO

Laminate mesh

Hernia repair and other fascial deficiencies

Ethicon Inc.

Ultrapro™ Partially Absorbable Lightweight Mesh

PGCL (poliglecaprone-25 fiber)

PP (50 wt%)

Mesh

Hernia repair

Ethicon Inc.

Parietex ProGrip™ Self-fixating Mesh

PLA

Mesh

Hernia repair

Covidien

Vicryl® (Woven Mesh)

PLGA (polyglactin 910)

Woven mesh

Hernia repair and for use as a buttress to provide temporary support during the healing process

Ethicon Inc.

Gore Bio-A® Hernia PLug

PGTMC

Bioabsorbable plug and nonabsorbable patch

Hernia repair

W.L. Gore & Associates Inc.

Bard® Sepramesh IP Composite

PGA fibers; chemically modified sodium HA, and CMC (coating)

PP fibers; PEG (coating)

Composite

Hernia repair

Bard Davol Inc.a

PVP™ Device

PDO

PVP

Laminate mesh path

Hernia repair and other fascial deficiencies such as those caused by trocar use

Ethicon Inc.

Brennen Medical Surgical Mesh Glucamesh®/ Glucatex®

β-Glucan (glucan II)

Mesh impregnated with β-glucan

Hernia repair

Genzyme Corp.b

Phasix® Mesh; Phasix®Plug and Patch

P4HB

Knitted mesh

Hernia repair; hernia repair in the groin

CR Bard

Continued

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Biopolymers: Applications and Trends

Table 7.5  Bioabsorbable Meshes—cont’d Commercial name

Biopolymer

Other components

Form

Function

Company

GalaFLEX Mesh

P4HB

Knitted mesh

Reconstructive surgeries: superficial musculoaponeurotic system (SMAS) rhytidectomy (face lift), brow lift, and platysmaplasty (neck lift)

Galatea Corp.

Tornier® Surgical Mesh

P4HB

Knitted mesh

Hernia repair and fascial defects

Tornier

TephaFLEX® Mesh; TephaFLEX® Surgical Mesh; TephaFLEX® Light Mesh; TephaFLEX® Composite Mesh

P4HB

Knitted mesh

Hernia repair and fascial defects; vaginal prolapsed repair, colon and rectal prolapse repair, reconstructions of the pelvic floor, and sacral colposuspension; hernia repair and fascial defects

Tepha Inc.

Knitted mesh Knitted mesh TephaFLEX® knitted surgical mesh/ TephaFLEX® film

CMC, Carboxymethyl cellulose; HA, Sodium hyaluronate; ORC, Oxidized regenerated cellulose; PDO or PDS, Poly(p-dioxanone); PEG, Poly(ethylene glycol); PGCL, Poly(glycolide-co-ε-caprolactone); PGTMC, Poly(glycolide-co-trimethylene carbonate); PLA, Poly(lactic acid), polylactide; PLGA, Poly(lactide-co-glycolide); PP, Polypropylene; PVP, Poly(vinyl pyrrolidone). aThe Genzyme Biosurgical licensed in 2007 the Sepra technology to Bard Davol, Inc. bGenzyme Corp. acquired in 2006 the Glucan Hernia product lines from Brennen Medical, Inc.

usually consists of a synthetic mesh material in the shape of a narrow ribbon but sometimes a biomaterial (bovine or porcine) or the patient’s own tissue. The ends are usually attached to a bone. Biodegradable slings can be used as implants to reinforce soft tissue where weakness exists. Examples of such procedures include pubourethral support and bladder support, urethral and vaginal prolapse repair, reconstruction of the pelvic floor, and sacrocolposuspension. The device can be used to treat female urinary incontinence resulting from urethral hypermobility or intrinsic sphincter deficiency. Examples of such state of the art devices include the Mentor Suspend™ Sling (Mentor Corp.). WO2007145974 A2 (2007, BOSTON SCIENT SCIMED INC.) discloses mesh materials adapted for use in an implantable sling. The mesh materials include biodegradable and nondegradable polymers that may

be adapted to facilitate scar tissue in-growth as the biodegradable components degrade (see Figure 7.12). Suitable bioabsorbable polymers include PLA, PGA, PLLA, poly(amino acids), and their copolymers.

7.2.11  Vascular Grafts and Patches The application of patches or grafts in cardiovascular surgery is a widely accepted surgical technique for repair or reconstruction of cardiovascular structures. Vascular grafts can be of either tubular or flat sheet forms. Tubular vascular graft repair materials are used to replace or bypass entire lengths of tubular veins or arteries while the flat sheet materials, conventionally called vascular patches or cardiovascular patches, are used to replace only a portion of the circumference of a vein or artery.

7: Medical, Dental, and Pharmaceutical Applications

Figure 7.12 Implantable supportive sling including both nonbiodegradable and biodegradable fibers (2007, WO2007145974 A2, BOSTON SCIENT SCIMED INC.). 100, Implantable supportive sling; 100a, Expanded section; 102a–102d, Nonbiodegradable sling fibers; 104a–104f, Biodegradable sling fibers.

The preferred procedure involves the use of an autograft, which entails a second traumatic surgical procedure to harvest a suitable artery or vein from the patient. In some cases, the harvested vessels can be unsuitable for use, and in other cases there can be a shortage of harvestable autografts particularly if the patient has previously had the same operation. When autologous vessels are not available, prosthetic vascular grafts made of synthetic polymers are often used. The use of prosthetic vascular grafts avoids the disadvantage of performing a second surgical operation to harvest a natural patch from another vessel in the body. Moreover, prosthetic vascular grafts have sufficient strength to avoid the dilation and rupturing problems associated with natural patches. The synthetic polymers most commonly used to form prosthetic vascular grafts are nonbiodegradable polymers such as polytetrafluoroethylene (PTFE), PET, and PU. The limitations of these nonbiodegradable polymers are well known. The implant becomes a permanent foreign body that has the potential to serve as a nidus for infection, and in constrained geometries this foreign body will prevent desirable tissue growth and remodeling. It is notable that the primary polymeric materials used in reconstructive procedures, PET (Dacron® cloth) and PTFE, are polymers developed for a broad array of commercial applications prior to adoption by the medical community (2008, US2008009830 A1, FUJIMOTO KAZURO L;

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TOBITA KIMIMASA; GUAN JIANJUN; WAGNER WILLIAM R; KELLER BRADLEY B). Synthetic grafts are used for applications involving the replacement of large and medium size vessels. Typically, these synthetic vessels remain open to blood flow for around 5 years before they begin to fail. Smaller diameter synthetic grafts, however, where blood flow rates are lower, generally fail rapidly, and thus are not used in procedures such as coronary artery bypass grafting (CABG), the most common open heart surgical procedure requiring smaller diameter vessels. A number of biodegradable polymers have been explored for the creation of cardiovascular patch grafting (pulmonary artery augmentation) such as P3HB [40]; cardiac patches such as poly(ester urethane) (PEU) [41], PGCL [42], and patch closure after endarterectomy such as PHAs (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.). Pericardial patches made of P3HB and PHBHV have shown very long in vivo degradation periods, of greater than 1 year for P3HB [43,44]. For many applications, this very long degradation time is undesirable as the persistence of polymer at a wound healing site may lead to a chronic inflammatory response in the patient. Slowly degrading P3HB patches used to regenerate arterial tissue have been found to elicit a long-term (greater than 2 years) macrophage response [40]. US2008009830 A1 (2008, FUJIMOTO KAZURO L; TOBITA KIMIMASA; GUAN JIANJUN; WAGNER WILLIAM R; KELLER BRADLEY B) discloses a biodegradable elastomeric patch that can be implanted on a heart or other portions of the cardiovascular system to repair tissue deficiencies or tissue damage. The biodegradable elastomeric patch may be engineered to have mechanical properties similar to those of soft tissue and to provide mechanical support to the damaged tissue. The biodegradable elastomeric patch may also comprise therapeutic agents to aid in the healing process. The biodegradable elastomeric patch comprises a polymer composition comprising one or both of a PEU urea elastomer or a poly(ether ester urethane) urea elastomer. The elastomer comprises a PCL, a PCL diol, a triblock copolymer comprising PCL, preferably a PCL-b-polyethylene glycol-b-PCL triblock copolymer, or comprises an isocyanate derivative, a triblock copolymer comprising PCL, and a diamine chain extender.

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Another approach involves the development of tissue-engineered vascular grafts formed by seeding cells onto a biodegradable polymer serving as a scaffold or patch. The main function of the biopolymer is to act as a vehicle for the delivery of cells to the damaged area, i.e., scarred tissue. Once cells are delivered to the desired region, the cells should integrate with the host tissue forming vascular neotissue. Biodegradable polymers used for cardiovascular tissue engineering are PLA, PGA, PLGA, poly(l-lactide-co-ε-caprolactone) (PLLCL), P4HB, and other PHAs. Vascular grafts particularly for CABG procedures, and below the knee grafting procedures, were developed from autologous cells seeded onto biodegradable polymers, which remain open to blood flow, as well as larger diameter grafts to improve patency rates. Tissue-engineered vascular grafts engineered from autologous cells and PLGA (polyglactin 910) functioned well in the pulmonary circulation as a pulmonary artery replacement. They demonstrated an increase in diameter suggesting growth and development of endothelial lining and extracellular matrix, including collagen and elastic fibers [45,46]. Tissue-engineered lamb heart valve leaflets (N—3) were constructed by repeatedly seeding a concentrated suspension of autologous myofibroblasts onto a biodegradable synthetic polymeric scaffold composed of fibers made of PGA and PLA [47]. WO2014044321 A1 (2014, TECNOLOGIAS AVANZADAS INSPIRALIA S L; UNIV MANCHESTER; INST CHIMII MACROMOLECULAR; USTAV EX MEDICINY AKADEMIE VED CESKE REPUBLIKY VEREJNA VYZKUMNA INSTITUC) discloses a biocompatible and biodegradable scaffold for cardiac patch comprising a porous PU base and a thick porous scaffold based on synthetic biodegradable polymer/natural polymers composites. Optionally, it can be combined with self-assembled gels composed of peptides and PUs. The scaffold is typically impregnated (seeded) with a patient’s cells before implantation. The synthetic biodegradable polymer is selected from PLA, PGA, PCL, or poly(glycerol sebacate) (PGS) and combinations thereof. The natural polymer is selected from chitosan, sodium alginate, and cellulose. The composition is useful in therapy for treating myocardial infarction, curettage, or transmural infarct, preferably myocardial infarction and cardiac tissue regeneration.

Biopolymers: Applications and Trends

Examples of biopolymers reported in the literature for cardiac patch application are listed in Table 7.6. Tissue-engineered vascular grafts have been formed by seeding autologous bone marrow cells onto a biodegradable polymer tube serving as a scaffold for the cells made of PLLCL. PGA woven fabric was used as reinforcement [48]. Although the tissue-engineered vascular grafts have shown promising functional results in early clinical trials, they have not proven physically strong enough for use in the arterial circulation (2014, WO2014008239 A1, HUMACYTE). WO2014008239 A1 (2014, HUMACYTE) discloses constructs including a tubular biodegradable scaffold made of PGA, optionally coated with extracellular matrix proteins, and being substantially acellular. The construct can further include non-PET supports at each end of the PGA tubular scaffold. The constructs can be utilized as an arteriovenous graft, a coronary graft, an arterial graft, a venous graft, a duct graft, a skin graft, or a urinary graft or conduit. The presence of sparse residual PGA fragments in extracellular matrix protein constructs at the time of implant is not of concern, as PGA is an FDA-approved biodegradable polymer with breakdown products that are readily metabolized. Furthermore, PGA has been used as a vascular graft component without any known negative effects on vascular remodeling [48]. The human cell-derived grafts produced in this study were an order of magnitude stronger than those described in previous reports that also used PGA as a support for tissue creation [49,50]. However, as noted by the inventors, these prior reports utilized human venous cells or commercially available human aorta cells at high passage [49,50]. In previous reports, use of dense PGA sutures to sew sheets of PGA into tubes left a substantial amount of residual PGA in extracellular matrix protein constructs, which diminished extracellular matrix protein construct strengths [51]. The methodology of seeding synthetic vascular grafts with autologous cells, however, is still problematic for many reasons. First, it requires an invasive procedure (biopsy) in addition to the need for a substantial period of time in order to expand the cells in culture that limited its clinical utility. This approach also faces the inherent difficulty in obtaining healthy autologous cells from diseased donors. The use of cell culture also results in an increased risk of contamination and even the potential for

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317

Table 7.6  Examples of Biopolymers Reported for Cardiac Patch Application (2014, WO2014044321 A1, TECNOLOGIAS AVANZADAS INSPIRALIA S L; UNIV MANCHESTER; INST CHIMII MACROMOLECULAR; USTAV EX MEDICINY AKADEMIE VED CESKE REPUBLIKY VEREJNA VYZKUMNA INSTITUC) Biopolymer

Comments

Application

Polyurethane

Support for cell sheet culture; biodegradable elastomer

Laminin-functionalized contention

1,3-Trimethylene carbonate copolymers

Able to sustain cyclic loads of heart muscle

Contention

Poly(ε-caprolactone)

Cell seeding

Support differentiated beating cardiomyocytes

Poly(n-isopolyacrylamide)

Nonimplantable polymer

Cell sheet engineering

Poly(glycolic acid)

Commercially available

Seeded with available extracellular matrix, support ventricular function

Collagen

Exogenous material;

In vitro electrosimulation

no improvement of ventricular function and immunogenicity issues when implanted in vivo;

Cell sheet engineering

mixed with Matrigel or glycosaminoglycan

Cell sheet engineering

Gelatin (Gelafoam)

Does not provide lead to any functional improvement of ventricular function

Support seeded beating cardiomyocytes

Alginate

Risk of necrosis and calcification

In vitro cell culture and seeded cardiac patch

Poly(ester-urethane)–collagen

Use of the mechanical properties of the synthetic components and cellular adhesion properties of the collagen

In vitro cell culture

Poly(ε-caprolactone)–collagen

Ditto

Ditto

Poly(d,l-lactide-co-caprolactone)– collagen

Ditto

Ditto

Poly(glycerol sebacate)–gelatin

No results regarding contractility or in vivo assays

Cell differentiation

dedifferentiation of the cultured cells. The use of autologous cells to seed the polymeric grafts also limits the off-the-shelf availability of tissue-engineered vascular grafts, thereby limiting their overall clinical utility (2009, WO2009089324 A1, UNIV YALE). WO2009089324 A1 (2009, UNIV YALE) discloses methods for increasing the patency of synthetic vascular grafts made from biodegradable polymeric scaffolds by administering one or more cytokines and/or chemokines that promote outward tissue remodeling of the vascular grafts and vascular

neotissue formation. The biodegradable polymers are selected from the group consisting of PLA, PGA, polyanhydrides, POEs, PUs, P3HB, poly(3-valeric acid) (P3HV), PCL, PLCL, their blends, or copolymers thereof. The disclosed biodegradable, synthetic vascular grafts do not require cell seeding to maintain patency of the grafts. This is advantageous, because it avoids problems associated with cell seeding, including the need for an invasive procedure to obtain autologous cells, the need for a substantial period of time in order to expand the cells in culture,

318

Biopolymers: Applications and Trends

the inherent difficulty in obtaining healthy autologous cells from diseased donors, an increased risk of contamination, and the potential for dedifferentiation of the cells. The disclosed biodegradable, synthetic vascular grafts are claimed to have a greater off-theshelf availability and increased overall clinical utility. WO2013154612 A2 (2013, UNIV PITTSBURGH) discloses a vascular graft including a biodegradable scaffold including (1) a biodegradable polyester tubular core made of PGS; (2) a biodegradable polyester electrospun outer sheath made of PCL surrounding the biodegradable polyester tubular core; and (3) a thromboresistant agent made of heparin coating the biodegradable scaffold, thereby forming a vascular graft. The vascular graft is used for forming a blood vessel of less than 6 mm, preferably less than 4 mm; and as a coronary or a peripheral arterial graft. Furthermore, in some embodiments, to encourage host cell migration into the scaffold, a porous scaffold is used. The biodegradable scaffold induces tissue

regeneration without cell seeding of the scaffold prior to implantation into a subject’s body. A list of exemplary patents describing vascular grafts is shown in Table 7.7.

7.2.12  Vein Valves Venous leg ulcers occur on the lower leg and are caused by venous insufficiency or poorly functioning valves in the veins of the legs. Currently, there is no treatment available to repair defective vein valves. It is desirable to provide replacement vein valves, which preferably can be implanted by a minimally invasive means, or by routine surgery. WO0180782 A1 (2001, HAVERICH AXEL) discloses the use of recipient-specific transformed natural or synthetic acellularized matrices made of a biodegradable polymer for the production of individual venous valve prostheses. The biodegradable polymer is preferably PLGA, and even more preferably a multilayered PDO. The recipient-specific transformation

Table 7.7  Exemplary Patents Describing Vascular Grafts made of Biopolymers Function

Biopolymer

Other polymer(s)

Patent

Vascular grafts, absorbable fabric in hernia repair, supporting damaged liver, kidney

PLA, PLGA

PTFE, nylon

US3636956 A (1972, ETHICON INC.)

Vascular grafts, absorbable fabric in hernia repair, supporting damaged liver, kidney

PLA, PLGA

PTFE, nylon

US3797499 A (1974, ETHICON INC.)

Arteriovenous graft, coronary graft, arterial graft, venous graft, duct graft, skin graft, urinary graft conduit

PGA

Extracellular matrix proteins

WO2014008239 A1 (2014, HUMACYTE)

Venous, arterial, artero-venous conduits (for congenital heart surgery, coronary artery bypass surgery, peripheral vascular surgery, and angioaccess)

PLA, PGA, PAs, POEs, P3HB, P3HV, PCL, PLCL, PUs

WO2009089324 A1 (2009, UNIV YALE)

Coronary, peripheral arterial graft

PGS and PCL

WO2013154612 A2 (2013, UNIV PITTSBURGH)

P3HB, Poly(3-hydroxybutyrate); P3HV, Poly(3-hydroxyvalerate); PA, Polyanhydride; PCL, Poly(ε-caprolactone); PGA, Poly(glycolic acid), polyglycolide; PGS, Poly(glycerol sebacate); PLA, Poly(lactic acid), polylactide; PLCL, Poly(lactide-co-ε-caprolactone); PLGA, Poly(lactic acid-co-glycolic acid), poly(lactide-co-glycolide); POE, Poly(ortho ester); PTFE, polyfluorotetraethylene; PU, Polyurethane.

7: Medical, Dental, and Pharmaceutical Applications

involves colonization of the matrix with recipientcompatible cells, especially autologous cells (e.g., fibroblasts of the prosthesis recipient), all natural cells preferably being removed from xeno- or allogeneic matrices prior to recipient-specific transformation. Vein valves can be derived from PHA polymers, wherein these polymers are formed into a valve structure. The polymers may be used alone, coated, or modified with another agent, such as a biological factor. They may be combined with other materials, and/or made porous. Alternatively, the polymers may be formed into scaffolds, which can optionally be cell seeded prior to implantation (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.).

7.2.13  Heart Valves The unidirectional flow of blood through the entire circulatory system is controlled by the heart’s four valves. However, natural heart valves may become dysfunctional from a variety of pathological causes such as stenosis and incompetence. A stenotic heart valve does not open fully due to stiffening of the valve tissue, thus more work is required for the heart to force blood through the valve. An incompetent valve causes inefficient blood circulation by permitting the flow of blood back into its originating chamber. Once valvular disease progresses to the point at which the heart’s ability to pump blood is significantly impaired, surgery is usually recommended to repair or replace the diseased valve. There are currently two primary types of artificial valve prostheses: mechanical heart valves and tissue heart valves. Each type has benefits and drawbacks. Mechanical valves, for example, are noted for their durability and reliability. However, a major drawback is the need for the recipient to be placed upon a lifelong anticoagulant therapy, which involves continuous monitoring of anticoagulant levels. Current tissue valves, derived from heterologous sources (cows and pigs), on the other hand, do not require anticoagulant therapy, they are quiet, provide physiological flow patterns, and typically have slowly developing rather than catastrophic failure modes. The major problem associated with these valves is their lack of durability. Most of the current tissue valves generally last between 5 and 15 years before they need to be replaced due to a gradual deterioration of the (nonliving) tissue (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.).

