A 5-year in vitro and in vivo study of the biodegradation of polylactide plates

J Oral Maxillofac 56:604-614,

Surg 1998

A S-Year In Vitro of the Biodegradation Riitta

and In Vivo Study of Polylactide Plates

Suuronen, MD, DDS, PbD, * Timo Pohjonen, Jarkko Hietanen, MD, DDS, PbD, MS& and Christian

Lindqvist,

f

MD, DDS, PbDJ

The purpose of this study was to investigate the long-term tissue response and duration of degradation of self-reinforced poly-L-lactide (SR-PLLA) multilayer plates in vivo. Purpose:

and Methods: Mandibular osteotomies in sheep were fixed with SR-PLLAmultilayer plates. The animals were followed for 1, 2, 3,4, and 5 years, after which histologic studies were performed. Materials

Results: The foreign-body reaction was mainly mild, and the osteotomies were well united. After 5 years in vivo, the material was almost completely resorbed, but small particles of polymer could still be detected at the implantation site. SR-PLLAplates were also incubated in vitro for 5 years. The material degraded considerably faster in vivo than in vitro. Molecular weight, melting temperature, and crystallinity of the plates remained at a constant level after 2 years in vitro, indicating very slow degradation of the oligomeric (molecular weight [Mw], 3500 daltons), highly crystalline (heat of fusion, 70 J/g), PLLA residue solely asa result of hydrolysis. Although the plates became increasingly fragile asthey degraded, they retained their macroscopic form until the end of the 5-year follow-up. Lossof massof the plates was 52% _t 8% after 5 years of incubation in vitro.

Although the long degradation period may seem to be a minor drawback to the use of such plates, it does not appear to affect the healing process. Conclusions:

Biodegradable devices have been used in orthopedic and oral and maxillofacial surgeryl-3 for over 10 years. The three most commonly used materials are polyglycolide (PGA), polylactide (PLA), and polydioxanone (PDS). Use of these materials started with sutures but has since spread to many fields of surgery. Today, biodegradable pins, plates, screws, tacks, and membranes are used. Biodegradable materials such as PGA have been reported to be osteoconductive. This property could be of use in several diIferent applications.* Drugs and growth factors could be added to the polymer and

*Resident, Department of Oral and Maxillofacial University Central Hospital, Helsinki, Finland. TSenior Researcher,

Institute

of Biomaterials,

of Technology, Tampere, Finland. *Head, Department of Oral Pathology, University

of Helsinki,

Helsinki,

Surgery, Tampere

Institute

accordingly exert their effects directly where they were needed, for example, in the healing tissue.5The property that makes biodegradable fracture fixation devices more attractive than metal ones is that no removal operation is needed after tissue has healed. Although PLA has been studied over the past 30 years, the number of clinical applications of the PLA-based materials has increased only recently.@ A possible explanation for the limited number of clinical applications of PLA in the past is that nonreinforced implants have been mechanically weak. In our department, self-reinforced poly-L-lactide (SR-PLLA) has been of special interest becauseof its good mechanical properties3 Use of SR-PLLA implants in fracture fixation started in the area of orthopedic surgery.8 It has since been gradually extended to the field of oral and maxillofacial surgery. It is now known that an osteotomy or a sagittal mandibular fracture can be fixed with SR-PLLAscrews, allowing free mandibular movement after operation.7T9A wide variety of other applications for small PLA plates and screws in oral and maxillofacial surgery also may be developed in the near future. In aqueous media, PLLA degrades into acid compounds through hydrolytic reactions. These may in turn be autocatalyzed by carboxylic end groups of degradation products. *0x11 PLLA also degrades because of simple hydrolysis, but tissue enzymes also may be involved in this degradation, especially during the

Helsinki University

of Dentistry,

Finland.

$Professor, Chairman, Departments of Oral and MaxiRofacial Surgery, Institute of Dentistry, Department of Surgery, Helsinki University, Address Department

Helsinki, Finland. correspondence and reprint requests to Dr Suuronen: of Oral and Maxillofacial Surgery, Helsinki University

Central

Hospital,

0 1998

American

Rasarmikatu Association

11-13, 00130 of Oral and Maxillofacial

Helsinki,

Finland.

Surgeons

0278.2391,‘98/5605-0011$3.00/0

604

SUURONEN

605

ET AL

later stages of the process. 12,is Very different mechanical protiles14 and rates of resorption of PLLA have been reported. Resorption times have ranged from 40 weeks to over 5.7 years. 15,16These discrepancies in resorption time have been accounted for in a number of ways, such as by reference to the different species and implantation sites studied and to the different properties of the materials, such as morphology (crystallinity and orientation), molecular weight and molecular weight distribution, presence of unreacted monomer and other impurities, size and shape of the implants, and processing and sterilization methods.r5J7 The presence of unreacted monomer and other impurities is the most critical factor affecting the rate of degradation and, consequently, the mechanical profile of PLLA.14 However, low molecular weight impurities accelerate the degradation initially, before they are leached out. Leaching takes place relatively quickly in the case of as-polymerized PLLA, which is known to be microporous.ls Crystallinity plays a key role in relation to absorption of polylactides.12J9~20 Increases in the number and size of the crystals, either formed during hydrolysis or left over because of preferential degradation of amorphous zones, retard the final biodegradation of PLLA.**,19,20 Factors that influence the degradation of PLLA also tiect the biologic reaction of the host tissues to the implanted polymer. 23 It has been shown that macrophages play a prominent role in resorption by phagocytosing particles of PLLA debris.** This process can induce cell deathz3 and an acute inflammatory response that can become evident even after a considerable time has elapsed since implantation.20 Such late adverse reactions caused by the implant degradation or host immunologic factors, or both, need further investigation. Results of recent studies suggest that the mechanism of degradation varies between subcutaneous and intraosseous PLLA implants, and that the histologic reaction induced by the implant may also may vary20*21 A general hypothesis, based on the experience of several research groups, is that the tissue response to degrading polymer correlates with the relationship between local accumulation of polymer debris from degrading implants and the capacity of surrounding tissues to eliminate such debris, and with host buffer capacity and increased osmotic pressure. 19,21 The differences between the reported resorption times of PLLA indicate a need for a long-term in vivo study whenever any material properties crucial to the degradation rate of an implant are changed. The aim in this study was to investigate long-term degradation of SR-PLLA multilayer plates in vitro and in vivo. Follow-up times were 1, 2, 3, 4, and 5 years, after which chemical and histologic studies were performed.