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EP1600182 A1 (2005, CORDIS CORP.) discloses in one of its embodiments a vascular or cardiovascular device useful as a heart valve for controlling pH levels in patient’s body comprising a structural component made of biodegradable polymer and a buffering agent on the polymer. The biodegradable polymer is a bulk-eroding polymer selected from the group consisting of poly(imino carbonates), POEs, PPHOs, and poly(ester amides). Figure 7.13 shows a cross-sectional slice taken from a sphere of a medical device having a first biodegradable polymer and a second biodegradable polymer, whose degradation is triggered by degradation products produced by degrading of the first biodegradable polymer. DE19613048 A1 (1996, GRUNZE MICHAEL) discloses the use of phosphazene-based polymers, specifically including poly[bis(trifluoroethoxy)phosphazene], to treat (coat) biological materials used for bioprosthetic heart valves. US2008086205 A (2008, CELONOVA BIOSCIENCES INC.) discloses a bioprosthetic heart valve comprising a biological tissue and a PPHO, preferably poly[bis(trifluoroethoxy)phosphazene]. The PPHO is coated, diffused, impregnated, grafted, or any combination thereof, into or onto the biological tissue. The heart valve is useful for reducing tissue calcification and imparting an anticalcification property. WO2013019756 A2 (2013, UNIV CARNEGIE MELLON) discloses in one of its embodiments an artificial heart valve composed of a conduit, a heart valve leaflet structure, and one or more conduit sinus structures incorporating one or more biodegradable structures having a first side affixed to the first conduit breach edge and a second side affixed to the second conduit breach edge. Such a heart valve structure is

Figure 7.13 Cross-sectional view of a biodegradable medical device (2005, EP1600182 A1, CORDIS CORP.). 50, Biodegradable medical device; 70, Coating; 80, First biodegradable polymer; 90, Second biodegradable polymer; 95, Buffering agent or a neutralizing agent.

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conveniently referred to as a hybrid tissue-engineered valved conduit (hybrid TEVC). The hybrid TEVC includes a conduit constructed of synthetic material and having a cross-section forming a partially closed circle, and a biodegradable structure, which is incorporated into the conduit wall to form an enclosed tubular structure. The hybrid TEVC may also include one or more heart valve leaflet structures, and one or more conduit sinus structures disposed within the conduit. The biodegradable structure in the hybrid TEVC is made of PGS encapsulated by a sheath of PCL. The biodegradable structure in the hybrid TEVC may be replaced over time by autologous tissue, thereby allowing the heart valve structure to enlarge as the patient grows. In such an embodiment, cells from a patient may migrate into the biodegradable structure over time to replace the material from which the biodegradable structure is made. Tissue-engineered heart valves, and components of heart valves such as leaflets (cusps) or supports can be made of PHAs, in particular P4HB, by constructing porous heart valve scaffolds from these polymers alone or with other materials. Preferably, these scaffolds are derived from foams and/or fibrous PHAs (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.).

7.2.14  Vascular Closure Devices There are three main categories of vascular closure devices (VCDs): collagen-based, suture-based, and staples and clips. VCDs have been demonstrated to reduce time to hemostasis, facilitate ambulation, and potentially decrease length of stay [52]. WO2009114868 A1 (2009, CORDIS CORP.) discloses a VCD comprising a porous fibrous structure formed from randomly oriented fibers such as a nonwoven PGA felt. The porous structure works very effectively to seal a puncture site with optimum porosity, absorbent capacity, and perfect anatomical fit. The plug density and other fiber properties (total denier; number of filaments, etc.) have provided an efficient structure that allows instantaneous absorption of blood during deployment. A commercial VCD is Exoseal® Vascular Closure Device (Cordis Corp.) composed of a bioabsorbable felt-like plug and a Plug Delivery System. The bioabsorbable plug is made entirely of PGA. The PGA plug material is constructed to stop bleeding from the puncture site. The Plug Delivery System is designed to position and release the bioabsorbable plug to seal

Biopolymers: Applications and Trends

the puncture site in the main artery of the thigh (femoral artery).

7.2.15  Adhesion Barriers Adhesion formation following surgery or trauma is generally considered to be undesirable. For example, adhesions that form in relation to intestinal surgery, for example, bowel resection, and hernia repair, may cause obstruction of the intestine. Adhesions that form near the bone fracture site may reduce or hinder the normal movement of the area of repair by restricting the natural movement of tendons over the adjacent bone. Adhesions may also form in the vicinity of nerves and disrupt nerve transmissions with a resultant diminution of sensory or motor function. An approach to prevent adhesion involves application of a physical barrier at the area of surgical injury. Among the materials used to form physical barrier, biodegradable adhesion barriers are particularly preferred. GB2174909 A (1987, AMERICAN HOSPITAL SUPPLY CORP.) discloses the interposing of a barrier layer of soft biological tissue, such as collagen, collagen-fabric films, collagen membranes, or reconstituted collagen or Dacron™ mesh, at the interface of a bone fracture and the surrounding tissue. US4603695 A (1986, NIPPON MEDICAL SUPPLY) discloses an adhesion prevention barrier made of a biodegradable polymer for preventing adhesion of vital tissues such as skin, blood vessels, or organs. The biodegradable polymer is selected from PLA, PGA, PDO, their copolymers, collagen, poly(amino acids), and chitin and may be placed, where there is a possibility of adhesion setting in. JP2000197693 A (2000, JMS CO LTD.; GUNZE KK) discloses a porous antiadhesion membrane for the tendon made of PLCL. The porous membrane is claimed to be capable of physically isolating the tendon and the peripheral tissue thereof, while having suitable bioabsorbability and has excellent operability, permeability of nutrient materials and safety. WO2004089434 A1 (2004, TEIJIN LTD.) discloses a honeycomb film made of a biodegradable polymer, to be used as an adhesion inhibitor for preventing postoperative adhesion. Moreover, in order to simply prepare a honeycomb structure with good reproducibility, a phospholipid may be added as a surfactant in addition to the foregoing biodegradable polymer. The biodegradable polymer is preferably selected from PLA, PLGA and PCL.The honeycomb

7: Medical, Dental, and Pharmaceutical Applications

film is claimed to have favorable handling properties and exhibit a stable and excellent effect of inhibiting adhesion over a desired period of time. Examples of biodegradable adhesion barriers devices, which are available in the market include: • Gynecare Interceed® Adhesion Barrier (Ethicon Inc.) made of oxidized cellulose for preventing formation of pelvic adhesions;

• Seprafilm™ (Genzyme Corp.) adhesion barrier is a bioresorbable membrane made of chemically modified sodium hyaluronate and carboxymethyl cellulose;

• Repel-CV® Bioresorbable Adhesion Barrier (SyntheMed Inc.) made of PLA/PEG. It prevents interconnection (adhesions) between two opposing surfaces, specifically between the chest wall and the pericardium (a membrane that covers the heart). It is used for reducing the severity of postoperative cardiac adhesions in pediatric patients, who are likely to require reoperation via sternotomy.

Physical barriers are also used to cover and protect wound sites. WO9210218 A (1992, GORE & ASS) discloses a surgical article having a bioabsorbable fibrous matrix laminarly affixed to one surface of a bioabsorbable cell-barrier sheet material. Both the cell-barrier sheet material and the fibrous matrix are made preferably of PGLA. EP0334046 A2 (1989) and US5092884 A (1992) of AMERICAN CYANAMID CO disclose a surgical composite structure having absorbable and nonabsorbable components, which may be useful for repairing anatomical defects, for example, preventing hernia formation in an infected area. The bioabsorbable polymer is preferably PGTMC. The nonabsorbable portion of the composite acts as a reinforcement material. Ingrowth of natural tissue is said to be enhanced by controlled degradation of the absorbable portion. EP0371736 A2 (1990, MITSUBISHI CHEM IND) discloses a wound-covering composition having a sheet of biopolymer and a film of PU. The biopolymer is selected from the group consisting of collagen, gelatin, alginic acid, chitin, chitosan, fibrin, dextran, and poly(amino acids). An antibacterial agent may be provided between the PU film and the sheet of biopolymer, thereby forming a three-layer woundcovering material. With the cure of the wound, it is said that the biopolymer is taken in a living tissue and

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the PU film can be peeled off from the sheet without hurting the surface of a wound.

7.2.16  Articular Cartilage Repair Articular cartilage is a tough, elastic tissue that covers the ends of bones in joints and enables the bones to move smoothly over one another. When articular cartilage is damaged through injury or a lifetime of use, however, it does not heal as rapidly or effectively as other tissues in the body. Instead, the damage tends to spread, allowing the bones to rub directly against each other, thereby, resulting in pain and reduced mobility. A number of treatment techniques for the repair of articular cartilage defects in clinical use are lavage and debridement, abrasion chondroplasty, microfracture techniques, subchondral drilling, transplantation of periosteal or perichondrial grafts, transplantation of osteochondral autografts or allografts, and autologous chondrocyte transplantation. Treatment techniques for the repair of articular cartilage defects include the implantation of biodegradable polymer matrices, alone or incorporating cell and/or bioactive molecular growth factors. Biodegradable polymers tested in vitro and in vivo for their efficacy in articular cartilage repair are shown in Table 7.8. Table 7.8  Biodegradable Polymers for the Repair of Articular Cartilage Defects Biodegradable polymer

References

Agarose

[53]

Alginate

[54]

Chitosan

[55]

Fibrin

[56]

Collagen

[57]

Gelatin

[58]

PLA, PLLA

[59,60]

PDLLA

[61]

PLCL

[62]

PGA

[63–65]

PHBHHx

[66,67]

PDLLA, Poly(d,l-lactide); PGA, Poly(glycolic acid), polyglycolide; PHBHHx, Poly(3-hydroxybutyrate-co-3-hexanoate); PLA, Poly(lactic acid); PLCL, Poly(lactide-co-ε-caprolactone); PLLA, Poly(l-lactic acid), poly(l-lactide).

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Embedding human mesenchymal stem cells (MSCs) in agarose or alginate gels could be cast into various shapes and caused substantial chondrogenesis. However, the alginate or agarose scaffolds for cartilage tissue engineering have the following drawbacks: poor cell adhesion and uncontrollable degradation of alginate following the diffusion of divalent cations into the surrounding medium. In addition, alginate matrices were used in the in vivo applications and reported to have severe foreign body giant cell reactions and immunological responses when implanted to treat full-thickness defects in cartilage in experimental animals [68]. Hence, alginate matrices have not been employed in human patients for articular cartilage repair. A large number of patents has been disclosed describing the implantation of biodegradable polymer matrices with chondrocyte cells (which generate cartilage, under proper conditions) into damaged knees and other joints. US4846835 A (1989, GRANDE DANIEL A) describes techniques for growing chondrocyte cells in vitro, seeding the cells into a collagen matrix, and implanting the matrix and cells in the knee. WO9012603 A1 (1990) and US5736372 A (1998) of MASSACHUSETTS INST TECHNOLOGY disclose biodegradable polymers for use as a matrix material instead of collagen. The biodegradable polymers are selected from the group consisting of polyanhydrides, POEs, PGA, PLA, PLGA, and blends thereof. US2003050709 A1 (2003, NOTH ULRICH; TUAN ROCKY S) discloses an in vitro-engineered osteochondral graft comprising a porous matrix block, more particularly, a porous PDLLA block, presscoated with mesenchymal stem cells (MSCs), wherein a cartilage layer is formed on the surface of the matrix block. This invention may be used for treating articular cartilage defects. US5876452 A (1999, UNIV TEXAS) discloses in one of its embodiments a osteochondral implant made of PLGA (d,l-lactide/glycolide = 50/50) and used as a controlled release delivery system for recombinant human transforming growth factor.

7.2.17  Bone Graft Substitutes Bone grafts are used in spinal fusions, trauma fractures, and in periodontal surgery. In a typical procedure, bone graft material is harvested surgically from the patient’s own hipbone and then inserted into the grafting site where bone regrowth is desired.

Biopolymers: Applications and Trends

The graft material contains a variety of bone-promoting agents, which help stimulate the formation of new bone and healing. This procedure frequently provides good results, but requires a second operation to harvest the autograft. To avoid the harvesting procedure, surgeons may use other types of bone graft substitutes including cadaver bone products and composites containing calcium phosphate and calcium carbonate. The latter materials generally do not perform well, and disease transmission issues always accompany the use of cadaver-derived materials. Because of these limitations, alternative techniques for making bone graft substitutes were developed based on the use of osteoconductive (bone scaffolding) and/or osteoinductive (new bone from biological stimulation) materials. It has become increasingly apparent that these materials require a carrier vehicle for optimum performance. JPH01232967 A (1989, OSMED INC.) discloses a bone graft substitute composed of PLA having voids which are connected to each other and filled with hyaluronic acid velour carrying an active substance such as bone morphogenetic protein (BMP). However, neither particular production example nor test example thereof is described therein, and thus its bone-forming capability is unknown. Because a PLA sponge was employed, the composition disclosed is rigid and fragile, it is poor in plasticity and elasticity, and thus insufficient in formability and workability at implantation (1996, WO9610426 A1, YAMANOUCHI PHARMA CO LTD.). US5133755 A (1992, THM BIOMEDICAL INC.) discloses a device and method for treating mammalian bone deficiencies, defects, voids, and conformational discontinuities produced by congenital deformities, osseous and/or soft tissue pathology, traumatic injuries, and functional atrophy. The device is a one piece-molded body member composed of four components: PLA, hyaluronic acid BMP, and bone-derived growth factor. Working together, the four components provide the following five biological functions prerequisite to the processes of osteoneogenesis: structural competence (PLA), chemotaxis (hyaluronic acid), electronegative field (hyaluronic acid and physical–chemical electrokinetic events), osteoinduction (BMP), and osteogenesis (bone-derived growth factor). When the single body member is implanted into a bone defect has the capacity to restore functional architecture and mechanical integrity, initiate osteoinduction and osteogenesis, and maintain the biological processes

7: Medical, Dental, and Pharmaceutical Applications

of bone formation and remodeling, while the host organism is simultaneously biodegrading the body member. The ultimate result of the functioning is formation of healthy, viable bone tissue where there was not bone before, while simultaneously, the entire device is hydrolyzed and completely metabolized by the host organism. Figure 7.14 shows the infusion of a biologically active agent (i.e., the osteoinductive agent known as BMP) solution into the device and the dispersion of this solution throughout the entire volume of the porous body member, enveloping all of the fibrils of the chemotactic ground substance velour and coating all surfaces of the structural polymer (PLA). WO9610426 A1 (1996, YAMANOUCHI PHARMA CO LTD.) discloses an osteoplastic graft comprising a bone inducer supported on a composite porous body comprising a porous structure of a bioabsorbable hydrophilic material and a surface layer of a bioabsorbable polymer. Preferably, the hydrophilic material comprises at least one member selected from the group consisting of gelatin, hyaluronic acid, collagen, chitosan, and triethanolammonium alginate and derivatives thereof, while the bioabsorbable polymer comprises at least one member selected from the group consisting of PLGA, and poly[bis(p-carboxyphenoxy)propane)-co-(sebacic acid anhydride]) [69,70]. As the graft is excellent in moldability and operability and has an internal structure suitable for in vivo bone neogenesis, bone grafting occurs not only at the periphery of the graft, but also within the graft. A number of different surgical procedures employ bone marrow transplants. In many cases, bone marrow is taken from the iliac crest and used at another location to aid in the repair of tissues and organs, usually bone. Examples include the use of bone marrow in the repair of bone fractures, such as a

Figure 7.14 Sectional view of the device being infused with a solution of biologically active agent (1992, US5133755 A, THM BIOMEDICAL INC.).

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tibial plateau fracture, spinal procedures, as well as treatment of abnormalities in the maxillofacial and craniofacial regions requiring surgery. In certain cases, large amounts of bone marrow are required for these procedures, but the amount of bone marrow available is limited, particularly in young and small patients. WO02082973 A2 (2002, UNIV TEXAS) discloses an osteoinductive protein synthesizing implant system comprising a porous scaffold or support structure formed from (1) a biodegradable polymer in combination with (2) a 3-hydroxy-3-methylglutaryl coenzyme A (HMG Co-A) reductase inhibitor. Bone cells are impregnated in the porous scaffold, with the HMG Co-A reductase inhibitor included in the scaffold in an amount effective to stimulate the formation of bone tissue from the bone cells (e.g., when the scaffold is implanted in a subject). The biodegradable polymer is preferably PLA, PGA, or copolymer thereof. The bone cells may in general be bone marrow cells, and more particularly may be osteoblasts, stem cells, progenitor cells, or combinations thereof. WO2006009452 A2 (2006, TECHNOLOGIESTICHTING STW) discloses an artificial bone implant, comprising a bioabsorbable porous matrix wherein an osteoinductive composition comprising calcium phosphate and stem cells from adipose tissue is present. The porous matrix comprises PLLA or PLDLLA; preferably PLDLLA is used. These biodegradable polymers have a tensile modulus which resembles that of vertebral bone. An additional advantage is that these materials are radiolucent, so that follow-up by means of X-rays and/or computed tomography and MRI scans is well possible. An advantage of the use of stem cells from adipose tissue over the use of stem cells from bone marrow is that, for obtaining the former, prior to the actual spinal column surgery, no separate surgical operation, i.e., a very painful and uncomfortable bone marrow collection, is necessary and neither is a labor-intensive, risky, expensive, and timeconsuming proliferation of stem cells in the laboratory for in vitro expansion of the bone marrow stem cells. The artificial bone implant is used particularly in making spinal fusions.

7.2.18  Dural Substitutes The dura mater, a thick membrane occurring between the cranial bones and brain and covering

324

the spinal cord, protects the brain and spinal cord and inhibits leakage of cerebrospinal fluid (CSF). Following neurosurgical operations, cadaveric dura mater grafts have commonly been used to repair dural defects. However, because of the risk of transmitting Creutzfeldt–Jakob disease through these grafts, the World Health Organization has recommended that cadaveric dural grafts no longer be used. Although polytetrafluoroethylene can be used as an alternative permanent synthetic material for dural repair, concerns relating to the material’s biocompatibility have been raised, increasing interest in the development of a bioabsorbable dural substitute. Artificial dural substitutes of biodegradable materials such as collagen [71] and gelatin [72] were produced, but they were not in practical use because of strength-related problems, i.e., because of the insufficiency in the suture strength to be sutured integrally with the internal dura mater. US5861034 A (1999, GUNZE KK) discloses an artificial dura mater comprising a sheet made of a biodegradable and bioabsorbable polymer, for example, PLCL and, further, an artificial dura mater comprising introducing a biodegradable and bioabsorbable polymer made of a material different from that of the sheet as a reinforcement between the sheets and integrally molding the sheets and the reinforcement. The PLCL has a molar ratio of about 25–60% of lactic acid and about 75–40% of ε-caprolactone. WO9917815 A1 (1999, GUNZE KK) involves an improvement in the artificial dura mater of the previous patent application disclosing an artificial dura mater having a three-layer structure comprising sandwiching between the sheets a reinforcement of a biodegradable polymer different from that of the sheets and integrally molding the sheets and the reinforcement. The biodegradable polymer of the reinforcing material is selected from PLCL, PLA, PGA, and PGLA. The other two sheets are made of PLCL. EP1738780 A2 (2007, CODMAN & SHURTLEFF) discloses a dural graft material comprising: at least one collagen layer defining a plurality of pores, wherein a majority of the plurality of pores have a diameter of less than about 10 μm; and at least one reinforcement layer joined to the collagen layer, the reinforcement layer being substantially fluid impermeable. The reinforcement layer is formed from a bioabsorbable polymer is selected from the group consisting of PLA, PGA, PCL, PDO, TMC copolymers, or blends thereof.

Biopolymers: Applications and Trends

There have been reports on the efficacy of a PGA mesh and fibrin glue as a substitute for dural repair for preventing elevated CSF leakage in spinal surgery [73].

7.2.19  Spinal Fusion Cages Spinal fusion cages are used to treat various forms of degenerative disc disease, a condition in which the spinal discs, located between each vertebra, are no longer able to cushion and protect the vertebra during movement. This can result in severe, and occasionally, crippling back pain, as the vertebrae rub against adjacent spinal nerves. The condition results from the wearing down of the shock-absorbing cartilage that separates the vertebrae of the spine, and can be due to aging or injury. Degenerating discs also become dehydrated losing height, and thereby bringing the vertebrae closer together. Surgeons are using a relatively new procedure involving spinal fusion cages to fuse two or more vertebrae into one stable bony mass. In this procedure, a cage, which comprises a hollow cylinder, is implanted in the disc space, following removal of the defective disc, and packed with bone graft material. Fusion occurs as new bone grows into the fusion cages through holes in the cylinder. The cages also serve to restore disc space height while the spine heals. EP1138285 A1 (2001, IMPLANT DESIGN AG) discloses a surgical cage for insertion between two vertebrae, comprising a closed frame with an internal cavity for receiving a bone replacement material made of PLDLLA. WO2006009452 A2 (2006, TECHNOLOGIESTICHTING STW) discloses in one of its embodiments the use of osteoinductive compositions comprising stem cells from adipose tissue and osteoinductive calcium phosphates in PLLA and PLDLLA cages as part of an artificial bone implant for stabilization and renewed alignment of spinal column segments (see also Section 7.2.17). EP1591133 A1 (2005, SYNTHES) discloses an implant comprising an abrasion-resistant, bioabsorbable coating containing a biodegradable polymer and a substance (e.g., statin) inducing the secretion of bone and/or cartilage growth stimulating peptides. The biodegradable polymer is preferably PDLLA. The implant is a fracture- and/or osteotomy-fixation device selected from the group consisting of plates, screws, nails, pins, wires, threads, and cages. Example: Method for making spinal fusion cage with a coating that induces the secretion of BMP-2.

7: Medical, Dental, and Pharmaceutical Applications

15 wt% of simvastatin is added to a solution of 100 mg/ml of PDLLA with an inherent viscosity of 0.5 dl/g in ethyl acetate. The clear solution is filtered with a 0.2 μm filter and transferred into a cooled vessel. A polyetheretherketone cage for spinal fusion procedures is dipped into the solution and extracted at a controlled speed. In order to avoid the accumulation of solution in the internal spaces of the implant, the implant is spun around its central axis immediately after extraction from the coating solution. After a drying time of at least 2 min this procedure is repeated. The procedure yields an implant with a coating that has the ability to stimulate the secretion of bone growth factors in osteoblasts. The thickness of the coating is preferably 5–10 μm. The obtained implant can be used for spinal fusion procedures, and bone formation is increased by the release of simvastatin from the coating. Implantation occurs by the standard surgical procedure used also for uncoated spinal cages.

7.2.20  Nerve Guides Biodegradable devices may be used as guides (conduits) to facilitate the regrowth and reconnection of severed or damaged nerves. The devices are generally fabricated as tubes [74]. Various biopolymers have been tested as candidates for nerve channel guides, and some have been used clinically. The tested biopolymers include a number of aliphatic polyesters such as PLLA [75,76], PDLLA [77], PGA [78–82], PLGA (polyglactin 910) [83], PCL [84], and PLCL [85]. It has been reported, however, that there are significant shortcomings with devices prepared from these materials (1988, WO8806866 A1, UNIV BROWN RES FOUND INC.); they include inflammatory responses, formation of scar tissue, and loss of sensory or motor function. Several bioabsorbable nerve guides are approved by the US Food and Drug Administration (FDA) and Conformity Europe (CE) for clinical repair of peripheral and cranial nerves. The available FDA- and CEapproved absorbable nerve guides are two collagen and two aliphatic polyester-based guides [86]. Two companies, Integra Lifesciences and Neuroregen, LLC, have commercialized nerve channel conduits made of collagen (NeuraGen® Nerve Guide) and PGA (GEM Neurotube®), respectively to bridge small nerve gaps. GEM Neurotube® is the first biodegradable nerve guide on the market cleared for use in the US and EU [87]. The GEM NeuroTube® is an absorbable woven PGA mesh tube, which is

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designed for single use in patients with an injury to a peripheral nerve in which the nerve gap is ≥8 mm, but ≤3 cm [78–82]. The GEM NeuroTube®creates a tensionless repair and offers return of sensation with no donor-site morbidity [87]. NeuraGen® Nerve Guide is an absorbable collagen tube designed to be an interface between the nerve and the surrounding tissue and to create a conduit for axonal growth across a nerve gap. The level of functional recovery achieved with the NeuraGen® Nerve Guide is equivalent to direct suture repair [88]. There were no reports of scar tissue or inflammatory response and episodes of compression or neuropathy of this material in animal studies [89,90]. The structural stability of the NeuraGen® Nerve Guide is increased with a crosslinking agent that also determines the rate of in vivo absorption [91]. P3HB has also been investigated as a material for nerve regeneration containing growth factors and Schwann cells to prevent nerve cell death and promote regeneration [92–94]. WO8806866 A1 (1988, UNIV BROWN RES FOUND INC.) discloses tubular piezoelectric nerve conduits including a device formed from P3HB. WO03041758 A1 (2003, WIBERG MIKAEL) discloses a nerve repair unit comprising P3HB and an alginate matrix containing human Schwann cells. WO0154593 A1 (2001, GEN HOSPITAL CORP.; CHILDRENS MEDICAL CENTER; MASSACHUSETTS EYE & EAR INFIRM) also discloses P3HB conduits that include Schwann cells. WO2005020825 A1 (2005, TEPHA INC.) discloses a nerve regeneration device comprising P4HB in the form of a porous conduit. Growth factors, drugs, or cells that improve nerve regeneration may be incorporated into the device. The device is claimed to be flexible, strong, not crushing the regenerating nerve, easy to handle, reduce surgical time by eliminating the need to harvest an autologous graft, and allow the surgeon to repair the nerve without a prolonged delay. Copolymers of 1,3-trimethylene carbonate (TMC) and ε-caprolactone have also been investigated for the preparation of porous artificial nerve guides. Structures comprising very high molecular weight TMC (Mn = 290, 000 and MW = 552, 000) exhibited improved mechanical properties probably due to strain-induced crystallization (Tm = 36 °C) [95].

7.2.21  Bulking and Filling Agents Bulking and filling agents are commonly used for augmenting dermal support, at the site of bone

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fractures for wound healing and bone repair, and within sphincter tissue for sphincter augmentation (e.g., for restoration of continence). A typically used material is collagen (1994, US5306500 A, COLLAGEN CORP.). Available dermal fillers comprise biodegradable natural substances (e.g., collagen, gelatin, hyaluronic acid, dextran, and dried acellular particulate dermal matrix) and biodegradable polymers derived from renewable resources (e.g., PLLA and carboxymethyl cellulose). WO9003768 A1 (1990, SOUTHERN RES INST) discloses the production and use of biodegradable polymers for providing syringeable, in situ forming, solid, biodegradable implants. Two types of biodegradable polymeric systems are described: thermoplastic polymers dissolved in a biocompatible solvent and thermosetting polymers that are liquids without the use of solvents. Preferred biodegradable thermoplastic polymers are PDLLA, poly(d,l-lactide-co-glycolide) (PDLGA), and poly(d,l-lactideco-ε-caprolactone) (PDLLCL). Preferred solvents are N-methyl-2-pyrrolidone, 2-pyrrolidone, dimethyl sulfoxide, and acetone because of their solvating ability and their compatibility. In one envisioned use of the thermoplastic system, the polymer solution is placed in a syringe and injected through a needle into the body. Once in place, the solvent dissipates, the remaining polymer solidifies, and a solid structure is formed. The implant will adhere to its surrounding tissue or bone by mechanical forces and can assume the shape of its surrounding cavity. Thus, the biodegradable polymer solution can be injected subdermally like collagen to build up tissue or to fill in defects. Unlike collagen, the degradation time of the implant can be varied from a few weeks to years depending upon the polymer selected and its molecular weight. The injectable polymer solution can also be used to mend bone defects or to provide a continuous matrix when other solid biodegradable implants such as hydroxyapatite plugs are inserted into bone gaps. Although the biodegradable thermoplastic polymers used in the above patent application are biocompatible, the solvents necessary to dissolve them for injection into tissue appear to be less than acceptable. Additionally, the thermoplastic and thermosetting polymers have limited utility in filling soft tissue because they solidify. Similar commercially available materials exhibit ultimate yield stresses of approximately 10,000 psi (68.9 MPa); in comparison, human

Biopolymers: Applications and Trends

skin exhibits ultimate yield stresses of 500–2000 psi (3.4–13.8 MPa). Therefore, due to palpability concerns, the proposed thermoplastic and thermosetting polymers appear to be too hard for use in soft tissue augmentation or repair and especially in dermal augmentation or repair (1996, EP0711794 A1, ETHICON INC.). JPH0323864 A (1991, GUNZE KK) discloses a filler composed of a biodegradable composite material consisting of collagen sponge and a fibrous PLLA embedded therein. By mixing a PLLA having a slow degradation rate within the collagen the pores in the sponge structure can be kept over a long period of time. The propagation of a fibroblast is accelerated by the compounding with fibrous PLLA and the strength and shape of the filler can be kept over a long period of time required in healing. WO9402184 A1 (1994, MEDINVENT SA) discloses the use of a biodegradable polymer for manufacturing of a composition for treating an impaired tissue by injecting the composition into the tissue for the purpose of joining and/or augmenting the same, said composition being in liquid state at the time of injection, and in solid state following injection. Exemplary biodegradable polymers are PLA and PLGA. Soft tissue repair or augmentation has also been proposed using blends of amorphous oligomers and crystalline polymers based on d- or l-lactic acid (1993, EP0544097 A1, BOEHRINGER INGELHEIM KG; BOEHRINGER INGELHEIM INT). These blends were developed to provide a pasty to waxy material which could be used as a bioabsorbable implant to replace the brittle copolymers of lactic acid and glycolic acid for use as bone waxes. However, these blends do not appear to be suitable for use as injectable soft tissue defect fillers, because they are too viscous to be injected through a needle, which significantly limits the utility of these blends. Furthermore, the low molecular weight liquid oligomers are slightly soluble in body fluids, which means that these oligomers will quickly diffuse out of the site of implantation to other areas of the body (1996, EP0711794 A1, ETHICON INC.). EP0711548 A1 (1996, ETHICON INC.) discloses injectable, bioabsorbable microdispersions suitable for use as a soft tissue repair or augmentation material in animals comprising a fluid carrier that is a liquid lactone copolymer; and a particulate material selected from the group consisting of PCL and its copolymers. Preferred liquid lactone

7: Medical, Dental, and Pharmaceutical Applications

copolymers are PLCL, poly(ε-caprolactone-co-pdioxanone), poly(ε-caprolactone-co-trimethylene carbonate), poly(glycolide-co-ε-caprolactone-co-pdioxanone), and poly(lactide-co-ε-caprolactone-cop-dioxanone). The microdispersion is used in facial tissue repair or augmentation, treating vesicoureteral reflex by subreteric injection, or especially for repairing or augmenting sphincter muscle (e.g., to restore or improve sphincter function for treating stress urinary incontinence). It is also worth mentioning the use of hydroxyapatite (Ca5(P04)3(OH)) as a biocompatible ceramic skin augmentation. Hydroxyapatite comprises the mineral constituent of bone, therefore rendering it biocompatible and nonimmunogenic when introduced into the body of a patient. Hydroxyapatite is biodegradable following the same metabolic pathways as bone debris resulting from common bone fractures, yet is semipermanent, as it lasts up to 3 years when implanted into a patient. The hydroxyapatite can be mixed with a biodegradable carrier selected from PEG, PLGA (polyglactin 910 or Lactomer™ 9-1), PGCL (poliglecaprone 25), PDO, poly(glycolide-co-p-dioxanone-cotrimethylene carbonate) (Glycomer 631), PGTMC (Polyglyconate), and a combination thereof (2014, WO2014041531 A1, AVRAHAM AMIR). Exemplary commercial products of bulking and filling agents are:

• Sculptra® (Dermik Laboratories, Sanofi-Aventis) an injectable PLLA implant containing PLLA microparticles in the form of a sterile, freeze-dried preparation. Sculptra works by temporarily adding volume to facial tissue claiming to restore shape and contour deficiencies in areas of facial lipoatrophy. After the initial treatment series, repeat treatments may be needed to maintain the effect. The effect lasts for about 1 year.

• Evolence® Collagen Filler (ColBar LifeScience Ltd.) a sterile, injectable, biodegradable, yellowish, opaque gel filler that is injected into the inner layers of facial skin (mid to deep dermis) in order to provide a temporary improvement to moderate deep facial wrinkles and folds such as those around the nose and mouth (nasolabial folds).

body. Skin is composed of two main layers: the surface epithelium or epidermis, which contains keratinocytes as one type of epidermal cells, and the subjacent connective tissue layer or dermis, which contains fibroblasts as one type of dermal cells. The functions of skin include protecting an organism from injury and desiccation by serving as a barrier to infection, perceiving or detecting environmental stimuli, excreting various substances, regulating body temperature, and helping to maintain water balance. The health and integrity of skin may be compromised by congenital or acquired pathologic conditions for which normal skin regeneration and repair processes may be inadequate. These conditions include burns, wounds, ulcers, infections, diseases, and/or congenital abnormalities. Patients who are burned over a large surface area often require immediate and extensive skin replacement. Skin substitutes derived either ex vivo or in vitro may be used to treat these or other conditions. Desirable properties of skin substitutes are ready availability, a minimum requirement for donor skin, relative simplicity to produce, and cost-effectiveness of fabrication and use. Several approaches to fabrication of skin substitutes, which satisfy some or all of these requirements, have been attempted with varying degrees of success. However, no skin substitute has yet regenerated all of the structures and functions of skin. A potential solution to these problems lies in the development of human skin substitutes based upon cell-seeded, or tissue-engineered, matrices. The matrices may be derived from bioabsorbable polymers, which can provide a wide range of properties and fabrication options needed to produce suitable skin substitutes. Several materials have been manufactured for therapeutic use in skin repair. These materials contain different components replacing or substituting the structures and functions of the dermis and/or epidermis. Examples of these materials include: • EpiCel™ (Genzyme Corp.), which lacks a dermal component and uses the patient’s own cultured keratinocytes; • Integra™ (Integra LifeSciences Corp.), which uses a collagen–glycosaminoglycan matrix to provide an acellular dermal component and uses a thin autograft;

7.2.22  Skin Substitutes Skin is one of the largest organs in the body and covers substantially the entire outer surface of the

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• AlloDerm® (LifeCell Corp.) and a thin autograft;

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• DermaGraft™ (Organogenesis Inc.), which uses a PGA/PLA matrix and allogeneic human fibroblasts for the dermis; •  Hyaff/LaserSkin™, which uses hyaluran and fibroblasts for the dermis, and hyaluran and the patient’s own keratinocytes for the epidermis. Materials to either temporarily cover wounds, or to stimulate permanent skin repair processes, include:

• ApliGraf®

(Organogenesis), which uses collagen gel and allogeneic fibroblasts for the dermis, and cultured allogeneic keratinocytes for the epidermis;

• OrCel™ (formally Composite Cultured Skin) (Ortec Int. Inc.), which uses collagen and allogeneic fibroblasts for the dermis, and cultured allogeneic keratinocytes for the epidermis. US5273900 A (1993), US5711172 A (1998), and US5976878 A (1999) of UNIV CALIFORNIA disclose a composite skin replacement to be used for therapeutic treatment of skin wounds. It is applied surgically in a single procedure, and contains a layer of cultured epidermal cells, an acellular polymeric dermal membrane component, and a substantially nonporous lamination layer on one surface of the dermal membrane component. The dermal membrane component is formed from collagen, or collagen and a mucopolysaccharide compound, and is laminated with the same collagen-, or collagen and mucopolysaccharide-containing solution with a volatile cryoprotectant. The substantially nonporous lamination layer may be located between the dermal component and the layer of cultured epidermal cells, promoting localization of epidermal cells on the surface of the dermal component and movement of nutrients to the cells of the cellular epidermal component. This composition can also be used to deliver biologically active molecules to the site where it is applied. WO03076604 A2 (2003, UNIV CINCINNATI; SHRINERS HOSPITALS CHILDREN) discloses a device for surgical grafting of skin wounds, or for a model of skin in vitro and in animals. The device has a cellular, biocompatible, reticulated proteinor polypeptide-containing matrix such as collagen to provide an attachment substrate for one or more layers or populations of cultured dermal and/or epidermal cells. The protein can be naturally occurring or synthetic and may be less than a full protein, for

example, it may be a polypeptide. In various embodiments, cells used to populate the matrix may be from the recipient (autologous), another human (allogeneic), from another species (xenogeneic), or from multiple sources (chimeric). The epidermal cells include keratinocytes, melanocytes, immunocytes, and/or stem cells.

7.2.23  Wound Dressings and Hemostats Wound dressings are used in the treatment of a variety of wound types, including pressure sores, decubitus ulcers, venous stasis ulcers, infected wounds, deep and open surgical wounds and incisions, sealing of percutaneous incisions or punctures, and burns. The purposes of wound dressings include mechanical protection of the wound, prevention of microbial contamination, prevention of wound dehydration, removal of wound exudate, and delivery of high local levels of a therapeutic agent. They are prepared as fiber mats, sponges, foams, nets, and fibrous substrates. The dressings can be prepared to have a range of different pore sizes and densities. Wound dressings and hemostat devices should possess a number of properties including an ability to remove excess exudate from the wound, protect the wound from mechanical injury, and reduce the risk of infection. The wound dressing must be free of toxic substances, and it should not adhere to the wound which would disturb the healing process. Commonly used dressings include cellulosic dressings such as cotton lint, cotton gauze, cotton wool pads, and cotton/rayon wool pads faced with nonwoven materials. Other dressings contain PUs, PU-polyols, and/ or natural polysaccharide or protein polymers such as collagen. These dressings may be impregnated, coated, or otherwise contain agents such as alginates, which raise the absorptive capacity of the dressing and can stimulate the clotting cascade for bleeding wounds, and/or other agents such as silver salts, antiseptics, analgesics, and/or preservatives devices (1998, WO9851812 A1; 1999, WO9932536 A1; 2000, WO0056376 A1, METABOLIX INC.). JPS61149160 A (1986, NIPPON MEDICAL SUPPLY; GUNZE KK; BIO MATERIARU YUNIBAASU KK) discloses a sponge made of PLA having continuous pores. Such a sponge is produced by a method in which PLA is dissolved in benzene or dioxane, and the solvent is sublimed by freeze-drying

7: Medical, Dental, and Pharmaceutical Applications

the polymer solution. This sponge is generally used for the blood stanching at the time of surgical operation or as a prosthetic material at the time of the suture of a soft tissue (e.g., a body organ) in the living body. However, regarding a porous article produced by a freeze-drying method such as the case of the above sponge, it is difficult to remove the solvent completely because it requires a prolonged period of time for the sublimation, and since it has a thin thickness of 1 mm or less, it is difficult in reality to produce a porous article having a thickness of several millimeters or more (2003, WO03045460 A1, TAKIRON CO). US3875937 A (1975, AMERICAN CYANAMID CO) discloses a surgical dressing comprising a sterile tissue contacting fabric susceptible to hydrolytic degradation to tissue absorbable components consisting essentially of sterile filaments of PGA. EP0560014 A1, US5632727 A, and US5725491 A (1998) of ATRIX LAB INC. disclose a biodegradable film dressing with or without additional therapeutic agents and a method for formation of the film dressing on a human or animal tissue. The film dressing is made of a liquid composition of at least one biodegradable polymer such as PLA, PGA, PCL, or copolymers thereof (e.g., PLCL) in a pharmaceutically acceptable solvent such as N-methylpyrrolidone, alcohols, or alkyl esters. The film is formed by dispensing, preferably by spraying, the liquid composition onto a tissue site and contacting the liquid composition with an aqueous-based fluid such as water, saline solution, or body fluid to coagulate or solidify the film onto the human or animal tissue. The biodegradable film can be used to protect and to promote healing of injured tissue and/or to deliver biologically active agents such as growth. EP1093824 A (2001, MERCK PATENT GMBH) discloses a biodegradable polymer, useful as wound dressing, including dye to make it easily detectable in the wound (e.g., to ensure that the wound is fully covered). The biodegradable polymer is preferably PLCL. Example: A solution of 0.075% D and C Green no.6 in ethyl acetate was treated with a 50/50 mixture of (1) a PLCL with a ratio of d,l-lactide/εcaprolactone = 85/15 and (2) a PLCL with a ratio of 82/18 of the same components, to total dissolved polymer content 20 wt%. The resulting solution was processed to a film, using standard doctor blade

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techniques, to give a colored film that was biodegradable, biocompatible, and clearly detectable in a wound. WO2008057600 A2 (2008, KCI LICENSING INC.) discloses a bioabsorbable dressing to be used in conjunction with reduced pressure therapy for treatment of a wound site. The bioabsorbable dressing comprises a casing and bioabsorbable microspheres in the form of a rope shape. Further, the casing of the dressing comprises pores formed by a porogen system that may be activated in situ by wound fluids to initiate pore formation. The bioabsorbable polymer is a biocompatible polymer. In one preferred embodiment the bioabsorbable polymer is a PLCL copolymer, wherein the ratio of lactide/ε-caprolactone (LA/ CL) is about 90/10. In other embodiments, the LA/ CL copolymer ratio is about 80/20. In yet another embodiment, the LA/CL copolymer ratio is about 70/30. US2011280919 A1 (2011, ETHICON INC.) discloses a reinforced bioabsorbable hemostat comprising at least one hemostatic agent such as thrombin and fibrinogen in a single layer of nonwoven synthetic fabric having a mixture of compressed fiber staples of PLGA and PDO. GB2166354 A (1986, ICI PLC) discloses wound dressings made of P3HB. P4HB–chitosan hemostats comprising P4HB films adhered to chitosan were shown to be effective at controlling parenchymal hemorrhage and sealing the renal collecting system [96]. The P4HB acts as a temporary waterproof backing and serves to reinforce the chitosan sponge as well to reduce its tendency to adhere to surgical instruments [10].