Materials IMPLANTS High molecular weight, medical-grade, purified PLLA polymer in ground form was obtained from PURAC/ CCA Biochem B.V., The Netherlands. The specifications of the polymer as given in the supplier’s certiticate of analysis were as follows: Lot number Batch number Molecular weight (determined by viscosity measurement) Intrinsic viscosity (in chloroform, T = 25°C) Specific rotation (in chloroform, T = 20°C) Melting point (“C) Melting range (“C) Heat of fusion (J/g) Residual solvent (%)

PLLA(EOO9) 89-A-l 3 720,000 10.3 -158 186.9 174.3-186.9 65.1 0.02

The plates were manufactured in the Tampere University of Technology using a self-reinforcing technique. The raw material was melt-extruded with an AXON BX-15 extruder (Axon plastmachiner AB, Astorp, Sweden) into rectangular billets with a cross-section of 3 X 12 mm and allowed to cool. The billets were then reheated and drawn uniaxially in the solid state through a tapered steel die at 150°C to a draw-ratio of 4 (cross section, 1 X 9 mm). During such die-drawing, spherulitic crystals are partially transformed to fibrillated structures orientated parallel to the long axis of the plate. After drawing, each billet was cut into pieces 50 mm long. These were compression-molded to a 0.5~mm-thick plate at 130°C and 30 MPa. During such compression-molding the structure is stretched perpendicular to the long axis to achieve orientation of the molecules and, consequently, also an increase in strength in the transverse direction. The plates were cut to 0.5 X 12 X 4Omm size and slightly curved as four-plate assemblies at 90°C and 5 MPa to adjust to the bone. Hence, the total thickness of the plate assembly was 2 mm. Each plate had a hole (diameter, 2.7 mm) in one corner. The plates were attached to each other with a wire to maintain the correct order. This wire was removed before the operation (Fig 1). One SR-PLLA plate has an initial tensile strength of 130 (It 10) MPa (with screw holes 100 ‘-+ 10 MPa), and shear strength of 100 (L 20) MPa perpendicular to the long axis and 60 (+- 10) MPa parallel to the long axis. After six holes have been drilled, an individual plate can resist 550 (2100) N tensile force. A four-layer plate assembly can therefore tolerate at most a 2200 (2400) N tensile force, assuming the force to be divided equally between the plates. At its weakest

WHAT HAPPENS TO PLLA PLATES IN 5 YEARS?

Operative

Methods

PREOPERATIVE PROCEDURES Preoperatively, each sheep was given 1 mg atropine (Atropin, 1 mg/mL, Orion, Espoo, Finland) subcutaneously, 1,200,OOOIU benzylpenicillin procaine (Procapen, 300,000 IU/mL, Orion) intramuscularly (IM), and 500 mg tinidazole intravenously (IV) (Tricanix, 5 mg/mL, Orion). The sheep were anesthetized with IM medetomidine (Domitor, 1 mg/mL, L;dHkefarmos, Turku, Finland), 0.025 mg/kg, and IM ketamine hydrochloride (Ketalar, 50 mg/mL, Parke-Davis, Barcelona, Spain), 1.Omg/kg. Every 30 to 60 minutes, each sheep received IV 50% of the original amount of medetomidine and ketamine hydrochloride. OPERATIVEPROCEDURE

FIGURE 1.

Multilayered

structure

of SR-PLLA

plate

point (ie, at the site of a screw hole), an individual plate can initially tolerate a 550 + 150 N shearing force perpendicular to the long axis.3 The plate assemblies were sterilized by gamma radiation (dose, 25 kGy) (Kolmi-Set Ltd, Ilomantsi, Finland). This does not significantly affect the initial strength of the device. After processing and sterilization, the molecular weight of the plates decreased to 49,800 (intrinsic viscosity, 1.46 dL/g). Their peak melting temperature was 179.1”C, and heat of fusion was 50 J/kg (degree of crystallinity, 53%). The advantage of such assembliesis that each plate is fairly flexible, and the plates can therefore be adjusted easily, one by one, to fit the contour of bone. When all of the plates have been attached to the bone (using plate forceps), five additional holes can be drilled through the plate assembly and bone, after which the assembly stiffens to the form desired. The PLLA-plate assembly gradually loses strength. After 24 weeks in vivo, the plates still exhibited more than half of their initial strength.3 Use of these assembliesfor experimental fracture fixation has been described in an earlier paper. l The plate assemblywas fixed with titanium screws. The diameter of the core and the thread of the screws used were 2.0 mm and 2.7 mm, respectively. The length of the screws ranged from 6 to 10 mm.

The left side of the face was shaved and scrubbed with povidone iodine (Betadine, 100 mg/mL, Leiras, Tammisaari, Finland) and chlorhexidine gluconate (Klorheksidos, 5 mg/mL, LZikefarmos) solutions. An incision was made below the lower border of the body of the mandible from the second molar to the third incisor region. The periosteum was incised to expose the bone. The mental foramen or foramina were identilied. A transverse osteotomy was made behind the mental foramen with an oscillating saw (blade width, 14 mm). The neurovascular bundle was also cut. The osteotomy was then reduced and hemostasisachieved by attaching the SR-PLLAplates one by one on top of each other to the buccal side of the osteotomy using plate forceps. When all of the plates were in place, five additional screw holes (2.7 mm) were drilled through the plate assembly(one hold had already been drilled through the plates), and six screw channels were drilled into the bone with a 2.0~mm drill, penetrating both cortices. The average width of the body was 8 mm. To avoid thermal damage, bone preparation was done under irrigation with saline solution. Threading was performed with a 2.7~mm metal tap. The plate was then fixed with 2.7~mm titanium screws, which were tightened manually (Fig 2). The incisions were closed in layers using absorb-

osteotomy

EXPERIMENTAL ANIMALS Five adult Finnish sheep each weighing 30 to 50 kg underwent operation. Osteotomies on the left side of the body of the mandible, at the site of the diastema, were fixed using a four-layer SR-PLLA plate and six titanium screws (Fig 2).

FIGURE

2. Schematic

drawing

of the operative

procedure.

607

SUURONEN ET AL

able sutures (Vicryl, Johnson & Johnson, Sollentuna, Sweden). POSTOPERATIVEPROCEDURE Postoperatively, each sheep was given 400 mg phenylbutazone (Reumuzol, 200 mg/mL, Lalkefarmos) and 1,200,OO IU benzylpenicillin procaine daily for 5 days. All sheep started to eat soft food 1 day after operation. On the sixth day after operation, they were given hay, and all ate and ruminated normally. Follow-up times were 1, 2, 3,4, and 5 years.

Examination

Methods

RADIOLOGIC AND HISTOLOGIC METHODS Three weeks after operation, all sheep were anesthetized as previously described, and intraoral occlusal radiographs were taken (Siemens Polyphos 30 M; Kodak T-MAT G film size 18 X 24 cm, intensifying screen Kodak Lanex medium, 44 kV, 8 mAs, 120~cm focus-object). The radiographs were evaluated for the condition of the reduction, any dislocation, relapse of fixation, extent of external callus, visibility of the osteotomy, and union of the osteotomy. The unoperated right side served as a control in comparing the thickness and bony structure of the mandible. All radiographs were analyzed by the same person. After each sheep had been killed, the mandible was carefully dissected, inspected, and photographed, and radiographs were taken in occlusal and lateral projections (Siemens Polyphos 30 M, Kodak T-MAT G film size 18 X 24 cm, intensifying screen Kodak Lanex medium, 44 kV, 8 mAs, 120-cm focus-object). The metal screws were removed, and the specimens were fixed in 70% alcohol and embedded in methylmethacrylate. Histologic studies were made of 5-urn sections cut with a Jung Polycut S (Reichert-Jung, Nussloch, Germany) microtome and stained using the MassonGoldner method. All histologic specimens were analyzed by the same person. DEGRADATION STUDIES IN VITRO The plates were arranged as four-layer assemblies similar to the in vivo test and placed in a Na,HPO* (2.78 g/L)-KH,P04 (13.58 g/L) phosphate-buffered aqueous solution (PBS) at pH 6.10 and 0.12 mol/L, and kept at 37°C. The medium (1.3)was changed periodically. Two four-plate assemblies were removed at 1, 1.5, 2, 3, 4, and 5 years and changes in thermal properties and in molecular weight were determined. Scanning electron micrographs were also obtained. THERMAL ANALYSIS AND MOLECULAR WEIGHT MEASUREMENTS A Perk&i-Elmer DSC 7 (Perkin Elmer, Norwalk, CT) differential scanning calorimeter (DSC) calibrated with

indium standards was used to determine the heat of fusion of intact SR-PLLA plates and of such plates after in vitro exposure. The DSC was operated at a heating rate of 20”C/min. Dry, 6 t l-mg, samples evacuated at room temperature for 3 days were used in each case. The samples were heated in a nitrogen atmosphere from room temperature to 200°C (about 40°C above melting temperature to ensure melting of all crystallines) and heat of fusion was estimated from the area enclosed by the DSC curve and the baseline. The level of crystallinity was estimated from the heat of fusion, assuming the value of 93.7 J/g calculated by Fisher et allo to apply to perfectly crystalline PLLA. Triplicate samples were used in each case. Solution viscosity (according to ASTM D 445-88) of intact SR-PLLA plates and of such plates after in vitro exposure were measured in chloroform at 25°C using an Ubbelohde (Schott, Hofheim, Germany) capillary viscometer (type Oa, according to ASTM D 446). Intrinsic viscosities (in dL/g) were obtained by linear regression analysis from dilution series (0.1 g/dL, 0.2 g/dL, 0.3 g/dL, 0.4 g/dL, 0.5 g/dL), and viscosityaverage molecular weights (in g/mol) were calculated using the Mark-Houwink equation and values for the parameters K = 5.45 X lo-* and a = 0.73.