7.2.24 Stents Stents are generally cylindrically shaped devices, which function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. The stent may be placed using endoscopic surgical techniques or percutaneously. The percutaneous approach is commonly used in vascular procedures. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. Stents are typically delivered to the target area within the body lumen using a catheter. With balloon-expandable stents, the stent is mounted to a balloon catheter, navigated to the appropriate area, and the stent is expanded by inflating the balloon. A self-expanding stent is delivered to the target

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Biopolymers: Applications and Trends

area and released, expanding to the required diameter to treat the disease. Stents are used not only for mechanical intervention but also as vehicles for the local administration of a therapeutic substance (2011, US2011065825 A1, ABBOTT CARDIOVASCULAR SYSTEMS INC.). The structure of a stent typically comprises scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts, links, and rings. The scaffolding can be formed from wires, tubes, or sheets of material rolled into a cylindrical shape (see Figure 7.15). A stent (100) may be fabricated from a tube by forming a pattern in the tube with a technique such as laser cutting or chemical etching. The scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment). Stents can be made from various materials, particularly metallic and/or polymeric materials, and may be nondegradable, biodegradable, or be formed from both degradable and nondegradable components. One of the main problems in using metallic stents in cardiovascular applications is the subsequent restenosis caused by excessive growth of the endothelial wall, which is believed due, at least in part, to inflammation caused by the metallic stent on the vessel wall. A disadvantage of polymer stents compared to metal stents of the same dimensions, is that polymer

105

100

Figure 7.15  An exemplary structure of a stent made up of struts (2011, US2011065825 A1, ABBOTT CARDIOVASCULAR SYSTEMS INC.). 100, Stent; 105, Struts.

stents typically have less radial strength and rigidity. A low radial strength potentially contributes to high recoil of polymer stents after implantation into an anatomical lumen. “Recoil” refers to the undesired retraction of a stent radially inward from its deployed diameter due to radially compressive forces that bear upon it after deployment. Another potential problem with polymer stents is that struts can crack or fracture during crimping, delivery, and deployment, especially for brittle polymers (2010, US2011066222 A1, WANG YUNBING, GADA MANISH B; 2011, WO2011031872 A1, ABBOTT CARDIOVASCULAR SYSTEMS). On the other hand, a stent made from, or coated with a biodegradable polymer produces reduced or no inflammation. Furthermore, in many treatment applications, the presence of a stent in the body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. A polymer scaffold made of a biodegradable polymer can bioabsorb, bioresorb, or bioerode away from an implant site, which is particularly desirable in applications such as urological applications, since a second procedure is not required to remove the stent. It is also believed that biodegradable scaffolds allow for improved healing of the anatomical lumen as compared to metal stents, which may lead to a reduced incidence of late stage thrombosis. Biodegradable polymers used to make stent scaffolding include PLLA, PLGA, scPLA, and PLLAbased polyester block copolymer containing a rigid segment of PLLA or PLGA and a soft segment of PCL or PTMC. PLLA with less than 10% d-lactide is attractive as a stent material since it remains stiff and rigid at human body temperature (about 37 °C). This property facilitates the ability of a PLA-based stent scaffolding to maintain a lumen at or near a deployed diameter without significant recoil (e.g., less than 10%) (2012, WO2012154842 A2, ABBOTT CARDIOVASCULAR SYSTEMS). But these materials can exhibit a brittle fracture mechanism at physiological conditions (approximately 37 °C) in which there is little or no plastic deformation prior to failure. A stent fabricated from such polymers can have insufficient toughness for the range of use of a stent. As a result, cracks, particularly in high strain regions, can be induced, which can result in mechanical failure of the stent (2011, WO2011031872 A1, ABBOTT CARDIOVASCULAR SYSTEMS).

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In comparison to metals conventionally used to form stent scaffolding, a suitable polymer has a low strength to weight ratio, which means more material is needed to provide an equivalent mechanical property to that of a metal. Therefore, struts in polymeric scaffolding must be made thicker and wider to have the strength needed. Polymeric scaffolding also tends to be brittle or have limited fracture toughness. The anisotropic and rate-dependent inelastic properties (i.e., strength/stiffness of the material varies depending upon the rate at which the material is deformed) that are inherent in the material only compound this complexity in working with a polymer, particularly, a biodegradable polymer such as PLLA and PLGA (2011, US2011270384 A1, ABBOTT CARDIOVASCULAR SYSTEMS). CA2025626 A (1991, SQUIBB BRISTOL MYERS CO) discloses a bioresorbable infusion stent made of a biodegradable terpolymer of l-lactide (45– 85 wt%), glycolide (5–50 wt%), and ε-caprolactone (15–25 wt%). The terpolymer has a minimum tensile strength of at least 500 psi (3.45 MPa), preferably 650 psi (4.48 MPa); elongation at break of greater than 10%, preferably greater than 100%; and Shore A hardness equal to 50–100%, preferably 75–95%. The invention includes a method for incorporating radiopaque materials such as finely ground barium sulfate into the polymer in an amount of 5–30 wt%. The bioresorbable infusion stent is used to treat ureteral obstructions. The mechanism of biodegradation is described as hydrolysis resulting in degradable products excreted in urine or reabsorbed into tissues. The duration of functional life of the stent is estimated at about 3–7 weeks. US5464450 A (1995) and US5500013 A (1996) of SCIMED LIFE SYSTEMS INC. discloses a stent including a main body of a generally tubular shape for insertion into a lumen of a vessel of a living organism. The tubular main body includes a substantially biodegradable matrix having collagen IV and laminin that enclose voids within the matrix. The tubular main body also includes a biodegradable strengthening material in contact with the matrix to strengthen the matrix. The tubular main body is essentially saturated with drugs. EP0689807 A2 (1996, ADVANCED CARDIOVASCULAR SYSTEM) discloses a biodegradable mesh-and-film stent for use in maintaining the patency of blood vessels, said stent comprising: a mesh layer (10) having first and second sides, formed from a plurality of fibers of a first biodegradable

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polymer (12) heat-bonded together with a plurality of a second biodegradable adhesive polymer fibers having a melting temperature below that of the first biodegradable polymer (14); at least one layer of a film of a biodegradable polymer (28 or 29) bonded to said mesh layer on at least one side of said mesh layer to form a sheet of biodegradable mesh-and-film material; the sheet of biodegradable mesh-and-film material having first and second ends (62, 64), and a main body portion (66) between said first and second ends, said sheet of biodegradable mesh-and-film material being rolled up into a cylindrical configuration whereby said first end overlaps said second end (see Figure 7.16). The first biodegradable polymer (12) is selected from the group consisting of PGA, PLLA, POEs, polyanhydrides, polyiminocarbonates, and inorganic calcium phosphate. The second biodegradable polymer (14) is selected from the group consisting of PCL, PDLLA, a combination of PLLA and PCL, POEs, aliphatic polycarbonates, PPHOs, and combinations thereof. US5670161 A (1997, BIOENGINEERED MATERIALS INC.; GEN VASCULAR DEVICES LTD.) discloses an expandable, bioabsorbable stent for use within a body lumen comprising a hollow tube made from PLLCL that is not plastically expandable at normal body temperatures, and that is expandable using thermomechanical means at a temperature of 38–55 °C using a balloon catheter (see Figure 7.17). The invention also relates to a method of deploying such a stent within the body. WO03034940 A2 (2003, SCIMED LIFE SYSTEMS INC.) discloses a method for forming a stent of a biodegradable polymer including the steps of (1) forming a generally tubular stent; (2) radially expanding the stent to produce an expanded diameter stent; and (3) annealing the expanded diameter stent to shrink its diameter and reduce its size. The biodegradable polymer is selected from the group consisting of poly(α-hydroxy acid), PLA–polyethylene oxide copolymers; modified cellulose; collagen or other connective proteins; adhesive proteins; hyaluronic acid; polyanhydrides; PPHOs; poly(amino acids); copolymers thereof; and mixtures of any of said materials. The poly(α-hydroxy acid) is selected from the group consisting of homopolymers and copolymers of PLA, PLLA, PDLA, PGA, PDO, PCL, P3HB, PGTMC (polyglyconate), and mixtures thereof. US2009248147 A1 (2009, ABBOTT CARDIOVASCULAR SYSTEMS; WANG YUNBING)

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Figure 7.16 Biodegradable mesh-and-film stent. (a) Biodegradable mesh layer, (b) mesh-and-film sheet, (c) top plan view of the biodegradable meshand-film stent, and (d) Perpsective view of the biodegradable mesh-and-film stent (1996, EP0689807 A2, ADVANCED CARDIOVASCULAR SYSTEM). 10, Mesh layer; 12, Fibers of a first biodegradable polymer; 14, Fibers of a second biodegradable polymer; 28 and 29, Films of biodegradable polymer; 62 and 64, First and second ends of biodegradable meshand-film sheet; 66, Main body; 68, Apertures; 70, Slot; 72, Loop-shaped stent; 74, Widened portion of second end; 76, Serrations; 78, Side of the main body.

Figure 7.17 View of an unexpanded stent (1997, US5670161 A, BIOENGINEERED MATERIALS INC.; GEN VASCULAR DEVICES LTD.).

Biopolymers: Applications and Trends

discloses a bioabsorbable stent made of a polymer mixture comprising: PLGA, wherein the l-lactic acid/glycolic acid ratio is in the range of about 80/20 to about 99.99/0.01 and a nucleating agent (0.1– 10 wt%). The nucleating agent is selected from the group consisting of PGA, PLGA (with less than about 10 wt% l-lactide), Mg silicate hydrate, ethylene bis (1,2-hydroxystearylamide), boron nitride, hydroxyapatite, decamethylenedicarboxylichydrazide, dibenzoylhydrazide, dioctyl phthalate, ethyl lactate, citric acid esters, lactic acid esters, lactide esters, triphenyl phosphate, glycerol, acetin, and butyrin. The resulting stent is claimed to have increased crystallinity, decreased crystal size, increased mechanical properties, and faster degradation times. WO2008008416 A1 (2008) and US2013085563 A1 (2013) of ABBOTT CARDIOVASCULAR SYSTEMS disclose stent scaffolds made of a bioabsorbable polymer composite comprising a high toughness polymer dispersed within a matrix polymer, the matrix polymer being glassy at physiological conditions, wherein the high toughness polymer enhances the fracture toughness of the composite at physiological conditions. The matrix polymer comprises a biodegradable polymer selected from the group consisting of PLLA, PGA, poly(ester amide), and copolymers thereof. The high toughness polymer comprises a biodegradable polymer selected from the group consisting of poly(butylene succinate) (PBS), P4HB, PCL, PTMC, PDO, and copolymers thereof; an exemplary commercial product of PBS is Bionolle® 3000 (Showa Highpolymer Co Ltd.). The high toughness biodegradable polymers are hydrophobic. The stent scaffolds may be formed by extruding polymer tubes made of the rubber-toughened material and laser cutting the tubes to form a scaffold. The rubbery component increases the fracture resistance of the stent scaffold by forming discrete rubbery domains, plasticizing, or both. In some embodiments, the rubbery polymer is completely miscible in the stiff polymer and the increase in toughness is due to plasticizing. Some biodegradable polymers have a degradation rate that is slower than desired for certain stent treatments. As a result, the degradation time of a stent made from such polymers can be longer than desired. For example, a stent made from a semicrystalline polymer such as PLLA can have a degradation time between about 2 and 3 years. In some treatment situations, a shorter degradation time is desirable, for example, less than a year. Amorphous biodegradable

7: Medical, Dental, and Pharmaceutical Applications

polymers (less than 10% crystallinity) degrade faster than crystalline polymers but are weaker than crystalline polymers and hence are not typically suitable for vascular implants, such as stents, which need sufficient strength to provide support to the blood vessel. US2008177374 A1 (2008, ELIXIR MEDICAL CORP.) discloses a method for fabricating a bioabsorbable stent including the following steps: (1) providing a tubular body having an initial diameter, wherein said tubular body is composed at least partially of a substantially amorphous biodegradable polymer, while the diameter remains substantially unchanged; (2) heating the tubular body to a temperature above the Tg and below the Tm of the polymer; (3) cooling the tubular body to increase the crystallinity of the polymer; and (4) patterning the tubular body into a structure capable of radial contraction and expansion. The amorphous polymer is selected from the group consisting of PDLLA, PLGA PLCL, PLTMC, PTMC, P3HB, P3HV, POEs, polyanhydrides, polyiminocarbonates, and their copolymers. Preferably, PLGA is used (l-lactide/glycolide = 85/15). The use of an amorphous polymer is desirable since it provides relatively short periods of biodegradation, usually less than 2 years, often less than 1 year, frequently less than 9 months, and sometimes shorter than 6 months, or even shorter. The present invention relies on modifying the amorphous polymers to introduce a desired degree of crystallinity. The crystallinity of a highly amorphous polymer is below 10% prior to modification. After modification (i.e., after heating and cooling) the polymer has an increased crystallinity by at least 20% of original crystallinity. By introducing crystallinity into the amorphous polymer increases the strength of the polymer, so that it is suitable for use as a stent or other endoprosthesis without substantially lengthening the period of biodegradation after implantation. US2011066222 A1 (2011, WANG YUNBING; GADA MANISH B) discloses a method for forming a stent by deforming a tube made of PLLA. The deforming process includes the following steps: (1) maintaining fluid pressure in the precursor tube at a process pressure of 110–150 psi (0.758–1.034 MPa); (2) heating the precursor tube to a temperature of 160–220  °F (71–104  °C), radially expanding the precursor tube according to a radial expansion ratio of 300–450% during the maintaining of fluid pressure, and the heating; (3) axially extending the precursor tube according to an axial extension ratio of 20–100% during the maintaining of fluid pressure

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and the heating; and (4) forming a network of stent struts from the deformed tube. Stents cut from the deformed PLLA tube is claimed to have improved fracture toughness upon deployment while maintaining sufficient flexibility for crimping and delivery and sufficient radial strength to prevent undue recoil. US2014265060 A1 (2011, ADVANCED CARDIOVASCULAR SYSTEM) discloses a method for fabricating a stent from a tube made of a biodegradable polymer such as PLLA comprising: inducing crystallization in the tube by treating at least a portion of the tube with a solvent; radially and axially deforming the tube including radially deforming the tube between 100% and 400% while the tube has a temperature equal to or above the Tg of the polymer of the tube; and fabricating the stent from the deformed and treated tube. The crystallization is induced, for example, by exposing the polymer to a solvent, before the deforming step, which lowers the Tg of the polymer. US2011065825 A1 (2011, ABBOTT CARDIOVASCULAR SYSTEMS INC.) discloses a method for fabricating a bioabsorbable polymer stent comprising: exposing a stent to a dose of e-beam radiation between 20 and 30 kGy, the stent comprises a bioabsorbable stent scaffolding made from a polymer formulation that contains PLLA and polymandelide (PM), or PLLA chemically modified with PM (see Chapter 1; Section 1.2.1.1.2). PM stabilizes the polymer formulation by reducing the amount of weight average molecular weight (Mw) degradation of PLLA to less than 30% due to radiation and high temperatures. PM is more brittle at body temperature than PLLA. Therefore, it is important for a polymeric stent body to have high fracture toughness since brittle behavior can result in fractures and premature failure of a stent. Additionally, PM can reduce crystallinity of PLLA which may reduce strength. Also, PM can decrease the in vivo degradation rate of PLLA. Therefore, it is critical that a polymer formulation contains enough PM to increase the radiation and thermal stability without adversely effecting stent performance, i.e., strength, and in vivo degradation rate. In a preferred embodiment, the polymer formulation contains 2–5 wt% PM. An amount of PM greater than 5 wt% is likely to adversely affect the fracture toughness, strength, and degradation rate of the stent body. The Mw of the PM in the blend can be 50–500 kg/mol and the Mw of the PLLA in the blend can be 300–600 kg/mol.

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Stents may be coated with a polymeric carrier impregnated with a drug or therapeutic substance. A conventional method of coating includes applying a composition including a solvent, a polymer dissolved in the solvent, and a therapeutic substance dispersed in the blend to the stent by immersing the stent in the composition or by spraying the composition onto the stent. The solvent is allowed to evaporate, leaving on the stent strut surfaces a coating of the polymer and the therapeutic substance impregnated in the polymer (2010, US2010004735 A1, ABBOTT CARDIOVASCULAR SYSTEMS INC.). DE19539449 A1 (1997, BIOTRONIK MESS & THERAPIEG) discloses a stent made of P3HB. The stent construct, which is reported to bioresorb rapidly, contains a large amount of plasticizer. However, the plasticized P3HB approach fails to work, as greater than 90% stent stenosis was shown at 4 weeks [97]. A list of commercial biodegradable stents is given in Table 7.9.

7.2.25 Wraps Sterilization wraps are mainly used in CSSD (Central Sterilization & Supplies Department) of hospitals and health care facilities for the packaging, the sterilization and the maintenance of the sterile state of reusable, freshly washed, and disinfected medical devices such as scalpels, pliers, scissors, endoscopes, bedpans, tongue depressors, or stents. Sterilization wraps are made of a mixture of cellulosic wood pulp, binders, fibers, and hydrophobic additives. Except the cellulosic wood pulp, those compounds are usually nonbiodegradable and made from nonrenewable raw materials (materials of fossil origin). Those nonwovens are known as “wet-laid nonwovens,” or 100% synthetic fossil fuel-based fibers such as polypropylene. Those nonwovens are known as “SMS” (Spunbond–Meltblown–Spunbond) nonwovens (2012, EP2450487 A1, ARJOWIGGINS PALALDA SAS). A nonwoven sterilization wrap shall conform with the general and specific requirements of ISO 116071:06 and European Standard EN 868-2:09. EP2450487 A1 (2012, ARJOWIGGINS PALALDA SAS) discloses a sterilization wrap made of biodegradable materials. The sterilization wrap defines a closed inner volume containing a sterilized medical device. The sterilized medical device is, in particular, chosen from sterilized scalpels, pliers, scissors, endoscopes, bedpans, tongue depressors, or stents. The biodegradable material is in the form

Biopolymers: Applications and Trends

of a nonwoven sheet and comprises (1) >15 wt% cellulose fibers, 1–80  wt%, preferably 2–20  wt% synthetic biodegradable fibers; (2) 20–10 wt%, preferably 55–85 wt% natural biodegradable fibers; and (3) 2–30 wt%, preferably <25 wt% a biodegradable fiber such as potato starch. The natural biodegradable fibers are preferably selected from at least one of the following ones: bleached wood pulp, semibleached wood pulp, unbleached wood pulp, cotton, abaca, straw, bamboo, hemp, jute, sisal, flax, kenaf, or esparto. The synthetic biodegradable fibers are preferably fibers of PLA, PHA, PHHx, PCL, PBS or poly(butylene succinate-co-adipate), viscose fibers, fibers of plastified starch, such as corn starch, wheat starch, or potato starch optionally modified with copolyesters, or fibers of plastified cereal flour polymer. The disclosed nonwoven sheet presents a relatively low grammage (preferably 50–75 g/m2) and an enhanced biodegradability while still meeting the requirements of EN 868-2:09 and presents satisfying barrier properties, in particular to microorganisms.