Results IN VITRO The results of thermal analysis and molecular weight measurements are shown in Table 1. Although high molecular weight PLLA raw polymer (molecular weight, 720,000) had been used for the SR-PLLA plates, the molecular weight decreased to 220,000 because of thermal degradation during melt-stage and solid-stage processing. Sterilization by gamma irradiation, using a minimum dose of 25 kGy further decreased the molecular weight to 49,800. The degree of crystallinity of intact SR-PLLAplates was 53%. After 1 year, the molecular weight was 18,100, approximately 36% of the original molecular weight after sterilization by gamma radiation. The plates still had a tensile strength of 67 (t 12) MPa.3 The degree of crystallinity had increased to 66%. After 1.5 years, the molecular weight had decreased to 6300, and the degree of crystallinity had increased to 74%. After 2 years, the plates had degraded significantly. The molecular weight was 3300, and the degree of crystallinity was 74%, indicating that most of the amorphous phase of the PLLA had degraded or crystallized. As compared with intact SR-PLLA plates, the peak melting temperature had decreased 10°C to 169.l”C. From 2 years on, the molecular weight and the thermal properties of the PLLA remained essentially unchanged, indicating that the PLLA crystals were extremely resistant to degradation through solely

608

WHAT HAPPENS TO PLLA PLATES IN 5 YEARS?

Time (v-1

Peak Melting Temperature (“Cl

Heat of Fusion U/kg)

Crystallinity w>

0 1 1.5 2 i

179.1 177.8 176.0 169.1 166.5 169.7

50.0 64.2 69.1 69.5 68.4 70.4

2:

5

169.9

70.5

simple hydrolysis (Table 1). On the basis of the Gibbs-Thompson equation, which relates the melting temperature and crystal thickness of a polymer, the constant melting temperatures of the PLLA crystals may be regarded as an indication that the size (thickness)of the crystals remained unchanged after 2 years of exposure. As the material degrades,it is probably an undetectable lactic acid or other low molecular weight degradation product that will be slowly washed away from the sides, rather than the fold surfaces’of the lamellar crystals, and therefore not seen in material analysis. This assumption reflects the finding that the pH of the medium in in vitro studies decreased from an initial value of 6.10 during each periodic change, even after 3 years of exposure, because of slow release of soluble degradation products (Pig 3). Release of acid and, consequently, the decrease in pH of the medium, was greatest between 1.5 and 2 years and subsequently slowed down. The average loss of mass of the plates was 52% +- 8% after 5 years of exposure in vitro. A rough surface on the side of the plate could be seen in scanning electron micrographs (Fig 4). IN WV0 After 1 year of follow-up, the multilayered structure of the implanted material was clearly discernible,

o pH measurement -.---.pHoffreshPBS __._____.._._____._...........................-.......

1.50

2.00

2.50 Exposure

FIGURE 1.5

and

3.00

Intrinsic Viscosity (dL/g)

Molecular Weight (g/m00

74 74 73 75

1.46 0.70 0.32 0.20 0.20 0.22

75

0.21

49,800 18,100 6,300 3,300 3,300 3,700 3,500

although disintegration had already started (Pig 5). The implant was surrounded by a thin layer of connective tissue in which occasional mononuclear inflammatory cells were visible. In the bone in the vicinity of the implanted material, resorption lacunae with osteoclastscould be seen. After follow-up of 2 years, the implanted plates could clearly be seenpig 6). The original morphology of was clearly discernible, but they had started to disintegrate. This was most marked in the outermost plates. The plates were surrounded by thin layers of fibrous connective tissue containing a mainly mild, occasionally moderate, inflammatory infiltrate of mononuclear cells. The thin bone between the implanted plates was rimmed by osteoblasts, but occasional osteoclasts could also be seen. Multinucleated foreignbody giant cells were often seen in close contact with the plates. Compared with the situation after 2 years, after 3 years very little implanted material could be detected. The implanted plates had almost completely disappeared (Pig 7). Lamellar bone replaced the area of the implants. However, in some marrow spaces, fragments of implanted material were still present. The fragments were angular, of varying size, and surrounded by mononuclear inflammatory cells and mul-

and change of PBS

I 4.00

3.50

time (years)

3. Changes in pH of the phosphate 4 years of in vitro exposure at 37°C.

buffer

solution

between

FIGURE 4. become

brittle

After and

5 years in vitro, the surface rough (original magnification

of the PLL4 x3.500).

plate

had

SUURONEN

609

ET AL

the connective tissue covering the bone. The tiny fragments of implant had occasionally elicited chronic inflammatory cells and mononucleated foreign-body giant cells in the surrounding area. In some marrow spaces without an implant on the outer cortex, osteoblasts and occasional osteoclastswere present.

FIGURE follow-up. radiolucent.

5. The osteotomy site can still be identified The metal screws can be seen while

the

after 1 year plate itself

of is

tinucleated foreign-body giant cells. Osteoid and osteoblasts were visible in some bone areas facing the partly resorbed implants. Some osteoid also was detectable on the outer cortex. After 4 years the remains of the implanted material were dilficult to find. They were angular and surrounded by a mild chronic intIammatory infiltrate. Osteoblasts rimmed the bone facing the implants. After 5 years, remnants of the implanted plates could only be found with the aid of polarized light (Fig 8). Small, rounded, or angular fragments were seen in both the marrow spaces of the outermost bone and

FIGURE pieces polarized

6. After 2 years, the lamellated structure of the Implanted of implanted material are surrounded by fibrous tissue with light, original magnification x 125).

material occasional

Discussion Degradation of PLA in an aqueous medium proceeds by random bulk hydrolysis of the ester bonds in polymer chains. During hydrolysis, the strength of PLLA devices decreasesto zero, after which fragmentation takes place. Several research groups have shown that the amorphous phases in semicrystalline PLLA implants degrade first. Factors governing degradation of PLLA have been discussed in the introductory section. As mentioned there, crystallinity plays a key role in relation to absorption of polylactides. Factors that Influence the degradation of PLLA also affect the biologic reaction of the host to implanted polymer.23 It hasbeen shown that macrophages play a prominent role in the resorption process by phagocytosing PLLA debris, a process that also can induce an acute

(I) is still visible. multinucleated

However, its disintegration foreign-body giant cells

has already started. Small [arrow). (Masson Goldner,

WHAT HAPPENS TO PLLA PLATES IN 5 YEARS?