7.2.26  Ocular Cell Implants Two monolayers of cells, known as retinal pigment epithelium and corneal endothelium are essential for normal vision. In age-related macular degeneration, the function of the retinal pigment epithelium is believed to be altered leading to visual loss. Replacement of this altered epithelium with a healthy retinal pigment epithelium can potentially provide a treatment for this debilitating condition. Transplantation of donor cell suspensions has been attempted but is problematic, and has led to several attempts to use synthetic bioabsorbable polymers and protein polymers as tissue engineering scaffolds to deliver retinal pigment epithelium and corneal endothelium into the eye. PLLA and PLGA have been tested as scaffolds to culture retinal pigment epithelial (RPE) and corneal endothelial cells for potential use in monolayer transplantation in the eye [111]. US2002183844 A1 (2002, UNIV LELAND STANFORD JUNIOR) discloses methods and related products for treating retinal diseases such as age-related macular degeneration (AMD), retinitis pigmentosa (RP), and other retinal diseases. A therapy for AMD is to transplant suspensions of either RPE cells or iris pigment epithelial cells to rescue the diseased retina. The methods include removal of membranous tissue, such as a lens capsule or an

7: Medical, Dental, and Pharmaceutical Applications

335

Table 7.9 Commercial Biodegradable Stents [98,99] Commercial product

Strut material

Coating

Eluted drug

Absorb® BVS Revision 1.0 Revision 1.1

PLLA PLLA

PDLLA PDLLA

Everolimus Everolimus

Amaranth BRS

PLLA

None

ART 18Z

PLLA

DESolve®

Company

Reference

Abbott Vascular

[100–103]

None

Amaranth Medical, Inc.

[102,104]

None

None

Arterial Remodeling Technologies

[102,105]

PLLA

PLLA

Myolimus, novolimus

Elixir Medical, Corp.

[102–104]

IDEAL™ Generation I Generation II

Polylactide anhydride + polymer of salicylic acid with sebacic acid linker

Salicilate linked with adipic acid

Sirolimus

Xenogenics Corp.

[102,104]

Igaki-Tamai®

PLLA

None

None

Kyoto Medical Planning Co.

[103,106,107]

On-ABS

PLLA + PDLLA + PLLCL

NA

Sirolimus

OrbusNeich Medical, Inc.

[102,103,108]

REVA Generation I

Tyrosine-derived PC

None

Paclitaxel

REVA Medical, Inc.

[103,109]

ReZolve®

Tyrosine-derived PC

Sirolimus

REVA Medical, Inc.

[102,103]

Xinsorb BRS™

PLLA, PLLCL, PLGA

Sirolimus

Huaan Biotechnology Group

[102,110]

NAa

PC, Polycarbonate; PDLLA, Poly(d,l-lactide); PLGA, Poly(l-lactide-co-glycolide); PLLA, Poly(l-lactic acid), poly(l-lactide); PLLCL, Poly(l-lactide-co-ε-caprolactone). aNA, not applicable.

inner limiting membrane, from an eye, flattening the membranous tissue onto a polymer substrate, such as a coverslip, submersed in phosphate-buffered saline, or flattening the membranous tissue onto a temporary dissolvable polymer for ease of surgical transplantation. The modified tissue provides a suitable substrate for cells, and may be exposed to cells which may attach and grow. The modified tissue, with adherent cells, if any, were applied to and grown on the tissue and/or with polymer, may next be transplanted into a desired location within the body of an animal. Following transplantation, where the modified tissue has been prepared with a dissolvable polymer, the polymer will dissolve and be absorbed by the body of the patient leaving the transplanted tissue and cells

in place. The substrate comprises a biodegradable polymer selected from the group consisting of PLA, PGA, PLGA, POEs, polyanhydrides, PPHOs, PEGPLA copolymer, and blends thereof. US5713955 A (189, DURETTE JEAN-FRANCOIS) discloses an orbital implant (10) for coupling with an ocular prosthesis to replace an eye in a patient following evisceration or enucleation, said implant comprising a generally spherical-shaped member having an anterior surface and a posterior surface, and a cap (26) of substantially uniformly thin material positioned on and covering only the anterior surface of the spherical-shaped member, the cap being attached to the spherical-shaped member and made of a bioabsorbable polymer forming a matrix having

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Biopolymers: Applications and Trends

Figure 7.18 Cross-sectional view of an orbital implant (1989, US5713955 A, DURETTE JEANFRANCOIS). 10, Orbital implant; 14a, Passageway; 22, Head; 24, Peg; 26, Cap.

a plurality of continuous random voids that promote in-growth into the cap of surrounding tissue after the implant is in place in the patient (see Figure 7.18). The bioabsorbable polymer is PGA or PLA. P4HB shunts were evaluated for their ability to reduce fibroblast proliferation and fibrosis. Aqueous shunts are sometimes used to treat glaucoma patients as an alternative to the use of medication. The shunts drain liquid from the anterior chamber of the eye into the connective tissue and can become blocked as a result of fibrotic reactions. Preliminary results indicated that the P4Hb shunt causes moderate vascularization in the cornea and conjunctiva [96].

7.2.27  Ear Implants US4650488 A (1987, RICHARDS MEDICAL CO) discloses a prosthetic device made of a biodegradable polymer being useful as a ventilation tube for insertion between the middle ear and the outer ear. The biodegradable polymer is selected from the group consisting of PDLLA, PLGA, PCL and PLCL. The degradation rates of various portions of the implant are varied by adjusting the molecular weight of the polymer, for example, with irradiation. The device has not hitherto been applied in the field inasmuch as it is potentially susceptible to exposing the patients in which it is implanted to numerous risks. First, the PLA-based copolymers involve the risk of growth of granulation tissues consequent upon their imperfect absorption by the tissues. The formation of granulation tissues is particularly risky in auricular treatments. Moreover, both through the typical degradation of the material used and through the frustoconical configuration of the tapered body, it involves the risk that the epithelial growth with which the hole for application of the prosthesis tends to close itself up again develops with invagination of the

Figure 7.19 (a) View in longitudinal section of a prosthetic device implanted in a tympanic membrane and (b) perspective view of the prosthetic device implanted in a tympanic membrane (2003, WO03059406 A2, D’EREDITA RICCARDO). 1, Auricular prosthetic device; 2, Tympanic membrane of the ear; 2a, Hole made in the tympanic membrane; 3, Outer ear; 4, Middle ear; 10, Cylindrical portion of the tubular body of (1); 11, Flange; 12, Flange; 14, Duct.

keratinized squamous epithelium, of the actual outer ear, along the conical portion and toward the middle ear, where mucous epithelium is present. There is in practice a risk of migration of the keratinized squamous epithelium toward the middle ear which is a possible cause of cholesteatoma. All the problems in question have hitherto led to delays in the marketing of devices of the aforesaid type (2003, WO03059406 A2, D’EREDITA RICCARDO). WO03059406 A2 (2003, D’EREDITA RICCARDO) discloses a biodegradable auricular prosthetic device (1) for the treatment of otitis media, comprising a tubular body (10) having axially opposed ends and flanged (11, 12) at least at one of the opposed ends and at least a portion of which is produced from a material subject to biological degradation in the presence of organic liquids, wherein at least the portion made of material subject to biological degradation is produced from a biodegradable polymer selected from the group of PPHOs (see Figure 7.19).

7: Medical, Dental, and Pharmaceutical Applications

7.2.28 Adhesives Biological and synthetic tissue adhesives are used as alternatives to sutures and staples for adhering biological tissue. Examples of biological tissue adhesives include fibrin glues. Examples of synthetic tissue adhesives include cyanoacrylates, urethane prepolymers, and gelatin–resorcinol–formaldehyde. Applications of adhesives to biological tissue range from soft (connective) tissue adhesion to hard (calcified) tissue adhesion. Soft tissue adhesives are used both externally and internally for wound closure and sealing. Hard tissue adhesives are used in applications that include bonding prosthetic materials to teeth and bone (2009, WO2009014886 A2, COHERA MEDICAL INC.). JP2000160125 A (2006, SHIMADZU CORP.) discloses an adhesive material for biological absorption soft tissue and surgical operation application, comprising a blend of biodegradable block copolymers (A) and (B) or (C) and (D). (A) is a block copolymer of PLLA and poly(alkylene ether), and (B) is a block copolymer of PDLA and poly(alkylene ether); (C) is a block copolymer of at least one kind of PLA with a poly(alkylene ether) and (D) is PLA. The adhesive strength is increased greatly without reducing hydrophilicity, since there is no crystallinity in the copolymer. WO2009014886 A2 (2009, COHERA MEDICAL INC.) discloses a moisture-curable and biodegradable adhesive that includes the reaction product of (1) an isocyanate component having an average functionality of at least 2; (2) an active hydrogen component having an average functionality greater than 2.1; and (3) an ionic salt component having an average hydroxyl or amino functionality, or combination thereof, of at least 1. The isocyanate component (1) is selected from the group consisting of lysine diisocyanate and derivatives thereof, lysine triisocyanate and derivatives thereof, and combinations thereof. The active hydrogen component (2) consists of components having primary hydroxyl groups and/or primary amine groups; preferably the active hydrogen component is selected from glycerol, diglycerol, erythritol, pentaerythritol, xylitol, arabitol, fucitol, ribitol, sorbitol, mannitol, hydroxyalkyl, saccharides, oligosaccharides, polysaccharides, and/or esters. The ionic salt component (3) is selected from the group consisting of ammoniates, halides, sulfonates, phosphonates, carboxylates, and combinations thereof. The adhesive is useful in a variety of medical applications

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for adhering biological tissue to (1) soft tissue (e.g., muscle, fat, skin, fascia, or tendons); (2) hard tissue (e.g., bone or cartilage); (3) biologically derived materials (e.g., extracellular matrices, small intestine submucosa, collage or tissue-engineered scaffolds); (4) synthetic materials (e.g., films and meshes made of polymers); (5) metal and inorganic implants (e.g., orthopedic hardware, joint replacements, or hydroxyapatite bone substitutes); and (6) tissue transplants (e.g., allografts, xenografts, or autografts). WO2012030819 A1 (2012, SURMODICS PHARMACEUTICALS INC.) discloses an implant device comprising a substrate having a biodegradable pressure-sensitive adhesive (PSA) adhered to a surface thereof. The PSA comprises a poly(d,llactide-co-glycolide-co-ε-caprolactone) terpolymer. The terpolymer exhibits tackiness or stickiness and thickness over extended periods of time, hence it can adhere to a variety of substrates, even with the application of light pressure and without the need of a solvent, and allows the implant device to be secured to a particular location within a subject. The implant device can be any type of medical implant including, for example, implants for drug delivery (e.g., drug delivery pumps); orthopedic implants (e.g., spinal implants, implants for osseointegration or bone repair); medical stents (e.g., stents with inherent drug delivery capability); prosthetic implants (e.g., breast implants or muscle implants); dental implants; ear implants (e.g., cochlear implants and hearing devices); cardiac implants (e.g., pacemakers, or catheters); space-filling implants; bioelectric implants; neural implants; internal organ implants (e.g., dialysis grafts; defibrillators; monitoring devices; recording devices; stimulators including deep brain stimulators, nerve stimulators, bladder stimulators, and diaphragm stimulators; implantable identification devices and information chips; artificial organs; drug administering devices; implantable sensors/ biosensors; screws; tubes; rods; plates; or artificial joints).

7.3  Drug Delivery Matrices or Vehicles Biodegradable polymers form the basis for drug delivery, either as a drug delivery system alone or in conjunction to functioning as a medical device. The use of biodegradable polymers to provide sustained or controlled release of drugs has been known since

338

the 1960s [112]. Biodegradable implants for the controlled release of hormones, particularly contraceptive hormones, were developed in the 1970s [113]. Known drug delivery systems for controlled or sustained drug release are, for example, particles such as micro- and nanoparticles, fibers, rods, films, or coatings. These drug delivery systems often comprise a drug dispersed in a biocompatible polymer matrix, which can be implanted, administered orally, or injected. The polymer matrices break down in the body by various endogenous substances such as enzymes and water. Biodegradable polymers most often used are, for example, PLA, PLGA, and PCL. A number of other biodegradable polymers have been used for the controlled release of drugs such as polyanhydrides [114,115], PPOEs [116–118], PPF (2004, US2004023028 A1, MAYO FOUNDATION FOR MEDICAL EDUCATION AND RESEARCH), and poly(ester amides) (2011, WO2011045443 A1, DSM IP ASSETS BV). Biodegradable polymer microparticles or microspheres made of PLGA and PLA, are effective delivery vehicles for the controlled release of therapeutic compositions such as polypeptides, proteins, nucleic acids, and vaccines. FR2070153 A1 (1973, DU PONT) discloses various methods for incorporating drugs into PLA, for example, by coating, mixing the drug with melted polymer, or mixing with a common solvent and drying to form powder. US4622244 A (1986, UNIV WASHINGTON) discloses a method for preparing microcapsules comprising active substances coated with a bioabsorbable polymer such as PLA useful for injection. US4818542 A (1989, UNIV KENTUCKY RES FOUND) discloses porous microspheres having a network of interconnecting channels containing pore-incorporated drugs. Typically, the porous micropsheres are made of PLGA (polyglactin 910). EP0302582 A1 (1989, SOUTHERN RES INST) discloses a composition capable of delivering an effective amount of constant dose of an active ingredient (peptide) to an animal over a preselected, prolonged period of time, comprising a blend of effective amounts of an active ingredient encapsulated in at least two bioabsorbable copolymer excipients to form first and second microcapsules, each excipient degrading at a different rate. The bioabsorbable copolymer excipients are preferably PLGA copolymers with different lactide/glycolide molar ratios in the range of 40/60– 100/0. GB2209937 A (1989, DEBIOPHARM S A) discloses the use of microparticles made of PLA or

Biopolymers: Applications and Trends

PLGA for the continuous release of a water-insoluble peptide salt selected from the group consisting of pamoate, tannate, and stearate salts. DE4021517 A1 (1991, SANDOZ AG; NOVARTIS AG) discloses microparticles formed by a method in which a peptide hormone (somatostatin such as an octreotide pamoate salt) and PLGA are dissolved in different phases. Studies of such microspheres reveal that proteins carried therein may be released before the polymer itself is degraded [119]. Nevertheless, microencapsulation with PLGA does appear to slow down diffusion of locally administered antibiotics [120]. Microspheres made of PLGA with a ratio of lactide/ glycolide = 75/25 and molecular weight of 14 kDa or less appear to have released proteins over a period of 4–6 weeks [121] (2000, US6013853 A, UNIV TEXAS). In general, the manufacturing of polymer/ drug particles in which the drug is a protein causes difficult issues to deal with, and is currently limited to a few techniques. Oil/water emulsion method and a neat mixing method at elevated temperature were used for the preparation of polymer/protein particles comprising PLA and either trypsin or insulin [122]. A disadvantage of the last method is that the particles prepared by the neat mixing method lost a significant fraction of protein activity, possibly due to the heating step. The particles also suffered from a large initial burst of protein release. The particles prepared by the oil/water emulsion method lost an even greater amount of protein activity, possibly caused by protein lability with respect to the oil. This means that the processes for manufacturing of drug delivery systems comprising the incorporation of thermally sensitive or solvent-sensitive bioactive agents or drugs, such as many proteins, peptides, and polynucleotides is limited due to the process conditions, which are often elevated temperatures (>45 °C) and/or aqueous/ organic emulsions. These process conditions result in a significant loss of drug activity. On the other hand, processes utilizing a simple mixture of polymer with a solid bioactive agent or drug do often not yield a fine microscopic homogeneous dispersion of the bioactive agent or drug within the polymer matrix. This results in a more erratic drug release in vitro and in vivo (2013, WO2013087903 A1, DSM IP ASSETS BV). The shape of a polymer particle can be of utmost importance to its suitability for a particular use. As such, there are many methods for preparing polymer particles of a particular morphology, and also

7: Medical, Dental, and Pharmaceutical Applications

for processing a particle (e.g., grinding, cutting, and milling) to obtain a desired shape. Often, the result of preparing a particular polymer particle or of processing that particle is a rough and/or jagged particle surface. Such irregularly shaped particles can have several disadvantages, such as great variability in the release profile and dosing of a drug within the particle. Other problems can relate to delivery (e.g., clotting and clumping) (2008, US2008103277 A1, BROOKWOOD PHARMACEUTICALS INC.). US2008103277 A1 (2008, BROOKWOOD PHARMACEUTICALS INC.) discloses a method for producing a spheroid particle from biodegradable polymers comprising: (1) providing a mixture comprising a liquid medium and a nonspheroid polymer particle comprising one or more secondary components; (2) heating the mixture above the Tg or Tm of the polymer; and (3) cooling the mixture to below the Tg or Tm of the polymer, thereby producing a spheroid polymer particle. The biodegradable polymer is selected from PLGA, PLA, PGA, PCL, and mixtures thereof. Aliphatic polyesters such as polymers based on lactic acid, glycolic acid, and their copolymers appear to degrade by a bulk erosion process, so that rate of drug release from monolithic devices may not be either linear or predictable [123]. These copolymers displaying bulk or homogenous erosion may display significant degradation in the matrix interior, which may result in “dose dumping” in contrast to surface-eroding systems such as those composed of polyanhydrides [124] and POEs. Efforts have been made to improve the predictability of drug release from PLGA microspheres. WO9109079 A1 (1991, ERBA CARLO SPA; PHARMACIA SPA) discloses the use of supercritical fluids to manufacture biodegradable porous microspheres from PLA and PLGA for drug delivery. WO9211844 A1 (1992, ALKERMES INC.) discloses a method for incorporation of proteins (such as factor VIII) in the form of specific noncovalent complexes with biological polycationic reagents into sustained release systems, which are polymeric microcapsules made, preferably, of PLGA. Biodegradable nanoparticles have received increasing attention as versatile drug delivery scaffolds to enhance the efficacy of therapeutics. Effectiveness of delivery, however, can be influenced by the particle size and morphology, as these parameters can greatly affect the biological function and fate of the material [104,105]. Narrowly dispersed particles

339

are highly preferred for use in delivery or sensing applications with respect to monitoring and predicting their behavior as they exhibit a more constant response to external stimuli [106]. One disadvantage of conventional methods is the irreproducibility in the size and shape of the particles, since these can be profoundly influenced by the stabilizer and the solvent used [125]. Another major drawback of conventional biodegradable nanoparticles, based on PCL and other aliphatic polyesters, is the lack of pendant functional groups, which can make physicochemical, mechanical, and biological properties difficult to modify [126,127]. The availability of functional groups is a desirable means of tailoring the properties of a particle, including hydrophilicity, biodegradation rate, and bioadhesion (2009, WO2009061854 A2, UNIV VANDERBILT). Delivery forms other than microspheres utilizing PLA and PGA polymers have also been formulated. DE2207635 A1 (1973, DU PONT; PFIZER) discloses the preparation of PLA films with active ingredients (e.g., anti-inflammatory agents or antibiotics) incorporated therein, for local application of medicinal film on skin, wounds, burns, etc. US3887699 A (1975, YOLLES SEYMOUR) discloses the use of a shaped dispenser made of biodegradable polymers, preferably PLLA, which contains a drug (antiinflammatory agent or antibiotic) to be controllably released to the surface of the dispenser. The shape of the dispenser depends on its intended use. For example, the device can be in the shape of a film for placing it upon the patient’s body in contact with the skin, thin “spaghetti-like” or “fiber-like” configurations, which can be injected into the bloodstream, hollow tubing suitable for catheters, or spheroids useful for injection or oral ingestion. US3991766 A (1976, AMERICAN CYANAMID CO) discloses filaments comprising PGA and antibiotics. US4118470 A (1978, AMERICAN CYANAMID CO) discloses bioabsorbable films for drug delivery produced from the transesterification product of (1) a PGA composition and (2) a polyester of diglycolic acid and an unhindered glycol. An indicative drug is pilocarpine hydrochloride. Hollow fibers of PLA have also been described [128]. FR2537980 A1 (1989, SANDOZ SA) discloses PLGA partially in the form of amides or esters with a sterol as carriers for active ingredients in the form of microparticles, 1 mm rods and films. EP0290891 A1 (1988, MASSACHUSETTS INST TECHNOLOGY) discloses the use of 3 mm

340

diameter discs with pinholes therein for linear release of drugs. The medical device is formed by encapsulation within an implantable biodegradable polymer one or more compounds, which have the effect of replacing or stimulating functions of the nervous system. A plethora of neurally active substances can be used to treat or manipulate neurological disorders, including Parkinson’s disease. The biodegradable polymer is preferably selected from the group of polyanhydrides, PLA, PGA, POEs, polyanhydrides, and combinations thereof. EP0251476 A1 (1988, SYNTEX INC.; SYNTEX LLC) discloses PLGA films and coated wires, which release growth promotant or growth inhibitory factors over a period of up to 100 days. The polymers comprise a microsuspension of polypeptide and other water-soluble components in which the particles have a diameter of 10 μm or less. GB2234169 A (1991, DEBIOPHARM SA) discloses a method for preparing a ground polymeric product containing salts of peptides including growth hormones comprising compressing, heating, and extruding mixed polymeric powder and salts and grinding the resultant product. WO9320859 A1 (1993, UNIV WASHINGTON) discloses a biodegradable film comprising PLGA, a therapeutically effective amount of a polypeptide growth factor (insulin-like growth factor I or transforming growth factor beta) and a carrier (e.g., albumin). The film may be affixed to the outer surface of an implantable or prosthetic device such as a screw, pin, plate, rod, or artificial joint component. The films and rods are useful therapeutically, such as within methods of enhancing repair of bone fractures. Delivery matrices provided in the form of a coating, for example, on a medical device, are often referred to as drug-eluting coatings. The major driver for the use of drug-eluting coatings is to improve the performance of the medical device, i.e., more successfully treating the disease and/or preventing or reducing undesired side reactions, such as inflammation or infection. Drug-eluting coatings allow the controlled release of biologically or pharmacologically active compounds due to which a therapeutically effective drug concentration can be achieved over a certain period of time. Drug-eluting coatings further allow local site-specific drug delivery. The drug can be delivered locally, thereby allowing the achievement of high concentrations of the active compound at the site where it is most needed. Total drug doses may be significantly lowered thereby preventing the high

Biopolymers: Applications and Trends

systemic concentrations associated with oral administration of the frequently highly toxic drugs. Amorphous PLGA copolymers and PDLLA homopolymers have a number of disadvantages when used in controlled drug release applications. Both polymers are relatively hydrophobic and do not provide an optimal environment for encapsulated proteins. Proteins may adsorb to the polymer resulting in slow and incomplete release, protein unfolding, and/or aggregation. Due to their high subbody temperature Tg’s (>37 °C), both PLGA and PDLLA are rigid matrices. The ability to manipulate the release of an encapsulated drug, especially if the drug has a high molecular weight such as polypeptides, is limited since diffusion of these molecules within the relatively rigid and nonswellable PLGA and PDLLA matrices is negligible. The release of drugs from PLGA and PDLLA matrices, therefore, is initially solely governed by diffusion of dissolved drug molecules through pores. Typically, the encapsulated protein remains entrapped in the polymer matrix until the latter has degraded to such an extent that it loses its integrity or dissolves, resulting in bi- or triphasic degradation-dependent release profiles. Only in a later stage, when hydrolytic degradation has lowered the molecular weight sufficiently, or when (parts of) the polymer matrix start to dissolve, degradation of drug molecules through the polymer matrix becomes possible, generally leading to dose dumping of the encapsulated drug. Furthermore, during degradation of PLGA and PDLLA, acidic degradation products (lactic and glycolic acid) are accumulating in the polymeric matrices due to their glassy character, resulting in the formation of an acidic microenvironment in the polymer matrix with in situ pH that can be as low as 1–2. The pH reduction is caused by a combination of the amount of acidic degradation products, the rate at which they are formed, and the permeability of the polymeric matrix for these degradation products. Under such acidic conditions encapsulated proteins may form aggregates leading to incomplete protein release. Moreover, the low pH may have a deleterious effect on the structural integrity and biological activity of the encapsulated peptide (2013, WO2013015685 A1, INNOCORE TECHNOLOGIES B V). Random PLCL copolymers yield less acidic degradation products. Moreover, these copolymers are not associated with significant pH reductions in the polymer matrix if the polymer matrix is rubbery under body conditions, i.e., Tg < 37 °C. Under these

7: Medical, Dental, and Pharmaceutical Applications

conditions, the polymer matrix is permeable to the degradation products that are released, thereby preventing accumulation and as a result preventing the generation of an acidic environment. Random PLCL copolymers in specific ratios may fulfill the above criteria; however, these materials are very sticky due to which processing into free flowing microspheres, which is a typical prerequisite for the formulation of injectable particulate drug delivery systems, is rather challenging. For the same reason, they are also difficult to handle when used as drug-eluting coatings on medical devices. Sticking can be reduced by increasing the lactide content, but more acidic degradation products will then be formed and the tendency to accumulate will increase as the polymer matrix is more rigid and less permeable for these degradation products (higher Tg). Sticking can also be reduced by increasing the ε-caprolactone content, but then the overall degradation rate of the polymers becomes so low that accumulation of the polymer material at the site of the injection might occur upon repeated injections. Because the reactivity of glycolide, lactide, and ε-caprolactone toward ring opening is very different and the high temperatures that are usually required for complete monomer conversion, it is difficult to obtain a controlled monomer distribution in this type of copolymers. Therefore, also randomly polymerized terpolymers, which are built of these monomers are not suitable enough to create a matrix of polymers with a wide range of polymer properties (2005, EP1555278 A1, INNOCORE TECHNOLOGIES B V). US2004023028 A1 (2004, MAYO FOUNDATION FOR MEDICAL EDUCATION AND RESEARCH) discloses microspheres for controlled release of a bioactive agent including PPF, a polymeric material other than PPF (e.g., PLGA) and a bioactive agent. In bone regeneration applications, the bioactive agent may be selected from osteoinductive agents, peptides, growth hormones, osteoconductive agents, cytokines, and mixtures thereof. The microspheres have a diameter in the range of 1–300 μm and release the bioactive agent in a sustained manner after an initial burst release. The microspheres may be covalently attached to a PPF scaffold for tissue regeneration applications in which the bioactive agent is released from the scaffold. WO2004075781 A2 (2004, MEDIVAS LLC) discloses a stent with a surface coating of a biodegradable bioactive polymer, wherein the polymer comprises at least one bioactive agent covalently bound to the polymer, and wherein the at least one

341

bioactive agent produces a therapeutic effect in situ. The biodegradable polymer is selected from PLA, PGA, PCL, P3HB, poly(amino acid), poly(ester amide), PU, or copolymers thereof. WO2012150255 A1 (2012), WO2012175746 A1 (2012), and WO2013087903 A1 (2013) of DSM IP ASSETS BV disclose an ocular polymer delivery composition sized for injection via a pharmaceutical syringe needle having a bore of 18–30 Gauge, comprising at least one ophthalmologic agent dispersed in biodegradable bis(α-amino-diol-diester) containing poly(ester amide). These specific poly(ester amides) provide unexpected properties in terms of release of ophthalmologic agents over a period of at least 3 months without showing a high burst release in the first 24 h. WO2011045443 A1 (2011, DSM IP ASSETS BV) discloses a coating composition comprising a bis-(αamino-diol-diester) containing poly(ester amide) and a bioactive agent (e.g., rapamycin) from which the release and release rate can be easily tuned. The coating composition is used on a medical device such as a stent. Drug delivery systems utilizing biodegradable polymers are listed in Table 7.10.

7.4 Dentistry Biodegradable polymers have found also use in dental applications. Porous biopolymer particles are packed into the cavity of an extracted tooth to assist in quicker healing. Guided tissue regeneration barrier membrane formed from a biodegradable polymer keeps unwanted gum tissue away from the tooth and underlying bone, for example, in cases of severe paradontitis [136], and permits the undisturbed regrowth of new bone and thus additional stability of the tooth [15]. An ideal material for guided tissue regeneration should be biodegradable because such a material is bioabsorbed by the body, accordingly, no removal operation is necessary. WO9007308 A1 (1990, PROCORDIA ORATECH AB) discloses a barrier for guided or controlled tissue regeneration to be utilized for selective influence on the healing process during regeneration of supporting tissues adjacent to teeth and dental implants as well as during healing after periapical surgery. The barrier can also be used in connection with bone surgery to control bone fill of bone cavities resulting from cysts and malformations and of diastases following bone

342

Biopolymers: Applications and Trends

Table 7.10  Exemplary Biodegradable Controlled Release Formulations Biopolymer

Form

Active ingredient

References

PLA, PLGA

Pellets, microparticles

Endocrine agent (fertility control agent)

FR2070153 A1 (1973, DU PONT)

PLA, PLGA

Microcapsules

Narcotic antagonist, antibiotic, cardiovascular agent, alcohol-sensitizing agent, drug

US4622244 A (1986, UNIV WASHINGTON)

PLA, PGA

Microparticles

Water-insoluble peptide salts (pamoate, tannate, and stearate salts)

GB2209937 A (1989, DEBIOPHARM S A)

PLGA

Porous microspheres

Diagnostic, pharmacologically active drug

US4818542 A (1989, UNIV KENTUCKY RES FOUND)

PLA, PLGA

Porous microspheres

Peptides (e.g., fibroblastic growth factor), antitumor agents (e.g., treatment of the glioblastoma multiform in the brain)

WO9109079 A1 (1991, ERBA CARLO SPA; PHARMACIA SPA)

PLGA

Microcapsules

Peptide

EP0302582 A1 (1989, SOUTHERN RES INST)

PLGA

Microparticles

Peptide hormone (somatostatin, such as an octreotide pamoate salt)

DE4021517 A1 (1991, SANDOZ AG; NOVARTIS AG)

OLGA derivative

Microparticles, rods (1 mm), films

Polypeptide (e.g., salmon calcitonin)

FR2537980 A1 (1989, SANDOZ SA)

PLGA

Microspheres

Complexed proteins or peptides stabilized with biological polycations

WO9211844 A1 (1992, ALKERMES INC.)

PLGA

Microspheres

Progesterone

[129]

PLGA

Microcapsules, microspheres, liposomes

Contraceptive hormones

[113]

PLGA

Microparticles

Salts of peptides including growth hormones

GB2234169 A (1991, DEBIOPHARM SA)

PLA, PGA,PA, POEs

Discs (3 mm) with pinholes

Neurally active substances (e.g., for treatment of Parkinson’s disease)

EP0290891 A1 (1988, MASSACHUSETTS INST TECHNOLOGY)

PPF and PLGA

Microspheres

Osteoinductive agents, peptides, growth hormones, osteoconductive agents, cytokines

US2004023028 A1 (2004, MAYO FOUNDATION FOR MEDICAL EDUCATION AND RESEARCH)

PA

Microspheres

Insulin

[110]

Patent

7: Medical, Dental, and Pharmaceutical Applications

343

Table 7.10  Exemplary Biodegradable Controlled Release Formulations—cont’d Biopolymer

Form

Active ingredient

References

Patent

PCPSA

Pellets

Anesthetic (bupivacaine-HCl)

[130]

POEs

Microspheres

Benzodiazepine (e.g., midazolam)

PLLA

Hollow fibers

Contraceptive hormone (levonorgestrel)

PGA

Filaments

Antibiotics

US3991766 A (1976, AMERICAN CYANAMID CO)

PLGA

Film on implant (screw, pin, plate, rod)

Polypeptide growth factor (insulin-like growth factor I or transforming growth factor beta)

WO9320859 A1 (1993, UNIV WASHINGTON)

PLA

Film

Anti-inflammatory agents (e.g., 1,4-pregnadiene3,20-dionellp,17a,21triol), antibiotics

DE2207635 A1 (1972, DU PONT; PFIZER)

PLLA

Films, hollow tubing, spheroids, “spaghetti-like,” “fiber-like”

Contraceptive steroid (e.g., progestin), opioid antagonists (e.g., cyclazocine)

US3887699 A (1975, YOLLES SEYMOUR)

PLGA

Films, coated wires

Polypeptide (cytokines, lymphokines, monokines, and interferons)

EP0251476 A1 (SYNTEX INC.; SYNTEX LLC)

PA

Film

Anti-inflammatory agents (e.g., salicylates)

WO0141753 A2 (2001, UNIV RUTGERS; UNIV NEW JERSEY MED)

PA

Microparticles

Water-soluble chondrogenic or osteogenic proteins (TGF-beta, EGF, FGF, and PDGF)

[115]

PDLLA

Sandwich device

Insulin

[131]

PLCL

Coating on metal stent

Paclitaxel

[132]

PLCL

Stent

Heparin

[133]

PDLLA, PLTMC

Coating on metal stent

Dexamethasone

[134]

PLLA

Stent

Dexamethasone

[135]

US2009202604 A1 (2009); WO2009129519 A2 (2009, MEDTRONIC INC.) [128]

WO9009783 A1 (1990); US5356630 A, (1994, MASSACHUSETTS INST TECHNOLOGY)

Continued

344

Biopolymers: Applications and Trends

Table 7.10  Exemplary Biodegradable Controlled Release Formulations—cont’d Biopolymer

Form

Active ingredient

References

Patent

PLA

Coating on stent

Tacrolimus (for graft rejection resistant to conventional immunosuppressive regiment)

EP1254674 A1 (2002, INFLOW DYNAMICS INC.; ALT ECKHARD)

PEA

Coating on stent

Rapamycin

WO2011045443 A1 (2011, DSM IP ASSETS BV)

OLGA, Oligo(lactide-co-glycolide); PA, Polyanhydride; PCPSA, Poly(1,3-bis-(p-carboxyphenoxy) propane-co-sebacic acid); PDLLA, Poly(d,l-lactide); PEA, bis-(α-amino-diol-diester) containing poly(ester amide); PGA, Poly(glycolic acid) or polyglycolide; PLA, Poly(lactic acid), polylactide; PLCL, Poly(lactide-co-ε-caprolactone); PLGA, Poly(lactide-co-glycolide); PLLA, Poly(l-lactic acid); PLTMC, Poly(lactide-co-trimethylene carbonate); POE, Poly(orthoester); PPF, Poly(propylene fumarate).

Figure 7.20  Bioabsorbable fabric used in teeth and jaw reconstruction (1993, DE4226465 A1; 1993, JPH05309103 A, GUNZE KK). 1, Fabric; 2, Tooth-root fixing device; 3, Plate; 4, Screw; 5, Lower tooth-root; 6, Upper side toothroot; 7 and 8, Mother bones; 9, Metallic wires.

fractures. Suitable biodegradable polymers are PGA, PLA, scPLA, and copolymers of PLA. Other examples include hyaluronic acid. Still other examples are P3HB and PHBHV and polyesters of succinic acid. DE4226465 A1 (1993), JPH05309103 A (1993), and JPH0542202 A (1993) of GUNZE KK disclose a method for the recovery of the shape and the function of a lost teeth and jaw bone arrangement by installing an artificial tooth-root fixing device integrally on a bioabsorbable fabric (1), which is for cancellous bone comminuted bone marrow transplantation (see Figure 7.20). The fabric (1) is formed of a plain weave fabric made by weaving a monofilament thread of a molten spun PLLA, PDLA, PGA, scPLA, PLGA, PCL, or PLCL in a coarse grid form. An artificial tooth-root fixing device (2) is fixed by inserting a woven thread in a hole provided to the plate (3) of the artificial tooth-root fixing device (2), and fixed to a lower tooth-root (5) made of pure titanium or the like with a screw (4) screwed to the fixing device (2).

An upper gum (6) is screwed up to the lower toothroot (5) with a screw formed at the lower side of the upper side tooth-root (6). The fabric (1) is prepared to fit to the defected part of the teeth and the jaw bone arrangement between mother bones (7) and (8), and it is fixed to the mother bones (7) and (8) with metallic wires (9). WO02089694 A2 (2002, POLYDENT MEDICAL DEVICES LTD.) discloses a dental implant comprising a multilayer structure of more than one polymeric composition; wherein at least one external layer is located adjacent to the bone and is substantially composed of elastic polymer, suitable for either local or systemic delivery of compounds selected from drugs and other substances; and wherein at least one internal layer is composed of relatively nonelastic, nonshape memory, and nonbiocompatible polymers, designed to anchor effectively the abutment (see Figure 7.21). At least one of the external polymeric layers is at least in part a biodegradable polymer

7: Medical, Dental, and Pharmaceutical Applications

Figure 7.21  Cross-section view of a polymeric dental implant comprising gradually homogenous polymeric composition and an abutment (2002, WO02089694 A2, POLYDENT MEDICAL DEVICES LTD.). 1, Polymeric composition; 2, Abutment; 2a, Proximal rim; 2b, Distal rim; 2c, Open bore; 3, Alveolar bone; 4, Gum.

selected from PLLA, PDLA, PDLLA, PGA, PLGA, PDO, PU, blends, and copolymers thereof. At least one of the internal polymeric layers is selected from methacrylate, at least partly crosslinked acrylates, poly(methyl methacrylates), and derivatives thereof. Most attempts to utilize biodegradable barriers have failed due to disturbances of the healing process. For instance, well-known natural nonbiostable materials like collagen and catgut, which are at least partially biodegradable, are not suitable for guided tissue regeneration because they induce rapidly (in 1–2 weeks) inflammatory reactions in living tissues [137,138]. The biodegradable and nonbiodegradable barriers utilized so far have neither been able to predictably differ regeneration of periodontal ligament tissue from regeneration of surrounding alveolar bone tissue nor have they been fully reliable as far as their space-making properties are concerned. Furthermore, these barrier membranes are difficult to handle and implant. The surgical process is time-consuming, cumbersome to the patient and involves considerably high costs (2014, EP1433489 A1, DEGRADABLE SOLUTIONS AG). EP1433489 A1 (2004, DEGRADABLE SOLUTIONS AG) discloses a biodegradable, biocompatible implant (1) for the treatment of a bone defect or a tooth extraction wound, comprising an open porous scaffold (3), and further comprising a membrane (4) which is sealed to a surface portion of the scaffold such that the scaffold and the membrane form a single piece of matter (see Figure 7.22). The scaffold is composed of fused solid, porous or hollow granules (2) (500–1000 μm) selected from bioceramics (e.g.,

345

Figure 7.22  Schematic drawing of a biodegradable, biocompatible implant into a tooth extraction wound (2004, EP1433489 A1, DEGRADABLE SOLUTIONS AG). 1, Biodegradable, biocompatible implant into a tooth extraction wound; 2, Granules of a biocompatible and biodegradable material; 3, Scaffold of interconnected granules; 4, Barrier membrane.

CPPs), or biopolymers such as PLA, PGA, PLGA, PCL, PGTMC, PDO, POEs, poly(ether esters), polyanhydrides, PPHOs, poly(ethylene fumarate), PPF, poly(ester amides), poly(amino acids), polysaccharides, polypeptides, P3HB, P3HV, PUs, poly(malic acid), or copolymers, terpolymers thereof, or blends of those polymers. The membrane is made of a biodegradable polymer such as PLA. Exemplary commercial bioabsorbable membranes include: • Guidor™ (Procordia Oratech A. B.); • Gore Resolut XT™ (W. L. Gore & Associates Inc.) • Vicryl™ Periodontal Mesh (Ethicon Inc.); and • Atrisorb™ Bioabsorbable GTR Barrier (Atrix Laboratories).