FIGURE 7. After 3 years, only small particles inflammatory cells and multinucleated polarized light, original magnification

foreign-body x250).

of implanted material [I) could be detected. They were angular and surrounded by mononuclear giant cells [arrow). Lamellar bone (BJ has replaced the area of the implant [Masson Goldner,

inflammatory response.22Highly crystalline as-polymerized PLLA disintegrates into very small, almost perfectly crystalline, needle-like debris, which was still present even after 5.7 years in vivo.16J8,22The structure of these crystals may be more perfect than that of melt-crystallized PLLA, because they are formed directly during the bulk polymerization process of melted L-lactide monomer at low temperature into high molecular weight PLLA polymer.‘* When PLLA polymer is subjected to an injection molding process and rapidly cooled in the mold cavity, amorphous samples are obtained. Gutwald et all3 have reported that amorphous PLLA was totally or nearly totally absorbed in 2 years in vivo, whereas as-polymerized crystalline PLLA remained almost stable in form and structure over the whole period of observation.12s*9 Unfortunately, the amorphous PLLA samples had too low a mechanical strength for use in osteosynthetic devices. ‘9 Li et al” studied melt-molded, rapidly cooled and annealed, crystalline PLLA in vitro, and showed that initially amorphous PLLA samplescrystallized asdegradation proceeded. l l After the integrity of the polymer mass had been lost, the residual crystalline matter

initially present or formed during degradation appeared very resistant and was still present in powder form after 2 years of incubation in PBS.The molecular weight of crystalline samples remained constant at about 5000 daltons after 90 weeks of exposure, a finding similar to our long-term in vitro results relating to SR-PLLA plates. As far as we know, the in vitro exposure time of 5 years used by us is the longest reported for PLLA implants. The results of our study indicate that PLLA crystals are practically insoluble in water. Gogolewski et all2 studied the degradation of injection-molded PLLA and P(L/D)LA 2 X 15-mm copolymer discs and 4 X 7-mm cylinders in the subcutaneous tissuesof mice. AU polylactides were well tolerated by the tissue and had degraded significantly (56% to 99% in terms of molecular weight) after 6 months. PLLA degradation rate in vivo decreased with increasing initial molecular weight of the polymer. PLLA samples with an initial viscosity average molecular weight of 76,400 and a level of crystallinity of 52% degraded to a molecular weight of 17,400, similar to the degradation kinetics found by us for SR-PLLAplates over 24 weeks of implantation in sheep (molecular weight decreased

SUURONEN

FIGURE 8. After 5 years of follow-up, inflammatory

611

ET AL

very small mainly rounded particles of implanted material cell infiltrate is visible in the conective tissue above bone (Masson-Goldner, polarizing

to lS,OOO).s Several samples showed pronounced bimodal molecular weight distributions. We suggest that the bimodality arose because of differences in degradation rate resulting from variability in distribution of crystalline and amorphous regions within the samples.Another explanation is that two mechanisms nonenzymatic and enzymatic, may have been involved in the degradation process, the latter being more extensive during the later stage when partially degraded polymer is present. A significant lossof mass (29% to 50%) was found in the case of poly(L/Dlactides) containing 10% and 20% of D-lactide units, and racemic PDLLA (60%). As expected, the highly crystalline PLLA exhibited a far smaller loss of mass (0% to 10%). It has been suggestedthat degradation kinetics also depend on the site of implantation. A plate and screw system made from injection-molded, intrinsically amorphous, racemic poly-DL-lactic acid was used for rigid fixation of experimental nasal bone fractures in 20 rabbits, In addition, prebent plates (30”, 45”, GO”) were placed in subcutaneous pockets in the backs of the animals. The plates were removed after 7, 14, 28, and 42 days, and bending angles, plate stability, molecular weights, and histologic findings were stud-

(I) are visible in the lamellar light, original magnification

bone (B). A moderate x2.50).

ied. A significant decrease in molecular weight over time, and a difference in loss of molecular weight indicating that degradation was faster subcutaneously, were observed. After 40 days, the molecular weight was less than 70% of the original molecular weight in vitro, and it was only approximately 25% of the original molecular weight in the subperiosteal plates and less than 15% of the original molecular weight in the subcutaneous plates. These results suggest that degradation rates differ according to site of implantation. The sterilization method was not described.** Matsutsue et alz5 reported that drawn PLLA rods lost their molecular weight most rapidly in the medullary cavity, followed by the subcutis, and most slowly in vitro. However, there has been some controversy about the mechanism of degradation of polylactides. In a short-term study (over 11 weeks), ISo@ found no difference between molecular weight loss of aspolymerized PLLA plates in vivo and in vitro. The long-term results of the study described in the present article indicate that there is a marked difference between the loss of massof crystalline PLLA in vitro and in vivo. One evident explanation for the faster loss of massof PLLA in vivo is that enzymatic degradation