7.5  Diagnostic or Therapeutic Imaging Biodegradable polymer particles are effective delivery vehicles for the controlled release of contrast and imaging agents in the human body finding applications in diagnostic and therapeutic imaging. The imaging contrast agents are typically attached to the surface of the polymer particle separately or incorporated along with a bioactive agent. Examples of imaging contrast agents include quantum dots,

346

gold nanoparticles, or Gadalinium–diethylenetriaminepentaacetate (Gd-DTPA) (for use in MRI), which simultaneously image the particles or scaffold in the body as they deliver and/or release the bioactive agent. Among other things, this allows the evaluation of the efficacy of the particle, for example, in reaching the target cells, intracellular uptake, and subsequent bioactive agent release. However, before therapeutic compositions or contrast and imaging agents can be loaded onto a particle’s surface, the surface must be functionalized. Examples of surface fictionalization include the addition of negative charges or amine (NH2) radical groups to a particle’s surface (2007, WO2007008755 A2, UNIV TEXAS). EP0327490 A1 (1989, SCHERING AG) discloses an ultrasonic contrast agent comprising: microparticles made of amylose or an aliphatic polyester of α-, β-, γ-, and ε-hydroxycarboxylic acid (e.g., PLA or PCL); and a gas and/or organic liquid of boiling point below 60 °C. The microparticles are useful for diagnosis and therapy, normally given by injection. They are claimed to have controlled and reproducible volumes; significantly longer in vivo life compared with known products; be well tolerated without allergenic activity; and after intravenous administration they become concentrated in the reticuloendothelial system (including liver and spleen). However, the particles that are produced according to the examples of this patent application have only a relatively small backscatter coefficient. The diagnostic action of contrast medium preparations that are prepared from them is, therefore, not satisfactory in all cases (1996, WO9628191 A1, SCHERING AG). WO9628191 A1 (1996, SCHERING AG) discloses a method for the production of improved gaseous microparticles for ultrasonic diagnosis, whose wall material is built up also from aliphatic polyesters of α-, β-, γ-, or ε-hydroxycarboxylic acids, including the following steps: (1) dissolving the polyester(s), and optionally a surface-active substance, in an organic solvent or solvent mixture to obtain a polyester solution, of which at least one solvent is readily water-miscible; (2) dispersing a liquid perfluorinated compound, which is not a solvent for the polymer, in the polyester solution to obtain a dispersion; (3) dispersing the dispersion in water that contains a surfactant using a stirring mechanism; (4) removing the solvent by pumping in gas and applying a vacuum to obtain a suspension; and (5) mixing the suspension with a suitable pharmaceutically acceptable

Biopolymers: Applications and Trends

cryoprotector and freeze-drying to yield gaseous microparticles. The microparticles are rapidly broken down in vivo, and the breakdown products are nontoxic. The microparticles are easy to produce and are well tolerated by the body. They are stable enough to survive passage through the lung and are thus suitable for contrasting of the left side of the heart. They are taken up by the reticuloendothelial system and are thus useful for contrasting of the liver and the spleen. The microparticles have a higher backscattering coefficient than the particles described in the previous patent application. The aliphatic polyesters are preferably selected from PGA, PLGA, PLLA, PDLA, PDLLA, PLDLLA, P3HB, PHBHV, PDO, or poly(3-hydroxypropionate). JP2000226337 A (2000, YOKOGAWA MEDICAL SYST) discloses a microballoon used as contrast medium in an ultrasonic apparatus having an outer covering (22) containing copolymer of PBS and poly(ethylene succinate), copolymer of PBS and poly(butylene adipate), or copolymer of PBS and poly(butylene sebacate) (see Figure 7.23). The microballoon contrast medium injected into the apparatus is irradiated with ultrasonic wave and the image is picked up based on an echo. WO2008029599 A1 (2008, CANON KK) discloses magnetic particles used as MRI contrast medium and drug delivery system carrier in diagnostic and pharmaceutical field, containing biodegradable polymer and ultrafine magnetic particles of various ferrites, preferably magnetite, having an average particle diameter in the range of 10–1000 nm. The biodegradable polymer is preferably an aliphatic polyester such as PLA and PLGA. WO2012133295 A1 (2012, JMS CO LTD.) discloses a bioabsorbable antiadhesive material useful for medical device for X-ray image and MRI image diagnostics comprising a biodegradable polymer

Figure 7.23 Microballoon contrast medium (2000, JP2000226337 A, YOKOGAWA MEDICAL SYST). 22, Outer covering; 24, Gas.

7: Medical, Dental, and Pharmaceutical Applications

containing a contrast agent. A preferred biodegradable polymer is PLCL. A preferred contrast agent is barium sulfate. The bioabsorbable antiadhesive material has a multilayered structure comprising a layer containing the contrast agent, and a layer which does not contain the contrast agent, where the layer which does not contain the contrast agent of two layers is

347

stacked through the layer containing the contrast agent (see Figure 7.24). The bioabsorbable antiadhesive material can be externally detectable in vitro in a patient body, hence the arrangement state of the bioabsorbable antiadhesive material can be checked without incising a patient, and has improved physical intensity.

Figure 7.24  (a) and (b) perspective views of multilayered structures of the bioabsorbable antiadhesive material (2012, WO2012133295 A1, JMS CO LTD.). 11, Lowest layer; 12, Middle layer; 13, Uppermost layer; 21, Lowest layer (contrast agent-not-containing layer); 22 and 24, Contrast agent content layer; 23, Contrast agent-notcontaining layer; 25, Uppermost layer (contrast agent-not-containing layer).

Patents Patent number

Publication date

BE654236 A

Family members

Priority numbers

Inventors

Applicants

Title

19650409

US3297033 A 19670110; GB1043518 A 19660921

US19630320543 19631031

SCHMITT EDWARD EMIL; POLISTINA ROCCO ALBERT

AMERICAN CYANAMID CO

File de suture chirurgicaux. “Surgical sutures”

CA2025626 A1

19910328

US5085629 A 19920204; JPH03205059 A 19910906; EP0420541 A2 19910403; EP0420541 A3 19920422; AU6364390 A 19910411

US19890414651 19890927

GOLDBERG JAY R; SINCLAIR RICHARD G

SQUIBB BRISTOL MYERS CO

Biodegradable stent

DE19539449 A1

19970430

US5935506 A 19990810; EP0770401 A2 19970502; EP0770401 A3 20000419; EP0770401 B1 20021120

DE1995139449 19951024

SCHMITZ KLAUS-PETER; BEHREND DETLEF

BIOTRONIK MESS & THERAPIEG

Verfahren zur Herstellung intraluminaler Stents aus bioresorbierbarem, Polymermaterial. “Method of manufacturing intraluminal stents made of bioresorbable polymer material”

DE19613048 A1

19961002

DE19613048 C2 19971218

DE1996113048 19960401; DE1995111718 19950330

GRUNZE MICHAEL

GRUNZE MICHAEL; SCHRENK MONIKA

Künstliche Implantate mit antithrombogenen Eigenschaften und Verfahren zu deren Herstellung. “Artificial implant, with antithrombogenic properties and method of its production”

DE2207635 A1

19720831

US3755558 A 19730828; GB1351409 A 19740501; FR2126270 A1 19721006; FR2126270 B1 19751128; CA950828 A1 19740709

US19710118081 19710223

SCRIBNER R; BENZOTHIENO; HESS H; DE ANGELIS G

DU PONT; PFIZER

Polylactide–drug mixtures for topical application

DE3445711 A1

19860619

WO8603671 A1 19860703; EP0511686 A2 19921104; JPS61501034 A 19860522; JP2620227 B2 19970611; EP0511686 A3 19921216; EP0204786 A1 19861217; EP0204786 B1 19921202; DE3445731 A1 19860619; AT82844 T 1992121

DE19843445711 19841214; DE19843445731 19841214

DRAENERT KLAUS

DRAENERT KLAUS

Knochenersatzwerkstoff und seine Verwendung. “Bone replacement material and utilization thereof”

DE4021517 A1

19910117

SE512992 C2 20000612; SE9002364 L 19910108; PT94628 A 19910320; PT94628 B 19970630; NZ234384 A 19940526; NO903001 A 19910108; NO302928 B1 19980511; NO983923 A 19910108; NO320444 B1 20051205; NL9001537 A 19910201; NL195027 B 20030801; NL195027 C 20031202; LU87764 A1 19920311; KR100442931 B1 20040723; JPH07285853 A 19951031; JP2931773 B2 19990809; JPH0368511 A 19910325; JPH0832624 B2 19960329; JPH07309897 A 19951128; JPH08198771 A 19960806; IT1241460 B 19940117; IT9048113 A1 19910108; IL131880 A 20011223; IL131881 A 20011223; IL94983 A 19990817; IE902435 A1 19910213; IE64216 B1 19950726; IE64411 B1 19950809; HU211602 A9 19951228; HU221294 B1 20020928; HUT54037 A 19910128; HK197496 A 19961108; HK97695 A 19950623;

US19890377023 19890707; US19890411347 19890922; DE19904042746 19900706; DE19904042753 19900706; DE19904042752 19900706

BODMER DAVID; FONG JONES W; KISSEL THOMAS; MAULDING HAWKINS V; NAGELE OSKAR; PEARSON JANE E

SANDOZ AG; NOVARTIS AG

Formulierungen mit verlangsamter Freisetzung wasserlöslicher Peptide. “Sustained release formulations of water soluble peptides”

Continued

—cont’d Patent number

Publication date

Family members

Priority numbers

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Applicants

Title

DE19914106823 19910304

METZLER RICHARD; SCHMIEDING REINHOLD; SCHMID PETER M

LIEBSCHER KUNSTSTOFFTECHNIK; ARTHREX MED INSTR GMBH

Knochendübel zur Fadenfixierung. “Bone dowel for fixing threads”

GR90100513 A 19911210; GR1001121 B 19930428; GB2265311 A 19930929; JP2001233897 A 20010828; GB2265311 B 19940209; FR2649319 A1 19910111; FR2649319 B1 19941209; FI108611 B 20020228; FI20000060 A 20000112; FI109334 B 20020715; FI20000059 A 20000112; FI109543 B 20020830; DK162590 A 19910108; DK175849 B1 20050329; DE4021517 B4 20090409; DE4042752 B4 20090507; CY1965 A 19970704; CH685230 A5 19950515; CA2020477 A1 19910108; CA2020477 C 20001121; BE1004486 A3 19921201; AU2332195 A 19950907; AU5874690 A 19910110; AU641407 B2 19930923; AU687553 B2 19980226; ATA144090 A 19990815; AT406225 B 20000327 DE4106823 C1

19920625

US5336240 A 19940809; EP0502509 A1 19920909; EP0502509 B1 19960911; AT142454 T 19960915

DE4226465 A1

19930211

DE4226465 C2 20031204

DE19924226465 19920810

IKADA YOSHITO; SHIMURA KAIZO; KINOSHITA YUKIHIKO; KIRIGAKUBO MITSUHIRO; KOBAYASHI MASARU; SUZUKI MASAKAZU; NISHIYA KOJI

GUNZE KK

Kieferknochenreproduzierendes Material. “Material for maxillary reconstruction”

EP0052998 A1

19820602

US4338926 A 19820713; JPS57117852 A 19820722; JPH0229333 B2 19900628; EP0052998 B1 19840808; CA1166108 A1 19840424;

US19800208906 19801121

KUMMER FREDERICK J; COUTTS RICHARD D

HOWMEDICA

Bone prosthesis with controlled stiffness

EP0108635 A2

19840516

ZA8308283 A 19850626; US4539981 A 19850910; NZ206057 A 19860221; JPS5997654 A 19840605; JPH0316866 B2 19910306; GR81278 A1 19841211; ES8605831 A1 19860916; EP0108635 A3 19850529; EP0108635 B1 19880608; EP0108635 B2 19960110; CA1230195 A1 19871208; AU2104383 A 19840517; AU561150 B2 19870430

US19820439962 19821108

TUNC DEGER C

JOHNSON & JOHNSON PROD INC

Absorbable bone fixation device

EP0159502 A2

19851030

ZA8503112 A 19851224; US4838884 A 19890613; US4633873 A 19870106; SG9590511 A2 19950818; PT80340 B 19870529; PH23747 A 19891103; NO851663 A 19851028; NO171485 B 19921214; NO171485 C 19930324; KR930000688 B1 19930129; JPS60234667 A 19851121; JPH0554352 B2 19930812; JP2000093496 A 20000404; IL74752 A 19910816; HK67895 A 19950512;

US19840604104 19840426

DUMICAN BARRY L; KAGANOV ALAN L; RITTER THOMAS A

AMERICAN CYANAMID CO

Absorbable surgical repair mesh

Continued

—cont’d Patent number

Publication date

Family members

Priority numbers

Inventors

Applicants

Title

GR850997 A1 19851125; FI851633 A 19851027; FI86257 B 19920430; FI86257 C 19920810; ES8606919 A1 19861016; EP0159502 A3 19870715; EP0159502 B1 19930908; DK186385 A 19851027; DK169400 B1 19941024; DE3587563 T2 19940414; CA1322915 C 19931012; AU4164185 A 19851031; AU574457 B2 19880707 EP0185453 A2

19860625

ZA8508620 A 19870624; US4646741 A 19870303; SG96591 G 19920117; MX163348 A 19920427; KR920007983 B1 19920921; JPS61179140 A 19860811; JPH0634812 B2 19940511; IE852796 L 19860509; IE57500 B1 19930310; HK107491 A 19920103; GR852705 A1 19860310; EP0185453 A3 19871209; EP0185453 B1 19911009; DK517785 A 19860510; CA1241888 A1 19880913; BR8505611 A 19860812; AU576403 B2 19880825; AU4970385 A 19860717

US19840670105 19841109

SMITH CARL ROBERT;

ETHICON INC

Surgical fastener made from polymeric blends

EP0209371 A1

19870121

ZA8605313 A 19880224; US4741337 A 19880503; JPS6219174 A 19870127; JPH0793940 B2 19951011; EP0209371 B1 19900131; CA1255044 A1 19890530; AU6024486 A 19870122; AU585038 B2 19890608

US19850755888 19850717

SMITH CARL ROBERT; GATERUD MARK TURNER; JAMIOLKOWSKI DENNIS DOUGLAS; NEWMAN JR HUGH DAMIAN; SHALABY SHALABY WAHBA

ETHICON INC

Surgical fastener made from glycolide-rich polymer blends

EP0239775 A2

19871007

ZA8701489 A 19870821; US4792336 A 19881220; US4942875 A 19900724; NO870858 A 19870904; NO166614 C 19910821; KR950001375 B1 19950218; JPS62270152 A 19871124; FI870898 A 19870904; FI88259 B 19930115; EP0239775 A3 19890726; DK106687 A 19870904; CA1282205 C 19910402; AU6959887 A 19870910; AU594435 B2 19900308

US19860835493 19860303

DUMICAN BARRY L; HLAVACEK ROBERT A; MCCUSKER EDWARD J

AMERICAN CYANAMID CO

Surgical repair device

EP0251476 A1

19880107

US4962091 A 19901009; NZ220403 A 19900626; JPS632930 A 19880107; JP2594560 B2 19970326; HK38997 A 19970404; ES2039437 T3 19931001; ES2039437 T5 20030616; EP0251476 B1 19930203; EP0251476 B2 20021127; DE3783958 T2 19930609; DE3783958 T3 20030807; CA1293443 C 19911224; AU608225 B2 19910328; AU7333287 A 19871126

US19860866625 19860523

EPPSTEIN DEBORAH ANN; SCHRYVER BRIAN BLOOR

SYNTEX INC; SYNTEX LLC

Controlled release of macromolecular polypeptides

EP0290891 A1

19881117

US4883666 A 19891128; ES2088863 T3 19961001; EP0290891 B1 19951108; DE3854656 T2 19960620; AT129892 T 19951115

US19870043695 19870429

SABEL BERNHARD; FREESE ANDREW; SALTZMAN WILLIAM MARK

MASSACHUSETTS INST TECHNOLOGY

Controlled drug delivery system for treatment of neural disorders Continued

—cont’d Patent number

Publication date

EP0302582 A1

Family members

Priority numbers

Inventors

Applicants

Title

19890208

US4897268 A 19900130; PH24813 A 19901030; NZ225608 A 19901221; MX162898 B 19910708; JPS6442420 A 19890214; IL86221 A 19920818; HK1007490 A1 19990416; GR3026376 T3 19980630; ES2056915 T3 19941016; ES2056915 T5 19980401; EP0302582 B1 19940727; EP0302582 B2 19971217; DE3850823 T2 19941117; DE3850823 T3 19980709; CA1302260 C 19920602; BR8801242 A 19890221; AU1099288 A 19890209; AU611662 B2 19910620; AT109000 T 19940815

US19870081289 19870803

TICE THOMAS R; GILLEY RICHARD M

SOUTHERN RES INST

Drug delivery system and method of making the same

EP0326426 A2

19890802

US4950258 A 19900821; JPH01192367 A 19890802; JP2561853 B2 19961211; EP0326426 A3 19920115; EP0326426 B1 19941221; DE68920055 T2 19950511; CA1307885 C 19920929

JP19880017574 19880128

KAWAI TATSUYA; MATSUDA TAKASHI

NIPPON MEDICAL SUPPLY

Plastic molded articles with shape memory property

EP0327490 A1

19890809

WO8906978 A1 19890810; US6264959 B1 20010724; US6071496 A 20000606; US6177062 B1 20010123; PT89635 A 19891004; EP0586875 A1 19940316; NZ227869 A 19921125; NO970732 A 19901003; NO304412 B1 19981214; NO903443 A 19901003;

DE19883803971 19880205; DE19883803972 19880205

STEIN MICHAEL; HELDMANN DIETER; FRITZSCH THOMAS; SIEGERT JOACHIM; ROESSLING GEORG; SPECK ULRICH

SCHERING AG

Ultrasonic contrast agents, process for producing them and their use as diagnostic and therapeutic agents

NO301260 B1 19971006; KR0133132 B1 19980417; JPH03503634 A 19910815; JP2907911 B2 19990621; JPH08208524 A 19960813; JP3027326 B2 20000404; IL89175 A 19930221; IE940809 L 19890805; IE890343 L 19890805; IE66912 B1 19960207; HUT59322 A 19920528; HU221485 B 20021028; FI99086 B 19970630; FI99086 C 19971010; ES2068917 T3 19950501; EP0398935 A1 19901128; EP0398935 B1 19940810; DK186490 A 19900823; DK175832 B1 20050314; CN1035437 A 19890913; CN1033840 C 19970122; CA1336164 C 19950704; AU3035189 A 19890825; AU635200 B2 19930318; AT109663 T 19940815 EP0334046 A2

19890927

ZA8902235 A 19891129; NO169874 B 19920511; NO891140 A 19890925; NO169874 C 19920819; MX169561 B 19930712; JPH0211150 A 19900116; HK1007965 A1 19990430; FI891404 A 19890925; ES2103701 T3 19971001; EP0334046 A3 19900613; EP0334046 B1 19970618; DK145889 A 19890925; DE68928132 T2 19971016; CA1302140 C 19920602; AU3170389 A 19890928; AU616491 B2 19911031; AT154516 T 19970715

US19880172607 19880324

DEVEREUX DENNIS F; LANDI HENRY P

AMERICAN CYANAMID CO

Surgical composite structure having absorbable and nonabsorbable components

Continued

—cont’d Patent number

Publication date

EP0371736 A2

Family members

Priority numbers

Inventors

Applicants

Title

19900606

US5035893 A 19910730; JPH02147062 A 19900606; JPH06104116 B2 19941221; EP0371736 A3 19901227; EP0371736 B1 19940413; DE68914604 T2 19940804

JP19880301856 19881129

SHIOYA NOBUYUKI; KUROYANAGI YOSHIMITSU; KOUNAMI YASUO; KOBAYASHI TATSUHIKO

MITSUBISHI CHEM IND

Wound covering

EP0530585 A2

19930310

US5275601 A 19940104; US5275601 X6 19940104; EP0530585 A3 19940119; EP0530585 B1 19961218; JPH07178115 A 19950718; JP3323244 B2 20020909; DE69216009 T2 19970403; CA2076501 A1 19930304; AT146351 T 19970115

US19910753837 19910903

GOGOLEWSKI SYLWESTER; PERREN STEPHAN M

SYNTHES AG

Self-locking resorbable screws and plates for internal fixation of bone fractures and tendon-to-bone attachment

EP0560014 A1

19930915

US5792469 A 19980811; JPH067423 A 19940118; CA2091552 A1 19930913; AU3117493 A 19930916

US19920849896 19920312

TIPTON ARTHUR J; FUJITA SHAWN M; DUNN RICHARD L

ATRIX LAB INC

Biodegradable film dressing and method for its formation

EP0610731 A1

19940817

US5795584 A 19980818; EP0610731 B1 20020918; DE69431376 T2 20030515; CA2114290 A1 19940728; CA2114290 C 20060110

US19930009726 19930127; US19930153336 19931116

TOTAKURA NAGABUSHANAM; MUTH ROSS R; GRAVENER ROY D; HAIN MATTHEW; KOYFMAN ILYA S

UNITED STATES SURGICAL CORP.

Post-surgical antiadhesion device

EP0668083 A1

19950823

US5626611 A 19970506; US5578046 A 19961126; EP0668083 B1 20020717; EP0668083 B2 20050330; DE69527382 T2 20030403; DE69527382 T3 20051229; CA2140090 A1 19950811; CA2140090 C 20060620

US19940194766 19940210

LIU CHENG-KUNG; JIANG YING; ROBY MARK S; BENNETT STEVEN L; STEVENSON RICHARD P; PIETRANGELI JEFFREY P

UNITED STATES SURGICAL CORP.

Composite bioabsorbable materials and surgical articles made thereform

EP0689807 A2

19960103

US5766710 A 19980616; US5766710 X6 19980616; US5629077 A 19970513; JPH0824346 A 19960130; EP0689807 A3 19960807; CA2152647 A1 19951228; CA2152647 C 20001031

US19940266964 19940627

TURNLUND TODD HANSON; EURY ROBERT PAUL

ADVANCED CARDIOVASCULAR SYSTEM

Biodegradable mesh and film stent

EP0691359 A2

19960110

ZA9505549 A 19970106; US5869597 A 19990209; US6090908 A 20000718; US5611986 A 19970318; JPH0852205 A 19960227; JP3794733 B2 20060712; EP0691359 A3 19971112; EP0691359 B1 20021009; CA2153196 A1 19960106; BR9503089 A 19960423; AU2483495 A 19960118

US19940270712 19940705

JAMIOLKOWSKI DENNIS D; NEWMAN JR HUGH DAMIAN; DATTA ARINDAM; TANDON ROHIT; CHEN I-JEN; SURYADEVARA JOGENDRA; CANTERBERRY BETHANY ANN; FITZGERALD SCOTT EVANS

ETHICON INC

Medical devices containing high inherent viscosity poly(p-dioxanone)

EP0711548 A1

19960515

ZA9508770 A 19970417; US5599852 A 19970204; US5728752 A 19980317; JPH08206191 A 19960813; EP0711548 B1 19980128; DE69501540 T2 19980604; CA2160767 A1 19960419; CA2160767 C 20070731; AU3422395 A 19960502; AU696591 B2 19980917; BR9504447 A 19970520

US19940324543 19941018

SCOPELIANOS ANGELO G; ARNOLD STEVEN C; BEZWADA RAO S; ROLLER MARK B; HUXEL SHAWN T

ETHICON INC

Injectable microdispersions for soft tissue repair and augmentation

EP0714666 A1

19960605

ZA9510147 A 19970529; US5679723 A 19971021; US5747390 A 19980505; JPH08215299 A 19960827; CA2164045 A1 19960531; BR9505580 A 19971104; AU3795395 A 19960606

US19940346652 19941130

COOPER KEVIN; CHEN CHAO C; SCOPELIANOS ANGELO G

ETHICON INC

Hard tissue bone cements and substitutes

EP0763559 A2

19970319

JPH09157513 A 19970617; EP0763559 A3 19980826; CN1149596 A 19970514; CA2183923 A1 19970224

US19950518258 19950823

FARR MICHAEL PATRICK; LINCOLN DAVID MARSTON; MOYERS CHARLES GUTHRIE

UNION CARBIDE CHEM PLASTIC

Stabilized dioxanone polymers

Continued

—cont’d Patent number

Publication date

EP0908142 A2

Family members

Priority numbers

Inventors

Applicants

Title

19990414

US2001018599 A1 20010830; US6712838 B2 20040330; JPH11217753 A 19990810; JP4354031 B2 20091028; EP0908142 A3 19990915; EP0908142 B1 20060503; DE69834375 T2 20070315

US19970061721P 19971010; US19980159025 19980923

D’ AVERSA MARGARET; SCALZO HOWARD L JR; JAMIOLKOWSKI DENNIS D; BEZWADA RAO S; HUNTER ALASTAIR W; HILL DONALD G

ETHICON INC

Braided suture with improved knot strength and process to produce same

EP1093824 A2

20010425

JP2001178812 A 20010703; EP1093824 A3 20040102; DE19950646 A1 20010503

DE1999150646 19991021

BURKHARD INGRID; PFEFFERLE HANS

MERCK PATENT GMBH

Gefärbte Materialien aus bioabbaubaren Polymeren im medizinischpharmakologischen Bereich. “Dyed biodegradable polymeric materials for medicalpharmacological use”

EP1138285 A1

20011004

EP20010810261 20010315

BUEHLER MARKUS; SGIER FRIEDRICH

IMPLANT DESIGN AG

Spinal cage for insertion between the vertebraes of the spine

EP1254674 A1

20021106

US20010847626 20010502

ALT ECKHARD

INFLOW DYNAMICS INC; ALT ECKHARD

Endovascular stent with coating comprising tacrolimus

US2002165607 A1 20021107; US6613083 B2 20030902; US2004019376 A1 20040129; US7011680 B2 20060314; US2006122690 A1 20060608; US7625410 B2 20091201; EP1254674 B1 20040915; EP1254674 B2 20100818; DE60105554 T2 20051117; DE60105554 T3 20110224

EP1433489 A1

20040630

WO2004056405 A2 20040708; WO2004056405 A3 20040826; US2006136071 A1 20060622; US7731756 B2 20100608; JP2006515767 A 20060608; JP4873600 B2 20120208; EP1575636 A2 20050921; AU2003296713 A8 20040714; ES2369080 T3 20111125; EP1575636 B1 20110810; AU2003296713 A1 20040714; AT519509 T 20110815

EP20020406138 20021223

MASPERO FABRIZIO ALESSANDRO; RUFFIEUX KURT

DEGRADABLE Biodegradable porous SOLUTIONS AG bone implant with a barrier membrane sealed thereto

EP1555278 A1

20050720

WO2005068533 A1 20050728; US2013273284 A1 20131017; US8674032 B2 20140318; US2007155906 A1 20070705; US8481651 B2 20130709; JP2007522274 A 20070809; JP4907359 B2 20120328; EP1709103 A1 20061011; CN1910217 A 20070207; CN1910217 B 20120620; CA2553619 A1 20050728; CA2553619 C 20120814; AU2005205409 A1 20050728; AU2005205409 B2 20100513

EP20040075099 20040115

HISSINK CATHARINA EVERDINA; MEYBOOM RONALD; STEENDAM ROB; FLIPSEN THEODORUS

INNOCORE TECHNOLOGIES B V

Biodegradable multiblock co-polymers

EP1591133 A1

20051102

WO2005105168 A1 20051110

EP20040010309 20040430

MONTALI ANDREA; SCHLOTTIG FALKO

SYNTHES

Biologically active implants

EP1600182 A1

20051130

US2005278015 A1 20051215; US7803182 B2 20100928; KR20060046240 A 20060517; JP2005334653 A 20051208; JP5318318 B2 20131016; IL168359 A 20110831; HK1083469 A1 20110916; EP1600182 B1 20110126; CA2508588 A1 20051128; AU2005201883 A1 20051215; AU2005201883 B2 20101028; AT496642 T 20110215

US20040856459 20040528

DAVE VIPUL BHUPENDRA; LANDAU GEORGE; PATEL PREMAL

CORDIS CORP.

Biodegradable vascular device with buffering agent

Continued

—cont’d Patent number

Publication date

EP1738780 A2

20070103

EP2425865 A1

20120307

EP2450487 A1

20120509

FR2070153 A1

FR2439003 A1

Family members

Priority numbers

Inventors

Applicants

Title

US2005283256 A1 20051222; US8795710 B2 20140805; JP2007007423 A 20070118; JP5188685 B2 20130424; EP1738780 A3 20081231; CA2551366 A1 20061230; AU2006202593 A1 20070118; AU2006202593 B2 20120412

US20050171638 20050630

SOMMERICH ROBERT E; CODMAN & MACOMBER LAUREL R SHURTLEFF

Collagen device and method of preparing the same

EP20100172204 20100806

ODERMATT ERICH; BARGON RAINER

AESCULAP AG

Medicinal thread having a polyhydroxyalkanoate coating

WO2012059847 A1 20120510; AR083752 A1 20130320

EP20100290595 20101105

SIMON CHRISTOPHE; DEBRAY FLORENCE; LEBRETTE LAURENT

ARJOWIGGINS PALALDA SAS

Biodegradable sterilization wrap

19710910

US3773919 A 19731120; JPS5017525 B1 19750621; GB1325209 A 19730801; FR2070153 B1 19750110; DE2051580 A1 19710506; DE2051580 B2 19800710; DE2051580 C3 19810402; CA982479 A1 19760127

US19690868899 19691023; US19700079309 19701008

BOSWELL G; SCRIBNER R

DU PONT

Compositions pharmaceutiques libérant das quantités efficaces de médicament pendant une période de temps donnée. “Pharmaceutical compostions which liberate an efficient quantity of the drug over a controlled period of time”

19800516

US4279249 A 19810721; FR2439003 B1 19821217; EP0011528 A1 19800528; EP0011528 B1 19850220;

FR19780029878 19781020

VERT MICHEL; CHABOT FRANCOIS; LERAY JEAN; CHRISTEL PASCAL

ANVAR

Nouvelles pièces d’ostéosynthèse, ainsi que leur préparation et leur application. “New prosthesis parts, their preparation and their application”

FR2537980 A1

US4801739 A 19890131; IT1197759 B 19881206; HK27290 A 19900420; GB2135320 A 19840830; GB2135320 B 19870429; GB2165849 A 19860423; GB2166652 A 19860514; GB2166652 B 19870311; GB2165454 A 19860416; GB2165454 B 19870311; FR2537980 B1 19861219; DE3345314 A1 19840705; CY1529 A 19901116; CY1530 A 19901116

CH19820007372 19821217; CH19820007373 19821217

RANZ JOACHIM; PRIKOSZOVICH WALTER; BRICH ZDENEK

SANDOZ SA

Dérivés d’acides hydroxycarboxyliques oligomères, leur preparation et leur utilization. “Oligomeric hydroxycarboxylic acid derivatives, their production and use”

GB1048088 A

19661109

NL299454 A 00000000; NL147329 B 19751015; JPS412734 B1 19660221; DE1492427 A1 19691204; CH461032 A 19680815

US19620231860 19621019; US19630308688 19630913

KENNETH SCHNEIDER ALLAN

DU PONT (ETHICON INC)

Surgical articles

GB2069650 A

19810826

US4317451 A 19820302; JPS56151033 A 19811121; GB2069650 B 19831109; FR2475886 A1 19810821; FR2475886 B1 19841228; DE3105957 A1 19820107; CA1144445 A1 19830412

US19800122557 19800219

CERWIN ROBERT J; MCVAY WILLIAM P

ETHICON INC

Plastic surgical staple

GB2075144 A

19811111

JPS56168737 A 19811225; GB2075144 B 19831012; FR2481594 A1 19811106; FR2481594 B1 19850104; DE3117277 A1 19820624; CA1162127 A1 19840214

US19800146943 19800502

MERICLE ROBERT W

ETHICON INC

Plastics surgical staple

GB2166354 A

19860508

GB2166354 B 19871209

GB19850023824 19850927; GB19840025545 19841010

WEBB ANDREW; ADSETTS DR JOHN ROBERT

ICI PLC

Wound dressings

GB2174909 A

19861119

US4674488 A 19870623; DE3607075 A1 19860904

US19850707954 19850304

NASHEF AWS S; CAMPBELL TODD D

AMERICAN HOSPITAL SUPPLY CORP.

Method of treating bone fractures to reduce formation of fibrous adhesions Continued

—cont’d Patent number

Publication date

GB2209937 A

19890601

Family members

Priority numbers

Inventors

Applicants

Title

ZA8806827 A 19890530; US5192741 A 19930309; US5776885 A 19980707; SE8803321 L 19890322; SE503406 C2 19960610; PT88557 A 19881001; PT88557 B 19921130; NO2005012 I2 20081201; NO884154 A 19890322; NO178604 B 19960122; NO178604 C 19960502; NL8802323 A 19890417; NL193818 B 20000801; NL193818 C 2000124; LU87340 A1 19890406; JPH01121222 A 19890512; JPH0713023 B2 19950215; IT1225148 B 19901102; IL87790 A 19920525; IE882727 L 19890321; IE60608 B1 19940727; GR88100619 A 19890622; GR1002244 B 19960422; GB2209937 B 19910703; FR2620621 A1 19890324; FR2620621 B1 19930219; FI884297 A 19890322; FI96919 B 19960614; FI96919 C 19960925; ES2009346 A6 19890916; DK518988 A 19890322; DK175311 B1 20040816; DE122004000023 I2 20090507; DE3822459 A1 19890330; DE3822459 C2 19940707; CH675968 A5 19901130; CA1326438 C 19940125; BE1001685 A5 19900206; AU2232688 A 19890323; AU611944 B2 19910627; ATA223488 A 19930615; AT397035 B 19940125

GB19870022134 19870921

ORSOLINI PIERO; MAUVERNAY ROLLAND-YVES; DEGHENGHI ROMANO

DEBIOPHARM S A Water insoluble polypeptides

GB2234169 A

19910130

ZA9005654 A 19910529; US5134122 A 19920728; SE504279 C2 19961223; SE9002522 L 19910129; PT94842 A 19910320; PT94842 B 19970430; NO903264 A 19910129; NO300304 B1 19970512; NL9001646 A 19910218; NL194858 B 20030106; NL194858 C 20030506; LU87772 A1 19901211; JPH0366625 A 19910322; JPH0662427 B2 19940817; IT9021019 A1 19910129; IT1243357 B 19940610; IL95120 A 19941007; IE902592 A1 19910227; IE65397 B1 19951018; GR90100568 A 19911210; GR1001215 B 19930621; GB2234169 B 19930210; FR2650182 A1 19910201; FR2650182 B1 19920515; FI97688 B 19961031; FI97688 C 19970210; ES2020890 A6 19911001; DK180790 A 19910129; DK175495 B1 20041108; DE4023134 A1 19910131; DE4023134 C2 19970417; CH679207 A5 19920115; CA2021767 A1 19910129; CA2021767 C 19961022; BE1003093 A3 19911119; AU5910390 A 19910131; AU619996 B2 19920206; ATA154590 A 19930715; AT397197 B 19940225

CH19890002829 19890728

ORSOLINI PIERO

DEBIOPHARM SA A method for preparing a sustained release pharmaceutical peptide composition

JP2000160125 A

20000613

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Priority numbers

Inventors

Applicants

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Priority numbers

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Priority numbers

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Applicants

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AMERICAN CYANAMID CO

Polyglycolic acid prosthetic devices

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Continued

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Publication date

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19711116

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19720125

Family members

ZA7100808 A 19711124; US3636956 X6 19720125; SE361599 B 19731112; NO132784 B 19750929; NO132784 C 19760107; NL7103263 A 19711116; JPS4936597 B1 19741002; FR2088548 A1 19720107;

Priority numbers

Inventors

Applicants

Title

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SCHMITT EDWARD EMIL; POLISTINA ROCCO ALBERT

AMERICAN CYANAMID CO

Cylindrical prosthetic devices of polyglycolic acid

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SCHNEIDER ALLAN K

ETHICON INC

Polylactide sutures

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SCHMITT EDWARD EMIL; POLISTINA ROCCO ALBERT

AMERICAN CYANAMID CO

Polyglycolic acid prosthetic devices

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SCHMITT EDWARD EMIL; EPSTEIN MARTIN

AMERICAN CYANAMID CO

Reducing capillarity of polyglycolic acid sutures

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SCHMITT EDWARD EMIL; POLISTINA ROCCO ALBERT

AMERICAN CYANAMID CO

Surgical dressings of absorbable polymers

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YOLLES SEYMOUR

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Biodegradable polymeric article for dispensing drugs

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DODDI NAMASSIVAYA; VERSFELT CHARLES C; WASSERMAN DAVID

ETHICON INC

Synthetic absorbable surgical devices of polydioxanone

Continued

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Publication date

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19781003

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Family members

Priority numbers

Inventors

Applicants

Title

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CASEY DONALD JAMES; EPSTEIN MARTIN

AMERICAN CYANAMID CO

Normally-solid, bioabsorbable, hydrolyzable, polymeric reaction product

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ROSENSAFT MICHAEL N; WEBB RICHARD L

AMERICAN CYANAMID CO

Synthetic polyester surgical articles

19840131

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ETHICON INC

Plastic surgical staple

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TUNC DEGER C

JOHNSON & JOHNSON PROD INC

Absorbable bone fixation device

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JP19830122643 19830705

IKADA YOSHITO; GEN SHOKYU; SHIMIZU YASUHIKO; TAMURA KOICHI; NAKAMURA TATSUO; KIMURA SOSUKE; CHO TSUNEO; TADOKORO HIDEKI; HORI KAZUAKI

NIPPON MEDICAL SUPPLY

Use of molded polymeric material for preventing adhesion of vital tissues

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19861111

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LAPKA GALEN G; MASON NORBERT S; THIES CURT

UNIV WASHINGTON

Process for preparation of microcapsules

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RICHARDS MEDICAL CO

Biodegradable prosthetic device

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UNIV KENTUCKY RES FOUND

Porous microspheres for drug delivery and methods for making same

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KAPLAN DONALD S; HERMES MATTHEW E; MUTH ROSS R; KENNEDY JOHN J

UNITED STATES SURGICAL CORP.

Process of making an absorbable surgical device

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GRANDE DANIEL A

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BAYS F BARRY; TROTT ARTHUR F; MARCHAND SAM R

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Tack and applicator for treating torn bodily material in vivo

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19891226

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JAMIOLKOWSKI DENNIS D; GATERUD MARK T; NEWMAN JR HUGH D; SHALABY SHALABY W

ETHICON INC

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BAYS F BARRY; TROTT ARTHUR F; MARCHAND SAM R

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Method of joining torn parts of bodily tissue in vivo with a biodegradable tack member

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TUNC DEGER

JOHNSON & JOHNSON ORTHOPAEDICS

Absorbable bone plate

Continued

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Publication date

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Family members

Priority numbers

Inventors

Applicants

Title

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GREEN DAVID T

UNITED STATES SURGICAL CORP.

Inwardly biased skin fastener

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DEVEREUX DENNIS F; LANDI HENRY P

AMERICAN CYANAMID CO

Surgical composite structure having absorbable and nonabsorbable components

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ROSS RANDALL D; SNYDER STEPHEN J; MARCHAND SAM R

LINVATEC CORP.

Bioabsorbable tack for joining bodily tissue and in vivo method and apparatus for deploying same

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THM BIOMEDICAL INC

Apparatus for biodegradable, osteogenic, bone graft substitute device

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BOYCE STEVEN T

UNIV CALIFORNIA

Method and apparatus for preparing composite skin replacement

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Continued

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Applicants

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Priority numbers

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Applicants

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Continued

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Continued

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Applicants

Title

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Continued

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