612 processes and other cellular activity are also involved.2G29 Few publications report the total resorption of PLLA. The results of our study are in accordance with the long resorption times of crystalline PLLA observed in these studies. Rohowsky et alsO studied PLLA screws implanted in canine femora and tibiae for 5 years. After 2 years, the screw heads had fractured, and the material was no longer homogenous. There was also increased turnover of bone adjacent to the screws. After 3 years, a substantial reduction of homogeneity of PLLA was observed, with a loss of the outline of the screw, particularly in relation to the threads. After 5 years, only small fragments of PLLA were still observable in the medullary canal. In one animal of four, there had been complete resorption of the PLLA screw. Matsutsue et a131 studied drawn PLLA rods implanted in the subcutaneous tissue and the medullary cavity of rabbits. After 5 years 2 months of intramedullary implantation, the material had been almost completely resorbed and replaced by bone marrow cells. Five years 9 months after subcutaneous implantation, the material had been resorbed completely, without scar formation. In a clinical study, Pihlajamaki et aP* reviewed 27 patients whose small fragment fractures or osteotomies had been treated using internal fixation with SR-PLLA pins. Biopsy specimens from two patients 20 and 37 months after implantation showed no residual polymeric material. The biocompatibility of PLA has been the subject of numerous studies since Kulkarni et aP3 first reported the potential of polylactides as resorbable biomaterials for use in surgery in 1966. The biocompatibility of PLLA is generally accepted as excellent. As long as degradation has not progressed, PLLA is usually considered to be the best polymer biocompatibility class.l* However, the tissue reaction changes as biodegradation progresses. Mild to moderate inflammatory reactions are frequently seen around degrading PLLA implants.3* Typical tissue reactions to degrading polylactides described in the literature include fibroblasts, histiocytes, lymphocytes, mast cells, foreignbody giant cells, macrophages, plasma cells, eosinophils, and lymphoid cells. It has been reported that the fibrous tissue layer around implants, usually in the form of granulation tissue, decreases progressively with time.12 Late adverse reactions caused by degrading PLLA implants are rare in the published reports of animal studies. We are aware of only two exceptions, which we published recently. In a 1993 study, particles of polytetrafluoroethylene (PTFE; control) and “predegraded” particles of PLLA having diameters of less than 38 m were injected intraperitoneally into mice. After periods of up to 7 days, cells were harvested from the abdominal cavity. Microscopic examination showed evidence of cell damage and

WHAT HAPPENS TO PLLA PLATES IN

5 YEARS?

death caused by phagocytosed PLLA particles. Cell death appeared to be related to the amount of PLLA ingested. Cell death was not observed with phagocytosed PTFE particles .23 In a 1991 study, PLLA plates were implanted subcutaneously into rats. At 104 weeks, no adverse reactions were noted. However, at 143 weeks, one animal showed a histologically disputable, late, foreign-body reaction to the PLLA implants0 In contrast to the findings in the above-mentioned studies relating to the soft tissues, we are aware of no reports of late tissue responses to PLLA relating to bone. Only a mild inflammatory response was reported during stimulated late-phase degradation of a large PLLA intramedullary nail in the pig in a recent study.*l Rozema et al35 have reported a clinical study in which 4 of 10 patients returned on their own initiative because of intermittent, painless swelling in the operative site 3 years after fixation of zygomatic bone fractures with large, as-polymerized PLLA plates and screws. On recall, the remaining five patients available also showed similar patterns.16 After postoperative periods varying from 2 years 11 months to 3 years 8 months, four patients underwent reoperation, and the material was removed.sl Characterization of PLLA debris particles showed an average molecular weight of about 5,000. The heat of fusion of the harvested PLLA was 96 J/g and the melting point 184°C which means that the degraded PLLA was almost perfectly crystalline and of low molecular weight. On light microscopy, a foreign-body reaction was evident. On electron microscopy, macrophages surrounding the implant were seen to contain small particles of degraded material, which was also visible in the extracellular spaces between the collagen fibers. The authors concluded that a combination of changes in quantity, degradation characteristics, and associated morphologic alterations of the implanted material, and the characteristics of the tissue in which the material had been implanted, were responsible for the clinical reactions.35 Later, Bergsma et aP6 reported that the average molecular weight of the crystalline PLLA polymer debris still present in the subcutaneous tissue after 5.7 years in a zygomatic-fracture patient was 5400. The mean particle size seemed to be smaller than that of the material implanted for 3.3 years, and scarcely any PLLA was found in extracellular space. Most PLLA crystals had been internalized by phagocytosing cells, but no signs of cell damage were evident. No substantial loss of mass of PLLA crystals took place between 3.3 and 5.7 years. It was concluded that the degradation had taken place by washing away of oligomers by the tissue fluids and was therefore not detectable by material analysis. The implantation site also seemed to be important in determining the type and intensity of the inflammatory response to the

613

SUURONEN ET AL

PLLA residues. The trephined bone from this patient showed various differences from results relating to subcutaneously implanted material. No swelling was observed in the intraosseously implanted thread portion of the PLLA screws, and internalization of PLLA particles by phagocytic cells was very limited. One of possible mechanisms suggestedfor the subcutaneous swelling was increased osmotic pressure in the subcutaneous tissue. The increase in volume that accompanied disintegration of the PLLA, and the development of fibrous tissue, were also advanced asan explanation of the origin of the swelling.16 Adverse effects also have been reported in another clinical study in which as-polymerized PLLA plates were used to repair ankle fractures in 25 patients. In four patients, remnants of the plates were observed extruding through the skin.37In these cases,the bulky as-polymerized PLLA implants had been placed on bone directly under skin with minimal amounts of subcutaneous tissue. Low reaction rates (1 of 52 and 1 of 84) both cases being subcutaneous swellings because of protruding screw heads, have been observed in clinical long-term follow-up studies in which PLLA screws have been placed intraosseously for displaced ankle fractures.“J2 In a long-term study on the fixation of sagittal split ramus osteotomies with SR-PLLAscrews, despite the osteolytic changes in radiographs, no clinically manifest foreign body reactions were reported at the longest follow-up period of 5 years.38 The long duration of degradation of SR-PLLAplates in vivo would seem to be a drawback to their use in maxillofacial surgery because most fractures in which biodegradable implants are used heal within a few months. Nevertheless, the long degradation time did not affect the fracture healing because the stressprotection effect is lost much earlier.3 Small PLLA crystals of 10 to 100 pm were visible after 4 and 5 years of follow-up. However, no clinically manifested foreign-body reactions were seen in the animals concerned. The findings suggest that use of SR-PLLA plates should be carefully planned so that there is enough well-vascularized tissue surrounding the implants. In this way, the local tissue tolerance and transportation mechanisms are capable managing the foreign material.

References 1. Suuronen R, Pohjonen absorbable self-reinforced plates in the fixation experimental study in 1992 2. Suuronen R, Manninen osteotomy fixed with animal study. Br J Oral

T, Vasenius J, et al: Comparison of multi-layer poly-I-lactide and metallic of mandibular body osteotomies: An sheep. J Oral Maxillofac Surg 50:255, M, Pohjonen T, et al: Mandibular biodegradable plates and screws: An Maxillofacial Surg 35:341, 1997

3. Suuronen R, Pohjonen T, Taurio R, et al: Strength and strength retention of self-reinforced poly-l-lactide screws and plates: In vivo and ln vitro study. J Mater Sci Mater Med 3:426, 1992 4. Hatton PV, Walsh J, Brook IM: The response of cultured bone cells to resorbable polyglycolic acid and silicone membranes for use in orbital Boor fracture repair. Clin Mater 17:71, 1994 5. Garvin KL, Miyano JA, Robinson D, et al: Polylactide/ polyglycollde antibiotic implants in the treatment of osteomyelitis. J Bone Joint Surg Am 76:1500,1994 6. Bucholz RW, Henry S, Henley MB: Fixation with biodegradable screws for the treatment of fractures of the ankle. J Bone Joint Surg Am 76:319, I994 7. Suuronen R, Lame P, Lmdqvist C: Sagittal osteotomies fixed with biodegradable screws: A preliminary report. J Oral Maxlllofacial Surg 52:715, 1994 8. Rokkanen P, B&man 0, Valnionpti S, et al: Absorbable devices ln the fixation of fractures. J Trauma 4O:S123, 1996 9. Kallela I, Laine P, Suuronen R, et al: Skeletal stability after mandibular advancement and fixation with biodegradable screws. Int J Oral Maxillofac Surg 27:3, 1998 10. Fisher EW, Sterzel HJ, Wegner G: Investigation of the structure of solution grown crystals of lactide copolymers by means of chemical reactions. Kolloid-Zeitschrlft und Zeitschrift fiir Polymere 251:980, 1973 11. Ll S, Garreau H, Vert M: Structure-property relationships in the case of the degradation of massive poly (oc-hydroxy acids) in aqueous media. Part 3. Influence of the morphology of poly(llactic acid). J Mater Sci Mater Med 1: 198, 1990 12. Gogolewski S, Jovanovic M, Perren SM, et al: Tissue response and in vivo degradation of selected polyhydroxyacids: Polylactides (PLA), poly(3-hydroxybutyrate) (PHB), and poly(3hydroxybutyrate-co-3-hydroxyvalerate) (PHB/VA). J Biomed MaterRes 27:1135, 1993 13. Gutwald R, Pistner H, Reuther J, et al: Biodegradation and tissue-reaction in a long-term implantation study of poly@ lactide). J Mater Sci Mater Med 5:485, 1994 14. Zhang X, Wyss UP, Pichora D, et al: An investigation of poly(lactic acid) degradation. J Bioact Compat Polym 9:80,

1994 15. Majola A: Biodegradation and biocompatibllity of self-reinforced polylactide (SR-PLA) implants in vivo, in Rokkanen P, Tormala P (eds): Self-Reinforced Bioabsorbable Polymeric Composites ln Surgery. Tampere, Finland, Tampereen Pikakopio, 1995 16. Bergsma JE, de Bruijn WC, Rozema FR, et al: Late degradation tissue response to poly(l-lactide) bone plates and screws, Biomaterials 16:25, 1995 17. Vert M, Li S, Garreau H: New insights on the degradation of bioresorbable polymeric devices based on lactic and glycolic acids. Clin Mater 10:3, 1992 18. Bos RRM: Poly(L-lactide) osteosynthesis: Development of bioresorbable bone plates and screws. Thesis, University of Groningen, The Netherlands, 1989 19. Pistner H, Bendix DR, Miihling J, et al: Poly(I-lactide): A long-term degradation study in vlvo 1993. Part III. Analytical characterization. Biomaterials 14:291, 1993 20. Bergsma EJ, Rozema FR, Bos RRM, et al: In vivo degradation and biocompatibility study of in vitro pre-degraded as-polymerized polylactide particles. Biomaterials 16:267, 1995 21. van der Elst M, Kuiper I, Klein CPAT, et al: The burst phenomenon, an animal model simulating the long-term tissue response on PLLA interlocking nails. J Biomed Mater Res 30:139, 1996 22. Rozema FR: Resorbable poly&lactide) bone plates and screws. Thesis, University of Groningen, The Netherlands, 1989 23. Lam KH, Schankenraad JM, Esselbrugge H, et al: The effect of phagocytosls of poly(l-lactic acid) fragments on cellular morphology and viability. J Biomed Mater Res 27:1569, 1993 24. Tschakaloff A, Losken HW, von Oepen R, et al: Degradation kinetics of biodegradable DL-polylactic acid biodegradable implants depending on the site of implantation. Int J Oral Maxillofac Surg 23:443, 1994 25. Matsutsue Y, Yamamuro T, Oka M, et al: In vitro and in vivo studies on bioabsorbable ultra-high-strength poly(I-lactide) rods. J Biomed Mater Res 261553, 1992

614

DISCUSSION

26. Verheyen CCPM, de Wijn JR, van Blitterswijk CA, et al: Examination of efferent lymph nodes after 2 years of transcorticaI implantation of poly(L-lactide) containing plugs: A case report. J Biomed Mater Res 27: 1115, 1993 27. Bergsma JE, Bos RRM, Rozema FR, et al: Biocompatibility of intraosseously implanted predegraded poly(lactide): An animal study. J Mater Sci Mater Med 7: 1, 1996 28. Pohjonen T, Helevirta P, T&m%& P, et al: Strength retention of self-reinforced poly-Llactide screws: A comparison of compression moulded and machine cut screws. J Mater Sci Mater Med 8:311, 1997 29. van der Elst M: Biodegradable interlocking nails for fracture fixation. Thesis, Free University Hospital in Amsterdam, The Netherlands, 1997 30. Rohowsky MW, Chen EH, Clemow AJT, et al: A five-year study of absorbable screws implanted in canine femora and tibiae, in Leung K, Hung L, Leung P (eds): Biodegradable Implants in Fracture Fixation. Singapore, Department of Orthopaedics and Traumatology, The Chinese University of Hong Kong and World Scientific Publishing Co Pte Ltd, 1994 31. Matsutsue Y, Hanafusa S, Yamamuro T, et al: Tissue reaction of bioabsorbal$e ultra high strength poly (I-lactide) rod: A long term study in rabbits. Clin Orthop 317:246,1995 32. PihlajamM H, Biistman 0, Hirvensalo E, et al: Absorbable pins

J Oral

Maxillofac

56:614-615,

33. 34.

35.

36.

37.

38.

of self-reinforced poly-Llactic acid for lixation of fractures and osteotomies. J Bone Joint Surg Br 72:853, 1992 Kulkarni RK, Pani KC, Neuman C, et al: Polylactic acid for surgical implants. Arch Surg 93:839, 1966 Bos RRM, Rozema FR, Boering G, et al: Degradation of and tissue reaction to biodegradable poly(L-lactide) for use as internal fixation of fractures: A study in rats. Biomaterials 12:32, 1991 Rozema FR, de Bruijn WC, Bos RRM, et al: Late tissue response to bone plates and screws of poly(l-lactide) used for fracture fixation of the zygomatic bone, in Doherty PJ, Williams RL, Williams DF (eds): Biomaterial-Tissue Interfaces, Advances in Biomaterials 10. Amsterdam, The Netherlands, Elsevier Science Publishers BV, 1992 Bergsma EJ, Rozema FR, Bos RRM, et al: Foreign body reactions to resorbable poly(L-lactide) bone plates and screws used for the fixation of unstable zygomatic fractures. J Oral Maxlllofac Surg 51:666, 1993 Eitenmiiller J, Divid A, Muhr G: Treatment of ankle fractures with complete biodegradable plates and screws of high molecular weight polylactide. Third World Biomaterials Congress, April 21-25, 1988, Kyoto, Japan Kallela I, Laine P, Suuronen R, et al: Saglttal split osteotomies stabilized with biodegradable polylactide screws: A clinical and radiological study. Int J Oral Maxlllofac Surg (submitted for publication)

Surg

1998

Discussion A 5-Year In Vitro and In Vivo Study of the Biodegradation of Polylactide Plates JE. Bergsma, DDS, PhD Resident, Groningen,

Department of Oral The Netherlands

and

Maxillofacial

Surgery,

I would like to compliment the authors for the clarity of this article. It is well written and easy to read, and is a good survey of the problems encountered with long-term degradation of biodegradable materials. The main problem with semi-crystalline biodegradables such as poly-L-lactide (PLLA) and self-reinforced

(SR)-PLLA

implants

is summarized

at the

end of this discussion; these materials can only be used safely as long as the application is carefully planned with regard to having enough well-vascularized tissue surrounding the implant. This is critical because the local tissue tolerance and transportation mechanisms must be able to cope with the degradation products, which are highly crystalline foreign body materials. Normally, in the facial area, tissue vascularization is excellent. However, it is our finding that the local tissue tolerance was exceeded when using as-polymerized PLLA plates and screws for the duration of zygomatic fractures, resulting in a subcutaneous swelling at the site of implantation. l-3 This swelling was caused by the previously mentioned, very ing, crystalline breakdown

stable

and therefore

nondegrad-

products. In the current article, the authors describe that SR-PLLA also disintegrates into stable and highly crystalline particles. Although the Y-year results suggest that the implanted

material has almost completely resorbed, in Figure 8 abundant amounts of intracellular birefringent PLLA material are seen. This internalization process is not mentioned in the results, as if detection and the presence of a resorbable material ends in the extracellular space. As mentioned in the discussion and in the literature, the internalization and cellular response to these highly crystalline fragments is crucial with regard to further degradation, overall tissue reaction, and the previously mentioned subcutaneous swelling.1-5 The authors state that the in vivo degradation and resorption were enhanced compared with in vitro degradation. I do not agree with the conclusion that this was probably caused by enzymatic degradation. In the current report, mass loss was not measured in the in vivo study; also, molecular weight loss was not measured or compared, and no enzymatic activity was shown. When a biodegradable material is used in vivo, internalization by cells, spread by lymph vessels, and the histologic techniques make it very diEcult to estimate real mass loss. In the literature, only a limited number of studies mention possible enzymatic degradation of PLLA.h More recent studies do not confirm this hypothesis for both intracellular and extracellular degradation.1,7zs The absence of enzymatic degradation, in combination with the conclusion that the SR-PLLA crystals are practically

insoluble

in water,

these crystals in macrophages factors

in the slow

degradation

and the abundant

presence

(Fig S), could of semi-crystalline

therefore its clinical applicability. As previously stated, the slow degradation

of

be crucial PLLA

and

and the forma-